Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes

Transcription

Membranes for the Life Sciences
Edited by
Klaus-Viktor Peinemann and
Suzana Pereira Nunes
Membranes for the Life Sciences
Volume 1
Edited by
Klaus-Viktor Peinemann and Suzana Pereira Nunes
The Editors
Dr. Klaus-Viktor Peinemann
GKSS Research Center
Institute of Polymer Research
Max-Planck-Strasse 1
21502 Geesthacht
Germany
Dr. Suzana Pereira Nunez
21502 Geesthacht
Germany
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ISBN: 978-3-527-31480-5
V
Contents
Preface XI
Contributors ((Peinemann Vol. 1))
1
1.1
1.2
1.3
1.3.1
1.3.1.1
1.3.1.2
1.3.1.3
1.3.1.4
1.3.1.5
1.3.1.6
1.3.2
1.3.2.1
1.3.2.2
1.3.2.3
1.3.2.4
1.4
1.4.1
1.4.2
1.4.3
1.4.4
1.5
1.6
1.6.1
1.6.2
XIII
Membranes in Hemodialysis 1
Jörg Vienken
Introduction 1
Historical Achievements 2
Membranes for Hemodialysis: Polymers and Nomenclature 7
Membranes from Regenerated Cellulose 8
Modified Cellulosic Membranes 9
Cellulose Acetates 9
DEAE-Modified Cellulose, Hemophan 10
Benzyl-Modified Cellulose (Synthetically Modified Cellulose, SMC)
PEG-Grafted Cellulose 11
Vitamin E-Modified Cellulosic Membranes 12
Synthetic Membranes 12
Polyacrylonitrile (PAN) 13
Polymethylmethacrylate (PMMA) 14
Polysulfone (PSu) 15
Polyamide (PA) 17
Dialyzer Constructions 17
Hollow Fiber Dialyzers 17
Housing 18
Potting Material 18
Fiber Bundle 19
Dialysis Membranes and Performance: Principles of
Membrane Transport 20
Dialysis Membranes and Biocompatibility 24
Some Basic Information on Membranes and
Biocompatibility Parameters 24
Thrombogenicity of Different Types of Dialyzers and Filters 26
Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes
Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved
ISBN: 978-3-527-31480-5
11
VI
Contents
1.6.3
1.6.4
1.6.4.1
1.6.5
1.6.5.1
1.6.6
1.6.6.1
1.6.7
1.6.8
1.6.9
1.7
2
2.1
2.2
2.2.1
2.3
2.4
2.4.1
2.4.2
2.5
2.6
2.6.1
2.6.2
2.7
2.7.1
2.7.2
3
3.1
3.2
3.3
3.4
3.4.1
3.4.2
3.5
3.6
3.7
Complement Activation by Different Dialyzers and Filters 29
Cell Activation by Different Types of Dialyzers and Hemofilters 31
Apoptosis 31
Oxygen Species Production – Induction of Oxidative Stress 32
Degranulation of Neutrophils 34
Stimulation of Cytokine Generation by Different Types of
Dialyzers and Hemofilters 35
The Impact of Membrane Types on LPS-Stimulated IL-1b
Secretion 36
The Impact of Large-Pore Dialysis Membranes on the
Inflammatory Response in HD Patients by Cytokine
Elimination 36
The Effect of Different Dialyzers on the Acute Phase Reaction 37
Activation of the Kinin System by Different Types of Dialyzers
and Hemofilters 37
Conclusion 39
Membranes for Artificial Lungs 49
Frank Wiese
Introduction 49
History of Blood Oxygenation 49
Membrane Oxygenators 50
Principle of Gas Transfer 53
Membranes and Membrane Properties 55
Microporous Membranes 55
Dense Membranes/‘‘Diffusion Membranes’’ 57
Membrane Production 59
Operational Modes and Membrane Makeup in Oxygenators
Microporous Capillary Membranes, Blood Inside 62
Microporous Capillary Membranes, Blood Outside 62
Extracorporeal Circulation 65
Cardiodiapulmonary Bypass (CPB) 65
Lung Support Systems 65
Membranes for Blood Fractionation/Apheresis 69
Frank Wiese
Introduction 69
History of Plasmapheresis 70
Principles of Plasmapheresis 73
Membranes and Membrane Properties 76
Plasma Separation Membranes 76
Plasma Fractionation Membranes 79
Membrane Production 81
Operational Modes in Plasmapheresis Procedures 83
Medical Indications for Blood Plasma Treatment 88
61
Contents
4
4.1
4.2
4.2.1
4.2.2
4.2.3
4.2.4
4.2.5
4.2.6
4.2.6.1
4.2.6.2
4.2.6.3
4.3
4.3.1
4.3.1.1
4.3.1.2
4.4
4.4.1
4.4.1.1
4.4.1.2
4.4.1.3
4.4.1.4
4.4.1.5
4.4.2
4.4.2.1
4.4.2.2
4.4.2.3
4.4.2.4
4.4.2.5
4.5
4.5.1
4.5.2
4.5.2.1
4.5.3
4.5.3.1
4.5.3.2
4.5.4
4.5.4.1
4.5.4.2
Membranes in the Biopharmaceutical Industry 91
Anthony Allegrezza, Todd Ireland, Willem Kools, Michael Phillips,
Bala Raghunath, Randy Wilkins, Alex Xenopoulos
Introduction 91
Microfiltration Membranes Used in the Biotech Industry 93
Introduction 93
Microfiltration Membranes: Development of Industrial
Membranes 94
Effect of Membrane Structure on Properties 95
Aspects of Cartridge Design 97
Membrane Surface Modification 98
Sterilizing Filters 99
Retention 99
Permeability 101
Capacity 102
Practical Membrane Considerations for Sterile Filtration by
Microporous Membranes 102
Sterile Filtration Process Considerations 103
Filter Selection 104
Device Selection 110
Ultrafiltration and Virus Filtration Membranes for
Biopharmaceutical Applications 113
Ultrafiltration Membranes 113
Membrane Suppliers 114
Membrane Selection 114
Membrane Structures 115
Characterization 115
Devices 116
Virus Filtration Membranes 117
Membrane Suppliers 117
Membrane Structures 118
Devices 118
Characterization 120
Applications of Ultrafiltration Membranes in Biopharmaceutical
Manufacturing 121
Ultrafiltration Theory 122
Typical Ultrafiltration Process 123
Process Development and Optimization 124
Processing Plan Optimization 128
Mode of Operation 128
Diafiltration Mode/Strategy 129
Scale-up Considerations 131
Process Implementation Considerations 132
Process Robustness 133
VII
VIII
Contents
4.5.5
4.5.5.1
4.5.5.2
4.6
4.6.1
4.6.2
4.6.3
4.6.4
4.6.5
4.6.6
4.6.7
4.7
4.7.1
4.7.2
4.7.2.1
4.7.2.2
4.7.2.3
4.7.3
5
5.1
5.2
5.2.1
5.2.2
5.2.3
5.2.4
5.3
5.3.1
5.3.1.1
5.3.1.2
5.3.2
5.3.2.1
5.3.2.2
5.3.3
5.3.4
5.3.5
5.3.6
5.3.7
5.3.8
5.4
5.5
System Considerations 134
Equipment Options 134
Process Control Options 134
Practical Aspects of Virus Filtration Process Design and
Implementation 135
Process Development and Optimization 138
Capacity 140
Small-Scale Simulation 141
Pilot-Scale Studies 141
Virus Validation Studies 142
Implementation 143
Membrane Adsorbers 145
Membrane Chemistries 146
Current Applications 147
Flow-Through Polishing 147
Flow-Through Precapture 148
Large Molecule Bind–Elute Purification 149
Future Trends 149
Membrane Applications in Red and White Biotechnology 155
Stephan Lütz, Nagaraj Rao
Introduction 155
Types of Membrane Processes in Red and White
Biotechnology 156
Bubble-Free Aeration 156
Filtration Processes 157
Dialysis and Electrodialysis 157
Adsorption of Microorganisms 158
Examples of Membrane Processes in Biotechnology 158
Bubble-Free Gassing 158
Hydrogen 158
Oxygen 159
Membranes for Cell Retention 161
Higher Cells/Red Biotechnology 161
Whole-Cell Biotransformation 162
Membranes for Enzyme Retention 162
Membranes for Cofactor Retention 166
Application of Dialysis and Electrodialysis in Biotransformations
Application of Pervaporation and Stripping in
Biotransformations 168
Nanofiltration and Ultrafiltration in Biotechnology 170
Bioelectrochemical Applications 170
Summary 172
Acknowledgment 172
167
Contents
6
6.1
6.2
6.2.1
6.2.2
6.3
6.3.1
6.3.2
6.3.2.1
6.3.2.2
6.4
7
7.1
7.2
7.2.1
7.2.2
7.2.3
7.2.3.1
7.2.3.2
7.2.3.3
7.2.3.4
7.3
7.3.1
7.3.1.1
7.3.1.2
7.3.1.3
7.3.1.4
7.3.2
7.3.2.1
7.3.2.2
7.3.2.3
7.4
8
8.1
8.1.1
8.1.2
8.2
Membranes in Controlled Release 175
Nicholas A. Peppas, Kristy M. Wood, J. Brock Thomas
Introduction 175
Controlled Release Kinetics 176
Diffusion in Membrane-Controlled Release 176
Physical Parameters of Controlling Release 180
Membranes and Solute Transport 180
Characterization of Membranes 180
Solute Transport in Network Membranes 182
Structural Parameters of Membranes 182
Determination of Molecular Pore Sizes 183
Applications in Drug Delivery 184
Drug Delivery Through Skin: Overcoming the Ultimate
Biological Membrane 191
Dimitrios F. Stamatialis
Introduction 191
Human Skin – Fundamentals of Skin Permeation 192
Human Skin Structure 192
Stratum Corneum – Main Drug Barrier 193
Drug Transport Through the Skin 195
Passive Diffusion 195
Iontophoresis 198
Electroporation 204
Other Methods 207
Transdermal Drug Delivery System – Structure/Design 212
Passive TDD Systems 212
Types 212
Materials 213
Skin or Device-Controlled Delivery 214
Commercialization – Patents 215
Active TDD Systems 216
Types 216
Materials – Devices 216
Commercialization – Patents 218
Conclusions and Outlook 221
Application of Membranes in Tissue Engineering
and Biohybrid Organ Technology 227
Thomas Groth, Zhen-Mei Liu
Introduction 227
Application of Membranes in Blood Detoxification 227
Requirements to Support Adhesion and Function of Cells
Application of Membranes in Tissue Engineering 231
228
IX
X
Contents
8.2.1
8.2.2
8.2.3
8.2.4
8.3
8.3.1
8.3.2
8.3.3
8.4
8.5
9
9.1
9.1.1
9.1.2
9.2
9.2.1
9.2.2
9.2.2.1
9.2.3
9.2.3.1
9.2.3.2
9.2.3.3
9.2.4
9.2.5
9.2.5.1
9.3
9.4
Introduction to Tissue Engineering and Membrane
Applications 231
Tissue Engineering of Skin 232
Tissue Engineering of Bone 236
Further Tissue Engineering Applications of Membranes
Membranes in Biohybrid Organ Technology 242
Organ Failure and Biohybrid Organ Technology 242
Biohybrid Liver 245
Biohybrid Kidney 249
Summary and Conclusions 253
Acknowledgments 254
240
Membranes in Bioartificial Pancreas – An Overview
of the Development of a Bioartificial Pancreas, as a Treatment
of Insulin-Dependent Diabetes Mellitus 263
Ana Isabel Silva, António Norton de Matos, I. Gabrielle M. Brons,
Marı́lia Clemente Velez Mateus
Introduction 263
Diabetes and Its Treatment 263
The Bioartificial Organ Concept 265
Bioartificial Pancreas 266
Immunoprotection and Biocompatibility of Implanted Devices
Vascular Devices 268
Biocompatible Materials in Vascular Devices 269
Extravascular Devices 272
Implantation Sites 272
Macrocapsules 273
Microcapsular Devices 280
Influence of Recipients Sensitization to Donor Antigens in
Graft Survival 308
Islet Oxygenation Studies 308
Oxygenation of Ba-Alginate Microencapsulated Islets 311
Final Comments 312
Acknowledgment 313
Index
321
267
XI
Preface
Life without membranes has never been possible. They are essential components
of our body. But in the past 50 years the membrane scientists managed to
incorporate artificial membranes in medical and engineering processes, which
improved the life quality of a great part of the world population. Just to mention
two examples, there are currently around 1.5 million hemodialysis patients
worldwide, who decades ago would have been condemned to die at an early
age. Two billion people are affected by water shortages in over 40 countries.
Reverse osmosis desalination plants assure water supply to a large area of the
planet, which barely has access to alternative sources. In the past decades the
membrane technology established its place in the chemical industry. Rapidly
growing applications are the nanofiltration of fine chemicals and the recovery of
monomers in polymerization plants. The electronic industry depends on clean
water obtained by membrane systems. The globalization of the food market
requires active packaging ‘‘membrane’’ systems to allow the transport of fruits
and vegetables for long distances and storage times. Membrane processes have
been for a long time the integrated steps for the cheese and wine production,
caring for the improvement of the product quality and for the introduction of
new products. But this is just the start. Recently, the concern about global
warming and the contribution of the greenhouse effect has increased dramatically. As a consequence, there is a growing demand for renewable energy carriers
with low emission of CO2. Membranes are the core components of fuel cells and
electrolysers. Apart from this, the implementation of the membrane processes in
a modern society that is dependent on hydrogen technology and renewable
energy is a long-term process, and in this transition period membranes are
expected to play an important role. Fossil fuels are still the predominant source
of hydrogen. Membranes can make the hydrogen production and processing in
refineries much more effective. Processes such as steam reforming and water
shift reaction can be significantly improved with the use of membranes. Membrane technology can also be involved in the CO2 emission reduction in refineries and coal plants.
After publishing the book ‘‘Membrane Technology in the Chemical Industry,’’
we clearly saw a demand for more information on membranes and their
ISBN: 978-3-527-31480-5
XII
Preface
applications in different fields. A series of six books with the following topics was
then initiated: Membranes in Life Science, Membranes for Energy, Membranes in the
Food Industry, Membranes for Water Treatment, Membranes for Chemical Technology,
and Membrane Preparation.
Suzana Nunes, Klaus-V. Peinemann
XIII
Anthony Allegrezza
Millipore Corporation
Bioprocess R & D
80 Ashby Road
Bedford, MA 01730
USA
Willem Kools
Bioprocess Applications Engineering
900 Middlesex Turnpike
Billerica, MA 01821
USA
I. Gabrielle Brons
University of Cambridge
Department of Surgery
Cambridge Institute of Medical
Research
CIMR Box 139
Addenbroke’s Sie
Cambridge CB2 2XY
UK
Zhen-Mei Liu
Martin-Luther-Universität HalleWittenberg
Institute of Pharmacy
Kurt-Mothes-Strasse 1
06120 Halle (Saale)
Germany
Thomas Groth
Martin-Luther-Universität HalleWittenberg
Institute of Pharmacy
Kurt-Mothes-Strasse 1
06120 Halle (Saale)
Germany
Todd Ireland
Bioprocess Sales Management
Billerica, MA 01821
USA
Stephan Lütz
Forschungszentrum Jülich GmbH
Institut für Biotechnologie 2
Leo-Brandt-Str.
52425 Jülich
Germany
Marilia Mateus
IBB-Institute for Biotechnology and
Bioengineering
Centre for Biological and Chemical
Engineering
Instituto Superior Técnico
Av. Rovisco Pais
1049-001 Lisboa
Portugal
ISBN: 978-3-527-31480-5
XIV
António Norton de Matos
University of Porto
ICBAS and HGSA-Santo António
General hospital
Division of Angiology and Vascular
Surgery
Largo Professor Abel Salazar
4099-001 Porto
Portugal
Ana Isabel Silva
IBB-Institute for Biotechnology and
Bioengineering
Centre for Biological and Chemical
Engineering
Institute Superior Técnico
Av. Rovisco Pais
1049-001 Lisboa
Portugal
Klaus-Viktor Peinemann
21502 Geesthacht
Germany
University of Twente
Faculty of Science and Technology
Biomedical Technology Institute
Membrane Technology Group
PO Box 217
7500 AE Enschede
The Netherlands
Nicholas Peppas
The University of Texas
Departments of Chemical and
Biological Engineering
University Code C 0400
Austin TX, 78712
USA
Michael Phillips
Bioprocess R & D
80 Ashby Road
Bedford, MA 01730
USA
Bala Raghunath
Bioprocess Applications Engineering
Billerica, MA 01821
USA
Nagaraj Rao
Rane Rao Reshamia
Laboratories Pvt. Ltd.
Plot 80, Sector 23,
CIDCO Industrial Area
400705-Navi Mumbai
India
J. Brock Thomas
Austin TX, 78712
USA
Jörg Vienken
BioSciences Department
Fresenius Medical Care
Else Kroenerstrasse 1
61346 Bad Homburg
Germany
Frank Wiese
Consulting Membrane Technology/
Application
Starenstrasse 100
42389 Wuppertal
Germany
Randy Wilkins
Bioprocess Field
Marketing
Billerica, MA 01821
USA
Kristy M. Wood
Austin TX, 78712
USA
XV
1
1
Membranes in Hemodialysis
Jörg Vienken
1.1
Introduction
The treatment of kidney patients by hemodialysis with tubular sheet or capillary
membranes represents a convincing success story in medical therapy. When John
Abel in the United States and Georg Hass in Germany started to investigate the
application of membranes at the onset of the twentieth century, they would never
have dreamt that nearly a hundred years later, in 2005, about 1.9 million kidney
patients would undergo treatment for end-stage renal disease and nearly 1.5 million
patients would owe their lives to hemodialysis and related devices such as membranes, tubing systems, and so on [1] (Figure 1.1). One reason for this overwhelming
success may be the successful miniaturization of dialyzers that facilitates the mass
production of these devices. Further, and as of today, the exclusive use of capillary
membranes considerably contributes to better performance properties of dialysis
membranes. Therefore, the following figures may not be surprising: In total, in 2005,
dialysis patients all over the world are currently in need of about 150 million dialyzers
annually. Compared to 2002, this figure represents an increase of 15%. Taking into
account that one dialyzer contains a length of about 3 km of capillary membranes,
nearly 450 million km of these materials have to be manufactured annually. The
figure reflects the unimaginable dimension of an annual capillary membrane
production being equivalent to nearly three times the distance between the earth
and the sun.
The technical skill of manufacturing capillary membranes in such large quantities
represents half a century of an effective research and development. A reproducible
high capillary membrane quality in terms of constant geometry, adequate performance, and high-level blood compatibility guarantees the safety of patients. Further,
membrane production at a large scale allows for a reasonable cost of production and
consequently for a facilitated access to these devices all over the world. Membranes
are the centerpiece of dialyzers and more than 650 different types of dialyzers are
currently available on the market. How they are characterized and how they differ,
will be the scope of this chapter.
ISBN: 978-3-527-31480-5
2
1 Membranes in Hemodialysis
Fig. 1.1 The ESRD patient population is
continuously increasing and adds up to 1 900 000
patients worldwide in 2005. A majority of 1 297 000
patients is treated by hemodialysis.Source: FMC
database 2005.
1.2
Historical Achievements
Following Gekas [2], a membrane is both an intervening phase to separate two phases
and a barrier to the transport of matter between phases adjacent to it. The separation
of water and toxins from blood, as observed in the glomeruli of the kidney, is also a
membrane process. Therefore, it is of no surprise that physicians searched for ways
to treat renal failure through the application of artificial membranes [3] (Table 1.1).
The first artificial membranes were handmade from collodium, a cellulose-nitrate
derivative. These collodium membranes are considered to be the forerunners of
today’s capillary hollow fiber membranes. They were used in the early experiments by
Abel [4] on rabbits and dogs. Abel termed for the first time the devices containing
collodium membranes as ‘‘artificial kidneys.’’ The membrane transport process
applied in these experiments was dialysis, that is, transport of low-molecular-weight
compounds such as urea and creatinine through a dense membrane driven by a
concentration gradient (Figure 1.2)
Tab. 1.1
1748
1861
1913
1923
1932
1938
1943
1968
1969
1974
1983
1987
1997
2000
2001
2005
Chronicle of capillary membranes and persons/companies behind.
Pig bladder as a membrane
Definition of term ‘‘dialyzer’’
Collodium (cellulose nitrate) tubes in dog dialysis
Collodium tubes for hemodialysis in humans
Continuous production of tubular membranes
Cellophane, Heparin
Rotating drum with cellophane
Capillary dialyzer with cellulose acetate
Polyacrylonitrile (PAN, AN69)
Cuprophan hollow fiber
Polysulfone (PSu), high flux, biocompatible
DEAE-modified cellulose Hemophan
Vitamin E-bonded membrane
Polyamide–PES membrane
Helixone, PSu, nanocontrolled spinning
Vitabran E, vitamin E-bonded PSu
Jean Antoine Nollet
Thomas Graham
John Abel
George Haas
Richard Weingand
William Thalhimer
Willem Kolff
Richard Stewart
Hospal Company
Werner Bandel
Ernst Streicher
Enka Company
Terumo Company
Gambro Company
ASAHI Medical Ltd
1.2 Historical Achievements
Fig. 1.2 The famous Chemist and Scotsman Thomas
Graham (1805–1869) was the first to define the term
‘‘dialyzer’’ in 1861.
The credit for pioneering the first hemodialysis in humans, however, goes to the
German physician Hans Georg Haas (1886–1971), who developed a system for
hemodialysis in 1923 independent of Abel in Baltimore, USA. Like Abel, Georg Haas
used collodium tubes [5–7]. Haas commented on his experimental efforts in
membranes by saying, ‘‘I’d like to say, it was a way of the Cross, because once,
one obstacle had been removed another followed.’’ and ‘‘Above all, it was necessary to
find a suitable dialysis membrane. I have tried a series of different dialyzers from a
variety of materials, animal, vegetable membranes and paper dialyzers. The best
performances were obtained by dialyzers from Collodium with respect to fabrication,
dialysis effects, safe sterilization and, because they can be obtained in any geometric
shape’’ [7]. Haas could profit from the results published earlier in 1921 by Arnold
Eggerth from the Hoagland Laboratories in New York. Eggerth had shown that it was
possible to achieve a tailor-made membrane clearance by just choosing the appropriate dilution of the alcoholic solvent [8].
In 1932, Weingand [9] proposed an apparatus for a continuous manufacturing of
membrane tubes from cellulose solutions. The cellulose solution issues from the
annular nozzle and flows – guided by gravity – along a guide member. ‘‘The solidified
tube is treated in a series of baths: wasting, bleaching, desulfurizing, and plasticizing
baths.’’ The application of a desulfurizing bath in Weingand’s experiments leads us to
the conclusion that the cellulose solution was a viscose solution (viscose – cellulose
xanthate in NaOH).
Progress in the development of dialysis membranes and dialyzers was not made
until 1938, when William Thalhimer [10] discovered that a membrane used in the
sausage industry could be employed in removing solutes from the blood. This manmade cellulose hydrate material called cellophane was manufactured either by the
Visking Company in Chicago, USA, or the Kalle Company in Germany. Cellophane
membranes turned out to be uniform in thickness and strength, and one that could
be produced in large quantities. Thalhimer experimented with dog blood and used
for the first time heparin as an anticoagulant in hemodialysis [11].
3
4
Fig. 1.3 Dialysis became a reality when Kolff
developed his ‘‘Rotating Drum.’’ Kolff’s
concept of a rotating drum with wound tubular
membranes from Cellophane/Cuprophan was
the basis of dialysis therapy until the early
sixties of the last century. The right-hand photo
shows Kolff (right, and 89 years old in 1999)
explaining his ‘‘Rotating Drum’’ to Jörg Vienken
(left) and Horst Klinkmann (center) during a
visit in Bad Homburg in 1999. A modified Kolffdialysis machine in use in Glasgow until 1965 is
shown on the left.
This was the first contribution to the long lasting success story of cellulosic
membranes in hemodialysis that were later considered to be the golden standard.
The next milestone in membrane application for hemodialysis was the development of the ‘‘Rotating Drum’’ by Willem Kolff. Supported by the local industry,
Kolff [12–14] treated his first patient in a hospital in Kampen, Netherlands. His
dialyzer contained a tubular membrane made from cellophane. By this means,
Kolff was able to separate 40 g urea from blood in 6 h (Figure 1.3).
The first dialyzer of Kolff [12] contained membrane tubes (Cellophane, later
Cuprophan) with a sufficiently large membrane surface area and made an efficient
clinical dialysis possible. Kolff commented at that time: ‘‘It had an additional
advantage: one could repair a leak in the cellophane tubing.’’ Kolff reported later
that ‘‘Once a nurse handed me a pair of scissors over the artificial kidney. Between of
us, we dropped it and there were light holes in the cellophane. We then stopped the
rotating drum, cut out the damaged spots of cellophane tubing, spliced the two
ends over a glass tube coated with soft rubber, and went on with the dialysis’’ [14]
(Figure 1.4).
According to Klinkmann, Falkenhagen, and Courney [3], the Swede Niels Alwall
can be considered to be the inventor of ultrafiltration. Ultrafiltration needs a pressure
difference as a driving force and can separate molecules with molecular weights up to
1 million by convectional forces. Alwall [15] used a device that allowed to exert a
negative pressure on the dialysis membrane by inserting the membrane into a
pressure-stable housing. By this means, he could control ultrafiltration and consequently overhydration in uremic patients.
1.2 Historical Achievements
Fig. 1.4 The chemical formula of the major membrane
polymers. Cellulosic membranes are mainly hydrophilic by
means of their hydroxyl groups, whereas most synthetic
membranes are basically hydrophobic and have to be
rendered partially hydrophilic by adding hydrophilizing
agents, such as polyvinylpyrrolidone.
The first artificial kidney with tiny capillary membranes was described by Richard
Stewart in 1964 [16]. The membrane material used for the production of membranes
at that time was cellulose acetate. Later, cellulose acetate hollow fiber membranes
were accompanied by membranes from unmodified regenerated cellulose. Based on
the traditional German Bemberg experience with the Cuprophan/Cuoxam process
[17], the ENKA Company in Wuppertal, Germany (later Membrana GmbH,
Wuppertal, Germany) developed the first cellulosic tubular membranes in cooperation with the American nephrologist L. Bluemle, followed by the development of
capillary membranes. For both the membrane types, cuoxam was used as a solvent for
the membrane polymer.
Cuprophan hollow fiber membranes represent the classical dialysis membranes
with a dense morphology. Its ultrastructure consists of cellulosic microfibrils, which
5
6
leads to a high mechanical stability of the tiny capillaries, while maintaining its high
flexibility. As a consequence, it was possible to produce capillary membranes with
very small wall dimensions: thickness between 5 and 11 mm is the standard till date.
These small dimensions allowed the manufacturing of compact and handy dialyzers
with a small extracorporeal blood volume.
In 1974, the ENKA Company started the routine production of Cuprophan
hollow fiber membranes [18] in Europe followed by the first clinical studies in
1977. Since that time and as a standard, dialyzers contain about 11 000 capillaries
representing a membrane surface area of >1 m2 [17]. In contrast to ENKA, CD
Medical Inc. (USA), the inventor of capillary membranes for dialyzers, still
continued to use cellulose acetate as a raw material for the following reason: by
saponification of cellulose acetate the morphology of the cellulose membrane could
be rendered more spongeous. As a result, membrane transport properties could be
improved and resulted in the formation of a membrane with an increased
ultrafiltration profile. This finally leads to the development of the so-called
high-flux dialysis [19].
Today, standardly applied high-flux membranes are typically made by new synthetic materials such as polysulfone (PSu), polyethersulfone (PES), polyacrylonitrile
(PAN), polyamide (PA), or polymethylmethacrylate (PMMA).
In 1983, Fresenius [20] started to produce hollow fibers made of polysulfone for
high-flux dialysis in parallel to the development of dialysis machines with ultrafiltration control. Ultrafiltration control is the prerequisite for high-flux dialysis; it allows
to run membranes under high convective forces. In parallel to Fresenius, the Gambro
Company [21] with its subsidiary in Germany started to manufacture polyamide
capillary membranes for hemofiltration.
In 1985, Henderson and Chenoweth published an article entitled ‘‘Cellulose
membranes, time for a change?’’ [22]. Based on investigations on blood/membrane
interaction, they challenged cellulosic membranes for their alleged lack of blood
compatibility in terms of cell and complement activation.
Previous experiments had shown that blood compatibility depends on the chemical composition of the membrane surface. An optimal surface modification would
then improve blood membrane interaction. Investigations lead to the development of
membranes from both modified cellulose and synthetic polymers, which showed
domain-like surface structures with lipophilic/hydrophilic areas [23]. Membranes
with this composition proved to show reduced direct interaction with blood and its
components and thus a better biocompatibility. Modified cellulosic membranes such
as cellulose acetate and cellulose-ether Hemophan, as well as the synthetic counterparts of cellulosic materials, polysulfone, polyamide, or polymethylmethacrylate are
examples for such biocompatible membranes.
During recent years, the performance of dialysis membranes has been further
improved through the introduction of nanotechnology into the production process.
With the nanocontrolled spinning technology [24] defined porosities in membranes
could be achieved as found in the new Helixone membranes in the FX Series of
dialyzers from Fresenius Medical Care. A further improvement for the Helixone
membrane bases on a special Moiré-like ondulation structure of the fiber, facilitating
1.3 Membranes for Hemodialysis: Polymers and Nomenclature
the easy entry of dialysis fluid into the center of the membrane bundle and thus
optimizes the concentration gradient of diffusible uremic toxins.
1.3
Membranes for Hemodialysis: Polymers and Nomenclature
A wide spectrum of hemodialyzers and filters combined with a multitude of different
membranes are currently offered in the market. In 2004, more than 650 different
dialyzer types with membranes made of 22 different polymers were commercially
available [25].
Chemical and physical behaviors of a dialysis membrane are primarily determined by its polymer composition. The sterilization mode of the final dialyzer,
resistivity against sterilizing agents possibly used for reprocessing of the dialyzer,
and biocompatibility are influenced by the type of polymer. Polymers for dialysis
membrane manufacturing typically originate from the textile industry due to its
spinning expertise. The ideal polymer suitable for dialysis should easily be
manufactured to a biocompatible membrane family whose members exhibit
different hydraulic permeabilities and display a considerable physical strength
and excellent diffusive properties, as well as the resistance to all chemicals and
sterilizing agents used in hemodialysis procedures, including steam (temperature
to stand: >121 8C). Additionally, some appreciate the ability of a hemodialysis
membrane to adsorb endotoxins at the outer surface because it is of substantial
benefit against microbial contamination of the dialysis fluid. However, only the
polysulfone and the polyamide membrane families fulfill these demands completely at the moment.
As can be concluded from historical details, hemodialysis membranes of today are
high-tech products that are tailored to the scientific-based demands of hemodialysis
therapy. Even ‘‘early’’ materials such as regenerated cellulose are subject to permanent improvement by their manufacturers, and actual types vary in several aspects
from their counterparts made 10 years ago, although they may not reach the maximal
demands formulated above. Two classes of materials are currently used for the
production of dialysis membranes: cellulose and synthetics that can be divided into
two and three subclasses, respectively.
Some structural differences exist between these two classes beside their different
polymers: cellulosic membranes have relatively thin walls (in the range of 6.5–
15 mm) and a uniform (symmetric) composition across their entire capillary wall in
order to achieve a high diffusive solute transport. This thin membrane does not
provide enough strength to withstand high ultrafiltration rates as are employed in
convective therapies. Therefore, most cellulosic membranes are not suitable for
convective dialysis treatments like hemodiafiltration or hemofiltration with the only
exception of cellulose triacetate (CTA) membranes. In contrast to the classical
cellulose membranes, synthetic membranes exhibit a membrane wall thickness
of 20 mm and more, which may be symmetric (e.g., PMMA) or asymmetric (e.g.,
Fresenius Polysulfone) (Figures 1.5 and 1.6).
7
8
Fig. 1.5 Cross section of a capillary membrane made from
the regenerated cellulose (Cuprophan). The membrane
wall exhibits a homogeneous structure with a tiny wall
thickness of 8 mm and an inner diameter of 200 mm.
1.3.1
Membranes from Regenerated Cellulose
For the production of cellulosic membranes, purified cellulose (e.g., linters) is solved
in an ammonia solution of cupric oxide (origin of the brand name Cuprophan). The
Fig. 1.6 Cross section of a capillary membrane made from
Fresenius Polysulfone. The polsulfone membrane is
represented by a small layer of about 1 mm thickness at the
luminal side of the capillary. A heterogeneous spongeous
wall of about 39 mm thickness guarantees mechanical
stability.
cellulosic polymer chains are newly arranged in the spinning process, therefore, the
term ‘‘regenerated cellulose’’ is used. The result of both the spinning processes is a
macroscopically homogenous structure that is extremely hydrophilic, sorbs water,
and thus forms a hydrogel. Actually, diffusion of solutes takes place through water
swollen amorphous regions. Regions of crystalline cellulose give the necessary
mechanical strength. Therefore, cellulosic membranes can be made very thin and
down to a thickness of 6.5 mm. It is considered as an advantage that the hydrophilic or
water attracting properties of the membrane allow only little protein adsorption and
thus avoid secondary layer formation.
The Japanese manufacturer Asahi Medical terms its cellulosic membrane
‘‘cuprammonium rayon’’ and the German manufacturer Membrana GmbH sells
its product under the brand name Cuprophan. Originally, cellulosic membranes
have all been of low hydraulic permeability. Today, dialyzers with regenerated
cellulose membranes and higher fluxes with ultrafiltration coefficients up to
50 mL/h mmHg (Bioflux, H 2000, IDEMSA, Spain) are available on the market.
In RC-HP 400 the average diameter of pores is increased from 2.76 nm for the
standard regenerated cellulose to 7.23 nm that was achieved in the production
process by the following ways: (i) an alteration of the concentration in the polymer
solution, (ii) the solvent content in the coagulation bath, and (iii) the velocity of
membrane formation through the coagulation bath [26]. In order to maintain the
physical strength, wall thickness also had to be elevated to 18.5 mm (RC-HP 400)
and 20 mm (Xanthogenate), respectively. The sieving coefficient (SC) for b2-m,
however, is only 0.3 for RC-HP 400, and elimination takes place solely by membrane
transfer. However, these more permeable membranes are not suitable for convective treatments, like hemofiltration or even hemodiafiltration, because their
mechanical strengths are not appropriate and their sieving coefficients for b2-m are
too low.
Membranes made of regenerated cellulose have – due to their low wall thickness – a
good low-molecular-weight clearance that is comparable and in a good dialyzer even
superior to low-flux synthetic membranes. The other advantage of unmodified
regenerated cellulosic membranes is their resistance to all sterilization modes. A
disadvantage, however, is the poor biocompatibility, and bacterial products are not
retained by some synthetic low-flux membranes.
1.3.1.1 Modified Cellulosic Membranes
As it became obvious that cellulosic materials and especially, the nucleophilic
hydroxyl groups of the cellulose polymer interact with humoral systems of blood,
for example, with complement proteins, modifications of the polymer were produced
mostly in the direction of substitution of the hydroxyl groups [23] in the molecular
backbone of the cellulose molecule.
1.3.1.2 Cellulose Acetates
In cellulose acetate (CA) membranes, at least two of the three hydroxyl groups of the
glucose monomer have been replaced by acetyl groups. In the presence of sulfuric
acid, cellulose is treated with pure acetic acid and acetic anhydride to form a cellulose
9
10
acetate ester. This ester bond is responsible for the thermolability – CA cannot be
sterilized by steam – and the chemolability of the CA polymer at pH >7.
Some reports exist about polymer degradation resulting in polymer fragments
such as acetylated carbohydrate derivatives that may lead to adverse reactions in the
patients within 24 h after cellulose acetate treatment, for example, anaphylactoid
reactions, scleritis, iritis, tinnitus, conjunctivitis, red eye syndrome, visual loss, or
hearing loss [27,28].
The acetate substitution makes the hydrophilic cellulosic backbone polymer more
hydrophobic and allows some protein adsorption on the inner surface. Several types of
CA membranes are currently on the market, differing in their degree of hydroxyl-group
substitution, their hydraulic permeability, and their manufacturing process (meltspun
or coextrusion). All these differences may lead to some small advantages of one CA over
the other [29]. Altogether, CA membranes exhibit the good performance characteristics
of their unmodified cellulosic counterparts but are improved in their biocompatibility,
although not reaching the excellent profile of some synthetic membranes.
Althin Medical AB, Sweden (acquisition by Baxter Healthcare Corporation, USA,
in the late 1990s) took over the Cordis Dow meltspun process from 1963 and
produced a symmetric cellulose diacetate membrane under the brand name Althane
that used to be available in a high-flux as well as in a low-flux version. Dialyzers with
this kind of membrane are held responsible for a number of deaths among dialysis
patients in Spain, Croatia, and Texas, USA, in 2001. Intensive investigations [30–32]
revealed that residual perfluorocarbon (PF5070), a processing fluid used for the
repair of leaky dialyzers, is the reason for these fatal reactions. The production of this
kind of Althane dialyzers had therefore been ceased in 2001, according to a press
release of Baxter Healthcare in 2001.
Cellulose triacetate with a thin uniform skin structure was first used only as highly
porous membrane for high-flux dialysis or hemodiafiltration (Nissho/Nipro) and was
the first and only cellulosic membrane capable for more convective therapies than
hemodialysis. Today low-flux types are also available. The highly permeable CTA
membrane is also suitable for continuous renal replacement therapy [33].
1.3.1.3 DEAE-Modified Cellulose, Hemophan
Assuming that cellulose acetate was the ‘‘oldest’’ polymer used for hemodialysis,
N,N-dimethyl-aminoethyl (DEAE) modified cellulose was the first cellulosic membrane that was modified on purpose in order to increase biocompatibility (brand
name Hemophan, Membrana Germany, was available during the years in dialyzers
from Baxter, Bellco/Sorin, B. Braun, Gambro AB, Haidylena, Cobe, IDEMSA, JMS,
Kawasumi, NephroSystems/Meditech, and Nikkiso). Only 1.5 % of all hydroxyl
groups of the cellulose molecule is replaced in this polymer by tertiary amino groups
through stable ether bonds [34]. These positively charged bulky groups first set
hydrophobic spots on a hydrophilic surface and then led to a steric hindrance of the
interaction between complement factors and the membrane. A marked improvement of the biocompatibility profile in comparison to unsubstituted regenerated
cellulose with equal performance characteristics is observed. In respect to C5ageneration and activation of leukocytes no statistically significant difference could be
observed during dialysis with DEAE-cellulose in comparison to ETO-sterilized
polysulfone low-flux dialyzers. But the positively charged DEAE groups are under
suspicion to lead to an increased platelet activation as well as to heparin consumption
during treatment. Heparin adsorption could be demonstrated with DEAE cellulose
from saline as well as from blood [35,36].
Advantageous effects are found in that the positively charged DEAE groups attract
the negatively charged phosphate molecules from plasma, improving phosphate
clearance along with that [37,38].
DEAE-cellulose exhibits the hydraulic properties of regenerated cellulose (Cuprophan) and is available as low-flux and high-flux types (Hemophan HP, UFcoeff 24 and
32 mL/h mmHg, respectively, g-wet Nikkiso [39]). Due to its stable ether modification of cellulose, the membrane is sterilizable by all current methods.
In early 2006, Membrana GmbH, Wuppertal, Germany, announced the termination of the production of cellulosic membranes by the end of 2006. Thus, membrane
polymers, which have influenced the development of kidney therapy to a high degree
over decades, will disappear from the market and leave the forum to their synthetic
counterparts.
1.3.1.4 Benzyl-Modified Cellulose (Synthetically Modified Cellulose, SMC)
Another example for the creation of hydrophobic domains on a hydrophilic surface is
synthetically modified cellulose) produced by Membrana GmbH Wuppertal [40]. It is
available in different housings from different manufacturers (under the brand
name Polysynthane from Baxter, as SMC from Bellco/Sorin, B. Braun and Kawasumi). In this polymer, less than 1 % of the hydrophilic hydroxyl groups (OH) of the
cellulosic backbone have been replaced by hydrophobic benzyl groups through ether
bonds. This modification leads to an improved biocompatibility in comparison to
unmodified regenerated cellulose, although it does not reach that of low-flux
polysulfone [41]. Furthermore, biocompatibility parameters such as elastase release,
complement activation, and leukopenia are timely delayed during dialysis with
benzylcellulose: the nadir of leukopenia is at 30 min rather than around 15 min
as it is with other modified or unmodified cellulosic membranes. Maximum
complement activation and elastase release are also observed after 30 min [42].
However, all the cellulosic properties are sustained in this low-flux membrane type,
which is available as for sterilizability with all current methods, mechanical strength,
and good low-molecular-weight clearances [43].
1.3.1.5 PEG-Grafted Cellulose
Improvement of biocompatibility with cellulosic membranes was also achieved by
grafting the cellulosic backbone of cuprammonium rayon with a polyethyleneglycol
(PEG) layer (AM-BIO membrane, Asahi Medical Co. Ltd, Tokyo, Japan). Alkylethercarboxylic acid (PEG acid) is esterified with its terminal carboxyl group to the hydroxyl
group of the cellulose [44]. These PEG chains form a so-called hydrogel layer on the
cellulosic surface (thickness 2.4 nm) that may act as a buffer zone between the
cellulosic backbone and blood hindering the direct contact of plasma proteins with
the membrane surface. This may lead to a reduction in platelet adhesion and
11
12
complement activation. Lower platelet adhesion could be observed with this kind of
membrane in vitro in comparison to unmodified and PEG-grafted cellulosic membranes [45].
1.3.1.6 Vitamin E-Modified Cellulosic Membranes
The first attempt to create a ‘‘bioreactive’’ dialysis membrane is the development of a
vitamin E-coated (D-a-tocopherol) cellulosic membrane with the brand name
‘‘Excerbane’’ from the Terumo Company in Japan (today in 2006: ASAHI Medical
[46]). Here, performance capabilities of a high porosic cellulosic membrane are
combined with the biocompatibility features of a synthetic copolymer together with
the therapeutic approach to reduce the oxidative stress during treatment by supporting the body’s own antioxidant defense mechanisms with the supplementation of
vitamin E [47].
The surface modification is carried out during the fiber spinning process: the
modifying solution consists of a hydrophilic acrylic polymer with reactive epoxy
groups. This polymer, based on fluorescein, possesses an inhibitory effect on
complement activation. A second polymer, an oleyl alcohol chain, shows inhibitory
properties with platelet aggregation. Both are dissolved into the core solution. The
original solved regenerated cellulose and the core solution, containing the modifier,
are both coextruded through the spinerette into the coagulation bath, where two
phases are formed. The outer circumference of the hollow fiber is composed of the
cellulosic membrane, and the primary hydrophilic inside is covered by a hydrophobic
layer of the modifier. This takes place by covalent bonding of the reactive epoxy
group of the modifier with the hydroxyl groups of the cellulosic membrane. The
amount of vitamin E immobilized via hydrophobic bonding to oleyl alcohol is around
150 mg/m2 [48].
Some clinical reports exist already with this kind of membrane showing an
improved biocompatibility in comparison to regenerated cellulose. The long-term
therapeutical effect of this kind of vitamin E supplementation has to be further
elucidated.
1.3.2
Synthetic Membranes
The majority of synthetic membrane polymers currently on the market are basically
hydrophobic and have to be made more hydrophilic by additives or copolymers
during the production process. The polymer mixture is extruded through a spinerette
followed by phase inversion and immersion. Partial evaporation of the solvent is
responsible for skin formation. Phase separation determines the structure of the
membrane as well as its crystallinity.
The main purpose to develop synthetic membranes was to create membranes with
higher porosity in order to mimic more the natural kidney filtration process [20] and
remove middle molecules and higher molecular weight uremic toxins like b2-m. As
an indication for a possible pore size dimension of high-flux dialysis membranes, the
molecular radius (Stokes radius) of b2-m is described as being 1.6 or 2.2 nm. It was
calculated that the pore radius of a b2-m removing membrane should be greater than
the double radius of the molecule, which is approximately 5 nm for b2-m removal
and lower than 8 nm to avoid albumin loss [49].
1.3.2.1 Polyacrylonitrile (PAN)
The company Rhone Poulenc in France has been the first manufacturer who brought
a highly permeable, symmetric, synthetic membrane on the market, which is a blend
of a copolymer of the hydrophobic polyacrylonitrile with the hydrophilic methallylNa-sulfonate. The medium-sized pores are distributed in high density over the
homogenous polymer, and glycerol is used as a pore filler. This membrane produced
and marketed under the brand name AN69 by Hospal (Gambro) till date has been of
great success over the years.
ASAHI Medical is the other PAN producer (PAN, PAN DX, ASAHI Medical,
Japan). Their membrane consists of the hydrophobic monomers, acrylonitrile and
methacrylate and is made hydrophilic by the addition of acrylic acid. Due to the
special process, the membrane is asymmetric and exhibits a skin layer with pores in a
wide range determining the sieving properties of the membrane.
Until the early 1990s AN69 was held as one of the most biocompatible dialysis
membranes, although it has one drawback: like all PAN membranes it is not
sterilizable by heat. Moreover, since 1990, reports were published reporting anaphylactoid reactions with AN69 in combination with the consumption of angiotensinconverting enzyme (ACE) inhibitors [50–53]. Contact phase activation of the
kallikrein–kinin system on the negatively charged surface of the AN69 membrane
resulting in bradykinin formation was identified as the underlying mechanism.
Surface electronegativity (zeta potential, see also Table 1.4) of the membrane, the pH
of the rinsing solution [54,55], as well as the dilution factor of the plasma were found
to be the influencing factors for the extent of reaction. These observations have also
been found true for the second polyacrylonitrile membrane currently on the market,
which also leads to bradykinin generation but to a lower extent: its zeta potential was
found to be lower, which was 60 mV in comparison to 70 mV of AN69.
Due to its microstructure and its surface electronegativity, AN69 exhibits through
negative charges of the sulfonate groups, a high adsorption capacity for proteins,
especially in positively charged proteins [58]. This was proven for complement factor
D, b2-microglobulin, and low-molecular-weight proteins [56,57]. This feature is of
advantage with respect to complement activation and b2-m elimination, but negative
with respect to high-molecular-weight kininogen adsorption resulting in contact
activation and higher residual blood volume in comparison to polysulfone membranes. Furthermore, therapeutic proteins like erythropoietin (EPO) are also
adsorbed, probably diminishing their effectiveness. In cases where EPO is administered subcutaneously after hemodialysis treatment, this is of no clinical importance.
But protein adsorption in substantial amounts may be a disadvantage, if dialyzers
should be reused, because they have to be cleaned with aggressive chemicals
like sodium hypochlorite that may damage their fibers in order to rebuild their
performance.
13
14
Both polyacrylonitrile membranes are available only with high hydraulic permeability suitable for high-flux dialysis, hemodiafiltration, and hemofiltration. AN69 is
additionally available in flat sheet format incorporated into plate dialyzers for acute
hemofiltration.
In order to overcome problems with anaphylactic reactions, Hospal developed a
new AN69 type by coating the polyacrylonitrile flat sheet membrane with polyethylenimine (PEI), a polycationic polymer [59]. During the manufacturing process PEI is
sprayed onto the membrane surface until an optimized concentration of 9 mg/m2 is
reached. In the resulting AN69ST (ST stands for surface treated [59]) the zeta
potential is reduced to around 0 mV. Negatively charged sulfone groups of the
AN69 polymer should be masked by the polycationic polymer and further provide a
steric barrier via its thickness of the layer. First reports exist about clinical experiences
with such a modified plate dialyzer (Crystal ST): No hypersensitivity reaction could be
observed during 3 years and 10 630 dialysis sessions, among them 3400 with ACE
inhibitor therapy. These data have to be confirmed by other centers to make sure that
this new membrane is safe for the patients even in combination with ACE inhibitor
therapy [60].
Coating the membrane with a positively charged polymer furthermore leads to
another effect, the binding of negatively charged heparin. First clinical trials show
that if heparin was brought onto the membrane during the priming procedure no
further systemic heparinization was needed during treatment. Heparin release from
the membrane was undetectable. This anticoagulation regimen was possible only
with unfractionated heparin, whereas low-molecular-weight heparin desorbs quickly.
It will be of great interest to know how the introduction of positive charges on the
surface will alter biocompatibility in comparison to the ‘‘old’’ AN69 type. That
biocompatibility will be different is quite sure because changes in electronegativity
of the surface have a great impact on protein adsorption.
1.3.2.2 Polymethylmethacrylate (PMMA)
PMMA membranes were introduced into the market in 1977 in Japan by Toray
Industries and have been the first g-ray sterilized synthetic membranes [49]. They
consist of a hydrophobic, nonpolar polymer produced from methylmethacrylate
monomers, homo-PMMA, or are in some type also copolymerized with the addition
of small amounts of p-styrene sodium sulfonate. The polymer forms a membrane
that is symmetric, almost homogeneous, and isotropic. In current PMMA membranes the pore radius varies between 2 and 10 nm, having a volume fraction of pores
(porosity ranging from 50 to 70 % [49]).
A membrane family consisting of nine different members was created with lowflux types (B2, B3 series), high-flux types (B1 series), which are mostly used for HD
today, and a version with increased b2-m removal (BK series U/P/F). The last was
achieved by increasing the pore size, resulting also in increased adsorption
properties for b2-m of the membrane [61]. The BK-F model is the most strongly
b2-m adsorbing HD-membrane currently on the market, where convective b2-m
removal is negligible [49,62]. The larger pores (10 nm) also allow the removal of
larger uremic toxins. An erythopoiesis-inhibiting fraction, KR4-0 and its subfraction YS-1 (MW 40 000), has been isolated from dialyzate of a PMMA BK-F HD
treatment [63]. A clinical benefit could be demonstrated through the fact that EPO
doses could be diminished to half in patients receiving BK-F treatment after 2 years
in comparison to HD with standard cellulosic membranes. Whether this is a
unique property of the BK-F membrane or it can also be achieved with other highly
permeable membranes has to be further elucidated. A multicenter, randomized,
controlled trial with 84 patients for 12 weeks could not confirm any effect on anemia
with highly porous membranes (BK-F PMMA) in comparison to low-flux cellulose
hemodialysis [64].
The strong adsorbing property of PMMA has also some disadvantages: it also
results in an undesired adsorption of platelets to the membrane with all its effects on
fibrin formation. A higher residual blood volume was also reported for PMMA
membranes. Due to the negative surface, anaphylactoid reactions occurred under
ACE inhibitor therapy. Therefore, with respect to biocompatibility, the membrane is
placed only at an upper level among the synthetic membranes. All dialyzers containing PMMA are sterilized by g-irradiation. The material is not steam sterilizable.
1.3.2.3 Polysulfone (PSu)
Not less than 18 suppliers of polysulfone dialyzers are currently on the market
underlining the great success of this membrane polymer. Polysulfone fits all
demands of a modern polymer: It is sterilizable with all methods (g-ray, b-ray,
ethylene-oxide, steam), biocompatible, has physical strength, and chemical resistance. The membrane material exhibits its good performance characteristics both in
its low-flux as well as in its high-flux versions that remove considerable amounts of
b2-m by filtration. Moreover, polysulfone is suitable as an endotoxin adsorber and
thus an active protection system for contaminated dialysis fluids (Table 1.2).
Because of all these advantages more and more membrane producers have
developed their own polysulfone, although it is sometimes hidden among difficult
nomenclature. Fresenius introduced the first high-flux polysulfone in 1983 [20],
followed by the low-flux version in 1989. Table 1.2 provides an overview about the
different manufacturers of polysulfones and some special features of the particular
polymer. Due to patent protection, all polysulfones developed till date have to be
different from the original Fresenius Polysulfone. They differ in their basic copolymer/polymer alloy, the addition of polyvinylpyrrolidone (PVP) (polysulfone alone is
hydrophobic and has to be made more hydrophilic, which happens in most cases by
blending the polymer with the hydrophilic PVP or not), and in their entire production
processes, resulting in different morphologies.
In chemistry, the terminus ‘‘polysulfone’’ comprises simply a group of polymers
containing sulfone groups and alkyl or aryl (e.g., arylether) groups. However,
according to chemical convention, all such polymers that additionally contain
isopropyliden groups are termed as polysulfones (Fresenius Polysulfone, Asahi
Polysulfone, Toraysulfone). Those dialysis membrane polysulfones that do not
contain isopropyliden groups are termed as polyarylethersulfones or shortly as
15
16
Membranes of the polysulfone (PSu) and polyarylethersulfone (PES) family on the
market and their characteristics.
Tab. 1.2
Manufacturer
Brand name
Sterilization
Flux
Fresenius
Polysulfone
InLine steam ETO
e-beam
Low
Helixone
ASAHI Medical
Gambro AB
Hospal/Cobe
Kimal
Membrana GmbH (Allmed, Baxter,
Bellco, Helbio, Haidylena,
IDEMSA, Kawasumi, Saxonia)
Minntech
Nikkiso
Saxonia (B. Braun)
Toray Industries
High
Low
High
High
APS
Vitabran E
Polyamix (PES)
PolyamideSa
Arylane (PES)
Polyethersulfone
DIAPES
g-wet
g-wet
Heat
g-dry
g-dry
b-dry
Low
High
High
High
Low
Purema (PES)
Minntech PS
Polyphen
PEPA
a-Polysulfone
Toraysulfone
g-dry, steam
ETO
ETO
g-wet
g-dry
g-wet
High
High
High
High
High
High
a
According to the manufacturer, PolyamideS is a copolymer of polyarylethersulfone and only a small
amount of polyamide. Therefore, it is mentioned under polysulfones [65].
Source: data from Ref. [1].
polyethersulfones (DIAPES, Arylane). This is a little bit confusing because, as
mentioned above, all dialysis membrane polysulfones include an arylether. PEPA
and PolyamideS contain another polymer in addition to polyarylethersulfone: PEPA
polyarylate and PolyamideS polyamide and PVP [66]. The polyamide in PolyamideS is
a matter of debate at the moment because some investigators could not find any
polyamide in PolyamideS [65].
PEPA is the only polysulfone membrane that does not contain PVP and therefore
exhibits some special characteristics: it adsorbs larger quantities of b2-m, which is
uncommon for all other membranes, but despite this adsorption removal rates for
b2-m are even lower than that of the other polysulfone membranes.
Recently an improved version of the original Fresenius Polysulfone, the Helixone
membrane, was developed [24,67]. The polymer is unchanged, but the wall
thickness (from 40 to 35 mm) and inner diameter of the fiber (from 200 to
185 mm) are reduced. By applying nanotechnology-based fabrication procedures
[24,67] for the first time in hemodialysis, the nominal average pore size has been
increased in Fresenius Polysulfone from 3.10 to 3.30 nm. With the advanced
production process, it was possible to create an almost uniform pore distribution at
the dense innermost layer and a homology in pore size that results in a sharper
1.4 Dialyzer Constructions
molecular weight cut-off. The sieving coefficient for b2-m was extended to 0.8,
whereas the sieving coefficient for albumin was preserved in the range between
0.001 and 0.01. First clinical studies revealed no difference in biocompatibility in
comparison to that of Fresenius Polysulfone, due to no changes in polymer
composition [68]. A considerably better performance was found with helixone
due to higher mass transfer coefficients, when compared with that of urea,
creatinine, phosphate, and b2-m clearances, which is explained by a new geometric
dialyzer, fiber design and pore size dimensions [69].
1.3.2.4 Polyamide (PA)
Polyamide membranes (in Polyflux dialyzers, Hemoflux hemofilters, and FH hemofilters, Gambro Hechingen, Germany) consist of a hydrophobic aromatic-aliphatic
copolyamide that is blended with hydrophilic polyvinylpyrrolidone [70]. This mixture
leads to a microdomain structure with alternating positively and negatively charged
regions on the surface. This is held responsible for the good biocompatibility of the
polymer. The membrane is asymmetric with three distinguishable regions: a thin
skin of 0.1–0.5 mm on the blood side followed by a sponge structure of 5 mm that is
supported by a finger structure of about 45 mm. Pore size increased dramatically
from the blood side to the dialyzate side, being smallest at the skin layer with around
5 nm.
Polyamide dialyzers and filters exhibit good b2-m removal with a SC of 0.6. This is
due to their performance characteristics and not due to adsorption because protein
adsorption is very low with this membrane. However, bacterial products are successfully rejected at the dialyzate side.
Polyamide dialyzers and filters are available only in ETO-sterilized form. In
order to overcome this disadvantage, polyarylethersulfone was added to the blend.
The new membrane PolyamideS is heat sterilizable and described under polysulfone
membranes.
1.4
Dialyzer Constructions
Biocompatibility, the size of removed particles, and the possibility of sterilization
mode are mainly determined by the dialysis membrane, and all other aspects of
effective dialysis are related to the design of a dialyzer. Current dialyzers are available
only as hollow fiber devices.
1.4.1
Hollow Fiber Dialyzers
A current hollow fiber dialyzer consists of a housing that contains one membrane
fiber bundle. This is fixed at its both ends in the housing by polyurethane (PUR). An
advanced cutting process is necessary to form a smooth surface after embedding in
order to minimize activation of humoral or cellular systems in the blood. This surface
17
18
is covered by caps that contain the ports for the inlet and outlet of blood respectively,
and in recent developments also for dialysis fluid. Normally the dialyzer housing
contains two ports for the inlet and outlet of the blood and the dialysis fluid,
respectively. Behind this relatively simple construction stand the 60 years of knowledge and development. The trend to minimize devices is evident, if Kolff’s artificial
kidney from 1943 [12] is compared to the latest dialyzer of today. Current dialyzers are
optimized with respect to nearly each component.
1.4.2
Housing
The size of each dialyzer housing is tailored to the performance characteristics of the
included fiber bundle. The blood compartment volume and resistance to blood flow
have to be as low as possible, and each fiber has to be equally washed round with
dialysis fluid. The number of fibers and the filling degree of the dialyzer housing
increase with surface area until an increase in surface is inefficient for a rise in
performance. Then the housing has to be increased. The comparison of dialyzers
from various manufacturers that contain the same membrane clearly demonstrates
the differences and what an advanced technology stands behind the creation of a good
housing design.
Current trends in device technology are the introduction of materials that can be
disposed in an environmentally friendly manner. Polypropylene is such a material
that is used in the newest housings instead of polycarbonate [69]. Furthermore
comparing dialyzer performances, housings are getting smaller and smaller since
the development of the first disposable hollow-fiber dialyzer. This is possible because
membranes with higher performance characteristics are available due to reduction of
wall thickness, reduction of inner surfaces of fibers, and a special bundle design.
Smaller dialyzers avoid waste, have a sparing effect on transport costs, and are
beneficial for the patient because blood contact surface area as well as blood volume in
the extracorporeal circuit is reduced. An advantage for the clinic is the saving of
storage space.
1.4.3
Potting Material
Composition of the potting compound changed over years in order to minimize
risks of toxic substances that may evolve after sterilization of the polyurethane.
Especially, the irradiation with b- or g-beams may lead to the fission product 4,40 methylene dianiline, a proven carcinogenic substance [71]. The amount of PUR was
considerably reduced over time, as it was obvious that it functions as a reservoir for
ETO leading to allergic reactions in the patients [72]. Furthermore, effective surface
area is increased with less potting area. Some manufacturers treat the reduced
surface of the potting area in a special manner in order to avoid blood activation
(e.g., Hospal Arylane series, FMC FX class series). Also, polycarbonate and/or
1.4 Dialyzer Constructions
silicon rings were introduced that allowed the PUR content of the hollow fiber
dialyzers to be reduced.
1.4.4
Fiber Bundle
Of considerable importance for the performance of a dialyzer is the fiber bundle
construction that has been dramatically improved till date. In early dialyzers, kinked
fibers resulting from insufficient potting technology were a big problem leading to
considerable thrombus formation. In modern high quality dialyzers, this problem
has been overcome due to sophisticated embedding and cutting procedures. Even
fiber distribution is nearly uniform in high quality products. In order to improve
dialysis fluid flow around the fibers, several bundle configurations have been
developed and tested. The dialysis fluid flowed across the fibers rather than along
them. Multiple bundles, with the use of a solid central core with fibers wound in a
spiral manner, or warp knitted hollow-fiber mats have been introduced into dialyzer
technology. Today, the newest fiber developments use an undulation of the fiber itself
in order to provide space for a continuous uniform flow of the dialysis fluid. Besides
this, fibers with fins and bundles with spacer yarns are also on the market [73–76]
(Figures 1.7 and 1.8).
Fiber bundle size and swelling of the membrane determine the priming blood
volume, which is an important parameter in the choice of the dialyzer for patients
with low blood volume, especially children. Today, the blood volume in most dialyzers
is smaller than that of the blood tubing sets.
Fig. 1.7 Dialyzer performance considerably
depends on the arrangement of capillaries in
the membrane bundle. In order to allow for an
easy access and flow of dialysis fluid to the
center of the bundle, capillary membranes have
to be either separated through the insertion of
spacer yarns (b in left-hand panel) or have to be
ondulated (a in left-hand panel). An ondulation
following the geometry of a Moiré structure, as
realized in FX-dialyzers with helixone
membranes (c) guarantees optimal clearance
characteristics through a homogeneous flow of
dialysis fluid.
Source: modified from Ref. [73].
19
20
Fig. 1.8 The new FX-class of dialyzers exhibits improved
clearance characteristics due to a modified PSu membrane
and a special ondulation configuration of the capillaries. As
a consequence, urea clearance is increased as compared to
traditional dialyzer configurations. This allows for reducing
dialysis fluid consumption at comparable blood flows.
1.5
Dialysis Membranes and Performance: Principles of Membrane Transport
Dialysis is a membrane separation process in which one or more dissolved species
flow across a selective barrier in response to a difference in concentration and a
difference in pressure. Concentration differences refer to the mode of transport of
diffusion, where separation occurs because small molecules diffuse more rapidly
than larger ones. In the absence of difference of pressure and temperature across the
membrane, A. Fick’s phenomenological description of diffusion, published in 1855
[77], states that solutes will move from regions of greater to regions of lesser
concentration [DC] and at a rate proportional to the differences (Figure 1.9).
J ¼ DADC=Dd;
ð1Þ
with J representing solute flow, D the diffusion coefficient, A the area, and d the
distance between the two separated compartments, that is, the membrane thickness
in the dialysis situation. The minus sign for the diffusion coefficient accounts for the
1.5 Dialysis Membranes and Performance: Principles of Membrane Transport
Fig. 1.9 Annual growth rates for dialyzer sales show that
high-flux and synthetic membranes have become the
dominating products in dialysis therapy.
Source: re. FMC market survey 2003.
convention that flux is considered positive in the direction of decreasing concentration. According to Albert Einstein, diffusion concentration decreases roughly in
proportion to the square root of molecular weight. In order to increase solute flow,
membrane manufacturers have tried for years to have a membrane thickness as small
as possible.
Complications may arise from the observation that blood in contact with artificial
surfaces, like a hemodialysis membrane, leads to the formation of a secondary
layer of proteins. This boundary layer affects diffusional transport across a
membrane.
A further effect for reduced diffusional clearance can be attributed to drug
administration to the patient, such as human recombinant erythropoietin (rhEPO).
Because the removal of solutes by the process of hemodialysis is dependent on the
flow of the solute and water from the blood compartment to the dialyzate compartment, it is comprehensible that the fraction of water in the blood is reduced by raising
the hematocrit (Hct). Indeed, increasing Hct from 20 to 40 %, leads to a reduction of
creatinine and phosphate clearances of 8 and 13 %, respectively. Urea removal,
however, is less affected [78]. Based on these observations, it appears prudent to
increase hemodialysis prescription, that is, treatment time, by 10–15 %, when Hct is
raised to near 40.
Dialysis membranes are primarily categorized according to their permeability for
water and solutes. The ultrafiltration factor (UF) and sieving coefficients are the
appropriate parameters. Ultrafiltration factors refer to the amount of filtered water in
mL/h in relation to the applied transmembrane pressure (TMP) in mmHg, which is
exerted through the blood pump of the dialysis machine. The sieving coefficient is
equivalent to the amount of a given solute removed (Figure 1.10):
SC ¼
2CF
þ CBo ;
CBi
ð2Þ
21
22
Fig. 1.10 The weekly performance of the human kidney
overrides by far the performance of classical dialysis
therapies, such as low- and high-flux dialysis. A closer
bridging the gap between the performance of conventional
therapy and the human kidney will certainly be based on
convective therapies such as hemodiafiltration.
whereby CF refers to the concentration of solute in filtrate and CBi and CBo to the solute
concentrations in blood inlet and blood outlet, respectively.
Consequently, the maximum value for a sieving coefficient is ‘‘1’’ representing a
100 % sieving. Following Table 1.3 and a general agreement, low-flux membranes
are characterized by an UF factor of less than 10 mL/h mmHg, whereas high-flux
membranes have an UF factor greater than 10 mL/h mmHg. In addition to their UF
factors, hemofilters and high-flux dialysis membranes are characterized by their
sieving coefficient for b2-microglobulin (b2-m). b2-m is a small protein of a
molecular weight of 11.818. High serum levels of b2-m are considered to be
causative for dialysis-related amyloidosis. The molecular weight cutoff of high-flux
membranes, which refers to the molecular weight of molecules that cannot pass
through the dialysis membrane, is typically around 60 000. It implies that albumin
is retained in the patient’s blood.
Tab. 1.3
Categories for hemodialysis membranes.
Category
Low flux
High efficiency
High flux
Hemofilter
UF factor
(mL/h mmHg)
Cut-off molecular
weight
Sieving coefficient
(for b2-m)
Surface
area (m2)
<10
>10
>10
>20
–—
–—
<60 000
>60 000
–—
–—
>0.6
>0.6
<1.5
>1.5
–—
–—
Low- and high-flux refer only to the hydraulic permeability.
1.5 Dialysis Membranes and Performance: Principles of Membrane Transport
By increasing the transmembrane pressure (e.g., by means of the peristaltic
pump), filtration flow will increase. This increase is almost linear as long as aqueous
solutions are used. With whole blood, the increase of filtration flow tends to become
nonlinear because of the formation of a secondary layer protein coat by protein
adsorption, which is a fast process. Albumin, one of the proteins with the highest
blood concentration, is already deposited in 50 ms after the contact with the dialysis
membrane [79].
These observations have to be taken into account when fixing ultrafiltration rate
and dialysis time. The amount of ultrafiltered water, for example, in the treatment of
acute renal failure, may determine the mortality of patients as shown by Storck et al.
[80], who compared pump-driven and spontaneous continuous hemofiltration in
postoperative renal failure. Increasing the ultrafiltration volume per day improved
mortality by about 20%.
Recent investigations in the treatment of kidney patients have focused on the
removal of large molecular weight solutes such as cytokines, complement proteins,
and proteins modified either by glucose (advanced glycation end products [AGEs]) or
by oxidative stress pathways.
Transport of large solutes across a dialysis membrane is improved if convectional
transport mechanisms are applied. Convective clearance is defined by
CConvection ¼ SC QF;
ð3Þ
and thus determined by the membrane’s sieving coefficient (SC) and the filtrate
flow (QF). It is possible to increase the sieving coefficient for a given solute by
switching from a low-flux to a high-flux membrane. Filtrate flow depends on
internal filtration, which is defined as the total water flux across the membrane
within the closed blood and dialyzate compartment of a dialysis filter. Internal
filtration depends on the pressure gradient characteristics (Dp) along the length of
dialyzer and may be given by a simplified approach through the application of
Hagen–Poiseulle’s law:
Dp ¼
8hLQB
;
NpR4
ð4Þ
where h is the blood viscosity, L is the length of dialyzer membrane, QB is the blood
flow, N is the number of capillaries in the dialyzer, and R is the internal radius of a
capillary membrane.
Consequently, a better convective clearance (higher Dp) directly depends on blood
viscosity to be adapted through the administration of erythropoietin, the length of the
dialyzer, and the blood flow. Applying higher blood flows, that is, >300 mL/min is
therefore mandatory in high-flux dialysis or in hemodiafiltration. Convective clearance
depends indirectly on the number of capillary membranes in a dialyzer and on
the membrane’s luminal radius to the power of four. Thus, reducing the inner
diameter of a capillary membrane would considerably increase the convective clearance [73,81].
23
24
Fig. 1.11 Data sheets of dialysis membranes refer to
ultrafiltration in relation to the applied transmembrane
pressure. However, the amount of ultrafiltrate (QF)
depends on the exclusive presence of water/saline or
blood. Due to a secondary layer formation by blood
proteins, QF is leveling off at higher TMPs.
Modern capillary membranes, for example, Helixone polysulfone membranes
(Fresenius Medical Care, Germany, inner diameter: 185 mm) have a reduced
inner diameter as compared to the classical diameter dimension of 200 mm for
typical dialysis membranes. Decreasing the inner diameter from 250 mm over
200 mm down to 175 mm has no impact on small solute clearances such as urea, but
shows considerable consequences for the removal of larger solutes such as
b2-microglobulin, for which convective clearance increases threefold [81]
(Figures 1.11–1.13).
1.6
Dialysis Membranes and Biocompatibility
1.6.1
Some Basic Information on Membranes and Biocompatibility Parameters
The biocompatibility of a dialyzer, particularly that of its membrane, is one of the
main criteria that influences dialyzer choice. During the ‘‘Consensus Conference on
Biocompatibility held in 1993, biocompatibility was defined as ‘‘the ability of a
material, or device, or system to perform without a significant host response in a
specific application’’ [82]. The definition was weakened by adding ‘‘significant’’ in
comparison to first explanations because it was realized over the years that any
foreign material would lead to some kind of reaction in the host.
In hemodialysis treatment, blood/artificial surface interactions (as special aspect of
biocompatibility also termed as hemocompatibility) are of special importance
because of its chronic treatment character.
1.6 Dialysis Membranes and Biocompatibility
Fig. 1.12 Reducing the inner diameter of a
capillary membrane leads to a high
pressure drop along the length of the
dialyzer and thus favors convective forces for
the removal of larger solutes. Here, a reduction
of the inner diameter of a polysulfone
capillary membrane from 200 to 175 mm
leads to a nearly 80 % increase of inulin- and
vitamin B12-clearance.
Source: modified from Ref. [73].
Fig. 1.13 Improved clearances can be obtained
for small proteins, such as b2-microglobulin
(b2-m), if the inner diameter of a capillary
membrane is reduced. b2-m clearance is found
to be up to threefold higher, when the inner
diameter is reduced from 250 mm over 200 mm
down to 175 mm. With this simple change in
fiber geometry convective therapies, such as
hemofiltration or hemodiafiltration, gain an
even better performance.
25
26
Parameters of biocompatibility are manifold and their number increases continuously with scientific knowledge. In the beginning of chronic hemodialysis (CRF)
treatment, thrombogenicity and hemolysis were the main matters of concern for
physicians [83].
Manufacturers focused on the development of less thrombogenic surfaces and
on the toxicity of the material that should not liberate plasticizers and chemical
additives. In the late 1970s, hypersensitivity reactions in response to ethyleneoxide (ETO) sterilization of the dialyzer were reported [84]. At the same time,
complement gained attention and was linked to the transient leukopenia and
sequestration of neutrophils in the lung during dialysis with cellulosic membranes [85,86]. The influence of chronic hemodialysis treatment on the immune
system was recognized in more detail over the years with the formulation of the
‘‘Interleukin Hypothesis’’ in 1983 being a milestone in this kind of research [87].
Stimulation of immune cells during dialysis, the release of mediators such as
cytokines, and the pathological consequences thereof have from then on been
topics of intensive research. In the early 1990s, the possible interaction between
extracorporeal treatment and pharmacological action of some drugs became
clinically evident: Anaphylactoid reactions had been observed in ACE inhibitor
patients in combination with the use of certain negatively charged dialysis
membrane, which until then had been considered ‘‘biocompatible’’ [88]. Investigations revealed that the kallikrein–kinin system was involved in this sort of side
effects leading the attention to the new biocompatibility parameter bradykinin
(Figure 1.14).
What are the biocompatibility parameters of today? Some parameters can be
measured relatively easily in the clinical setting, such as the differential cell count,
whereas factors of the complement and clotting cascade need specialized laboratories. On a more in-depth scientific level, an assessment of gene expression in
relation to cytokine or b2-microglobulin generation, and the stimulation of immune
cells and their reactions (receptor expression, generation of reactive oxygen species
[ROS], etc.) have become state of the art.
1.6.2
Thrombogenicity of Different Types of Dialyzers and Filters
The thrombogenic potential of a dialyzer or filter is strongly dependent on the type of
membrane polymer, its permeability, the design of the device influencing flow
conditions, and its manufacturing quality.
Surface free energy, charge, roughness, and chemical composition of a dialyzer
membrane have been identified as being responsible for the variable thrombogenic
potential of dialysis membranes [89]. As the extrinsic system of the clotting cascade
is activated by negatively charged surfaces, membrane polymers that exhibit an
anionic surface, such as polyacrylonitrile membranes (PAN DX, PAN; AN69) could
be considered more thrombogenic than more neutral polymers such as polysulfone.
Fig. 1.14 Contact of blood with negatively
charged biomaterial surfaces leads to the
activation of the contact phase and
subsequently to the formation of the
nonapeptide bradykinin via the kallikrein
pathway. If degradation of bradykinin by
angiotensin-converting enzyme is blocked
through the administration of ACE-inhibitors,
blood pressure is downregulated by
prostaglandin generation. Generally speaking,
synergistic effects between biomaterial or
membrane properties and the administration
of pharmaceutical drugs may not be neglected
in the future.
In fact, the highest factor XII adsorption and autoactivation to factor XIIa can
be observed with AN69 membranes in vitro [90]. However, the activation of the
coagulation cascade also depends on other plasma constituents, for example, highmolecular-weight kininogen, prekallikrein, plasma inhibitors of XIIa, and the action
of heparin [90]. High-molecular-weight kininogen (HMWK), another constituent of
the trimolecular complex initiating contact activation, could be eluated in its intact
form from the used cellulose acetate (Cordis Dow) and regenerated cellulose
(Cuprophan), whereas it was cleaved at PMMA and polyacrylonitrile (PAN;
AN69) surfaces suggesting activation [91]. Moreover, low-molecular-weight fragments of plasminogen were detected in eluates from polyacrylonitrile (AN69)
dialyzers indicating that the fibrinolytic system was activated in response to clotting
activation. Intact plasminogen was adsorbed also with all other materials [91]
(Figure 1.15).
Clinical studies investigating coagulation activation with different types of
membranes using the ex vivo formation of the thrombin–antithrombin III complex
(TAT) as clotting parameter showed a higher TAT formation with polyacrylonitrile
membranes (Asahi PAN and Hospal AN69) than with polysulfone (Fresenius
Polysulfone) and regenerated cellulose (Cuprophan) [92]. In another study higher
levels with PMMA than with EVAL or polysulfone [93] were shown. These findings
were confirmed in controlled ex vivo studies, which found higher TAT generation,
platelet factor 4 release, and platelet consumption with AN69 (Nephral 300,
27
28
Fig. 1.15 Bradykinin formation after in vivo
contact of patient blood with the surfaces of
different dialysis membranes. Bradykinin levels
in the venous line of negatively charged PAN
(AN69) dialyzers is far above normal and
explains adverse events in these patients. Note
that the concentration of bradykinin in these
cases is found to be in the femtomolar (fmol)
range.
Source: adapted from Ref. [157].
Hospal) in comparison to regenerated cellulose (Cuprophan, Renak RA-15U,
Kawasumi) and polysulfone (Fresenius Polysulfone, Hemoflow F6HPS) [94]. A
low TAT generation with regenerated cellulose is in agreement with the findings
that hydroxyl-bearing surfaces are relatively inert with respect to activation of the
intrinsic coagulation pathway; such membranes inhibit the development of the
trimolecular complex (HMWK, PK Factor XII) that is necessary for contact
activation [95].
DEAE-modified cellulose (Hemophan), a membrane that is a low level platelet
activator, led to more TAT generations than high-flux polyamide [96]. This is not the
result of a higher contact system activation by DEAE cellulose, but it is rather due to
the positive DEAE groups adsorbing negatively charged heparin from whole blood.
This adsorption inactivates heparin and, consequently, sufficient amounts of heparin
are not available for anticoagulation. This is normally reflected in an increased
heparin consumption for Hemophan during treatment [97,98]. However, another
in vivo investigation failed to find such an increased heparin requirement and also did
not observe any increase in TAT formation with Hemophan in comparison to low-flux
polysulfone [99].
Another, albeit less reliable, parameter for the assessment of thrombogenicity is
the residual blood volume (RBV) in the dialyzer postdialysis. This subjective parameter has a number of pitfalls that mostly concern the reproducibility of the method.
Although TAT measurements are preferable, RBV values have been reported in the
literature. Under conditions of low heparin dosage, residual blood volume was found
to be higher with Hemophan than with CA, PMMA and low with Fresenius
Polysulfone [100].
In contrast to coagulation factors, platelets adhere more to cationic charges and
hydrophobic materials [95]. Therefore, it is not surprising that a significant decrease
in the platelet count after 15 min of dialysis of about 9 % could be found with
hydrophobic cellulose acetate dialyzers (Cordis Dow), whereas only insignificant
decreases were observed with cuprammonium rayon (AM50-Bio, Asahi), DEAEcellulose (Hemophan, GFS, Gambro), polyacrylonitrile (AN69 Hospal), and high-flux
polysulfone dialyzers (Fresenius Polysulfone F60, FMC) [101]. In another study, a
similar ranking in platelet drop was observed with the same polymers but different
manufacturers: cellulose diacetate (CA 150 Baxter), Hemophan (Bio-Allegro, Cobe),
cellulose diacetate (Acepal 1500, Hospal), and low-flux polysulfone dialyzers (F6,
FMC) [38].
One reliable parameter for platelet activation is the expression of the glycoprotein GMP-140 at the platelet membrane surface. The highest GMP-140
expression was found for Cuprophan dialyzers, less expression was observed
with cellulose acetate and PMMA, and the lowest expression with polysulfone and
polyacrylonitrile dialyzers [102]. This ranking for platelet activation was also found
for another parameter, the expression of P-selectin on the platelet surface:
Expression was highest with Cuprophan, followed by cellulose diactetate and
cellulose triacetate, and was lowest with Hemophan and Fresenius Polysulfone
[103]. Although platelets are obviously activated by regenerated cellulose, this
activation does not result in or involve activation of the blood-clotting cascade, as
is obvious from previously mentioned investigations into the formation of the
TAT complex.
A number of publications have dealt with theoretical considerations about
artificial surfaces and their thrombogenic potential, but biological processes are
too complex to easily predict the thrombogenic behavior from charges, functional
groups, or other characteristics of current dialysis membranes. Differences
between membranes, and even between dialyzers of different manufacturers
containing the same membrane exist. This has been shown in controlled ex vivo
or in vivo studies using the same blood for all membranes compared. However, the
numbers of reports about membranes correlate with their frequency of use in the
market and not necessarily with their thrombogenic behavior. Therefore, although
no published data are available for some membranes this should not be interpreted
as absence of thrombogenic potential per se.
1.6.3
Complement Activation by Different Dialyzers and Filters
It is accepted today that membranes made of regenerated, unmodified cellulose
are the strongest complement activators among all dialysis membranes. This can be
partly explained by the binding of complement factor C3b to the hydroxyl groups
on the membrane surface. In fact, either partial substitution of these hydroxyl
groups with acetyl groups (cellulose diacetate or cellulose triacetate), DEAE
groups (Hemophan), and benzyl groups (SMC), or coating with polyethylene glycol
(PEG-grafted regenerated cellulose) resulted in considerable reduction of complement activation compared to unmodified regenerated cellulose. Surprisingly, the
29
30
degree of substitution does not necessarily correlate with the degree of complement
activation, demonstrating involvement of other factors. For example, in Hemophan,
less than 1 % of all hydroxyl groups are modified, but complement activation is
at least equal to or even better than that of cellulose acetate membranes where
about 60 (cellulose diacetate) to 90–100 % (cellulose triacetate) of all hydroxyl groups
are substituted [23]. Binding of regulatory proteins that downregulate the alternative
pathway such as ‘‘factor H’’ by Hemophan [104] or ‘‘factor D,’’ a rate-limiting enzyme
of the alternative pathway by AN69 [105] and PMMA [106], also plays a role in
the extent of complement generation. Factors H and B bind onto cellulose acetate
membranes resulting in an accelerated degradation of surface-bound C3b; neither of
these factors bind to regenerated cellulose [107]. However, all measures taken to
prevent the interaction of C3 molecules with nucleophilic groups, for example,
hydroxyl or amino groups, on a polymeric surface reduce activation of the alternative
pathway. This can be achieved either by avoidance of such groups as in synthetic
membranes or by masking of such groups with acetyl, DEAE, or benzyl groups (as in
cellulose acetates, Hemophan, or SMC) or by coating the whole surface with, for
example, PEG groups. In order to avoid activation of the classical pathway, none of the
IgG, IgM, C1, C2, or C4 complement components should be adsorbed to the
polymeric surface [108].
Parameters mostly used for the assessment of complement activation by hemodialysis membranes are the complement factors C3a and C5a. The amount of
detectable C3a or C5a in the effluent of a dialyzer depends on the capability of the
membrane to adsorb or to filtrate these factors besides its grade of generation: The
polyacrylonitrile membrane AN69 generates more C3a than regenerated cellulose,
but adsorbs C3a and C5a almost completely [109]. Furthermore, these anaphylatoxins are eliminated during dialysis by high-flux membranes due to their middle
molecular weight (MW C3a: 8900; C5a: 11 000) [110,111]. This is also the fact for
the regulatory factor D (MW 23 000). This upregulator of the alternative complement pathway is eliminated by glomerular filtration in the healthy individuals and
its concentration is, therefore, markedly elevated in ESRD patients. Factor D
removal is negligible in low-flux dialysis but can be significant in hemofiltration
[112]. Hence, detectable complement generation may differ between the low-flux
and high-flux versions of a particular membrane [113]. Furthermore, binding to
receptors of circulating blood cells may lower the concentrations of C3a and C5a
in blood.
What are the long-term consequences of complement activation in the long-term
perspective? Because complement generation peaks 15–30 min after the start of
treatment and returns to nearly normal levels until the end of treatment, it is of
interest whether this effect is important or detectable in the long-term perspective.
Unfortunately, little data exist on this aspect of complement activation. Predialysis
C3a levels were reported to increase slowly over a time period of 1 year with the low
complement activating membranes, Hemophan and polyamide [114]. In long-term
HD patients (around 8 years on dialysis), the C3 activity is altered in such a way that
both the alternate and the classical pathways are suppressed after stimulation. This
effect lasts more than 4 h after the end of a dialysis session and is more pronounced
in sera of patients treated with regenerated cellulose membranes (AM-SD18M,
Asahi Medical, Japan), than in the sera of those treated with polyacrylonitrile
membranes (PAN17DX, Asahi Medical). Therefore, it appears that due to the
chronic high stimulation three times a week, membranes made of regenerated
cellulose may induce a suppression of complement activation in the long term.
The mechanism has not been fully elucidated, but a contribution of increased
levels of the regulatory complement proteins, factor H and SP-40,40, has been
discussed [115].
1.6.4
Cell Activation by Different Types of Dialyzers and Hemofilters
Essential functions of polymorphonuclear leukocytes are disturbed in ESRD patients
and are additionally influenced by the dialysis procedure, for example, phagocytosis,
oxygen species production, upregulation of specific cell surface receptor proteins,
and apoptosis. With the exception of polymorphonuclear leukocyte degranulation,
complement and consequently, complement generating dialysis membranes have
the greatest impact on functional alterations of these cells.
Leukopenia, the most widespread used parameter to assess leukocyte activation, is
the assessment of the disappearance of leukocytes from the blood 15–30 min after
the start of hemodialysis. This dialysis-induced leukopenia is mainly induced by the
overexpression of receptors (CD11b/CD18, CD15s), leading to an increased adhesiveness and aggregation with subsequent sequestration in the lung. Because these
activation processes are mostly complement mediated, complement activation and
leukopenia directly correlate with each other [116]. Therefore, regenerated cellulose
membranes cause the strongest leukopenia, modified cellulosic membranes cause
an intermediate-to-low drop in leukocyte numbers, depending on the type of
modification, and synthetic membranes also induce a moderate-to-very low drop
depending on the polymer. Leukocyte drop is found to be similar to that of polyamide
(Polyflux 130, Gambro) and polysulfone (Fresenius Polysulfone F60, FMC) after
15 min at approximately 12 %.
Increasing the permeability of regenerated cellulose should, theoretically, result in
an increased removal rate of complement factors and, consequently, an improved
biocompatibility with respect to measurable complement generation and leukocyte
drop. Surprisingly, this was not the case with RC-HP 400 – a high performance
version of Cuprophan – which induces the same leukocyte drop as low-flux
Cuprophan [117].
1.6.4.1 Apoptosis
Dialysis membranes promote neutrophil apoptosis both directly and through interaction with monocytes. Reactive oxygen species seem to be strong mediators of this
process [118]. In in vitro experiments, low-flux cellulose diacetate induced significantly greater apoptosis than low-flux polysulfone [118]. An effect of HD membranes
31
32
Fig. 1.16 The effect of different dialysis membranes on
apoptosis of mononuclear cells. During the in vivo study
period, bacterial and endotoxin contamination was below
detection limit. Cell apoptosis was measured in circulating
MNCs isolated from heparinized blood.
on apoptosis was also found in vivo. The percentage of mononuclear cell (MNC)
apoptosis was high with Hemophan and Cuprophan dialyzers, was moderate with
cellulose diacetate, and was relatively low with both high-flux polyacrylonitrile and
polysulfone membranes. When seven of these patients were switched from Hemophan to Fresenius Polysulfone, apoptosis decreased markedly only after 8 weeks of
treatment [119]. This in vivo study was repeated in vitro with mini dialyzers of
the membranes used and with blood from healthy donors in order to eliminate the
influence of uremia. Interestingly, the results were quite similar implicating a role of
the type of dialysis membrane in inducing apoptosis beside uremia [119]
(Figure 1.16).
1.6.5
Oxygen Species Production – Induction of Oxidative Stress
Dialysis treatment enhances oxidative stress in chronic renal patients who additionally suffer from a chronic deficiency in the major antioxidant systems [120,121].
Detectable manifestations of this physiological condition are the increased plasma
levels of protein oxidation products (oxidation of plasma protein-associated thiol
groups) in HD patients [122]. Oxidative stress may contribute to atherosclerosis,
cardiovascular disease, dialysis-related amyloidosis, and anemia [123,124]. Four
main factors are held responsible: uremia and the comorbid status of the ESRD
patient, antioxidant and trace element losses during treatment, the bioincompatibility of the system, especially of the dialysis membrane, and contamination of blood
with trace amounts of endotoxins introduced by microbial contaminated dialysis
fluid [125,126] (Figure 1.17).
Fig. 1.17 Relative amount of IL-1b gene
expression in peripheral blood mononuclear
cells from healthy volunteers, nondialyzed
uremics and chronic hemodialysis patients
treated with different dialyzers. Maximal
accumulation of IL-1b mRNA was observed in
blood samples taken from the venous line after
5 min of dialysis. IL-1b gene expression could
be detected in cells from all HD-patients, but
not in healthy volunteers and nondialyzed
uremics.
The highest increase in reactive oxygen species production during dialysis was
reported in complement activating membranes [126,127]: Cuprophan induced a
significant production of reactive oxygen species 15 and 30 min after the start of
dialysis. A lower production was observed with Hemophan, cellulose diacetate and
low-flux Fresenius Polysulfone, whereas increases with SMC were not even statistically significant [127]. Another clinical study reported low ROS production with
polyacrylonitrile (AN69) and EVAL [128]. However, ROS production in response to
stimulation by endotoxins is suppressed in patients treated with Cuprophan membranes and normal with low-flux polysulfone or PMMA membranes [126,129].
Apart from measures regarding the improvement of treatment such as the use of
more biocompatible membranes and pyrogen-free dialysis fluid, another approach to
reducing the possible deleterious consequences of the prooxidant status has been
followed: supplementation with antioxidant vitamins such as vitamin E [130]. In
addition to oral prescriptions, some clinical experience has been made in the vitamin
E-coated membrane Excebrane: a 10 % predialysis increase of plasma vitamin E
concentrations (exclusively in HDLs not in LDLs) was observed after a 3-month
treatment period with this membrane, whereas plasma levels did not change when
conventional biocompatible membranes, AN69, Fresenius Polysulfone, or PMMA
membranes were used. This effect is believed to be due to the vitamin E sparing effect
by the Excebrane membrane [130]. The b-carotene content of plasma and the LDL and
HDL levels were about 26 % higher than when other biocompatible membranes were
used, a finding which can be explained by the secondary protective effect of vitamin E
toward b-carotene [131].
However, no improvement in the blood oxidative stress status could be found in
this study, which used determinations of thiobarbituric acid-reactive substances and
33
34
antioxidant defenses, like erythrocyte-Cu, Zn-superoxide dismutase and plasma, and
erythrocyte glutathione peroxidase [130]. In contrast, another investigation detected
an improvement in oxidative stress in the form of a threefold increase in plasma
glutathione concentration and a slight increase in erythrocyte numbers after three
months of treatment [132]. In this study, the improved antioxidant status (increased
vitamin E and glutathione peroxidase concentration in plasma) resulted in a 10-fold
increase in plasma levels of arachidonic acid, one of the most abundant polyunsaturated lipids and a useful marker for lipoperoxidation. Leukocyte function (responsiveness to chemical stimuli) was significantly improved after one month of
treatment with Excebrane, and the number of apoptotic cells was reduced in
comparison to dialysis with regenerated cellulose [132].
An increase in oxidative markers, like malondialdehyde, advanced glycation end
products, and 8-hydroxydeoxyguanosine was prevented with Excebrane but not
during dialysis with Terumo polysulfone [133]. Another study used plasma vitamin
C levels and constant ascorbyl free radical (AFR)/vitamin C ratio postdialysis as an
index of oxidative stress. After AN69-dialysis, basal vitamin C levels were decreased
and the AVR/vitamin C ratio was increased compared to predialysis levels and to
dialysis with Excebrane. Both of these oxidative stress parameters remained nearly
unchanged compared to predialysis values when the Excebrane membrane was
used [134].
Reviewing the available literature, the antioxidant characteristics of Excebrane
compared to other membranes from regenerated cellulose or polysulfone are
controversial. It seems an interesting new approach to offer a specific and timely
protection against oxygen free radicals at their site of generation. However, the
efficiency of membrane-associated effects must be compared with other treatment
forms, such as oral prescription of vitamin E or other antioxidant therapy, for
example, vitamin C supplementation. A comparison of vitamin E-coated membranes
with regenerated cellulose plus additional vitamin C infusion revealed that both
approaches had equal positive effects on thiobarbituric acid reacting substances [135].
It is further argued that vitamin E deficiency is rare in uremic patients because the
vitamin is not removed during dialysis. Lower vitamin E levels may be induced via
vitamin C loss, due to its low molecular weight (MW 176 Da). Vitamin C loss may be
considerably high during dialysis and result in less regeneration of vitamin E after
dialysis [135].
In 2005, Asahi Medical further introduced a vitamin E-bonded polysulfone
membrane into the dialysis market, and thus proved that it is possible to bind
biologically active compounds also to synthetic dialysis membranes and sterilize
them without loss of performance.
1.6.5.1 Degranulation of Neutrophils
During hemodialysis neutrophil degranulation occurs, which is independent of
complement activation but influenced by intracellular calcium and two neutrophil
degranulation-inhibiting proteins: angiogenin (MW 14 000) [136] and complement
factor D (MW 23 000) [137]. Both inhibitors are found to be up to 15-fold elevated in
end-stage renal disease patients protecting against degranulation and consequently
degranulation products like lactoferrin [137]. Both molecules are removed only in
small amounts by highly convective treatments, but some protein adsorbing membranes like PMMA and AN69 adsorb angiogenin [116] and factor D onto their
surfaces [105,106]. AN69 reduced plasma angiogenin levels by 66 % and factor D
level by 37 % during HD treatment, whereas Fresenius Polysulfone reduced plasma
angiogenin concentration by 36 % and had no influence on factor D [116]. The
removal of inhibitors during AN69 dialysis resulted in a more neutrophil degranulation and lactoferrin release than did treatment with polysulfone [116]. PMMA and
regenerated cellulose induced more neutrophil degranulation than Hemophan,
high-flux Fresenius Polysulfone, and polyamide [113].
1.6.6
Stimulation of Cytokine Generation by Different Types of Dialyzers and Hemofilters
Predialysis intracellular interleukin-1 (IL-1) and TNFa levels are higher in
HD patients than in healthy individuals. Furthermore, both cytokines may be
transiently generated during dialysis treatment with nonultrapure dialysis fluid
[138–140]. The same has been described for the soluble receptors, IL-1RA and
TNFsRp55 [138]. Predialysis IL-1b and TNFa in zymosan-stimulated PBMCs
were found to be higher in patients treated with regenerated cellulose than in
patients treated with PMMA and in healthy volunteers [141]. As complement is
involved in the activation process, cytokine mRNA production and complement
activation normally correlate [141,142].
However, a second stimulus is necessary for cytokine protein production and
release. In hemodialysis, this is mainly provided by microbial contamination
products [143] and, if used, by acetate from the dialysis fluid. Therefore, neither
higher IL-1 nor TNFa plasma levels could be observed with membranes from
regenerated cellulose [144] when pure, acetate-free dialysis fluid was used. In
studies using pure dialysis fluid, increased plasma levels of IL-1 were observed to
differ only insignificantly between normal subjects, nondialyzed chronic renal
failure patients, and hemodialyzed patients [143]. Therefore, comparison of results
from different clinical studies is only meaningful if the purity of the dialysis fluid is
similar. In cases of contaminated dialysis fluid, the permeability of a membrane as
well as its ability to adsorb microbial products at its outer surface is key factor for
cytokine release [145].
The level of circulating cytokines that can be determined with modern methods
may also be affected by patient-specific factors such as residual renal function,
different renal and comorbid diseases, drugs, and different cellular productions.
Dialyzer-related influencing factors are permeability/clearance, ultrafiltration
rate, adsorption of cytokines, and stimulation of generation by different materials
[140,146]. A comparison of plasma cytokine levels of circulating IL-1b, IL-6, and
IL-10 measured using hemodialysis over a period of 4 months with high-flux
polyamide and low-flux Hemophan revealed no differences between the membranes [147].
35
36
In contrast, in vitro experiments with isolated monocytes from these patients
reveal a higher state of preactivation of these cells with Hemophan. Preactivated
monocytes secrete high amounts of proinflammatory cytokines when exposed to a
second stimulus such as endotoxin [147]. A second study compared gene expression of IL-1b with different membranes using pure dialysis fluid: No differences in
plasma levels among dialysis patients using different membranes, healthy subjects, and nondialyzed ESRD patients were found, but gene expression of IL-1b was
threefold higher with regenerated cellulose in comparison to PMMA and polysulfone [144]. These data clearly demonstrate that gene expression of cytokines, but
not necessarily their release from cells, is triggered with regenerated cellulose
membranes. Therefore, plasma levels of circulating cytokines may remain stable
during treatment.
1.6.6.1 The Impact of Membrane Types on LPS-Stimulated IL-1b Secretion
Intracellular cytokine levels, such as IL-1b, are increased in hemodialysis patients,
especially during dialysis with regenerated cellulose [139]. This chronic stimulation
obviously leads to a suppression of the response to biological stimulants such as
bacterial products. Lipopolysaccaraides (LPS)-induced IL-1b secretion is reduced in
patients treated with membranes made of regenerated cellulose compared to that of
healthy controls, nondialyzed chronic renal failure patients, and patients dialyzed
with polyacrylonitrile membranes [148]. This functional change in PBMC response is
specific for IL-1b, and it does not happen with TNF-a, is hemodialysis membrane is
dependent and reversible. As possible mechanism may refer to increased prostaglandine E2 (PGE2) levels that suppress IL-1b secretion due to the following observed
effects: when patients dialyzed with regenerated cellulose were switched to PAN
membranes, total cell-associated and secreted IL-1b concentrations remained nearly
constant, but the secreted amount after LPS stimulation increased. In parallel, PGE2
synthesis decreased with PAN compared to regenerated cellulose. Evidence for the
role of PGE2 is further provided by the fact that with the addition of a PGE2 inhibitor,
IL-1b secretion after LPS stimulation improved in PBMCs from HD-patients treated
with regenerative cellulose.
1.6.7
The Impact of Large-Pore Dialysis Membranes on the Inflammatory Response
in HD Patients by Cytokine Elimination
Due to the presence of large pores and their ability to adsorb proteins, high-flux
dialyzers or hemofilters may eliminate some cytokines. Here some limitations of
extracorporeal removal have to be kept in mind: cytokines are molecules of molecular
size ranging from 15 000–30 000, which have a short half-life (in the range of
minutes) and may be bound to carrier proteins such as a2-macroglobulin [149].
Therefore, a detectable amount may only be efficiently removed when convective
treatments like hemofiltration with high ultrafiltration volumes and highly permeable membranes are performed. The adsorption capacity of the membrane for
cytokines is probably of more importance. In contrast to membranes from regenerated cellulose, membranes made from PAN (AN69) are able to adsorb considerable
amounts of IL-1 in vitro [146].
1.6.8
The Effect of Different Dialyzers on the Acute Phase Reaction
IL-6 is the major regulator of the hepatic acute phase response in inflammation: it
stimulates hepatic synthesis of C-reactive protein (CRP) and serum amyloid A (SAA)
up to several 100-fold [150].
Cultured PBMCs from patients dialyzed with Cuprophan spontaneously
released more IL-6 than PBMCs from healthy individuals or from patients dialyzed
with PMMA membranes [151], SMC, or cellulose diacetate [152]. The same effect
was observed with the soluble IL-6 receptor, which probably reflects more the
biological activity of IL-6 [153]. IL-6 release correlates positively with levels of
circulating CRP [150], and hemodialysis patients exhibit elevated levels of IL-6, C
reactive protein and serum amyloid A [150]. The reason for this is yet not clear, but
the role of dialysis membranes and contaminated dialysis fluid is implicated: IL-6
levels were found to be elevated after the third hour of treatment with cuprammonium rayon membranes (Asahi AM-UP-75) and correlated with the release of acute
phase proteins, albeit with different time schedules [150]. Increased concentrations
of CRP and secretory phospholipase A2 (sPLA2) were found 24 hafter the start of
hemodialysis with this membrane. In contrast, dialysis with polysulfone (F60S,
FMC) showed no marked variation at these time points neither for IL-6, CRP nor
sPLA2 levels [150].
1.6.9
Activation of the Kinin System by Different Types of Dialyzers and Hemofilters
The main factor determining activation of the kinin system with subsequent
generation of bradykinin is the electronegativity of the dialysis membrane [154].
Table 1.4 provides an overview of the zeta potentials –as parameter for membrane
electronegativity – of seven frequently used membranes. As is obvious from these
in vitro experiments, AN69 and PAN DX membranes are the most electronegative
membranes and generate the highest amounts of bradykinin, a mediator for
anaphylactoid reactions [154]. When the electronegativity of the polyacrylonitrile
polymer is reduced to near zero, as found for the PAN modification AN 69ST, in
vitro bradykinin generation is reduced 200-fold. This was achieved by coating of the
membrane with the polycationic polymer, polyethyleneimine [154].
Bradykinin generation with PAN membranes could be detected in vivo, especially
in patients under ACE inhibitor therapy. Bradykinin is normally rapidly degraded in
the body by the serine protease kininase II and therefore most patients are symptomfree under AN69 and PAN DX dialyzes. However, if this enzyme is blocked by ACE,
37
38
Zeta potential (surface electronegativity) plasma kallikrein activity and bradykinin
generation with different membranes in vitro.
Tab. 1.4
Membrane
Polyacrylonitrile
AN69
Polyacrylonitrile
PAN DX
Polymethylmethacrylate
Cellulose triacetate
Cuprophan
Fresenius Polysulfone
Polyacrylonitrile
AN69ST
Zeta
potential (mV)
Plasma
kallikrein (U/mL)
Bradykinin
generation (fmol/mL)
70 5
60 15
32 100 (26 500–41 200)
60 4
80 20
28 983 (22 600–36 150)
25 2
20 2
10 1
5 1
3 1
10 5
<5
<5
<5
<5
130
65
78
62
150
(50–130)
(25–100)
(25–150)
(25–120)
(30–450)
Source: adapted from Ref. [154]
inhibitors plasma levels may increase by more than 100-fold and patients will suffer
from severe anaphylactoid reactions [50–55,155].
The highest bradykinin generation reported was measured in vivo in sheep under
ACE inhibitor therapy and treated with poylacrylonitrile membranes (AN69). Lower
Fig. 1.18 What is the possible future strategy
for the treatment of kidney patients by means of
extracorporeal blood circuits? Recent
investigations have shown that a variety of
uremic toxins is found in the larger molecular
weight range. Conventional dialysis
membranes such as the F60S PSu-membrane
are determined by their molecular weight cutoff, that is, the sieving coefficient distribution.
Based on the experience with liver failure
therapy, recent observations have shown that
adsorber cartridges placed in a secondary
circuit may support the removal of largemolecular-weight toxins, given that a filter with
a higher cut-off (e.g., an Albuflow filter with a
highly permeable polysulfone membrane)
allows for the separation of plasma in the
primary circuit.
References
generation was observed when a PAN membrane with lower surface electronegativity
(SPAN-Membrana, not available any more on the market) was used [156]. In another
study, even symptom-free patients without ACE medication exhibited an approximately 10-fold higher bradykinin generation with PAN (AN69) plate dialyzers than
with Cuprophan, Hemophan, and high-flux Fresenius Polysulfone [157]. A very low
bradykinin generation by high-flux Fresenius Polysulfone and low-flux Hemophan
membranes was confirmed also for patients under ACE inhibitor therapy [158].
1.7
Conclusion
During the last 20 years of hemodialysis therapy, an intense research on membranerelated factors and properties has been performed. Understanding of membrane
properties and performance in the clinical setting has considerably improved such
that adverse events related to the membrane material have become rare. A general
trend can be seen in that cellulosic membranes have lost their dominating role and
synthetic membranes have taken the lead. Furthermore, membranes with high-flux
properties gain increasing importance due to new concepts of therapy, such as online
hemodiafiltration. Dialysis membranes of today have become bulk products and are
available without any restriction. Future therapies will possibly involve highly
permeable membranes combined in series with adsorber cartridges in order to
cope with newly identified uremic toxins of large molecular weights and under
hemodiafiltration treatments (Figure 1.18).
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Renaux, J., Thomas, M., Crost, T.,
Loughraieb, N., Vantard, G. (1999)
Activation of the kallikrein-kinin
system in hemodialysis: role of
membrane electronegativity, blood
dilution and pH. Kidney International,
55, 1097–1103.
Schaefer, R., Schaefer, L., Hoerl, W.
(1994) Anaphylactoid reactions
during hemodialysis. Clinical
Nephrology, 42 (Suppl. 1), S44–S47.
Krieter, D., Grude, M., Lemke, H.,
Fink, E., Bönner, G., Schölkens, B.,
Schulz, E., Müller, G. (1998)
hemodialysis in sheep are ACEinhibitor dose-dependent and
mediated by bradykinin. Kidney
Verresen, L., Fink, E., Lemke, H.,
Vanrenterghem, Y. (1994) Bradykinin
is a mediator of anaphylactoid
reactions during hemodialysis with
AN69 membranes. Kidney
Mannstadt, M., Touam, M., Fink, E.,
Urena, P., Hruby, M., Zingraff, J.,
Uhlenbusch-Körwer, I., Grassmann,
A., Lemke, H., Drueke, T. (1995) No
generation of bradykinin with a new
polyacrylonitrile membrane [SPAN]
in hemodialysis patients treated with
ACE-inhibitors. Nephrology Dialysis
49
2
Membranes for Artificial Lungs
Frank Wiese
2.1
Introduction
Membrane oxygenators are used inside heart–lung machines during open-heart
surgery or for lung support systems.
In the natural lung, oxygen and carbon dioxide are exchanged between the blood
and the body environment. The oxygen is needed in the body tissues to sustain life.
The carbon dioxide produced during the metabolism has to be removed.
Worldwide, oxygenators are used in surgical interventions during which the heart–
lung function is interrupted, for example, for cardiovascular bypass, heart valve
replacement, or when repairing congestive heart failure [30–32]. Furthermore,
extracorporeal membrane oxygenation (ECMO) is used in several configurations
for long-term lung support, for example, in cases of multiorgan failure (MOF) and in
different pulmonary diseases [7,33–35]. Implantable membrane devices for longterm lung support are under development [5]. Special newly developed membranes
were required for these long-term applications.
In general, at the very core of these devices are membranes that separate the patient’s
blood from gases and that allow oxygen to enter and carbon dioxide to leave the blood.
In 2005, around 1.2 million open-heart operations were carried out, corresponding to a
consumption of 3 million km of oxygenation membranes. Membranes for gas
exchange are also an important component in membrane bioreactors for biotechnology applications and in hybrid bioartificial organs (see Chapters 5 and 8).
2.2
History of Blood Oxygenation
Some milestones in the history of the development of blood membrane oxygenators
[6,30,37] are listed below:
1812: Le Gallois’ suggestion for the first extracorporeal
circulation
ISBN: 978-3-527-31480-5
50
2 Membranes for Artificial Lungs
1876: an isolated kidney was perfused with oxygenated blood
for the first time
1918: discovery of heparin
1944: Kolff had the idea of constructing an oxygenator
(beginning of membrane oxygenation)
1955: development of the first membrane oxygenator by Kolff
and Balzer
1956: first clinical application
1959: first heart surgery using a membrane oxygenator
1969: membrane oxygenators became commercially available.
The start of the historical development of extracorporeal circulation is strongly
related to César Julian Jean Le Gallois (1770–1814). He suggested that a part of the
body might be preserved by a mechanical heart replacement and some kind
of external perfusion devices. The important monograph ‘‘Experiences sur le
principe de la vie’’ was published in 1812 [1] and 1 year later in the United States
[2] (see Figure 2.1). Alexander Schmidt for the first time perfused an isolated dog
kidney with oxygenated blood in 1876 [3]. Many different types of disks, films, and
bubble oxygenators were developed for extracorporeal circulation experiments. All
these early oxygenators exposed blood directly to gas mixtures or oxygen to provide
oxygenation and remove CO2. The developments, started in the 1930s, resulted in
disposable single-use bubble oxygenators and opened the area of cardiac surgery in
the early 1950s. John Gibbon was a pioneer in this development [4,8,30].
In 1955, the DeWall–Lillehei bubble oxygenator was a major breakthrough [36,23].
The simplicity of the basic concept was linked to simple construction and disposability. The principle is: Venous blood is pumped into an oxygenation chamber where
oxygen bubbles are dispersed. The gas exchange occurs on the bubble surface. The
mixture of blood and gas that emerges from the bubble chamber has to be defoamed
in a defoaming compartment. The main disadvantage of bubble oxygenators is that
they are basically not biocompatible. The blood–gas interface represents a nonphysiological state resulting, for example, in hemolysis, coagulopathy, and so on.
Nevertheless, bubble oxygenators were used until 10 years ago in around 2–3 %
of open heart surgery worldwide.
2.2.1
Membrane Oxygenators
In 1944, Kolff and Berk found that venous blood was oxygenated while flowing
through a cellophane dialyzer and being in contact with oxygen containing
dialyzate. Oxygenation was carried out across the artificial membrane [24]. This
discovery stimulated the development of the use of gas permeable membranes in
order to separate the blood phase from the gas phase in an artificial lung. The first
membrane oxygenator was reported by Kolff in 1955 [45,46] (Figure 2.2). At this
time the available membrane materials were relatively impermeable to the respiratory gases.
2.2 History of Blood Oxygenation
Fig. 2.1 Front page of paper of Le Gallois, describing the
basic idea of extracorporeal circulation and oxygenation of
blood for the first time (1812).
The first membrane oxygenator built and used clinically was reported by Clowes
and Hopkins in 1956 [8]. It was constructed of ethylcellulose multilayer flatsheet
membranes and had a membrane area of 25 m2. These factors limited the clinical
applications until the 1970s and early 1980s, although the very first thin sheet
membranes of polyethylene and polytetraflourethylene were available. The development of polydimethylsiloxane with high permeability for oxygen and carbon
51
52
Fig. 2.2 Professor Kolf in 2004 – ‘‘The father of artificial
organs,’’ also invented artificial coil kidney and the coil
oxygenator 50 years ago [47].
dioxide brought a major advance in establishing the technical feasibility of membrane oxygenators in the 1960s and 1970s. The required membrane area could be
optimized at less than 6 m2.
The real breakthrough came with the development of hydrophobic microporous
flat sheet membranes with pore sizes <0.1 mm in the 1980s [26]. Later in the 1980s
microporous hollow fiber membranes [27] became commercially available. With the
introduction of extraluminal flow oxygenators [28,29] membranes, areas of less than
2 m2 were required to treat an adult patient.
2.3 Principle of Gas Transfer
2.3
Principle of Gas Transfer
The transport of gas into the blood is a process of convection and diffusion via the
membrane and the blood combined with O2 and CO2 reaction with the erythrocyte
hemoglobin. The hemoglobin in the erythrocytes has a specific chemical gas-binding
capacity for O2 and CO2 and assures the adequate transport of both these gases inside
the body. The driving force for the gas exchange is the partial pressure difference of
these gases between the blood side and the gas side (O2) of the membrane. According
to the pressure differences, gas diffuses through the membrane wall from the side of
high partial pressure to the side of low partial pressure (Figure 2.3).
The maximum concentration of chemically bound O2, called the oxygen capacity, is
calculated from the hemoglobin content of the blood and the binding capacity of
hemoglobin for oxygen. For a mean hemoglobin content of Hb ¼ 0.12 kg/L and a
binding capacity of Ha ¼ 1.34 L/kg, the oxygen capacity will be 0.1608 L O2/L blood.
This value will be reached at an O2 partial pressure of approximately
20 kPa ¼ 150 mmHg, which is the ambient partial pressure of oxygen. An additional increase in the O2 content is possible but nonphysiological. The oxyhemoglobin dissociation curve is usually calculated according to the mathematical model of
Hill [9] (Figure 2.4).
Nowadays the active area of membrane oxygenators is between 0.5 and 2.5 m2.
This is around 10 % of that of the natural lung of young children or adults. In order to
reach a sufficient gas exchange, the contact time has to be increased and/or the partial
pressure differences have to be increased or other suitable conditions have to be
adjusted. According to Fick’s law (1) the gas exchange efficiency can be increased
[30,37].
VO 2 ¼
P1 P2
KF:
L
ð1Þ
In Equation (1), VO2 is the amount of oxygen exchanged in a certain time, P1 P2 is
the partial pressure difference, K is the diffusion factor (adsorption coefficient,
turbulence), F is the surface, and L is the diffusion layer thickness.
Fig. 2.3 Partial pressure difference of oxygen and carbon
dioxide across an oxygenation membrane.
53
54
Fig. 2.4 Blood saturation according to Hill.
The transfer of oxygen into the blood is hindered by three main resistances:
Rtotal ¼ RB þ RM þ RG :
ð2Þ
In Equation (2), the total resistance to gas transport Rtotal is the sum of the
resistance on the blood side RB, the resistance across the membrane RM, and the
resistance on the gas side RG (Figure 2.5).
For a usual oxygenator type, containing microporous capillary membranes, as
described in Section 2.4, the blood boundary layer lies outside the capillaries. The
diffusion resistances that opposes the gas exchange consists of the resistance of
the gas boundary layer, the membrane, and the blood boundary layer. Here the
resistance of the gas boundary layer is infinitely small and approaches zero, and
accordingly, can be neglected. The resistance of the membrane is about
1.25 104 cm2 s cmHg/cm3. The resistance in the blood boundary layer is
approximately 100 times larger and should be optimized in such a way that this
layer becomes very small (see Section 2.6). The mass transfer in relation to the
geometry flow pattern and the hydrodynamics of an artificial lung has been
investigated in hundreds of papers [6,10–13].
Fig. 2.5 Mass transfer model for an extraluminal flow
oxygenator.
2.4 Membranes and Membrane Properties
2.4
Membranes and Membrane Properties
Nowadays membranes for oxygenators used in open-heart surgery are mostly
microporous hydrophobic capillary membranes with pore sizes <0.1 mm and outer
diameters between 300 and 500 mm.
For special applications, for example, in lung support systems, dense skinned
membranes are used because of better long-term blood plasma resistance.
The main membrane polymers used are polyolefins such as polypropylene (PP),
polyethylene (PE), and poly-4-methylpentene (PMP). Further, the polymers used on a
small scale are polyvinylidenfluoride (PVDF) and silicon rubber.
In flat sheet membrane oxygenators, the blood is separated from the gas by spiral
wound membranes or stapled membranes. Baffles placed in the area between the
membranes, in order to reduce the boundary layer and increase efficiency, require a
lot of effort. Therefore, only one type of microporous flat sheet membrane oxygenator
is available but without any market relevance.
Around 80 % of oxygenation membranes used worldwide are produced by
Membrana GmbH in Wuppertal. Other producers of oxygenation membranes are
Dainippon Inc. and Terumo.
2.4.1
Microporous Membranes
Gas exchange between the blood side and the gas side in microporous membranes
is performed by diffusion and convection of the respiratory gases via the open
pores. Thus, the resistance of the membrane is very low for these gas exchanges, as
explained in Section 2.3. Most microporous oxygenator membranes are made from
polypropylene by temperature-induced phase separation (TIPS) process (see
Section 2.5).
The sponge-like structure of this membrane is homogeneous and isotropic with
open surfaces on both wall sides (Figure 2.6). These membranes achieve reliable,
adequate performance (Figure 2.10), and the mechanical stability in terms of tensile
strength and elongation at break is combined with high reliability and handling safety
during the manufacturing process (see Section 2.6). Today, it is the most widely used
oxygenation membrane.
The second type of microporous membranes is produced by the ‘‘melt spin stretch
process’’ (see Section 2.5), resulting in fibrillar polymer structures (see Figure 2.7).
Polyolefins are again suitable for these processes because only polymers with a certain
crystallinity are suitable for creating such membrane structures. Performance wise the
membranes can reach a comparable level to that of TIPS membranes. Long-term
plasma breakthrough characteristics are a little more critical because of the structure.
Due to the hydrophobicity of the membrane and the pore size (<0.1 mm), gas
molecules can pass through the membrane wall, whereas liquids and corpuscular
components are retained. The polar part of the surface energy [15] has to be very small
so that no wetting of the pores with aqueous solutions, for example, blood, is possible,
55
56
Fig. 2.6 REM photograph of a
microporous PP membrane (OXYPHAN,
Membrana, Wuppertal).
and a water droplet on this surface has to create a high-contact angle 908 (see
Figure 2.8).
For the duration of an open-heart operation and for several days, these open
membranes work very reliably and are safe. However, during the treatment, plasma
proteins adsorb first at the membrane surface and then this protein layer also grows
Fig. 2.7 REM photograph of a microporous PP membrane
produced by a melt spin stretch process.
Fig. 2.8 Contact angle of hydrophilic and hydrophobic surfaces.
through the membrane pores. This results in an increase in the surface energy of the
pore ‘‘skin.’’ Moreover, due to this surfactant impact the pores become wet and
plasma breakthrough from the blood side to the gas side occurs [41,42].
2.4.2
Dense Membranes/‘‘Diffusion Membranes’’
Dense membranes, also called ‘‘diffusion’’ membranes, have no open pores in the
membrane wall or the outer skin of the wall [38,44,53]. In long-term ‘‘extracorporeal
membrane oxygenation’’ applications such membranes are used to prevent plasma
breakthrough.
The respiratory gases have to have a good solubility in the membrane polymer.
Driven by the concentration difference of the gases on both sides of the membrane,
the gases are transported by diffusion through the polymer matrix to the lower
concentration. In the past, silicon rubber or silicon-skinned microporous membranes were used because of the high specific permeability of this polymer. The
disadvantage of these membranes was their relatively high wall thicknesses due to
mechanical requirements. According to Fick’s law (Section 2.3), the diffusive gas
transport is reciprocally proportional to the thickness of the polymer layer. Thus, high
membrane areas were required to achieve sufficient gas exchange. To overcome this
problem, membranes made of poly-4-methylpentene were developed. The membrane with the best performance is produced by a thermally induced phase separation
process (ACCUREL) (Figure 2.9). The membrane is integrally asymmetric, as it has a
sponge-like microporous wall with a thin dense outer skin (Figure 2.9).
Due to the high specific permeabilities of PMP to oxygen and carbon dioxide
(Table 2.1) and the dense outer skin, which is kept very thin (0.1 mm), OXYPLUS
has gas transfer capabilities comparable to common microporous membranes.
Figure 2.10 shows a comparison of the gas transfer of the same oxygenator types
equipped with either PMP or PP membranes. The plasma breakthrough times for
OXYPLUS are very much longer than those of common microporous membranes: in
the clinical experience of different working groups the membrane has worked
properly in ECMO for up to 42 days without exchanging the device [18,19].
Additionally, the dense outer skin largely prevents entry of gas into the blood and
avoids direct blood–gas contact. As a trade off, the transfer of volatile anesthetics is
impaired [20] and the anesthetic protocol may have to be amended.
The permeability of PMP is higher for oxygen than for nitrogen. This membrane is
thus also suitable for gas separation or enrichment. Table 2.2 shows the fluxes of
57
58
Fig. 2.9 Gas transfer rates of two commercially
available oxygenators, equipped either with a
microporous PP membrane or with a PMP
membrane with dense outer skin [16,17].
Gas permeability coefficients of polypropylene and poly-4-methylpentene for oxygen and
carbon dioxide [14].
Tab. 2.1
Gas
PP
PMP
Oxygen
Carbon dioxide
2.2
9.2
32.3
92.6
Permeability (1 Barrer ¼ 1010 cm3 cm/(s cm2 cmHg).
2.5 Membrane Production
Fig. 2.10 REM photograph of a microporous PMP
membrane with dense outer skin (OXYPLUS, Membrana,
Wuppertal).
Gas permeabilities of OXYPLUS for different gases and the resulting separation
coefficients.
Tab. 2.2
Gas
Nitrogen
Oxygen
Carbon dioxide
Flux (L/m2 min bar)
1.68
5.90
16.90
Separation coefficient
related to nitrogen
–—
3.53
10.13
different gases through OXYPLUS and the resulting gas separation coefficients. With
this membrane, the enrichment of oxygen in air to a concentration of, for example,
25 %, is possible and effective in a single-step process.
2.5
Membrane Production
Oxygenation membranes are mainly produced by two different processes:
Temperature-induced phase separation process
Melt spin stretch process.
Seventy percent of oxygenation membranes worldwide are produced by a special
TIPS process, the so-called ‘‘ACCUREL’’ process [48–50] (Figure 2.11). A suitable
polymer and a solvent or solvent mixture such as natural seed oils (e.g., soybean and
castor) are heated together in an extruder to produce a homogeneously mixed
polymer solution. This solution is pumped through a spinneret together with a
gas as bore medium to create the right geometry of the hollow fiber. The preformed
hollow fiber geometry is then cooled in an air gap and in a cold spinning oil, initiating
phase separation of the solvent and the porous polymer matrix with pores still filled
59
60
Fig. 2.11 Scheme of the ACCUREL process for
manufacturing of membranes by thermally induced phase
separation.
with oil. After a washing step, for example, with hot alcohol, to remove the solvent, the
membrane is wound onto bobbins.
The following makeup steps will be explained in Section 2.6. The membrane
structure is adjustable by varying the composition of the polymer solution, the
cooling speed, and the process temperatures during spinning and cooling.
A homogenous sponge-like structure is formed, as shown schematically in
Figure 2.12.
In the melt spin stretch process, the membrane pores are produced in a completely
different way. First a suitable crystalline or semicrystalline polymer is molten in an
Fig. 2.12 Schematic picture of ACCUREL membrane wall.
2.6 Operational Modes and Membrane Makeup in Oxygenators
Fig. 2.13 Scheme of the stretching process to produce
microporous membranes.
Fig. 2.14 Scheme of microporous membrane produced by
melt spin stretch process.
extruder. The molten polymer is pressed through a spinneret and on cooling a dense
compact capillary a flat sheet is created. By annealing, stretching, and cooling steps with
a special temperature and tension profile in several repeated steps (see Figure 2.13),
microporous membranes with completely different structures are created.
This process has advantages in that no solvent and extraction are required but also
disadvantages due to the lower porosity, the limited range of pore size, and the
elongated pores (lower selectivity) of the product. The schematic structure of the
membrane wall is shown in Figure 2.14.
2.6
Operational Modes and Membrane Makeup in Oxygenators
The oxygenator construction and the arrangement of the membranes inside the
oxygenator are as important as the performance of the oxygenation membrane itself.
61
62
2.6.1
Microporous Capillary Membranes, Blood Inside
In the 1980s, the blood was pumped inside the capillary membrane (luminal flow)
whereas the membrane was flushed with gas from the outside. The disadvantages of
these oxygenators, which are no longer in use today, are
a large membrane quantity is required, as the gas exchange
surface is the smaller inner capillary surface
poor biocompatibility
large priming volume
high production costs
high pressure drop.
2.6.2
Microporous Capillary Membranes, Blood Outside
Because of the disadvantages of intraluminal blood flow described above, oxygenators with blood flowing outside the capillaries (extraluminal flow) were developed in
the mid-1980s [28,29,37,43]. The structure of these oxygenators with wound capillary
membranes or with knitted mats provides good mixing of the blood and thus an
improvement in the oxygen transfer rates, especially in the blood boundary layer (see
Section 2.3). The membrane surface used can be minimized. More than 95 % of the
membrane oxygenators used in hospitals today are oxygenators with blood flow
outside the capillaries.
The most effective construction uses cross-laid knitted mats (Figure 2.15). With
this capillary arrangement, the following features are achieved:
‘‘enhanced mass transfer efficiency’’ with ‘‘local laminar
mixing’’ (static mixer effect)
repeated renewal of the concentration in the boundary layers;
therefore, highest efficiency (see Section 2.3)
uniform distribution of capillaries
Fig. 2.15 Cross-laid knitted membrane mat.
2.6 Operational Modes and Membrane Makeup in Oxygenators
capillaries fixed by warp threads (no movement)
suitable for various oxygenator designs (axial, tangential,
crossflow, etc.)
blood film thickness (pressure drop) can be adjusted easily
easy handling during manufacture
high production yield
closed capillary ends for safe potting.
The process steps for the production of cross-laid knitted mats are shown in
Figure 2.16. Several capillaries are continuously laid down in parallel to a knitting
machine and connected with a warp thread at defined distances. During this step
multiple mats are cut, the membrane ends are closed, and the mat spools are wound
up in parallel. In the next step, two mats are pulled at each defined angle and wound
up together.
Fig. 2.16 Scheme of production of cross-laid knitted hollow fiber mats.
63
64
Fig. 2.17 Mass transfer coefficient versus membrane angle
in a cross-wound mat device.
For each oxygenator setup, a special distance between the membranes, and a
special warp thread distance and angle between the cross wound mats are possible
and can be optimized (Figure 2.18). In this way one can always create an optimal
blood stream. In the case of membrane contactors in technical applications, such
constructions are also an advantage.
The effects of design and operating variables on module performance have been
investigated with respect to oxygen transfer into water. It was found that the
membrane angle with respect to the main direction of the liquid flow had a significant
influence on oxygenator performance [37]. The mass transfer coefficient increased
proportionally to the membrane angle up to about 108 and started leveling off at about
20–258 (Figure 2.17).
Fig. 2.18 Scheme of cross-laid hollow fibers.
2.7 Extracorporeal Circulation
2.7
Extracorporeal Circulation
2.7.1
Cardiodiapulmonary Bypass (CPB)
In Figure 2.19, an extracorporeal circuit called a cardiopulmonary bypass, as used
during open-heart surgery, is schematically demonstrated. Venous blood is sucked
out by several cannulas from the patient. The blood is pumped, directly or via a
venous reservoir, through the oxygenator. If applied, it is mixed with a part of sucked
cardiotomic wound blood.
The blood flow through the oxygenator is between 3 and 6 L/min. Here the CO2 is
removed from the blood and the oxygen is transported into the blood. All the blood
from the wound is also brought back to the circuit (Figure 2.19).
2.7.2
Lung Support Systems
If the native lung is diseased, for example, by asthma, emphysema, or acute
respiratory distress syndrome (ARDS), or is injured and thus no longer able to
supply the body with sufficient amounts of oxygen, two treatment options exist,
depending on the severity and comorbidities: lung support with oxygen-enriched air
or blood oxygenation (ECMO). For the first treatment option, there is a need for a
mobile oxygen-enriching device, and for the second option, there is a need for an
oxygenation membrane suitable for long-term application (several days to weeks).
Some membranes described in Section 2.4.2 are suitable for both tasks (Figure 2.21).
In intensive care, artificial lungs with long-term oxygenation membranes, as
described in Section 2.4.2 become a treatment option for acute lung failure (e.g.,
Fig. 2.19 Scheme of an oxygenator in a heart lung machine setup.
65
66
Fig. 2.20 Example of an oxygenator setup with heat
exchanger, venous reservoir, and oxygenator and photo of
the oxygenator with cross-laid knitted mats inside.
Fig. 2.21 Nova lung oxygenator for an ECMO application.
from virus pneumonia, after blunt trauma [21], following pancreatitis, or in conjunction with sepsis or multiorgan failure) or for lung failure, for example, posttransplant, postoperative, or pediatric cases. An excellent example for a pumpless
ECMO device is the Novalung ‘‘membrane ventilator.’’ Two cannulas are connected
to large vessels of the patient and by the force of the patients’ heart the blood is
Fig. 2.22 Implantable lung support setup.
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The membrane can also be suitable for use in long-term paracorporeal or
implantable artificial lungs, for example, inside the vena cava [22,25]. In
Figure 2.22, an intravascular membrane oxygenator (IVOX) shows the first vena
caval device with membranes [5,40].
Comparable devices and oxygenation membrane setups, as described in Sections
2.6 and 2.7, are also used in biotechnology (see Chapter 5) and bioartificial organs
(see Chapter 8) [51,52].
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Alpha, D., King, E., Bicknell, D. A.
(1986) Proceedings of the American
Academy on Clinical Perfusion, 7,
32–34.
Lauterbach, G. (2002) Handbuch der
Kardiotechnik, Urban & Fischer,
München.
Tschaud, R. J. (1999) Extrakorporale
Zirkulation in der Theorie und Praxis,
Papst Science Publischers, Lengerich.
Taylor, K. M. (1986) Cardiopulmonary
Bypass, Lippincott Williams & Wilkins,
Philadelphia, PA.
Philipp, A., Foltan, M., Gietl, M.,
Reng, M., Liebold, A., Kobuch, R.,
Keyl, C., Bein, T., Müller, T., Schmid,
F.-X., Birnbaum, D. E. (2003)
Kardiotechnik, 1, 7–13.
Schmid, F.-X., Philipp, A., Link, J.,
Zimmermann, M., Birnbaum, D. E.
(2002) Annals of Thoracic Surgery, 73,
1618–1620.
Zwischenberger, J. B., Anderson, C.
M., Cook, K. E., Lick, S. D., Mockros,
L. F., Bartlett, R. H. (2001) ASAIO
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Wiese, F. (2000) Maku (Membrane),
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106.
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Schuster, O. (2004) ASAIO Journal, 50,
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J., Boone, Y., Seigneul, I., Jachertz, M.,
von Segesser, L. K. (2004) ESAO,
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extracorporal lung assist device.
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3
Membranes for Blood Fractionation/Apheresis
Frank Wiese
3.1
Introduction
The use of membranes in extracorporeal systems connected to the human body
started in the 1940s with Kolff’s first use of an artificial kidney. Since that time, the
application of membranes has been established in different areas (Figure 3.1).
Nowadays, the expression, apheresis or plasmapheresis, is used for the removal of
blood plasma from cells as well as for the removal of protein components or fractions
Fig. 3.1 Membranes in extracorporeal systems.
Fig. 3.2 Therapeutic hemapheresis classified according to technologies.
ISBN: 978-3-527-31480-5
70
3 Membranes for Blood Fractionation/Apheresis
from this plasma. The origin is the Greek word aphairesis – ‘‘to separate, to remove,
and to take away.’’ In comparison to 120 million artificial kidney treatments and
1.2 million open-heart operations using artificial lungs, there are only around
600 000 membrane apheresis treatments carried out per year. Nevertheless, these
treatments are an important therapeutic tool for special medical indications for
autoimmune diseases or diseases related to blood rheology. Membrane procedures
have significance for plasma separation and plasma fractionation processes – also
called cascade filtration or double filtration – as well as for specific adsorption
processes because of the requirement for cell-free plasma for these procedures. For
the functionality and the life-time of these modern adsorption processes, cell-free
plasma is a strict requirement (Figure 3.2). Blood plasma donation for medical and
pharmaceutical applications is dominated by centrifugal procedures. Procedures
such as cytapheresis or cell separation are not discussed in this chapter because of the
negligible use of membranes in these procedures.
3.2
History of Plasmapheresis
In the ancient world, the role of blood in health and illness was already known. The
name of Hippocrates (460–377 B.C.), the ‘‘Father of Medicine,’’ is linked to the
Fig. 3.3 Hippocrates Hiraclidae (460–377 B.C.) –
The ‘‘Father of Medicine.’’
Fig. 3.4 Bloodletting depicted by Peytel.
Fig. 3.5 ‘‘Plasmapheresis’’ described by Abel, Rowntree, and Turner [3].
72
Fig. 3.6 John Jacob Abel 1857–1938.
Fig. 3.7 Monsanto hollow fiber device as
experimental plasma separator 1972 [4].
3.3 Principles of Plasmapheresis
Fig. 3.8 Principle of double filtration first described by Agishi [5].
introduction of bloodletting as a preventative method as well as a remedy [1,2].
Although bloodletting should not be confused with plasmapheresis, it is its therapeutic precursor (Figures 3.3 and 3.4). The term plasmapheresis was first used by
Abel, Rowntree, and Turner in 1914. In their paper ‘‘Plasma Removal with Return of
Corpuscles (Plasmapheresis)’’ [3], they investigated the effect of plasma removal
from dog blood and reinjected the cells together with replacement solution
(Figures 3.5 and 3.6). This was the early beginning of therapeutic plasmapheresis.
The first hollow fiber membrane device for plasma separation was introduced by
Monsanto in 1972 [4]. This was the start of therapeutic plasma separation by
membranes (Figure 3.7). The first double filtration in place of two membrane filters
with different pore sizes was described by Agishi [5]. High molecular weight proteins
such as low-density lipoproteins (LDL) and triglycerides were removed and lower
molecular weight proteins such as albumin and high-density lipoprotein (HDL) were
returned to the body (Figure 3.8).
3.3
Principles of Plasmapheresis
The schematic setup of a plasma separation treatment is shown in Figure 3.9. The
blood is removed from the body and filtered in a plasma filter device. Then the filtered
plasma is removed and cells are returned to the body as concentrate, together with a
substitution fluid such as an isotonic electrolyte solution, an albumin-isotonic
electrolyte solution, or donor plasma [6–9]. Blood is a special streaming medium,
and this special characteristic has to be considered by selecting the right process
conditions. In capillary modules, the transmembrane pressure (TMP) is usually not
higher than 150 mmHg, and the adjusted filtrate flow is between 20 and 35% of the
blood flow rate of 50–250 mL/min.
In most classical plasma exchange procedures (TPE – therapeutic plasma
exchange), plasma is removed and replaced by donor plasma (FFP – fresh frozen
plasma). It is used not only for the treatment of several immunological diseases but
73
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Fig. 3.9 Membrane plasma separation set up.
also for the removal of metabolism products and exogenic and endogenic toxins [6,8].
In several countries, this unspecific exchange is carried out even now (mostly by the
use of centrifuges), but the frequency is decreasing because of the availability of more
specific adsorption methods and the many clinical disadvantages of this method.
To save mostly patients’ own specific proteins such as albumin and IgG and to
remove only the poisoning components, for example, LDL from, for example,
hypercholesteremic patients, membrane plasma separation is used in combination
with secondary membrane filtration called ‘‘plasma fractionation,’’ ‘‘membrane
double filtration,’’ ‘‘rheopheresis,’’ or ‘‘cascade filtration’’ (Figure 3.10) [29]. The blood
is divided into a cell-concentrated part and a cell-free plasma part in the first filtration
step. In the second step, the blood plasma is divided into high and low molecular
Fig. 3.10 Membrane double filtration set up (cascade
filtration, plasma fractionation).
3.3 Principles of Plasmapheresis
Tab. 3.1
Blood as a streaming medium – specific features during application.
Primary filtration
Secondary filtration
Blood viscosity depends on
Type of cells
Concentration of cells (hematocrit)
‘‘Cross-flow operation’’: plasma-rich layer
Hemolysis
Plasma viscosity depends on
Type of macromolecules/proteins
Concentration of macromolecules/proteins
‘‘Dead-end operation’’: secondary membrane layer
Blocking
weight fractions of plasma. The low molecular weight fraction and nearly the complete
plasma fluid volume are then recombined with the cell concentrate and returned to the
patient. The sieving characteristics of these membranes are described in Section 3.4.
The primary filtration step (plasma separation) is done as a ‘‘cross-flow operation,’’
and the secondary filtration as a ‘‘dead-end operation.’’ This has to be considered by
selecting the right process conditions to prevent blood damage and to keep the
process running. The right shearing profile in the primary filter prevents hemolysis,
and the right pressure in the secondary filer prevents blocking (see Table 3.1 and
Section 3.6).
The setup of plasma fractionation in combination with adsorbers is shown in
Figure 3.11. Again the plasma is separated with a plasma filter. In the second step, the
separated plasma is pumped through protein-specific adsorbers to remove special
proteins such as immunoglobulins or toxins according to the therapeutic objectives.
Finally, the treated plasma together with the blood cells is returned to the patient.
Different kinds of single-use adsorbers or regenerable double adsorbers for different
medical indications are on the market. In Figure 3.11, an example of a double-column
setup is shown. While one column is recirculated with plasma and loaded with
Fig. 3.11 Membrane plasma separation combined with
plasma fractionation by adsorbers.
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proteins to be removed, the second column is regenerated. The columns are switched
from time to time and are applicable for multiple patient treatments.
3.4
Membranes and Membrane Properties
The general profile of the requirements of plasma treatment membranes is summarized as follows [6,7,10,28]:
permeable for all plasma proteins/special sieving
characteristics
high surface porosity and volume porosity to reach high fluxes
hydrophilic, spontaneous, and wettable
low fouling behavior
low protein adsorption
smooth blood contact surfaces
low hemolytic potential
consistency in sieving behavior and filtration performance
during treatment
good blood compatibility [17–19,22,24]
suitable mechanical stability
stability against chemicals and all common sterilization
procedures (ETO, g-radiation, steam).
The plasma fractionation membranes have to have special sieving characteristics
for the removal of different protein fractions, depending on the medical indications.
Different producers of filters use different polymers such as cellulose acetate (CA),
polypropylene (PP), polyethylene (PE), polysulfone (PSU), polyethersulfone (PES),
and polymethylmethacrylate (PMMA) to fulfill these requirements as far as possible.
Hydrophobic polymers such as PP are hydrophilized during module manufacture or
PE is coated with polyethylvinylalcohol (EVAL), and PES is blended with polyvinylpyrolidone (PVP) [31–34].
3.4.1
Plasma Separation Membranes
Plasma separation capillary membrane filters from different suppliers are listed in
Table 3.2 [14].
All the used membranes have nominal pore sizes around 0.2–0.3 mm (microfiltration membranes). The module suppliers often do not explain whether the pore
size in their data sheets is theoretically calculated from the bubble point or measured
by a sieving curve, thus creating some confusion.
The size exclusion pore size of 0.2 mm guarantees that all plasma proteins can
pass through the membrane whereas all cells including the small thrombocytes
with 2 mm diameter are rejected.
Capillary membrane plasma separation filters from different producers.
Tab. 3.2
Supplier
Filter
Asahi
Plasmaflo OP-02
B. Braun
Bellco
Gambro
Fresenius
Kaneka
Kuraray
Toray
Edwards
Dideco
Plasmaflo OP-05
Plasmaflo OP-08
Hemoselect
MPS 02
MPS 05
MPS 07
PF 1000
PF 2000
Plasmaflux P1S
Plasmaflux P2S
Sulflux FS-03
Sulflux FS-05
Sulflux FS-07
Plasmacure
Plasmax PS-02
Plasmax PS-05
Microplas MPS 05
Hemaplex BT 900
Membrane
material
Membrane
area (m2)
ID/wall
(mm)
Sterilization
mode
Polyethylene/
EVAL coating
0.20
330/50
ETO
330/150
300/100
Steam
ETO
330/150
ETO
330/150
Steam
340/50
g-radiation
320/65
330/90
g-radiation
g-radiation
300/100
330/150
ETO, g-radiation
ETO
Polypropylene
Polyethersulfone
Polypropylene
Polypropylene
Polysulfone
Polysulfone
PMMA
Polyethersulfone
Polypropylene
0.50
0.80
0.20
0.28
0.45
0.68
0.16
0.35
0.25
0.50
0.30
0.50
0.70
0.60
0.15
0.50
0.45
0.50
The sieving coefficient (SC) for a special protein in a filtration experiment is
calculated according to Equation (1): CF is the concentration of the protein in the
filtrate, CBi the filter input concentration, and CBo the filter output concentration of
the protein in the blood or blood plasma.
SC ¼
2CF
:
CBi þ CBo
ð1Þ
The sieving characteristics of a typical plasma separation membrane are shown in
Figure 3.12(a). Nearly 100 % of all plasma proteins pass through the membrane
including the lipoproteins with molecular weights >2 000 000 Da. Fibrinogen is
slightly reduced. This can be related to the special shape (see Figure 3.13) [11] of this
molecule and adsorption effects.
The same membrane was also investigated with latex beads of defined different
sizes. The rampant sieving curve shows that it has a nominal pore size of 0.2 mm and
a narrow pore size distribution (Figure 3.12(b)).
These capillary membranes also meet the requirements of the HIMA test, which
means that the membrane can also be used for sterile filtration, for example,
intravenous filtration, because it is fully bacteria retentive against pseudomonas
diminuta.
Depending on the membrane polymer used, the composition of the spinning
solution and the production process, and the structure and wall thickness of the
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78
Fig. 3.12 Sieving coefficients of plasma separation
membrane tested with blood and latex beads (MicroPES
TF10, Membrana GmbH).
Fig. 3.13 Size comparison of noncellular blood components [11].
Fig. 3.14 SEM picture of cross sections (a, b) and blood
contacting inner surfaces (c, d) of two different plasma
separation membranes [25] (a, c) Plasmaphan (Membrana
GmbH) and (b, d) Plasmaflo (Asahi)).
membranes are different. The SEM pictures in Figure 3.14 illustrate this with
examples of two different plasma separation membranes produced from polyolefins
PP and PE. The PP membrane has a significantly higher wall thickness than the PE
membrane and different pore structure of the inside skin, whereas in vivo performance is comparable.
3.4.2
Plasma Fractionation Membranes
Plasma fractionation capillary membrane filters from different suppliers are listed in
Table 3.3.
The membranes used in the filters have different sieving characteristics, depending on the medical indications. For example, the different Evaflux membrane filter
types are used for the removal of different protein fractions. Evaflux 2A is used for the
removal of immunoglobulins IgG and IgA, Evaflux 3A for the removal of cytokines,
Evaflux 4A for the removal of IgM and LDL, and Evaflux 5A for the removal of LDL-C
(cholesterin) [26].
The sieving coefficient for a special protein in a filtration experiment is calculated
according to Equation (1).
In Figure 3.15, the sieving characteristics of a plasma fractionation membrane are
shown to exemplify the behavior of such a membrane.
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Tab. 3.3
Capillary membrane plasma fractionation filters from different producers.
Supplier
Filter
Membrane
material
Membrane
area (m2)
ID/wall
(mm)
Sterilization
mode
Asahi
Rheofilter
Cascadeflo EC-20
Cascadeflo EC-30
Cascadeflo EC-40
Cascadeflo EC-50
Evaflux 2A
Evaflux 3A
Evaflux 4A
Evaflux 5A
Albuafe FP 2
Cellulose diacetate
EVAL
EVAL
EVAL
EVAL
EVAL
EVAL
EVAL
EVAL
Polyethersulfone
2.0
2.0
2.0
2.0
2.0
2.0
2.0
2.0
2.0
1.7
220/80
175/40
175/40
175/40
175/40
175/40
175/40
175/40
175/40
200/35
g-radiation
Kuraray
Dideco
g-radiation
g-radiation, steam
The sieving coefficient for albumin is around 0.88. This is important because the
albumin has to pass through the membrane so that as much as possible of this
protein is given back to the patient. Most molecules with molecular weights up to
around 400 kDa, for instance HDL, pass through the membrane with the exception
of fibrinogen. Fibrinogen is a stretched long molecule and for this geometric reason
is partly rejected (see Figure 3.13). Larger molecules such as alpha-2-macroglobulin
or IGM are mostly rejected – the sieving coefficient of IGM is 0.2. The main goal
of the treatment is to remove these lipoprotein molecules. In conclusion, this
membrane can be used for hypercholesteremic as well as for rheological indications
in the clinical applications (see Section 3.7). The SEM picture of the cross section of
this membrane (see Figure 3.16) shows an isotropic and homogeneous pore
structure and a smooth inner surface. The inner diameter is 200 mm, and the wall
thickness is 35 mm.
Fig. 3.15 Sieving coefficients of a plasma fractionation
membrane for different blood plasma proteins –
FractioPES (Membrana GmbH).
3.5 Membrane Production
Fig. 3.16 SEM picture of the membrane
structure of the plasma fractionation
membrane FractioPES (Membrana GmbH).
3.5
Membrane Production
Plasma separation membranes are mainly produced by three different processes (see
also Section 2.5) (Figure 3.17–3.19):
Solvent-induced phase separation (SIPS) process
Temperature-induced phase separation (TIPS) process
Melt spin stretch (MESS) process.
The schematic course of an SIPS process, also called the wet spinning process, is
shown in Figure 3.17. From a suitable polymer and a suitable solvent or solvent
mixture, a solution is prepared. Here the ‘‘genome’’ of the membrane is created
Fig. 3.17 Schematic depiction of the solvent/nonsolventinduced phase separation process for manufacturing of
microporous membranes.
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because the membrane structure is predetermined (prefixed) in the solution. The
solution preparation, including proper filtration, is a complex and important step.
After this step, the solution is pumped into a spinneret for capillary membranes or
into a casting box in the case of flat membrane production. To give the capillary
membranes the correct geometry and to steer the inner skin pore structure, the bore
liquid can be an inert or a composition of solvent and nonsolvent. The next important
step is the precipitation, the ‘‘birth’’ of the membrane. The solvent is replaced by the
nonsolvent of the coagulation bath and creates the porous polymer matrix. This
diffusion process is influenced by the polymer solution composition as well as by
process conditions such as the temperature of the precipitation bath, the spinning
speed, the climatic conditions between the spinneret and the precipitation bath, and
so on. The coagulation is followed by the washing, drying, and bundle makeup steps
until the bundle is ready for insertion into a module housing. In special cases, a
postfinal extraction is carried out. Most membranes worldwide are produced by wet
spinning processes.
Plasmapheresis membranes from polyolefins (PP, PE) are mainly produced by two
different processes:
TIPS
MESS.
A special TIPS process, the so-called ‘‘ACCUREL’’ process, is described in
Figure 3.18 [15,16,38,44].
A suitable polymer and a solvent or solvent mixture such as natural seed oils
(e.g., soybean and castor) are heated together in an extruder to produce a homogeneously mixed polymer solution. This solution is pumped through a spinneret
Fig. 3.18 Schematic depiction of the ACCUREL process for
the manufacturing of membranes by thermally induced
phase separation.
3.6 Operational Modes in Plasmapheresis Procedures
Fig. 3.19 Schematic depiction of the pore building part in a
melt spin stretch process for producing of microporous
membranes.
together with a gas as bore medium to create the right geometry of the hollow fiber.
The preformed, hollow fiber geometry is then cooled in an air gap and in cold
spinning oil, initiating a phase separation of the solvent and the porous polymer
matrix, the pores of which are still filled with oil. After a washing step, for example,
with hot alcohol, to remove the solvent, the membrane is wound onto bobbins or into
bundles. The membrane structure can be adjusted by varying the composition of the
polymer solution, the cooling speed, and the different process temperatures during
spinning and cooling.
In the melt spin stretch process, the membrane pores are produced in a completely
different way. First, a suitable crystalline or semicrystalline polymer is melted in an
extruder. The molten polymer is pressed through a spinneret, and on cooling a dense
compact capillary or flat sheet is created. By annealing, stretching, and cooling steps
with a special temperature and tension profile in several repeated steps (see
Figure 3.19), microporous membranes with completely different structures are
created (see Figure 3.14(d) and Section 2.5). This process has advantages because
no solvent and extraction is required but also disadvantages because of the lower
porosity, limited range of pore size, and elongated pores (lower selectivity).
3.6
Operational Modes in Plasmapheresis Procedures
As mentioned in Section 3.3, suitable process conditions have to be adjusted in the
clinical application of plasma separation to reach a successful treatment without
negative side effects. The main design criteria are the wall shearing rate, the created
pressure drop along the filter housing, and the resulting plasma filtration rate. The
wall shearing rate is calculated according to Equation (2):
wall shearing rate : gw ¼
4QB
;
Npr 3
ð2Þ
where N is the number of capillary membranes with inner radius r and QB is the blood
flow distributed to the different capillary membranes. The blood flow changes by
reduction of the plasma content over the length of the membrane or the filter,
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84
respectively. This has to be considered when calculating the wall shearing rate.
Furthermore, the transmembrane pressure is further an important parameter for
steering the process in the right direction. The transmembrane pressure is calculated
according to Equation (3):
TMP ¼
PBi þ PBo
PF :
2
ð3Þ
The TMP is calculated by blood input pressure (PBi), blood output pressure (PBo), and
filtrate pressure (PF) in the plasma compartment.
In a laminar stream, the velocity profile will be distributed as shown in Figure 3.20.
The erythrocytes will be forced to go into the central stream.
As explained in Section 3.3, plasma separation is carried out in cross-flow filtration
mode. In the low-pressure range, the filtration performance increases linearly with
increasing transmembrane pressure. Later, a plateau is reached according to the wellknown concentration polarization phenomenon [7,30,36,37] in cross-flow filtration.
Blood is not a Newtonian fluid, but nevertheless it is well known that, according to
Hagen–Poiseuille’s law, the streaming volume is very sensitive to tube/capillary
diameter. Consequently, the blood flow rate, the plasma flux, and the corresponding
transmembrane pressure have to be in the right relation to each other [14,31]. As long
as the streaming velocity is high enough, the cells will be forced to move into the
center of the stream (see Figure 3.20).
If the wall shearing rate becomes too low and the membrane pressure is too high,
the risk of hemolysis increases. This means that the red blood cells (erythrocytes) are
destroyed and hemoglobin is released in the body circuit. The blood plasma becomes
Fig. 3.20 Velocity profile in a capillary
membrane [13].
Fig. 3.21 Schematic explanation of hemolytic behavior of
different membranes in blood contact depending on
transmembrane pressure and wall sheer rate.
red in color resulting in toxic behavior. This condition has to be prevented. The
starting phase of the treatment is particularly critical. For this reason, some filters
have to be processed at transmembrane pressures less than 50 mmHg. The user has
always to follow the care of cartridge producers’ manual.
In Figure 3.21, two membranes A and B are compared with respect to their
hemolytic dispositions. In general, the hemolytic tendency decreases with increasing
wall shear rate and decreasing pressure. Membrane B shows hemolysis related to
lower transmembrane pressures than membrane A. Above the dashed line is the area
where hemolysis takes place.
The schematic assembling of a plasma separation capillary membrane filter is
shown in Figure 3.22. The requirements according to module design for capillary
membrane plasma filters are not so high as for, for example, dialyzers [7,9,11,35].
Only a convective flow via the membrane wall takes place, and because the plasma is
only drained in the filtrate (plasma) compartment of the module, no defined
streaming conditions have to be created inside the bundle. Therefore, no fiber
makeup such as undulation is required. According to the above-explained criteria, the
filter area and the length have to be adjusted to the membrane. According to Table 3.2,
filter areas between 0.15 and 0.8 m2 are usual, depending on the application. A
special monitor steers the flows and pressures comparable to a dialysis monitor.
Plasma volumes between 2 and 4 L are filtered in one treatment session.
Flat membranes are also used in special filtration devices. Regarding the membrane area they are very effective, but need special monitors (Figure 3.23).
In the Autopheresis C system produced by Baxter/Fenval, IL, USA, a polyamide
(PA) membrane with an area of 70 cm2 is applied to a rotating cylinder. The blood
enters the top of the filter assembly and travels down the defined space between the
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86
Fig. 3.22 Schematic drawing of the module construction of
a capillary membrane plasma separation filter.
rotating cylindrical membrane and the outer cylindrical housing (stationary cylinder). The plasma filtered through the membrane flows out from the filters’ central
tube (axis). By rotation of the membrane, Taylor vortices are created, spiraling
downward. The rotation-induced Taylor vortex enhances the plasma flux [27].
In the device of Milttenyi Biotec GmbH, as described in Figure 3.24, less than
250 cm2 membrane area is required to filter several liters of blood. A polyethersulfone
Fig. 3.23 Flat membrane filter used in a rotating plasma
separation device. The inner rotating cylinder is
upholstered with a microfiltration membrane of polyamide
(Baxter/Fenwall, USA).
Fig. 3.24 Rotating disc plasma separator. Microfiltration
flat membranes of MicroPES (Membrana GmbH) with
rotating disk in between.
flat membrane MicroPES is fixed on both sides of the housing where the filtered
plasma is drained. A rotating disk in between and at a defined small distance from the
membrane ensures the right streaming force. This device is used to produce cell-free
plasma, that is, treated in regenerable specific adsorber columns.
Fig. 3.25 From sieving . . . to adsorption – from
diffusion . . . to convection membrane systems in
future with combined sieving and selective adsorption
feature inside the pores by grafted tentacles inside
the pores.
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88
View to the future: Classical adsorber technology uses beads packed in a column.
The beads carry the ligand and are exposed to the plasma or the blood stream.
Membrane technology is an alternative way to build adsorption devices. The inherent
advantage of membranes lies in the better perfusion of the plasma or whole blood.
This is seen in the convective transport of the target molecules to the ligand sites
without diffusive limitations. This is in contrast to the bead technology where the
adsorption kinetics is limited by the long diffusion pass (Figure 3.25).
3.7
Medical Indications for Blood Plasma Treatment
In Table 3.4, medical indications for plasma fractionation treatments are listed.
Nowadays, especially applications where the blood viscosity has to be reduced to
improve the microcirculation of blood have been successful in an increasing number
of cases [21,22,23,25]. For example, the membrane described in Figure 3.15 is well
adapted to this type of application [20]. The high-molecular-weight lipoproteins and
triglycerides are rejected completely, and the slight reduction in fibrinogen also
improves the viscosity behavior. Other illnesses require immunomodulation of the
blood by fractionation of immunoglobulins [8,11,32,38,39]. Special kinds of such
treatments are the so-called cryofiltrations where plasma proteins are precipitated
by cooling the plasma to 4 8C. The precipitated aggregates are then filtered by
Tab. 3.4
Medical indications for plasma fractionation by double filtration.
Indications belonging to top 10 of registry:
Hypercholesterolemia
Myasthenia gravis (MG)
Syst. lupus erythematosus (SLE)
Guillian–Barré syndrome (GBS)
Thromb. thrombocyt. purpua (TTP)
Cryoglobulinemia
New indications:
Age-related macular degeneration (AMD)
Sudden hearing loss
Diseases with rare incidence:
Burger’s arthritis
Cyclosporemia
Diabetic retinopathy
Gromerulonephritis
Polymyositis/dermatomyositis
Polyneuropathy
Small vessel disease
Vasculitis
Waldenstrom’s macroglobulinemia
W. granulomatosus
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membranes. In the so-called thermofiltration special proteins are precipitated by
heating the plasma to 42 8C and are then removed by filtration [11,12].
The relatively ‘‘old’’ technology of membrane plasma separation is now undergoing a renaissance because of new indications and the requirement for highly
sophisticated specific adsorber technologies for cell-free plasma [40–43].
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Bambauer, R. (1997) Therapeutischer
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Plasmaseparationsverfahren, Pabst,
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Microfiltration and Ultrafiltration:
Principles and ApplicationsMarcel
Dekker, New York.
Hörl, W. H. and Wanner, C. (2004)
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Malchesky, P. S. (2001) Therapeutic
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Nosé, Y. (1988) Artificial Organs, 12,
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(1998) Nieren und
Hochdruckkrankheiten, 27, 137–144.
Henderson, L. W., Koch, K. M.,
Dinarello, C. A., Shaldon, S. (1983)
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Rock, W. (1990) Apheresis, Wiley-Liss,
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Kaplan, A. A. and Halley, S. E. (1990)
Andrade, J.D. (1985) Surface and Interfacial Aspects of Boimedical Polymers,
Vol. 2, Plenum Press, New York.
Baeyer, H., Kochinke, F., Schwertfeger,
R. (1985) Journal of Membrane Science,
22, 297–315.
Jaffrin, M.J., Gupta, B.B., Ding, L.H.,
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Y., Baurmeister, U. (1991) International
56–60.
SMOLIK, G. (1989) Ph.D. thesis.
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(München).
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synthetischer Membranen, Steinkopf,
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38 Klingel, R., Fassbender, C., Fischer, I.,
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M. G., Schwandt, P. (1999) Therapeutic
Stegmayr, B. G. (2000) Blood
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Stoffel, W., Borberg, H., Greve, V.
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Braun, N. and Bosch, T. (2000) Expert
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91
4
Membranes in the Biopharmaceutical Industry
Anthony Allegrezza, Todd Ireland, Willem Kools, Michael Phillips, Bala Raghunath,
Randy Wilkins, Alex Xenopoulos
4.1
Introduction
Membranes for the biopharmaceutical industry operate in streams containing
relatively low concentrations of exceedingly expensive products in a highly regulated
industry. These considerations have determined the progress of membrane development for this industry and the manner of choosing and optimizing membrane
usage. In this chapter we will describe the present state of microporous, ultrafiltration
(UF) and virus removal membranes, and the developments that led to them.
Membrane use will be described from the point of view of a process engineer
who has to specify and optimize the various membrane separation and purification
steps in a biopharmaceutical manufacturing process.
From its inception, the biopharmaceutical industry has needed to concentrate and
purify the protein drugs it produced. While the first membranes used were those
available at the time, manufacturers have become more sophisticated and demanding in their requirements for membranes as the industry has grown and matured. In
addition, competition among membrane companies has increased. These trends
have led to new membrane products specifically designed for the needs of the
biopharmaceutical industry.
The response of membrane manufacturers to industry needs for microporous,
ultrafiltration, and membranes for virus removal has gone beyond providing
economical flux rate and satisfactory retention properties and has focused on other
needs. These include lowered extractables from the membrane device to the
permeating stream and the capability to withstand alkalies. More recently, the need
for increased capacity has resulted in newer asymmetric and composite membranes
and multilayered combinations of membranes and prefilters in filtration devices.
All membrane filtration processes are designed to remove some species and allow
the passage of solvent and other, usually smaller, species. The membrane can remove
large impurities and pass the product, or retain the product and pass solvent and salts.
Every manufacturing process is different, but some common themes prevail. The
ISBN: 978-3-527-31480-5
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4 Membranes in the Biopharmaceutical Industry
Fig. 4.1 Basic elements of a recombinant therapeutic manufacturing line.
process engineer seeking to develop and eventually optimize a process step must
identify the proper membrane through a combination of experience (e.g., manufacturer and reference literature) and laboratory experimentation, and then use
scaled-down testing to determine how the membrane will perform. Intermediate
sized testing is then used to determine operating parameters, and cleaning and
sanitizing regimes. Microfiltration, ultrafiltration and virus removal process development all follow these basics steps, with additional specific requirements for each.
Figure 4.1 shows a generic biopharmaceutical manufacturing line and will be
referred to throughout this chapter. The shaded boxes denote UF processes and the
shaded circle, the virus removal filtration step. All other filtrations are done using
microporous membranes. It is evident that microporous steps are numerous in the
overall process. With the exception of crossflow microfiltration sometimes used for
clarification, all microporous use is with dead-end disposable cartridges or capsules.
Microporous and UF membranes can be defined by what they do, and how they do
it. Microporous membranes are capable of removing particles larger than 0.1 mm and
being used primarily in dead-end filtration applications with disposable cartridges or
capsules. Ultrafiltration membranes are used for concentrating or diafiltering soluble
macromolecules that have a size in solution of less than about 0.1 mm and operating
continuously in a tangential flow mode for extended periods of time, usually more
than 4 hours and for up to 24 hours. We will see, however, that in the virus removal
application, specialized UF membranes are used in a dead-end mode.
4.2 Microfiltration Membranes Used in the Biotech Industry
Microporous membranes clarify the streams from the fermenter and the centrifuge output. They are used in the process train for aseptic filtration of media and
other streams entering the fermenter and for buffers used in chromatography steps.
Microporous membranes aseptically filter the final filling of dosage containers,
perhaps the most critical filtration step in the entire process train.
Ultrafiltration membranes concentrate and diafilter protein solutions. Diafiltration (DF) is used to change buffer solutions for subsequent process steps. Flat sheet
membranes dominate this area, although hollow fibers are used for some applications. Polyethersulfone (PES) and regenerated cellulose are the two major types of flat
sheet membranes. Hollow fibers are primarily PES. Regenerated cellulose finds
particular use where low protein binding of the expensive protein therapeutic drug
product to the membrane is desired. For uses where harsh cleaning conditions or
where protein binding is less important, PES membranes can be used.
Virus removal in the production process for biopharmaceutical drugs require
several separate robust steps with additive capabilities. Membrane processes have
been proven to be an important part in virus removal and are now found in virtually all
biopharmaceutical manufacturing processes. These membranes operate under
perhaps the most severe requirements in membrane technology. Not only must
they be capable of removing ‘‘3 logs’’ (99.9 %) of virus, but also they must pass almost
the entire protein product in the feed stream. As the desire to remove smaller
parvoviruses has developed, the difference in size between the protein and the virus
has diminished, making meeting these almost contradictory requirements more
difficult. In addition, virus removal membranes must be validatable.
The most prevalent virus removal membranes presently used are regenerated
cellulose hollow fibers, and flat sheet composite polyvinylidene fluoride (PVDF), PES
asymmetric, or PVDF symmetric membranes.
4.2
Microfiltration Membranes Used in the Biotech Industry
4.2.1
Introduction
In the early 1900s, Zigmondy [1] developed a process of making membranes that
could be used to retain bacteria. A century later, similar processes are still being used
to make the microfiltration membranes that are used throughout a typical biopharmaceutical process.
The more efficient pleated cartridge devices introduced in the 1970s accelerated
membrane development as stronger membranes were needed to withstand the
pleating process. Pleated cartridges made the process and sterile microfiltration of
liquids more economical. When the first biopharmaceutical products started moving
from the lab to commercialization in the early 1980s, these membrane devices were
available and eased process development. For decades they have served biopharmaceutical manufacturers with little change. A combination of increasing customer
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requirements has led to new microporous membranes and membrane structures. In
this section, we will discuss how present-day microporous membranes evolved.
This overview will attempt to rationalize why certain materials and specific
morphologies are chosen for certain applications and how one might choose
membranes for various applications. Although the original customer emphasis
was on the performance of the membranes (flux and rejection), these attributes
are now expected. Current emphasis is focused on the robustness of the membrane
device, and increased capacity.
Robustness refers to the control of variability of a filter device so that the device
operates successfully within the window of process variability. Capacity is the ability
to run longer at a reasonable rate and not become plugged. Increasing capacity has
been an empirical endeavor at best. Now, newer mathematical models more
accurately describe the plugging behavior of microfiltration membranes.
Final filtration applications require a high retention of Mycoplasma or the removal
of bacteria from process liquids. Although the guidance of Food and Drug Administration (FDA) is clear, the interpretation of retention assurance is less clear. Some
early and more state-of-the-art interpretations of retention assurance will be
presented.
4.2.2
Microfiltration Membranes: Development of Industrial Membranes
Ever since Zigmondy’s invention in 1914, membranes have been made by phase
inversion processes. In the Zigmondy process, a thin layer of cellulose nitrate
solution is evaporated and the resulting porous film used to filter bacteria while
passing other constituents of the feed. Zeman and Zydney [37] give an excellent
description of processes currently used to manufacture microporous membranes. In
most applications, microfiltration membranes are defined as those porous membranes that have the ability to retain bacteria but allow proteins to pass through. It was
only in the 1960s that through the availability of scanning electron microscopy, it
became clear that these films consisted of highly interconnected pores uniformly
distributed within the film.
In the 1970s, the need for new membrane materials for pleated cartridges led to
polyamide membranes (nylon) from Pall Corporation and AMF (now Cuno/3M), and
PVDF membranes from Millipore Corporation. These materials significantly
improved the mechanical properties of the membranes in part due to slightly lower
porosities. However, their strong mechanical properties allowed the fabrication of
robust pleated devices which more than made up for any loss in permeability.
Introduced as the biotech industry started, these devices are still being used in the
majority of the microfiltration applications.
Early microporous structures were web supported with woven or nonwoven fabrics
to make up for the inherent weakness of the membrane. Web-supported membranes
have the disadvantage that the web material is not compatible with certain posttreatments. For example, polypropylene (PP) is not resistant to gamma irradiation but
has a high melting point, whereas polyethylene (PE) is not resistant to high steaming
temperatures but is resistant to gamma. PP is therefore often the material of choice
for steamable cartridges while polyethylene can be used in gamma-treated cartridges.
With the development of more robust membranes in the 1970s for pleated cartridges,
a transition occurred from web-supported membranes to unsupported membranes,
which are presently used in the majority of the microfiltration applications.
The first microporous membranes were symmetric in structure, meaning that they
had a uniform pore structure through their depth or thickness dimension. In the
mid-1980s, the introduction of a new class of engineering plastics, first polysulfone
and then the more stable polyethersulfone resulted in asymmetric microporous
membrane structures developed at Brunswick (eventually US Filter, now part of Pall
Corporation) by Wrasidlo [2].
Further developments using polysulfone polymers came from Kraus [3] (assigned
to Gelman, now Pall Corporation and commercialized as Supor1) and Roesink et al.
[4] (assigned to X-Flow/Norit) who independently invented the use of hydrophobic–
hydrophilic polymer blends to create membranes that were intrinsically hydrophilic.
The Kraus patent, based on polyethersulfone, led to structures that are symmetric.
AKZO AG [5] (currently Membrana) and Fuji [6] filed patents on preparing asymmetric membranes based on a single-polymer solution that had a retentive portion of
the membrane located internally in the membrane, sometimes called an hourglass
symmetry from the profile of pore sizes through the membrane cross-section.
Despite being available for several years, asymmetric membranes were not successfully introduced into the biotech filtration marketplace until the introduction in 1997
of the Sartopore 2 from Sartorius, a polyethersulfone-based membrane with a
relatively low, hourglass-shaped asymmetry.
In 2003, Millipore followed with highly asymmetric composite membranes having
a superior permeability. In 2005, Pall Corporation introduced its Mach V technology.
These membranes resemble in microstructure the original Supor technology as
described by Kraus but show an asymmetric structure. Depending on the product,
low or medium asymmetries are available. Besides Millipore, Pall, and Sartorius,
other companies introduced products for this industry as well, that is, Membrana
(DuraPES) and Cuno/3M (BioAssure). Although most recent products are based on
polyethersulfone, the major market share in 2006 is still held by symmetric
membranes: Durapore (polyvinylidene fluoride) and Ultipor (Nylon 66).
4.2.3
Effect of Membrane Structure on Properties
Figure 4.2 below illustrates the different structures. These are scanning electron
photomicrografts of the cross-sections of the three types of membranes. The cartoon
below each is used, in some of the following figures, to help explain membrane
properties.
Asymmetric membranes have higher permeabilities for a given bubble point
compared to symmetric membranes. Symmetric or isotropic membranes can be
regarded as having a pore size that remains relatively constant throughout the
thickness of the membrane. Thus, if the membrane thickness were to be divided into
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96
Fig. 4.2 Types of microporous membrane structures.
equal fractions, each would contribute equally to flow resistance. For asymmetric
membranes, each equally thick fraction has a different pore size and different flow
resistance. For asymmetric membranes, because the retention fraction is only a small
part of the total thickness, higher permeabilities result. A new class of membranes,
composite asymmetric membranes having two zones within the membrane with
different pore size, show an increased flow performance compared to singlepore-size symmetric membranes and to its asymmetric variants.
The data shown in Figure 4.3 below illustrates how membrane structure affects
properties. The data are typical of commercial membranes used in the biopharmaceutical industry. For each type, as the water bubble point (BP) increases, indicating
smaller average pore size, permeability decreases. Bubble point is described in
Section 3.1.2.2. Asymmetric membranes with what is considered ‘‘low’’ asymmetry,
Fig. 4.3 Water permeability relationships for three types of
microporous membrane structures.
Tab. 4.1
Commercial microporous membranes by structure and polymer.
Brand name
Vendor
Membrane material
Membrane structure
Ultipor
Sartobran
Durapore
Supor
Fluorodyne IIP
Millipore Express
Sartopore 2
Supor Mach V
BioAssure
Pall Corp.
Sartorius Corp.
Millipore Corp.
Pall Corp.
Pall Corp.
Millipore Corp.
Sartorius Corp.
Pall Corp.
Cuno/3M
Nylon
Cellulose acetate
PVDF
PES
PVDF
PES
PES
PES
PES
Symmetric, web supported
Symmetric, web supported
Symmetric
Symmetric
Symmetric
Composite asymmetric
Low asymmetry, hour glass
Low asymmetry
Hour glass
where the ratio of pore size from one surface to the other is approximately 10, have
higher permeability than symmetric membranes for equivalent BP. The asymmetric
composites have the highest permeabilities.
Higher permeabilities are desirable for reducing the footprint of installed membrane devices (less membrane area needed ¼ less capital cost) for applications where
membrane pores are not plugged during filtration. For applications where the feed
stream plugs the membrane, capacity, that is, amount of liquid that can be filtered, is
more important, and permeability is less critical. Asymmetric membranes have
significantly higher capacity in plugging streams when used with the more open
pores at the feed side. This is more pronounced for streams that do not have any
prefiltration. Given this advantage, it is therefore also not surprising that all major
vendors of microfiltration normal flow filtration (NFF) devices are currently introducing devices that have asymmetric membranes.
Membranes from suppliers to the biopharmaceutical industry classified by membrane material and structure are given in the Table 4.1.
4.2.4
Aspects of Cartridge Design
For some steps in the manufacturing train, many commercial devices have a
combination of membranes layered in the cartridge, that is, a microporous guarding
layer and a sterilizing grade microporous layer (see Table 4.2). Especially in cases
where the loading on the guarding layer is not extremely high, using a guarding layer
instead of a separate prefiltration step is highly desirable due to process compression
and its associated reduction in footprint and hardware requirements.
Pleated cartridges have one or more layers of nonwoven support layers in addition
to the membrane. In pleated cartridges, cartridge output is the product of membrane
flux and membrane area. Membrane suppliers can make up for lower inherent
membrane flux by increasing membrane area in a cartridge, sometimes with very
imaginative pleat designs. See, for example, US Patents 5,543,047 and 6,315,130.
Pleated cartridge area (nominal 10 in. long by 3 in. diameter) ranges from 1 to 3 m2.
How effectively the membrane is used in such designs depends on the packing
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Tab. 4.2
Guarded microporous devices.
Brand name
Vendor
Pore size rating
combination
Sartobran
Multimedia Durapore
Multilayer Durapore
Supor EKV
Fluorodyne EX
Millipore Express SHC
Sartopore 2
Sartorius Corp.
Millipore Corp.
Millipore Corp.
Pall Corp.
Pall Corp.
Millipore Corp.
Sartorius Corp.
0.45/0.22
0.2/0.22
0.45/0.22
0.65/0.22
0.2/0.22
0.5/0.22
0.45/0.22
density and the type of support materials used, and the nature of the feed being
filtered.
Scaling studies in filtration process development for a particular stream will
quantitate device efficiency. The engineering scaling process (described later in
Section 3) starts with membrane disks, moving to small area devices and finally
large area manufacturing scale devices. The purpose is to determine inherent
membrane properties at small scale and then to determine any loss in effectiveness
for the conditions seen in the process being studied, due to cartridge packing or
design.
4.2.5
Membrane Surface Modification
In addition to the basic material and structure of the membrane, surface modification
approaches and chemistries differentiate membranes. The engineering plastics used
to make membranes are hydrophobic and exhibit high protein binding (adsorption).
Membrane manufacturers have devised many surface modification chemistries to
produce hydrophilic, low binding membranes. The patent literature contains a wide
variety of chemistries that go beyond the production of hydrophilic surfaces to
improved caustic stability, low binding capabilities for specific components and
increased temperature resistance. A review [7] by Ulbricht covers much recent
research.
Caustic stability refers to the desire of biotech producers to clean their system with
up to 1N alkali and not have filter properties changed by such cleaning. In final fill
operations, low binding is required for the product being filtered. For protein
filtration, an ultra low binding surface modification is needed. For other active
pharmaceutical ingredients (APIs), a low binding characteristic for those components is similarly useful.
The surface-modified membrane must be stable toward different sterilization
methods. Typical requirements for steam sterilization require temperatures significantly higher than 121 8C. Many sterile filters have product claims specifying
steaming temperatures of up to 136 8C. Stability toward steaming can be judged
by an integrity test (see Section 3.1.2.2) of these membrane devices. Such tests give
biotech manufacturers proof that the sterility assurance is maintained after steaming
and filtration. Typically, this integrity test is performed through either diffusion
testing or bubble point testing. Another method that is sometimes used is to steam
sterilize the membrane devices, and then determine the ability to obtain a sterile
filtrate when challenged with a high bioburden load.
Furthermore, steaming may change the surface chemistry resulting in a change in
the measured bubble point. Bubble point is directly proportional to the surface energy
of the test liquid, inversely proportional to the largest pore size in the most retentive
zone in the membrane and proportional to the cosine of the contact angle of the liquid
with the membrane surface. If steaming changes membrane contact angle sufficiently, interpretation of calculated pore size will be in error.
During membrane development, the likelihood of steaming having a deleterious
effect on wetting can be probed by measuring the critical surface energy of the
membrane surface before and after one or multiple steaming cycles. The critical
surface energy can be characterized by identifying the highest surface tension liquid
that instantly wets the membrane. Having a membrane with a critical surface energy
that is significantly higher than 72 dyn/cm is desirable to guarantee that proper
wetting occurs. Ideally, the critical surface energy is not changed by steaming or
autoclaving.
4.2.6
Sterilizing Filters
Sterilization by use of 0.2 rated filters in pharmaceutical and biotech processes is
more than just using a 0.2 rated membrane. A filter is qualified as sterilizing grade
based on its ability to completely retain microorganisms under the specific conditions and with a particular process feed stream composition. Test methods are
described in ASTM F-838-05 and HIMA tests. The FDA provides guidance in
the Guideline on sterile drug products produced by aseptic processing [8]. Although
the overall test conditions are clear, differences in some details still vary between
different manufacturers.
4.2.6.1 Retention
In the past, retention of sterilizing grade filters was expressed as the ability to retain
more than 107 microorganisms of a certain size. Since retention occurs through size
exclusion, good correlations were found between bubble point, a measure of the
largest pore size of the membranes, and retention. Results were expressed in log
reduction value (LRV) units.
LRV ¼ log10
colony forming units found in the filtrate
:
colony forming units challenged to the filter
Although the LRV gives a good indication what constitutes retention, it fails to
describe of what sterility means.
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100
The authors expect that with the implementation of the FDA’s new initiative [9]:
‘‘Pharmaceutical cGMP’s for the twenty-first century: a risk based approach,’’ membrane science in the near future will require a high level of understanding of where
certain risks originate. For sterilizing grade membranes, manufacturers will need to
better define how sterility assurance can be defined in a scientifically sound way.
Recently, a paper [10] was presented by researchers at Millipore Corporation in
which a statistical-based model was described by which sterility assurance can be
calculated. The researchers based the model using fundamental principles and
assumptions based on experimental studies of the retention of bacteria with microporous membranes. Rather than focusing on achieving a particular log reduction
value, this approach relied on calculating the assurance to obtain complete retention
downstream of a membrane filter as a function of bubble point. The method chosen
as most appropriate for the data was regression of logarithm of filtrate colony forming
units (CFU) versus bubble point, using retentive outcomes as censored (i.e., partial)
information. A linear model of isopropyl alcohol (IPA) bubble point versus retention
was fitted. From it, the probability of sterile filtrate was obtained by solving Probability
(retention) ¼ exp (10(LC)), where LC is the expected log count at a given bubble
point. This assumes that the probability of zero counts (i.e., retention) at a fixed
bubble point follows a Poisson distribution with parameter 10^(LC) when testing is
done under uniform conditions.
Retention confidence was obtained from the regression relationship between
bubble point and LRV (Figure 4.4). From this relationship, the minimum bubble
Fig. 4.4 Logarithm of colony forming units found on the
downstream side of the membrane as a function of alcohol
bubble point for a symmetric and asymmetric membrane. The
lowest bubble points with complete retention are indicated
with an upward arrow. The calculated minimum specifications
for bubble point with a 99.9 % assurance level that sterile
effluent is obtained is indicated with a downward arrow.
point of the membrane device is calculated for a confidence level of greater than
99.9 % that a sterile effluent is obtained when challenged with B. revundimonas
diminuta at >107 organisms per cm2.
4.2.6.2 Permeability
Membrane manufacturers are constantly trying to attain higher flux and capacity
while maintaining required sterility assurance. One illustration of how membrane
structure affects the balance between properties is shown in Figure 4.5. In this figure,
the bubble point corresponding to a sterile filtrate in B. diminuta testing is given for
the three types of membranes. To get the same retention assurance for an asymmetric
membrane, a minimum bubble point is defined, which is significantly higher than
that for the symmetric membrane such as are those currently being used commercially. Despite the higher bubble point, which generally reflects a smaller pore size, a
significantly higher flux is achieved for the asymmetric membranes as shown in
Figure 4.5.
Besides relying on the minimum bubble point for the sterilizing grade membranes
established by the membrane manufacturer, users of these membranes who sterile
filter liquids have to demonstrate that acceptable sterility is obtained in their process
fluid. This is described in detail in Section 3.
It is known that certain organisms could potentially penetrate through 0.2
sterilizing grade membranes, that is, Mycoplasma. For those cases, a smaller pore
size membrane filtration is recommended. Due to the smaller pore size, lower
permeabilities and capacities are typically obtained and a larger installed footprint is
needed to filter the same volume of process liquid.
Fig. 4.5 Bubble point corresponding to obtaining sterile
filtrate with B. diminuta for different membrane structures.
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4.2.6.3 Capacity
In many steps throughout the process, the membrane retains particulate matter
and other species. How these particles are retained whether on or within the
membrane structure will define the ‘‘dirt holding capacity.’’ Common practice is
to use a standard pore-blockage model or a caking model to predict, based on smallscale testing, what the total capacity is for candidate membranes. This will be
described in Section Membrane System Sizing – Capacity.
Currently, academic and industrial researchers are studying capacity with more
sophisticated models. Ho and Zydney [11] published a paper describing a transition
between a standard pore-blockage model and a caking regime. Bolton et al. [12]
combined a cake model with a complete pore-blockage model in a single equation to
capture this transition. Knowledge of what fraction of the volume filtered is governed
by complete blockage versus caking can be used to optimize membranes for its
capacity. Further research is underway in an attempt to describe capacity as a function
of asymmetry and as a function of prefilter layering.
4.3
Practical Membrane Considerations for Sterile Filtration by Microporous Membranes
Biopharmaceutical manufacturing is a membrane intensive industry and in particular
is a heavy user of microporous membrane devices. As illustrated in Figure 4.1 and as
discussed in this section, microporous devices are primarily used for sterile filtration for
input streams and for aseptic vial or other container filling at the end of the manufacturing process. In the following brief tour, we will start at the fermenter and show where
microporous membrane devices are placed in the various manufacturing steps.
The fermenter requires multiple filtration devices. Media and associated buffers,
which comprise the nutrients for the cells being grown, are sterile filtered and fed into a
bioreactor. Hydrophobic sterilization membranes filter any gas streams entering the
fermenter, and are on any vents during processing. Since they will not wet from liquid
splashing during the process, there is no opportunity for bacteria to grow through.
Except for microfiltration used for crossflow or tangential flow filtration (TFF)
clarification of the cell culture, all microfiltration unit operations focus on minimizing the risk for bacterial or Mycoplasma contamination of the product.
Biopharmaceutical manufacturers select media to maximize the expression levels
in the cells. In recent years, there has been a shift from animal derived, serum-based
media to plant derived media (i.e., soy digests) due to regulatory concerns with
Kreuzfeld–Jacob syndrome and other potential contaminants. A shift to chemically
defined media would be even more desirable but expression levels using these media
have not been optimized. As this trend continues, membrane filtration of media will
become easier and more economical.
To maintain sterility in the bioreactor, a pharmaceutical grade 0.2 or 0.1mm rated
membrane is used. For media where a high reduction in Mycoplasma is desirable,
different 0.1 mm rated devices are recommended. Many media, especially serumbased media, have a plugging character, and in order to obtain process economies,
microporous prefilter devices are commonly used prior to sterile filter devices.
4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes
Depending on the choice of expression system, proteins are expressed internally or
externally to the cell. For internally expressed proteins, the cell culture is harvested
when a predefined yield is achieved. The cells are usually concentrated by a
centrifugation step, although crossflow microporous membrane devices are being
used successfully as well. Once the cells are concentrated, the cells are lysed, most
typically in a high-pressure homogenizer. The cell lysate containing the product of
interest is then clarified to remove cell debris using depth filters and finally sterile
filtered into a holding tank.
Cell cultures which express protein internally, such as E. coli, are optimized to
yield high titer levels of expressed protein, which is typically associated with a low
viability of cells. Such cells are more easily ruptured during harvest. As a consequence, there is a high loading of cell debris that needs to be clarified. A second
clarification after the centrifugation or microporous TFF step is performed through
depth filtration. These are usually lenticular pad filters of bound cellulose fibers and
filter aids, such as diatomaceous earth. The number and size of colloids and other
particles exiting the depth filter depends on the attributes of the particular depth
filter being used and the feed stream. If this permeate stream is not clean enough,
the downstream sterile filter will become quickly plugged. An intermediate filter
may be needed.
A series of chromatography columns, that is, ion exchange, protein A, or other
affinity media, purifies the clarified product. Buffers used in chromatography are
usually sterile filtered to remove bioburden and other contaminants in order to
increase column life and economies.
For protein that is expressed external to the cells (i.e., Chinese Hamster Ovary cell
lines, NS0 cell lines), the cells are removed by centrifugation, TFF or depth filtration.
The separated product is then processed as above.
Figure 4.1 shows that large volumes of buffers are used throughout the process,
constituting a considerable usage of membrane filters. Since these buffers are clean,
membrane devices used to prolong column life are chosen primarily for high
permeability, to reduce the number of cartridges needed. It is expected that any
commercial sterilizing (0.2mm rated) membrane will have the retention needed.
Filtration of protein pools with high levels of protein can lead to undesirable
protein aggregation. In such cases, microfiltration membranes are used to
remove aggregates before the expensive virus removal filtration. This is described
in more detail in Section 6.
Once the proteins are concentrated to their final level, a final sterile filtration step
can be used to sterilize the liquids for vial or container filling. In order not to affect the
dosage level, a low protein binding characteristic of the membrane is required. In
most filtrations at this stage, permeability and capacity are less important than the
confidence that the membrane device can be validated.
4.3.1
Sterile Filtration Process Considerations
Properly done sterile filtration removes all living organisms from a fluid. It is a critical
application of membranes to the biopharmaceutical manufacturing process. To
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implement a sterile filtration process, the manufacturer must take great care in
material selection, system design, and scale-up to assure that the filtered product is
suitable for use. In this section, we will present a standardized methodology to assist
in this process.
We will cover membrane filter selection criteria including materials, retention
requirements and testing, flow and capacity performance testing, installation,
sterilization, integrity testing, and process validation.
4.3.1.1 Filter Selection
Generally, filter selection has four main considerations:
compatibility;
retention;
and two considerations which determine system sizing
flow rate;
capacity.
A logical approach to filter selection will consider these four criteria in the
order listed. For instance, there is no point in determining if a given filter has
sufficient flow for a given application if the filter materials are not compatible with the
feed stream.
Compatibility Compatibility covers how membranes and membrane devices withstand the conditions under which they are operated.
Chemical compatibility refers to solvation, swelling, or weakening of the membrane due to contact with the process fluids. Typical modern membranes used for
sterilizing filtration include PVDF, PES, polytetrafluoroethylene (PTFE used for
venting), and cellulosics. PTFE has the broadest chemical compatibility and cellulose
acetate is the most limited. Whatever the material, chemical compatibility guides can
be used to make an initial assessment of suitability. Guides can be found in
manufacturers’ literature or chemical engineering manuals.
Chemical compatibility can also be determined experimentally. Using membrane
coupons, key membrane properties such as water flux, weight, thickness, and
integrity can be measured before and after exposure to the product stream or
individual components of the feedstream. Changes in any of these key properties
indicate chemical incompatibility that could impact membrane performance.
Chemical compatibility also includes other feed-stream membrane interactions
such as adsorption of critical feed-stream components. Examples are proteins,
preservatives, and APIs. Here again, initial determination can be done by soak tests
with membrane disks and quantization done by spectroscopic or other analytical
methods. Pitt [13] provides some guidance and test methodologies for protein
adsorption. Hem [14] provides some guidance for preservative adsorption. In most
cases, however, the only true measure is with the actual formulation processed with
actual process parameters. One important reason for this is to be able to understand
the effect of nonmembrane components of the filtration device on adsorption and
filtration performance.
Heat resistance of sterilizing grade membranes is another key compatibility
question. Live steam is the most common method for sterilizing membranes prior
to sterile filtration. Most typical sterilizing membranes have very good thermal
compatibility. Typically, the limiting factor is the support materials used in cartridge
and capsule devices. Manufacturers’ recommendations should be followed.
Gamma irradiation is also typical for filters that are supplied sterile. PVDF and PES
are compatible with the levels of gamma irradiation required for sterilization,
typically 25 kGray. PTFE and cellulosics are not gamma compatible. AAMI provides
guidelines for qualifying and requalifying gamma sterilization processes [15].
A final consideration for materials compatibility is whether the membrane is
hydrophobic or hydrophilic. Hydrophilic membranes wet spontaneously with
water, and hydrophobic membranes repel water. Generally, hydrophilic membranes are used for aqueous streams, hydrophobic membranes for vent applications or gas filtration. Hydrophobic membranes can also be used to filter low
surface tension fluids such as alcohols, but hydrophilic membranes are suitable and
often used to minimize the number of different filters that must be stocked by an
end user.
Retention Retention requirements for sterilizing membranes are quite simple –
complete retention of microorganisms in the feed stream and complete passage of
the product components. Because this separation is primarily based on size, and
there is a generally a large difference in size between microorganisms and most
proteins or other APIs, membranes with pore size ratings of 0.2 mm or 0.1 mm will
provide the necessary bacterial retention and product/passage requirements.
When selecting a 0.2 or 0.1mm membrane, it is important to assure there is a
certified sterilizing grade claim. The general definition of a sterilizing grade filter is a
filter that when challenged with at least 107 cm2B. diminuta (ATCC19146) will
produce a sterile effluent. Standard methods are available (ASTM F-838-05) that
detail methods for B. diminuta retention testing. The ASTM method or close
derivatives are used by filter manufacturers for product development and quality
control tests. Some manufacturers use the test for membranes and some for
membranes and devices.
The ASTM test is designed to challenge each pore of the membrane with a single
bacterial particle. B. diminuta has long been accepted as a suitable ‘‘worst case’’
organism for this test [8,16]. It is suitable due to the small size, ability to grow
monodispersed and at high concentration. It is also desirable for its relative safety.
Although the ASTM test is considered a suitable QC release test for membrane
manufacturers, it is not sufficient for demonstrating the filter will provide suitable
retention in a specific drug manufacturing process. To qualify a filter’s retention in
actual production, a modification of the ASTM test must be used. The modified test
must take into account process operating parameters, which could impact bacterial
retention such as process time, differential pressure, and volume/area. In addition,
the modified test must take into account actual process bioburden and any impact the
process fluid may have on organism size. The result is a ‘‘custom’’ retention test
performed as a component of the validation process.
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The components of the custom retention test are as follows:
Identify native bioburden and compare it to B. diminuta to
determine worst case challenge organism. (Over 99 % of
existing custom retention validations have been performed
with B. diminuta.)
Determine if actual drug solution will inhibit microbial growth.
This would be the case with antibiotics or preservative containing solutions. A noninhibitory modified product must then
be used.
Spiked product or noninhibitory modified product with test
organism.
Challenge membrane under process scale down conditions,
keeping volume/area, time, and pressure consistent with fullscale process.
Acceptance criteria for the test are a challenge level of at least 107 organisms per
cm2 membrane area and 0 cfu downstream.
The results of this test are included with the process validation documentation
submitted to regulatory agencies. Results from the custom retention test, along with
documentation that membrane manufacturing processes are validated, and integrity
test processes are correlated to retention, are used to demonstrate that a specific
filtration process is sterilizing grade.
Membrane System Sizing – Flow Rate Requirements Compatibility and retention
define the type of filter to be used for a given application. The next step is
determining the necessary amount of filter area. Filter area requirements are
driven by two process parameters – flow rate and volume capacity. Examples of
process constraints that define flow rate requirements are as follows:
The need to filter a specific volume in a specific period of time
(e.g., one shift) or other processing time constraints. Tankto-tank transfers, such as buffer filtration processes, are
typically time based.
Minimum flow constraints, where the flow through the filter
must be maintained at a certain level in order to maintain
adequate process fluid to the next step. Filling operations are an
example where low flow means low fill volume and likely
process shutdown.
The driving force for flow is a critical component of flow rate determination.
Driving pressure can come from either positive pressure applied with gas pressure or
a pump. Or driving pressure can come from vacuum as is common in tank vent
applications.
When the flow requirements and driving force are defined, the first estimate of
filter size can be determined using the flow/DP (DP is the pressure drop across the
filter system) information provided by membrane manufacturers for the various
devices. These curves are given for standard fluids, usually water for hydrophilic
filters and air for hydrophobic. Information for water flow is usually given for
25 8C and 1 cps. Correction tables are available for water at temperature other than
25 8C.
A flow/DP curve, adjusted for feed stream viscosity, will provide a reasonable
estimate of membrane behavior at the start of the filtration. However, as the
membrane plugs or fouls, the flow/DP properties will change. The amount of
change varies with feedstream. In some cases, such as gases or purified water,
the flow/DP properties change very little and therefore the manufacturers’ flow/DP
curve will not only give an estimate of how the membrane performs at the start of the
filtration but also at the end. Sizing filters for these applications is simply a matter of
understanding the flow/DP curve.
Membrane System Sizing – Capacity However, there are many feedstreams that
have enough retained contaminants to significantly change the flow/DP properties of
the membrane. Examples include proteins, cell culture media, cell harvest, and so on.
In these cases, trials are generally required to determine how the flow/DP conditions
will change through the course of the filtration.
Capacity trials can be operated in one of two modes.
Abbreviated trials run at constant pressure, Vmax trials, can be
used when gradual pore plugging is the operative mode. A
major advantage of Vmax trials is the ability to screen multiple
membrane samples in a short time with minimum volume of
feedstock.
Constant flow trials can be used when gradual pore plugging is
not operative or for simulating processes run at constant flow.
Vmax is a method for predicting the throughput of filters based on a gradual pore
plugging model [17]. Gradual pore plugging occurs when colloids or suspended
matter collect on the sides of filter pores to gradually block them off, until a state of
total occlusion is eventually reached. This gradual blocking of the pores results in a
distinct filtration pattern.
In a Vmax test, the time and volume collected up to that time are recorded at regular
intervals. Test apparatus is shown in Figure 4.6.
A plot of time/volume versus time is made on linear/linear axes. If the results plot
as a straight line, the filter is considered to plug by the gradual pore plugging model
and the formulas of Vmax can be applied to predict filter life. If the results do not plot as
a straight line, it indicates the filter is plugging by some other model, such as cake
formation. In these cases, Vmax should not be used. A traditional flow decay method,
in which the filter is actually run with the feed stream or close surrogate until
completely plugged, should be used.
In the gradual pore plugging model, flow decay follows the equation
Q ¼ Qi ð1 kV=2Þ2 ;
ð1Þ
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Fig. 4.6 Vmax test apparatus.
where Q is the flow rate (L/min), Qi is the initial flow rate (L/min), k is the constant,
and V is the volume filtered (L).
At infinite time, Q ¼ 0, and Equation (2) results
Maximum volume filtered ¼ Vmax ¼ ðk=2Þ 1:
ð2Þ
When time/volume versus time is plotted and a straight line results, it follows the
equation
t=v ¼ At þ B;
where t is the time, v is the volume filtered up to that time, A is the slope, and B is the
y-intercept.
Solving for V gives the following equation:
V ¼ t=At þ B ¼ 1=ðA þ ðB=tÞÞ:
Letting time go to infinity, B/t becomes zero and
V ¼ Vmax ¼ 1=A:
Therefore, the inverse of the slope of the plot of t/v versus t equals the maximum
volume that can be filtered by the test filter. B is the y-intercept, or the value of t/v when
time is zero. 1/B has the dimensions volume per time, a flow rate. Therefore, the
inverse of the y-intercept is the flow rate through the test filter at zero time or the initial
flow rate.
Rearranging Equation (1) with Vmax = 2/k gives
Q=Q i ¼ ð1 V=Vmax Þ2 :
This equation can be used to determine the percentage of Vmax that has been used
when the flow rate has dropped to a percentage of the initial flow rate.
The process development engineer will use Vmax to test and compare multiple
membrane options in a short time with small volumes of feed solution. Vmax results
can be used to make initial choices with a small investment in time and product
solution. Since the trial does not simulate all the process parameters, such as process
flux, time, differential pressure, and volume/area, it is limited as to scaling accuracy.
For scaling purposes, it is highly recommended that process simulation trials be
run before finalizing the system size. The trial should simulate as closely as possible
the actual process mode (constant pressure or constant flow), actual process time,
volume per area, and membrane flux. Process simulation trials can be performed
with membrane coupons or small-scale capsules when sufficient quantities of feed
stock are available.
An alternative to Vmax testing is the constant flow Pmax test. Pmax is conducted with a
positive displacement pump feeding the test filter at constant flow. The differential
pressure is recorded and trended as a function of the volume filtered. The advantage
of Pmax tests is that they often more closely simulate the actual process. This is
especially true when process flux and maximum process operating pressure is
adhered to. In addition, unlike Vmax, the test is run until the filter is plugged.
The plugging volume is not predicted by extrapolation. Another advantage of Pmax
test is it is not dependent on any specific plugging mechanism. The disadvantage of
Pmax tests is it could potentially take a long time and large volume to run, as in cases
where filter fouling is very low. Pmax test apparatus is shown in Figure 4.7.
Fig. 4.7 Pmax test apparatus.
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Fig. 4.8 A typical pleated cartridge filter showing the pleat
components. Cartridges are inserted into permanently
installed housings.
4.3.1.2 Device Selection
There are two major categories of devices used in NFF – cartridges (Fig. 4.8) and
capsules (Fig. 4.9). Cartridges supplied by membrane manufacturers are inserted
into housings with an o-ring seal at the process site. Capsule filters include a plastic
housing with the membrane permanently bonded inside. Cartridges generally will
have lower cost per membrane area. Capsules can provide lower operating costs by
minimizing labor requirements.
Sterilization Considerations When deciding between cartridges or capsules, the first
key consideration is sterilization method. If live steam is to be used, a cartridge in
stainless steel housing will provide the highest degree of safety. Steam in place (SIP)
for capsules is not recommended for the vast majority of typical capsules. Capsules
with SIP claims require extensive safety precautions and must be operated with great
care.
On the other hand, autoclave sterilization is very common and suitable for both
cartridge and capsule type devices. Typically temperatures from 121 to 126 8C are used.
Gamma irradiation is the sterilization method of choice for ‘‘pre-sterilized’’
devices because it is highly lethal to all organisms. There is also the advantage of
being able to sterilize large numbers of devices in their final packaging, eliminating
issues of package exposure to steam or moisture.
Fig. 4.9 An all in one capsule filter. The pleated membrane/
support is sealed into the capsule and the entire device is
disposed of after use.
All sterilization methods must be qualified for effectiveness and membrane
compatibility. AAMI [11] provides guidelines for gamma irradiation dose qualification and requalification intervals.
Implementation Successful implementation of the chosen membrane, device and
sterilization method includes installation, sterilization, integrity testing, and running
the actual filtration.
An initial consideration for installation is to assure that there is sufficient space
above and around the filter installation to raise the housing dome, attach and detach
tubing and piping to inlet and outlet.
Before finalizing system installation, the process development engineer must
assure that the process scale system will perform as predicted from small-scale
testing. Typical mistakes include sizing based on membrane pressure drop, then
encountering flow limiting piping restrictions at full scale. For example, a common
mistake is to size a process based on pressure at source, with the membrane device
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being operated at a significant distance or head difference from the source, resulting
in less driving force than design.
The choice of sterilization method will have a significant impact on implementation. Easiest to use are disposable filter assemblies supplied pre-sterilized. These can
be obtained complete with whatever tubing, valves, containers, and so on, are
required for the application. All that is required is attachment to other processing
equipment.
Next easiest to implement are autoclaved assemblies of the membrane device and
associated tubing, and so on. These are assembled by the user on-site, wrapped in
protective material, and autoclave sterilized. Ideally, the assembly will be as complete
as possible prior to autoclaving in order minimize the number of aseptic connections
required post-sterilization.
Most difficult to operate are steam-in-place operations. First, there are safety considerations as operators must handle live steam lines. Second, the steam inlet pressure
and membrane differential pressure must be carefully controlled to avoid damaging the
membrane or device. Finally, as discussed in the microporous membrane section, the
membrane must be able to withstand steaming without a change of properties.
Integrity testing is performed to prove the membrane is free of oversized pores that
would compromise retention. Most common are nondestructive tests based on
capillary theory. The theory and practice of integrity testing have been covered
previously many times [18].
Nondestructive integrity tests are based on capillary forces as defined in the bubble
point equation
BP ¼ 4 kgcosu=d;
where BP is the bubble point, k is the shape correction factor, g is the surface tension, u
is the contact angle, and d is the pore diameter.
This shows the relationship between pore size and the pressure required to force
liquid out of membrane pores to initiate bulk gas flow. This is the fundamental basis
for all nondestructive membrane integrity tests. ‘‘Bubble Point’’ tests measure the
pressure required to initiate bulk gas flow. ‘‘Diffusion’’ tests measure the gas flow as a
pressure below the expected bubble point pressure. Excessive gas flow at a pressure
below the expected bubble point indicates the presence of oversized pores or defects.
Successful integrity testing is dependent on completely wetting the membrane
pores with test fluid. The most common test fluid for hydrophilic filters is water. But
because water has a relatively high surface tension, it can be difficult to get water to
fully penetrate all the pores. False integrity test failures due to poor wetting are
common and time consuming.
As discussed above, a common problem with hydrophilic modifications is they can
be removed or reversed by process exposure. Live steam will reduce the surface
energy of many modified polymers. In some cases, hydrophilic membranes
are rendered hydrophobic by steam sterilization. The result is limited flow or, more
commonly, false integrity test failure. Manufacturers’ recommendations for steam
sterilization should be followed to minimize this problem. Before final membrane is
4.4 Ultrafiltration and Virus Filtration Membranes for Biopharmaceutical Applications
specified into the process, trials should be run that simulate all process conditions,
including integrity testing and sterilization.
The information presented in this section is common to all sterile filtration applications in the biotech industry. A process engineer who is designing a specified one or
more sterile filtration steps will find that following the considerations given here and
cartridge manufacturers’ recommendations will result in a well functioning process.
4.4
Ultrafiltration and Virus Filtration Membranes for Biopharmaceutical Applications
Filtration membranes are generally classified based on their size exclusion properties. At the largest pore size, we have microfiltration membranes (typical pore sizes in
the range of 0.05–10mm). Ultrafiltration and virus filtration membranes, the subject
of this section, have pore sizes approximately in the range of 1–100 nm, corresponding to solute molecular weights in the 1000–1 000 000 Da.
Although there is overlap in the pore size range of ultrafiltration and virus filtration
membrane groups, their application uses are very different. As a result, different membrane types have emerged to dominate each application. For ultrafiltration, regenerated
cellulose and PES membranes are the two main types. Flat sheet membranes dominate
the market, but PES hollow fiber devices are also available. For virus filtration, PVDF
membranes are used, in addition to PES and regenerated cellulose. Both hollow fiber
(regenerated cellulose) and flat sheet devices are available.
We will cover ultrafiltration and virus filtration membranes separately.
4.4.1
Ultrafiltration Membranes
UF is widely employed in biopharmaceutical manufacturing to concentrate and
diafilter biological molecules, generally proteins [19]. Biopharmaceutical manufacturers use these process steps to attain a specified product concentration range and
buffer composition that is optimal for subsequent processing such as chromatography or formulation.
4.4.1.1 Membrane Suppliers
The principal suppliers of UF membranes for biopharmaceutical applications today
are GE Healthcare [20], Millipore [21], Pall [22], and Sartorius [23]. The first UF
membranes (nonwoven supported regenerated cellulose – YM series) were introduced commercially by Amicon – now part of Millipore Corporation – in the 1970s.
Millipore introduced nonwoven based PES membranes soon thereafter. Filtron –
now part of Pall Corporation – followed in the 1980s with Omega membranes (also
nonwoven based PES). Alpha PES membranes are also available from Pall for
applications where antifoam agents are a problem. The 1990s saw the introduction
of several new membranes. Millipore started offering cellulose (PL-series) and
Biomax (PES) membranes, a response to YM and Omega membranes, respectively.
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Manufacturers and brand names of ultrafiltration membranes for biopharmaceutical
manufacturing.
Tab. 4.3
Supplier
GE Healthcare
GE Healthcare
Millipore
Millipore
Pall
Pall
Sartorius
Sartorius
Cellulose
Ultracel PLC (5–1000 kDa)
Ultracel PL and YM (1–300 kDa)
Regen (10–50 kDa)
Hydrosart (10–100 kDa)
Cellulose triacetate (5–20 kDa)
PES
A/G HF (1–750 kDa)
Kvick (5–100 kDa)
Biomax (5–1000 kDa)
PT (10–300 kDa)
Omega (1–1000 kDa)
Alpha (10–30 kDa)
PESU (5–700 kDa)
PESU (5–700 kDa)
The novelty of Biomax membranes was that they were the first macrovoid-free
membrane structures. Sartorius introduced cellulose Sartocon Hydrosart as a
response to YM and PL, with the additional claim of sodium hydroxide cleanability.
Sartorius also offers PES and cellulose triacetate membranes. Millipore also introduced cellulose composite membranes, replacing the traditional nonwoven support
with a microporous membrane. The first hollow fiber devices were offered by A/G
Technology – now part of GE Healthcare – are made of PES and also claimed a
macrovoid-free structure. The most recent offering is Kvick PES flat-sheet membranes by GE Healthcare Biosciences AB.
Table 4.3 summarizes the key suppliers and membranes, along with the NMWCO
offered.
4.4.1.2 Membrane Selection
As is obvious from the above survey, PES and regenerated cellulose are the two
dominant types of ultrafiltration membranes. Regenerated cellulose is often selected
as a result of its low protein binding properties. This is especially important for
expensive protein therapeutic drugs. Low protein binding reduces fouling and
benefits a process with improved consistency, easier cleaning, and improved yield.
PES membranes are used in applications where harsher cleaning chemicals are used
or where low protein binding is less important. UF membranes are typically reused
several times, and for that reason cleanability and solvent compatibility is a key
consideration in membrane selection. Separate, but related, considerations apply for
membrane storage, as the devices need to be stored in appropriate preservative
solutions, in many cases, alkali, between uses.
Membrane selection for a given application is based on the following factors:
retention of the product protein;
retention consistency;
process flux and overall process economics;
scalability;
mechanical robustness;
chemical compatibility for cleaning/storage.
Selection of a particular pore size is usually done empirically and should always be
confirmed in the user’s particular application. A rule of thumb for selecting
membrane nominal molecular weight cut-off is to select one third to one fifth of
the product MW. For example, membranes rated at 30–50kDa should be used for
retention of a monoclonal antibody that has an MW around 150kDa. Even tighter
membranes should be used if anything above 0.01 % passage of the product is
unacceptable.
4.4.1.3 Membrane Structures
Flat sheet ultrafiltration membranes are typically used, cleaned, and sanitized,
sometimes stored, through many cycles. To obtain the required mechanical stability
and robustness for multiple uses, most UF membranes are made with a nonwoven
fabric support. Nonwoven polyoloefin supports have been used for a long time, as
they offer robustness and simplicity, as well as an open, permeable substructure.
Composite membranes (e.g., Millipore Ultracel PLC) are instead cast on microporous polyethylene membrane supports. The advantages are reduced microdefects that
could be caused by the rough surface of nonwovens and more uniformity and flatness
of the UF layer, resulting in high retention consistency. In addition, there is improved
adhesion of the UF layer to the substrate, resulting in a more mechanically robust
membrane.
The original UF membranes had large macrovoids. More recent membranes claim
a macrovoid-free structure. Although macrovoid-containing membranes have been
successfully used for many years, macrovoid-free structures provide higher assurance of retention and improved consistency. They have become the norm in most
current applications. Figure 4.10 shows examples of typical membrane structures.
4.4.1.4 Characterization
Ultrafiltration membranes are characterized in terms of a nominal molecular weight
cut-off rating, NMWCO. The rating corresponds to the retention to a certain level
Fig. 4.10 SEM images of typical ultrafiltration membranes:
macrovoids on nonwoven substrate (left), macrovoid-free
on nonwoven substrate (middle), composite, that is,
macrovoid-free on microporous substrate (right).
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(typically 90 %) of a marker molecule. Each membrane manufacturer has their own
characterization methods, and they should not be directly compared without knowledge of the methodologies used.
Originally, manufacturers characterized all UF membranes with a protein retention test. Each membrane was challenged with one or more protein molecules, and
specifications were set at a minimum required retention. Given the variability in
purity of commercially available proteins, the lack of relevance to specific protein
formulations used in the industry and the low robustness of a ‘‘minimum specification’’ approach, alternative methods were introduced. A mixed dextran solution with
molecules covering a broad MW range from 1 to 1 000 000 kDa was proposed and has
been adopted by the UF community [24]. Normal Molecular Weight Cutoff
(NMWCO) is typically defined as the dextran MW where 90 % of dextran molecules
are retained, a value known as R90. The advantage of such tests is that they provide a
common platform for testing a whole family of UF membranes, which would
otherwise require testing with different proteins.
In addition to retention properties, water permeability, back pressure (to
test integrity of the membrane–support interface) and air diffusion (to test for
the presence of large defects) are also usually measured and reported for UF
membranes.
4.4.1.5 Devices
Essentially all ultrafiltration membranes for bioprocessing are used in TFF devices,
with the exception of some virus removal devices. In TFF, also referred to as
‘‘crossflow’’ filtration, fluid flows across (tangentially) the filter membrane surface,
with a small fraction of the flow permeating the membrane (perpendicular to the
surface). The sweeping action of the tangential flow clears the surface of deposits
thereby minimizing membrane fouling and maximizing process flux and product
yield. The common device (module) formats used for biopharmaceutical applications
are flat sheet membranes used in plate and frame-type cassettes or spiral wound
module formats, and hollow fiber modules (Figure 4.11).
Cassettes can have an open channel or a turbulence promoting screen. Screened
cassettes have become the dominant module format in the biopharmaceutical
Fig. 4.11 Commercial ultrafiltration modules: hollow fibers
(left), spiral-wound (middle), cassettes (right).
industry, as they offer high mass transfer efficiency due to their controlled flow
channels, resulting in high fluxes at low tangential flows. Open-channel cassettes
are used for high solids or high viscosity feeds. A very important attribute of
cassettes is their ability to offer linear scale-up for reliability and speed of
implementation. Spiral wound modules are still being used in high-volume
pharmaceutical or industrial biotech applications. When cost is the main concern,
the high membrane packing density of spiral modules is an important benefit.
Hollow fiber devices are preferred when viscous fluids are being processed and an
open-channel device is needed.
4.4.2
Virus Filtration Membranes
Membrane processes have proven to play an important part in virus removal and are
now found in virtually all biopharmaceutical manufacturing processes. Virus
removal for biopharmaceutical drugs requires several steps with additive capabilities.
Virus membranes have severe requirements, as they must remove 99.9 % of virus
particles while passing almost the entire protein product in the feed stream [25]. As
smaller viruses (e.g., parvoviruses) have become important, the difference in size
between the protein and the virus has diminished, making size-based separations
even more difficult. The additional requirements (compared to UF) have led
membrane development in a slightly different direction. This section will only focus
on membranes that remove viruses by size exclusion. Membrane adsorbers also
claim some virus removal capabilities through electrostatic interactions, but they will
not be covered here.
4.4.2.1 Membrane Suppliers
The most prevalent virus removal membranes presently used are regenerated
cellulose hollow fibers and flat sheet polyvinylidene fluoride (PVDF) and PES
membranes. The principal suppliers of virus membranes today are Asahi [26],
Millipore [27], Pall [28], and Sartorius [29].
Filter manufacturers classify virus clearance filters into two broad categories based
on the removal needs of the biotech industry – filters that are capable of removing
viruses 50 nm or larger (retroviruses) and filters that can remove both small (20nm
parvoviruses) and large viruses.
Asahi Kasei Corporation (Japan) manufactures Planova filters, which were among
the first membranes specifically designed for virus filtration. They are dead-end
hollow fiber units with regenerated cellulose membrane made through the cuprammonium process (i.e., different from regenerated cellulose UF membranes). Pall
Corporation uses hydrophilic PVDF for its Ultipor DV50 and DV20 membranes.
Sartorius produces the Virosart CPV filter, which features a PES membrane.
For TFF removal of small viruses, Millipore offers membranes with hydrophilic
PVDF in two ratings, Viresolve 70 and Viresolve 180 membranes, designed to pass
proteins with 70 kDa or 180 kDa molecular weight, respectively, while retaining virus
particles. For NFF removal of small and large viruses, Millipore offers NFP and NFR
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Tab. 4.4
Commercial virus removal membranes and brand names.
Supplier
Large viruses only
Asahi (TFF/NFF)
Millipore (NFF)
Pall (NFF)
Small viruses
Asahi (TFF/NFF)
Millipore (TFF)
Millipore (NFF)
Pall (NFF)
Sartorius (NFF)
Cellulose
PES
PVDF
Planova 35N, 75N
NFR
Ultipor VF DV50
Planova 15N, 20N
V70, V180
NFP
Ultipor VF DV20
Virosart CPV
filters, respectively. NFP uses membrane similar to V180, while NFR features a
hydrophilic PES membrane.
Table 4.4 above summarizes available virus membranes, separating them according to the polymer used, the viruses removed and the filtration mode they are used in.
4.4.2.2 Membrane Structures
There are two broad families of virus membrane structures, skinned asymmetric
membranes (Millipore) and symmetric membranes (Asahi, Pall, Sartorius). Early UF
membranes of appropriate pore size were examined for virus removal, but it
was found that the pore size distribution was too broad to provide the necessary
removal [30]. Development of virus membranes, such as Viresolve V180, then
focused on eliminating macrovoids in the UF layer. The pore size in the ultrafiltration
layer increases gradually from the skin layer progressing towards the large-pore
microfiltration layer. The UF skin provides the selectivity needed to exclude viruses,
while the thicker microporous support layer provides mechanical support for the
membrane. The advantage of a thin retentive membrane is significantly higher
permeability.
Alternative approaches used a much thicker, dense retentive layer that traps virus
particles in its multilayer, interconnected pore structure [31]. Asahi Planova, for
example, has a 35 mm-thick filtration layer to assure high viral retention. Pall’s Ultipor
DV membranes are symmetric without a distinct skin layer. Commercial membranes
have average pore sizes 10 nm and larger in size, so that viruses, in some cases, as
small as 20 nm may be sterically excluded from the pores while still permitting
transmission of the desired protein product.
4.4.2.3 Devices
Both tangential flow (TFF) and normal flow (NFF) filtration devices are used for viral
clearance. In both cases membrane units are typically disposable and designed for a
single use, in order to avoid the strict validation requirements that would be linked to
cleaning a potentially infectious particle. Historically, TFF systems were more
Fig. 4.12 Commercial virus membrane modules. (Courtesy Pall
Corporation, Millipore Corporation, Asahi Kasei Medical, and Sartorius AG)
common, following UF experience. Recent improvements in NFF filters along with
their reduced complexity of operation and easier validation have led to a predominance of NFF filters. All NFF flat-sheet filters contain two to three layers of membrane
in order to increase virus removal assurance.
TFF has been covered in the UF section. In NFF, also referred to as ‘‘dead-end’’
filtration, fluid flows perpendicular to the filter membrane surface, with the flow
driven either with a pump (constant flow) or by pressure (constant pressure). NFF
virus filters are available in a variety of formats and filter areas to satisfy strict
requirements for scalability. In virus filtration, scalability affects not just process
economics but also validation of virus spiking studies. Small-area filters are available
for process development and virus validation studies, while for pilot and production
scale operations, filters with membrane area up to 6 m2 are available. Additionally,
multiple filter cartridges or module assemblies can be configured to achieve even
larger filtration areas. Figure 4.12 shows various commercial devices in use.
4.4.2.4 Membrane Selection
Criteria for selecting a virus filter tend to be product- and process-specific [32]. A
typical biotech process might contain more than one virus filtration steps. To
effectively evaluate the impact of these variables, in-house testing is typically
performed. Performance related criteria for selecting a virus filter include the
following:
virus retention capabilities;
protein product transmission;
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product throughput requirements;
overall process economics;
process compatibility;
thermal and hydraulic stress resistances;
extractables.
Compatibility with cleaning solutions is not an issue for virus filters, as they are
typically disposable. Detailed validation packages, including spiking studies, are
typically required by users of these filters.
4.4.2.5 Characterization
The key parameter needed to characterize a virus filter is its ability to retain viruses of
a certain size. This parameter is normally expressed in terms of a log reduction value,
LRV, defined as follows:
LRV ¼ logðCfeed =Cpermeate Þ;
where Cfeed and Cpermeate are the virus concentrations in the feed and permeate
streams, respectively.
For smaller viruses (such as parvovirus) membrane filters provide LRVs above 3–4.
LRVs above 6 are normally seen for larger viruses (such as retrovirus). These
parameters are typically determined during viral spiking studies, where appropriate
virus preps are used to challenge the membranes. Membrane manufacturers
have specific virus removal claims on their filters, but users generally perform a
separate spiking study, using their own product formulations. Regulatory agencies
currently require validation of virus removal steps using actual viruses.
For assurance of filter integrity, membrane manufacturers practice particle
challenge tests, using bacteriophage [33] or gold nanoparticles [34]. Compared to
viruses, the advantages of bacteriophages are increased prep purity, faster more
sensitive assays and safer operation. Gold particles offer much better control of size
and uniformity, and simpler detection. Particle challenge tests can be performed in
conditions that simulate more or less actual process conditions. The presence of
fouling streams, for example, might affect retention. Proper selection of the stream is
important to generate meaningful data.
To completely validate virus removal within a specific process run, novel
liquid porosimetric integrity tests have been developed that rapidly and nondestructively predict the intrinsic viral retention capabilities of virus-retentive membranes.
The CorrTest method [35] is essentially a ratio of two membrane permeabilities
measured at preselected operating conditions using a pair of mutually immiscible
fluids, one of which is employed as a membrane wetting agent and the other used as
an intrusion fluid. The CorrTest method can be used to verify pre- and post-use
membrane integrity and has been shown to correlate well with bacteriophage fX-174
retention.
Standard air diffusion tests are also performed on devices to assure against the
presence of gross leaks. In contrast to particle tests – that give information on the pore
4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing
size distribution of the membrane – air diffusion tests identify large defects on the
membrane surface.
In addition to integrity, adequate passage of the product of interest – or an
appropriately sized marker molecule – can be verified in a separate test. The
optimization of virus retention and product passage properties can yield a membrane
that is best suited for this application.
4.5
Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing
Biopharmaceutical manufacturers use ultrafiltration to concentrate and condition
biological molecules, generally proteins [35–39]. Ultrafiltration membrane processes
are used to provide optimum concentration ranges and/or buffer compositions for
subsequent processing steps.
An ultrafiltration process operates in a tangential flow (also known as crossflow)
mode to retain a biological molecule while letting small molecular components
such as buffer and water to freely pass through the membrane. In biopharmaceutical processing, ultrafiltration is primarily operated in a batch or a fed-batch
mode. Process operation times are typically between 3 and 8 hours. Manufacturing
scale system sizes range from being as low as 5 m2 of membrane filter area for
specialty therapeutics to as high as 300 m2 for high-dose monoclonal antibody
applications.
Figure 4.1 (see page 92) shows the typical location of ultrafiltration steps in the
downstream purification of a monoclonal antibody. There are broadly three
different areas where manufacturers use ultrafiltration in downstream biopharmaceutical applications:
1. Post-harvest concentration: The ultrafiltration step is used to
reduce volume and adjust the buffer conditions prior to
loading of the product on the first chromatography step. Low
feed concentrations (0.2–2 g/L), high process volumes (1000–
20 000 L), and higher level of impurities (DNA, lipids, host cell
protein, etc.) characterize this area. Bioburden and fouling play
an important role in process design in this process step. A
typical process will concentrate the feed stream 5–10-fold with
two- to threefold diafiltration.
2. Intermediate buffer exchange: This step is very similar to step
1 above except that the feed is more concentrated (5–30 g/L)
and process volumes (200–1000 L) and impurity levels are
lower. Typically, the feed is concentrated between 5 and 10 fold
with three- to fivefold diafiltration.
3. Formulation: The emphasis on the formulation UF step is
to carefully achieve a predetermined product concentration and
formulation buffer conditions. The product is very pure at this
step and may be concentrated up to 180–200 g/L. Diafiltration
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volumes are also relatively high, ranging between 8 and 12
depending on the extent of buffer exchange required.
The engineer faced with developing and operating an ultrafiltration process
operation will want to optimize the key elements of process consistency and
robustness, the ability of the process to operate under normal operating condition
variations, product yield, product purity, ease of use, scalability, validation, and
process economics.
The manufacturing process may impose constraints in the form of fitting a process
to an existing equipment, limitations on buffer or cleaning chemical usage, meeting
floor space requirements and product quality considerations at operating conditions
(pumping, process duration, etc.).
4.5.1
Ultrafiltration Theory
There are several excellent treatments of ultrafiltration theory in the literature
[36,37,41–43]. In this section, we examine the relationships among critical parameters that derive from the commonly employed theories to explain ultrafiltration
behavior.
In the polarization theory, the product concentrations may be related to the
permeate flux and tangential flow parameters as follows:
Cw Cp
J
¼ exp
;
Cb Cp
k
where Cw, Cb and Cp are the concentrations of the product at the membrane surface,
bulk feed, and permeate, respectively, J is the flux or permeate flow per unit area, and k
is the mass transfer coefficient, which is affected by tangential flow rate.
An empirical gel model is obtained from Equation (1) by taking Cp ¼ 0 and setting
Cg (or gel concentration) ¼ Cw to get
J ¼ k lnðCg =Cb Þ:
Equation (1) indicates that the ultrafiltration performance parameters, k and Cg, may
be obtained from a plot of J versus ln Cb, k is obtained from the slope, and Cg from the
x-intercept.
Finally, there is also an osmotic pressure model that relates process flux to
transmembrane pressure and osmotic pressure, P, resulting from concentration
of protein at the membrane wall. The osmotic model is expressed as
J ¼ L pðTMP DPÞ:
The method for determining the ultrafiltration performance parameters, k and Cw, are
detailed in [19].
Fig. 4.13 Characteristic flux versus TMP curve in an
ultrafiltration process.
Figure 4.13 shows a typical flux versus transmembrane pressure (TMP) curve. The
region at low flux/low TMP is called the linear region, where the flux is linearly
dependent on TMP. The region at high flux/high TMP is called the polarized region
and in this section, the flux is independent of TMP. In between these two regions lies
what is termed the ‘‘knee’’ of the flux curve. Increasing pressure beyond the knee
gives diminishing return in improving flux. The flux versus TMP curve is also
dependent on crossflow and protein concentration – a higher flux is generally
realized at higher crossflow rates or lower protein concentrations for a given TMP.
The intrinsic membrane sieving, also called passage, is defined as si ¼ (Cperm/
Cwall), while intrinsic membrane retention or rejection is defined as ri ¼ 1 si. The
intrinsic sieving is inherent to the membrane and solute, while an observed sieving as
so ¼ (Cperm/Cbulk) varies with polarization.
4.5.2
Typical Ultrafiltration Process
A typical schematic of an ultrafiltration process is shown in Figure 4.14. The process
starts with filling the protein product solution in the feed tank, recirculating this feed
using the feed pump, and removing virtually product-free permeate as a waste
stream. As the process proceeds, the tank volume drops and the concentration of
retained product in the tank increases. This concentration step proceeds until the
product concentration meets a certain target concentration. The buffer condition is
then modified using what is called the diafiltration step where a different buffer, or
diafiltrate, is added at the same rate as permeate is withdrawn (also known as constant
volume diafiltration). During diafiltration, the tank volume and retained protein
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Fig. 4.14 Ultrafiltration process schematic.
concentration remains constant (for r ¼ 1). Diafiltration is complete when the buffer
is changed to the desired composition. The solution then proceeds to the next step.
The relationship between product concentration and volumetric reduction factor
X ¼ (initial volume)/(final volume), the number of diafiltration volumes, N ¼ buffer
volume added)/(fixed retentate volume) and membrane rejection, r is
CRN ¼ C0 e½½ð1rÞNþrlnX :
Note that for freely passing solutes (r ¼ 0), >4 diavolumes are needed to achieve a
specification of <1% of the original buffer components.
The corresponding equation for retentate product yield may be written as
YieldRetentate ¼ eð1rÞðNþlnXÞ :
Permeate loss is expressed as 1 YieldRetentate.
Note that a membrane with 1 % sieving (99 % retention) can have process yield
losses much higher than 1 % because the protein is repeatedly cycled past the
membrane during the entire process with losses at every pass, that is, the higher the
N, the more is the yield loss. A high yielding process (<1 % total product loss) typically
requires sieving of <0.1 % (retention of >99.9 %).
4.5.2.1 Process Development and Optimization [38]
The process parameters that govern successful development, design and operation of
an ultrafiltration process are as follows:
1. membrane type and module characteristics;
2. operating parameter optimization;
3. processing plan;
4. system design implementation.
Membrane Type and Module (Device) Format Membrane selection for a given
application is based on factors such as chemical compatibility, mechanical robustness, and retention consistency for the product protein, as well as process flux and
overall economics (cost). Experience with vendors as well as their record of quality
manufacturing practices (ISO, cGMP) is also a consideration that influences membrane selection.
Operating Parameter Selection and Optimization A systematic process development
methodology is required to achieve a robust, consistent and optimized TFF step.
Candidate membranes are often screened using low membrane area (50–100 cm2)
linearly scaleable cassette devices – volume requirements for these devices are
between 50 and 500 cc. For more detailed optimization investigation, a 0.1 m2 device
is often used with process volumes ranging from 1 to 15 L. Pilot-scale processing,
which is often carried out to generate material for product stability
and characterization studies, is used to confirm operation scalability and
robustness. Volumes at pilot-scale range from 20 to 2000 L depending on the
application.
A typical process development sequence is described below:
1. Module installation and conditioning:
Flush and measure normalized water permeability (NWP).
Verify integrity using air diffusion.
Precondition membrane devices by flushing with buffer.
2. Initial fouling characterization:
Add product solution and measure flux and passage versus
time in permeate recycle mode (permeate and retentate
recycled to feed tank).
This is done to evaluate membrane fouling behavior and
allow for membrane conditioning in product solution.
(Note: The initial conditioning verifies that the module is integral and also establishes
as baseline permeability for the device. The fouling characterization study should
show asymptotic approach to steady-state flux and retention. Continued decline in
performance indicates a membrane compatibility issue.)
3. Flux and crossflow excursion in initial feed:
In a total recycle mode (permeate and retentate returned to
feed tank), measure flux, protein passage, and turbidity
versus TMP (10–60 psi) and crossflow (or DP of 10–30 psi).
(Note: Vary the TMP (10–60psi) at constant crossflow (say P ¼ 10 psi); repeat the
procedure at different crossflows (or P ¼ 20 psi, 30 psi, etc.).)
4. Concentration mode:
Measure flux, passage, turbidity, retentate concentration,
temperature, and permeate volume versus time while
removing permeate to achieve concentration.
5. Flux and crossflow excursion in final retentate:
Measure flux, passage, and turbidity versus TMP (10–50 psi)
and crossflow (P of 10–30psi) in total recycle mode.
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6. Diafiltration mode:
Repeat step 5 while withdrawing permeate and adding
diafiltrate at the same rate to the feed tank to achieve target
impurity concentration – this mode of diafiltration is
referred to as constant volume diafiltration.
7. Product recovery:
Recover retentate product by depolarizing and using a
plug-flow flush.
8. Measure yield.
9. Membrane CIP:
Flush with buffer, then WFI, add 0.1N NaOH solution,
recirculate for 45 min at 20 8C in permeate recycle mode,
and measure NWP.
NWP should be the same as the initial value. If not,
recleaning is required.
Additional development will include repeatability evaluation, cleaning optimization, and further exploration of particular conditions.
Selection of Optimum TMP and Crossflow A typical flux excursion behavior is shown
in Figure 4.15. The optimum TMP is located at the ‘‘knee’’ of the flux curve to obtain
reasonable flux and avoid formation of aggregates. If there is a difference in the flux
curves between the starting solution and the diafiltered solution, a conservative
approach for polarization would be to select the lower of the two TMP values.
Fig. 4.15 General flux trends in ultrafiltration.
Fig. 4.16 Process data for example scale-up calculations.
Choice of crossflow rate needs further consideration. Higher crossflow rates tend
to result in higher flux and reduced polarization, but this also leads to larger pump
size with a concomitant increase in holdup volume and pump passes (potential for
product degradation). For linearly scalable modules, the following approach may be
used for guidance in crossflow selection.
Figure 4.16 provides data for an example calculation. In this example, a device with
0.1 m2 membrane area is operated at two crossflow rates and the permeate flux plotted
versus TMP. The optimum point for each rate is chosen, and the permeate flux recorded.
For this example, the crossflow feed rates are
Q1 ¼ 0.65 L/min
Q2 ¼ 0.39 L/min
Crossflow feed rates normalized for membrane area are termed feed flux
JFQ1 ¼ 6.5 L/min/m2
JFQ2 ¼ 3.9 L/min/m2
The permeate flux at the optimum point for each case is
JfQ1 ¼ 150 L/m2/h
JfQ2 ¼ 86 L/m2/h
We assume V/T, permeate volume/process time is equal. Then, since
A ¼ Membrane area ¼ V/T permeate flux,
A1/A2 ¼ (1/150)/(1/86) ¼ 0.57
The area at the higher crossflow rate is 57 % of the lower crossflow rate.
Now we calculate the ratio of feed rates required at each feed crossflow.
Pump rate ¼ feed flow flux area ¼ JFQ A;
and the ratio is 6:5=3:9 0:57 ¼ 0:95:
In this example, the difference between pumping rates is not significant, so one
would chose the higher crossflow rate, because less membrane area would result in
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lower capital costs and smaller system size. In other cases, the shapes of the permeate
flux versus TMP curves may dictate otherwise. It is also important to note that this
method applies to systems that scale linearly (i.e., crossflow per unit area is the same
for all system sizes). For nonlinear system scaling, for example, some hollow fiber or
spiral wound systems, the procedure for selecting crossflow rates is based more on
experience than any analytical method.
Flux typically drops as protein concentration increases, so choose an average flux
for the above calculation. For a robust scale-up, always incorporate a safety margin
into the membrane area requirements to account for lot-to-lot variability effects on
membrane permeability, feedstock fouling characteristics, and batch volumes.
Typically, a safety margin of at least 20–50 % extra membrane area is used, but this
could increase or decrease depending on the expected variability in the process.
4.5.3
Processing Plan Optimization
In the development of a processing plan, the engineer has to consider choice of the
mode of operation, diafiltration strategy, process time, temperature, and an assessment of process robustness vis-à-vis lot-to-lot variability associated with feed and
membrane.
4.5.3.1 Mode of Operation
Most UF processes in the biopharmaceutical industry are operated in a batch mode
(as opposed to continuous mode). In a batch mode, the entire product feed is
recirculated within the system (no new feed is added to the system); permeate is
continuously removed from the system during this time resulting in an increase in
product concentration. The retentate, at its desired concentration, is available at the
end of the processing time.
When large concentration factors (>15–20) need to be achieved, a variation of
batch process called fed batch is often used (Figure 4.17). The configuration provides
some flexibility in processing a variety of batch volumes in a single skid. For a
Fig. 4.17 Fed-batch configuration.
Fig. 4.18 Fed-batch process concentrations.
fed-batch operation, the retentate is returned to a smaller tank, not the large feed tank.
Feed is added to the small retentate tank as permeate is withdrawn. The smaller
retentate tank can allow a smaller working volume without foaming.
Figure 4.18 shows the retentate concentration over the course of the process
as a function of the tank ratio ¼ VFeed tank/VRetentate tank. Notice that the fedbatch process operates at a higher overall concentration compared to the batch
process and consequently at a lower average flux. Typical fed-batch tank ratios are
chosen to be <4–5 to allow for process flexibility, while minimizing loss of
operating flux.
4.5.3.2 Diafiltration Mode/Strategy
Constant-volume diafiltration is commonly used control mode for adding diafiltration buffer. To perform a constant-volume DF, buffer is added to the recycle tank at
the same rate that filtrate is removed. The total volume of retentate remains constant
throughout the process.
Another important aspect of the diafiltration process is its placement (strategy) in
the overall process. The optimum point for placement of diafiltration may be derived
from the gel polarization model [44] as (Figure 4.19).
Copt ¼ Cg =e;
where Cg is the concentration at which the flux drops to zero and e ¼ 2.718, which is
the base of the Naperian logarithm.
For monoclonal antibody processing, Cg values of 200–300 g/L have been
reported, yielding an optimum diafiltration concentration of 70–110 g/L. Until
recently, final antibody concentrations have generally been in the 5–20 g/L range,
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Fig. 4.19 Optimum diafiltration point – polarization model.
where it is convenient to diafilter to the final formulation concentration. That is, it is
simpler to concentrate to 20 g/L and diafilter at this concentration (i.e., you would
not take the concentrate to the optimum range of 70–110 g/L in order to diafilter
followed by dilution to come back to 20 g/L). However, with a recent push towards
high formulation concentrations, in the range of 150–200 g/L, diafiltration at the
optimum diafiltration concentration in process applications is increasingly becoming a reality. Since, for example, if the final desired concentration is 150 g/L, one can
pause during the concentration step at 70–110 g/L (which is the optimum diafiltration concentration) in order to carry out the diafiltration, and once diafiltration is
complete, further concentrate to the desired final concentration of 150 g/L.
A more general approach for selecting the optimum diafiltration point is to use the
DF optimization parameter approach. In this approach, one calculates the diafiltration
optimization parameter, DF ¼ C J, at each protein concentration (from the plot of the
flux vs. protein concentration curve). A plot of the DF optimization parameter versus
protein concentration is constructed to find the optimum protein concentration where
the value of the DF optimization parameter is maximized (Figure 4.20). It is important
to plot these data with the protein in both the initial and final buffers, since flux can
often change significantly with different buffers. If the optimum is very different for
the two buffers, it is most conservative to choose the optimum based on the buffer
curve which results in the lower value for the product concentration.
Finally, the goal of a diafiltration step is to reduce buffer or contaminant species
from a product in the retentate. Since increasing the number of diavolumes that are
performed adversely impacts yield, it is best to minimize diavolumes consistent with
obtaining correct final composition. Figure 4.21 illustrates the relationship between
contaminant removal and number of diavolumes. The maximum effective diavolume is about 14 before the limits of the system or the chemistry are reached (system
deadlegs, mixing effects, etc.).
Fig. 4.20 Optimum diafiltration point – diafiltration
optimization parameter approach.
4.5.4
Scale-up Considerations
Once a process has been developed at lab bench scale, it must be translated to
industrial scale and validated, and this can present unanticipated surprises and
challenges. Often, there is little opportunity to perform intermediate-scale runs due
to time and material constraints. In addition, material from the initial industrial-scale
runs is usually required for time-sensitive clinical or market supply. Therefore,
accurate and dependable scale-up is critical for the success of a process. The simplest
way to ensure accurate and predictable translation of product yield and purity from
bench to industrial scale is to use linear-scale techniques.
Fig. 4.21 An example of contaminant removal by
diafiltration.
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To linearly scale a TFF process [45], all fluid dynamic and membrane module
parameters must be kept constant between scales of operation. Fluid dynamic
parameters, which are set by the user, include the ratio of feed volume to membrane
area, feed rate per membrane area, filtrate flux, and retentate and filtrate pressures.
Membrane module parameters are inherent in the specific filters that are chosen, and
include membrane material and pore size, turbulence promoter, channel height,
channel length, and feed and filtrate flow geometries. Therefore, true linear scaling
can only be done with modules from the same supplier.
4.5.4.1 Process Implementation Considerations
Once a protein processing procedure has been developed, it must be integrated into a
complete process. The typical sequence of steps in an ultrafiltration/diafiltration
process are set-up and pre-use cleaning, integrity and permeability testing, pre-use
equilibration, protein processing, product recovery, post-use cleaning and testing and
storage (Figure 4.22).
The integrity test helps to ensure that the installed membranes have not sustained
any damage during storage and handling. An integrity testing is normally carried out
prior to startup and after each post-use cleaning. The typical integrity test for an
ultrafiltration system involves measuring air diffusion across the wetted membranes
at a specified pressure. The diffusive flow is compared against the specification
provided by the membrane manufacturer. The air flow integrity test allows for
identification of problems such as macroscopic holes in the membrane, cracks in the
seals, or improperly seated modules.
Fig. 4.22 Typical operational sequence for an ultrafiltration
process.
It is also common practice to measure and compare the permeability (also referred
to as normal water permeability or NWP) of the membrane system to assess its
cleanliness from run-to-run. The normal water permeability is calculated as
NWP ¼ ðflux temperature correction factorÞ=TMP:
The units of NWP are [L/m2/h psi/g].
Product recovery strategy is a key component of the process operating sequence.
Product recovery is the process of removing the product from the TFF system into
a vessel appropriate for storage or further processing. The bulk of the product,
which is typically in the recycle tank, is pumped out using the feed pump.
However, some liquid remains held up in the piping and the modules. A welldesigned system has minimal deadlegs in the piping and is sloped to a collection
port at the lowest point in the piping to improve draining. Beyond simply draining
the system, however, one of the following methods can be used to increase
product recovery:
low-pressure air blowdown;
buffer displacement;
buffer flush;
buffer recirculation.
Overall product recovery in the final retentate should be >95 % and yields of
>99 % are common.
After each membrane use and product recovery, the system assembly is cleaned
using standard cleaning protocols [46,47]. Cleaning/sanitizing chemicals that are
commonly used in ultrafiltration processes include 0.1–0.5 N NaOH, 200–400 ppm
sodium hypochlorite, 0.1–0.3 N phosphoric acid, peroxiacetic acid, and so on. If
another protein processing run does not immediately follow the cleaning, recirculate an appropriate storage solution through the TFF assembly to prevent
organism growth. The membrane modules must remain filled with storage
solution until the next run to prevent drying of the membranes.
4.5.4.2 Process Robustness
A process implemented at large scale must be robust if it is to be successful. A robust
TFF process will perform well within the lot-to-lot feedstock and membrane variations that it encounters. While developing a process, it can be very useful to test
performance at the extremes of these variations, as far as possible.
These performance parameters characterize a well-designed, robust ultrafiltration
process:
Yield (overall)
Process flux consistency
Product retention (membrane)
NWP recovery (run-to-run)
95–98 %
10 % run-to-run
99.9 %
20 % cellulose membranes
20–35 % for PES membranes
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Typical flux
Typical sizing
Product and process specific; for example,
30–120 LMH for 30 kDa membrane
with MAb
Product and process specific; for example,
5–10 m2/KL for MAb
4.5.5
System Considerations
To implement a complete TFF process, the piping and equipment associated with the
membrane modules must be selected and a method for controlling the process
parameters at their set-points must be chosen.
4.5.5.1 Equipment Options
In addition to the membrane modules and holder, at minimum a working TFF
operation requires a recycle vessel, a feed pump, a retentate control valve, and
pressure sensors for the feed and retentate lines. Many systems also include feed and
filtrate flow meters, a filtrate pressure sensor, and sensors for temperature, pH,
conductivity, or UV absorbance. Most TFF systems used for protein processing are
operated in a sanitary manner, requiring sanitary fitting.
4.5.5.2 Process Control Options
Throughout a TFF process, as protein is concentrated or exchanged into different
buffers, the process parameters need to be adjusted so that they remain at their set
points. Several methods of process control are used to accomplish this. The tangential
flow can be controlled to maintain either
constant pressure drop;
constant crossflow rate.
Controlling to constant pressure drop is a simple and easy option in the sense that
it requires the use of only feed and retentate pressure gauges. The exact feed flow is
not known. Since pressure drop is uniquely related to crossflow (P ¼ AQn), a constant
pressure drop is assumed to connote a constant flow rate. However, in many cases a
constant pressure drop may not always ensure that the flow rate is maintained
constant. Differences in the internal membrane module resistance, pump slip, fluid
viscosity changes, and so on. during a process may contribute to changes in the fluid
crossflow without impacting pressure drop. Controlling to constant crossflow rate is a
more robust way of controlling the process. A flow meter is required in the feed or
retentate flow line for this control option.
The applied pressure can be controlled (by modulating the retentate valve) to
maintain a constant
Retentate pressure;
TMP;
4.6 Practical Aspects of Virus Filtration Process Design and Implementation
flux;
Cwall (protein concentration at the membrane surface);
mixed mode control.
Our process development engineer, having reached this point, will find that by
following the systematic methodology presented here will have an excellent understanding of the biopharmaceutical manufacturing process that results. Our engineer
will be able to produce the necessary SOP’s, validation documents and operating
instructions from detailed experience. In operation, the variance in day-to-day
running can be expected to be controlled based on the knowledge gained during
scale-up and implementation.
4.6
Practical Aspects of Virus Filtration Process Design and Implementation
Recombinant products such as monoclonal antibodies are expressed by mammalian,
bacteria or yeast cells in fermenters, and within ascites fluid, or fluids from
transgenic mammals. Mammalian cell lines may contain endogenous viruses that
are generated in the bioreactor. Endogenous retroviruses are expressed because the
retroviral genome is integrated into the cell line and cannot be screened out during
the creation of the Master Cell Bank. This causes retrovirus like particles, or RVLPs, to
be produced within the bioreactor. Products can also become contaminated by
adventitious viruses, which enter the process streams.
Although regulatory agencies do not mandate the inclusion of specific viral
clearance technologies, they do mandate the safety standard for final doses, and
require full validation of viral clearance steps by the manufacturer. Regulations and
good manufacturing practice (GMP) require two orthogonal and robust steps for
endogenous viral clearance [48]. Orthogonal steps rely upon different mechanisms to
achieve virus clearance. For adventitious viruses, these requirements typically result
in the use of multiple viral clearance steps [49].
Filtration removes viruses based on size exclusion, since the pores of the filter
are smaller than the virus. Virus filtration is considered a robust operation because the
removal efficiency is insensitive to normal variations in process conditions. Consequently, most well-designed downstream processes include a virus filtration step.
4.6.1
Membrane Selection
There are many commercially available virus filters, each with their own strengths
and weaknesses. Filter selection should be based upon the nature of the application
(product and process) and the performance of the filter as demonstrated in both
qualification and validation studies. A systematic methodology for developing a
robust virus removal step includes the following:
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virus filter selection;
process design and optimization;
process sizing and simulation;
virus validation studies;
manufacturing implementation.
Virus clearance filters are classified into two broad categories to meet the removal
needs of the biotech industry – filters that are capable of removing viruses 50nm or
larger (retroviruses) and filters that can remove both small (20nm Parvoviruses)
and large viruses.
NFF virus filters are available in a variety of formats and filter areas. Small-area
filters are generally available for process development and optimization studies as
well as virus validation studies. For pilot and production scale operations, commercial vendors offer individual virus filters ranging in size from 0.07 to 6m2.
Additionally, multiple filter cartridges or module assemblies can be configured to
achieve even larger filtration areas. Commercial filters that are currently available
for virus removal applications in the normal flow filtration mode are summarized
in Table 4.5.
Tab. 4.5
Vendors and brand names of commercially available virus filtration products.
Company type
Product name
Virus
Sizes
Millipore FF
NFP
>4 log X-174 bacteriophage
Scale-down 3.5 cm2;
process modules
0.08–1.5 m2
Millipore TFF
NFR
Viresolve 70
>6 log retrovirus
>4 log polio;
>7 log retrovirus
Sartorius NFF
Viresolve 180
Virosart CPV
>3 log polio; >6 log retrovirus
>4 log PP7 bacteriophage;
>6 log retrovirus
Pall NFF
DV20
>3 log PP7 bacteriophage;
>6 log PR772 bacteriophage
Asahi TFF/NFF
DV50
Planova 15N
>6 log PR772 bacteriophage
>6.2 log parvovirus; >6.7 log
poliovirus
Planova 20N
>4.3 log parvovirus;
>5.4 log Encephalomyocarditis
>5.9 log Bovine viral diarrhea
virus; >7.3 HIV
Planova 35N
Scale-down 150
and 1000 cm2;process
modules 0.75–1.4 m2
Scale-down module 5
and 20 cm2; process
module through
0.7–2.1 m2
Scale-down 14
and 140 cm2; process
modules 0.07–6 m2
Scale-down modules 10
and 100 cm2; process
modules 0.12–4.0 m2
Performance-related criteria for selecting a virus filter include virus retention
capabilities, protein product transmission/product recovery from the filtration step,
product throughput (rate) requirements, and overall process economics. These
criteria tend to be product specific, and to effectively evaluate the impact of these
variables, in-house testing is typically performed. Less obvious considerations for
selecting a virus filter revolve around process compatibility and system integration
issues. All materials of construction must be chemically compatible both with the
protein product as well as all relevant processing conditions.
Additionally, thermal and hydraulic stress resistances, extractibles, and the
effects of cleaning/sterilization/sanitization on the membrane devices should be
evaluated to determine if they are consistent with the proposed implementation
scheme. Additional information on these topics can be found in the PDA
Technical Report No 41 on Virus filtration [49] or from the various vendors listed
in Table 4.4.
To properly optimize a virus filtration process and establish process robustness,
one must consider all processing variables that impact virus retention (LRV), product
recovery, and product throughput. Additionally, from an economic point of view,
process optimization is extremely important. For large volume processes, such as
monoclonal antibodies, virus filtration can be one of the most expensive unit
operations. Virus filtration is more expensive than sterile filtration due to both
higher filter costs and lower product throughputs. Table 4.6 provides a typical range
for cost and performance parameters for virus and sterile filters.
Virus filtration can either be run in a NFF mode or a TFF mode [50]. Historically,
TFF systems were more common, but recent improvements in NFF filters have lead
to their predominance. Because of this predominance of NFF approaches, TFF
systems will not be discussed further here.
In NFF, also referred to a ‘‘dead-end’’ filtration, fluid flows perpendicular to the
filter membrane surface. NFF processes can be run either under constant flow
operation or constant pressure operation. Constant flow is more common in
manufacturing settings as most people would use pumps to drive the filtration
and it is simpler to run at a constant pump setting.
Tab. 4.6 Comparison of virus filter costs, process fluxes, and capacities compared to sterilizing
grade filters.
Filter type
Virus
Parameter
Sterile
NFF 20 nm
virus filter
NFF 50 nm
virus filter
TFF virus
filter (>20 nm)
Unit filter cost ($/m2)
Typical process flux (L/m2/h/psi)
Typical design capacity (L/m2)
200–300
200–500
>2000–4000
>2000–4000
0.5–6
60–500
>1000
20–30
800–1500
>2000–4000
5–15
250–500
NFF, normal flow filtration and TFF, tangential flow filtration.
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During operation, protein products or other components can accumulate at the
membrane surface or adsorb to internal surfaces. These two fouling mechanisms
will reduce the hydraulic permeability of the membrane and may impact virus
retention. Fouling results in a decrease in filtrate flow rate with time for constant
pressure operations or an increase in upstream pressure with time for constant
flow operations.
4.6.2
Process Development and Optimization
A schematic representation of the experimental set-up to conduct normal flow
filtration experiments at constant pressure is shown in Figure 4.23. The scale-down
test may be carried out in either a constant pressure or a constant flow mode. The
constant pressure set-up is often simpler, and the testing is easier to execute in that
the set-up does not require a pump to drive the filtration process.
The typical steps employed in a scale-down NFFprocess evaluation are described in
the following general test protocol.
System set-up.
Water flush.
Installation check: Flush the filter with water and carry out a
pressure hold test to confirm installation integrity of the
devices. (This step is optional for process development and is
typically carried out in virus validation studies.)
Fig. 4.23 Experimental set-up for virus removal testing.
Buffer conditioning.
Product filtration: During filtration, filtrate volume (V) collected
is measured and recorded at various filtration times (t). Filtration time may vary between 45 and 120 minutes. Assay
filtrate sample for product concentration.
Recovery: Assay buffer flush sample for product concentration
and calculate product recovery.
Installation check.
Calculate filter capacity and initial flux: The filter capacity and the
initial flux are generally obtained by fitting the experimental
data (V vs. t) to the gradual pore plugging model. (Vmax – see
Section 3.1.1.4.)
Calculate minimum required filter area: Once the filter capacity
and initial flux values are calculated, the minimum required
filter area can be calculated from the Vmax model. This value
should only be used as a comparative tool for different devices
during process optimization studies. The final design filter area
will be determined during process simulation and virus validation studies.
Optimization of a virus filtration process involves evaluating the effect of a variety
of process parameters to arrive at optimum conditions that would ensure robust,
consistent, and scalable operation. Figure 4.24 represents a generic approach to
optimization schematically.
Some of the key process development parameters that impact process performance are described in more detail below. There are typically three locations within
the downstream process train where a normal flow virus filtration step is implemented within a given downstream process. These locations are as follows:
following the low-pH inactivation step;
following the intermediate chromatographic operation;
or after the final chromatography step.
Since protein concentration, impurity concentration, and process volumes vary
dramatically throughout the downstream process train, it should come as no surprise
that the actual filtration requirements are highly dependent upon where in the
process the virus filtration step is located.
The benefit of increasing filter capacity and flow at lower product concentrations is
offset by an increase in process volume. The interplay of these two competing effects
can often result in a feed concentration that minimizes required filtration area [51].
An optimum feed concentration may exist that maximizes filtration performance
(minimizes filtration area (m2) and maximizes productivity (g/m2/h)). For high
concentrations (>10–15 g/L), it may be advantageous to dilute the product to improve
filterability.
The effect of filtration pressure is often best determined by conducting an
excursion study to evaluate filter capacity and flow as a function of pressure. It is
customary to evaluate pressure effects in the 10–50 psi range. In general, higher
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Fig. 4.24 General approach to optimization of virus filtration.
operating pressures increase the average process flux and decrease the required
filter area [40]. The magnitude of this impact is dependent upon several factors,
including feed solution condition, feed concentration, impurity profile, and virus
filter.
4.6.3
Capacity
Prefiltration of the feed solution can have a dramatic impact on filter performance.
Prefiltration removes various impurities or contaminants such as protein aggregates,
DNA and other trace materials. While larger size impurities can be removed by
prefiltering with a 0.2mm or 0.1mm rated microporous membrane, smaller
impurities such as protein aggregates that may only be marginally larger in size
compared to the protein product, are not easily amenable to size-based removal
methods. Prefiltration through adsorptive depth filtration has been observed to
provide significant protection for certain virus removal filters [52]. The impact of
prefiltration can be quite dramatic; with up to 10-fold reductions in required filter
area sometimes achievable. As these filters work by nonspecific multimode adsorption, product recovery should be confirmed to ensure good yield.
For some protein solutions, the freeze–thaw of a material can have a significant
impact on filtration performance. In fact, in some instances, it has been observed that
the required filter area is five- to sixfold higher when measured using material that
has been previous frozen compared to fresh feed [51]. While the actual purification
process may not have a freeze–thaw step, feed samples required for virus validation
testing are often conveniently submitted in a frozen form due to material stability/
availability considerations. In such situations, if freeze–thaw is observed to produce
an adverse impact of filtration performance, a prefilter is often used to restore
performance similar to the unfrozen material.
As discussed in Section 3.1.1.4, a gradual pore-plugging model is described by the
Vmax model. Filter sizing is impacted by the filter capacity, Vmax, the initial flow rate,
Qi, and the batch time, tb. For typical processing times less than 4 hours, a higher flux
membrane with a corresponding lower capacity often results in lower filtration area
compared to a high-capacity/low-flux filter. For processing times greater than
18 hours, the high-capacity/low-flux filter would result in a process with lower filter
area. It should be noted, however, that shorter processing times have the added
benefit of allowing for the possibility of in-line processing with other purification
steps as well as mitigating potential product stability issues.
4.6.4
Small-Scale Simulation
Once the optimum filtration conditions have been determined, it is recommended
that a simulation study be performed. This would initially be performed at the small
scale (3.5–14 cm2), then repeated at a larger scale as the process is scaled-up. This
would involve running the filtration to the desired end point, which may be a specific
filtration time, volume/area ratio, or percent flux decline.
One of the outcomes of successful process development is a process that is robust
and easy to implement when scaled-up to manufacturing. In order to demonstrate
scalability of the process, it is recommended that pilot-scale studies be conducted
using devices containing 100–1000 cm2 of filter area. This scale represents a 10–300fold scale-up from the initial simulation studies.
4.6.5
Pilot-Scale Studies
The objectives of the pilot-scale studies are twofold. A first objective is to obtain
confirmation that the process parameters (process loading, time, flux or pressure,
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and yield) are with predicted ranges and estimated bounds. Second, the pilot-scale
studies are used to obtain information on the entire operation (installation, flushing,
sterilization/sanitization, integrity testing, process, product recovery, etc.) so as to
enable drafting of SOPs and batch records for cGMP manufacturing. Information
obtained during the pilot-scale studies can also be used to establish appropriate
performance limits for water permeability (NWP), integrity testing, and other
secondary operations related to the virus filtration.
4.6.6
Virus Validation Studies
The purposes of the virus validation studies are to confirm the LRV claims for the
filtration step and to verify the filter sizing established in the scale-up phase.
These tests are run at a small scale, maintaining critical parameters such as
pressure, flux, and loading capacity at their commercial operation values while
mimicking other operating procedures such as pre-processing WFI flush-outs
and buffer equilibration. The tests are typically run concurrent with the manufacturing-scale consistency/validation batches. Due to the handling and assay
requirements for virus studies, the tests are typically conducted at specialized
labs.
The filter is challenged with representative feedstock containing a virus spike. The
concentration of the virus in the feed and the pooled permeate is measured to
calculate the LRV. Parallel control assays are run to correct for virus losses due to
artifacts such as dilution, concentration, filtration, and storage of samples before
titration. Typically, manufacturers will place a lower limit of LRV 3 on the (log)
reduction factors that will be combined to yield the overall reduction factor for the
manufacturing process.
To accurately represent manufacturing settings, the test feedstock must be
identical to the commercial-scale feedstock. Shipping or storage constraints may
require freezing feedstock, which can result in protein aggregates. Aggregates can
cause premature filter plugging that may alter scaling parameters such as loading
capacity. The problem can be obviated by either removing aggregates with microfiltration or generating fresh feedstock at the site of the virus spiking study.
The viruses used in the spiking studies depend on the specifics of the process
and virus contaminants. The regulations recognize ‘‘relevant’’ model viruses that
represent endogenous viruses, and nonspecific model viruses to validate ‘‘general
viral clearance’’ for adventitious contamination. Murine Leukemia Virus (MuLV) is
the generally accepted RVLP model for endogenous virus tests. If use of a relevant
virus is not possible, the manufacturer chooses the best specific model virus to
serve as a model for the relevant virus. To satisfy the ‘‘general viral clearance’’
objective, the study sponsor will generally evaluate two or three additional viruses.
The nonenveloped parvoviruses (20 nm) are often accepted as a worst case for
filtration. The test thus comprises a four- or five-virus panel that represents viruses
of different genomes (DNA and RNA), sizes and surface properties (enveloped and
nonenveloped).
Regulatory guidance states that ‘‘the amount of virus added to the starting
material for the production step that is to be studied should be as high as possible’’
[52]. However, so as not to unacceptably alter the product composition, the
volume of spike should be kept below 10 % and typically below 5 %. The
guidance also voices concerns over virus aggregation that could be induced by
deliberately concentrating the virus. The use of aggregated virus could lead to
underestimation of inactivation effectiveness and overestimation of size-exclusion
effectiveness [54]. The common practice is to use size-based prefiltration, such as
a 0.22 or 0.45 microporous membrane to remove virus aggregates from a spiked
feed stream prior to performing the clearance study. Asahi uses a Planova 35 or 70
as pre-filter.
Impurities contained within the virus spike may also foul the membrane, preventing the tests from reaching important scaling parameters such as loading
capacity. Methods of generating highly pure virus preparations are increasingly
being used to prevent fouling due to spike impurities [55,56]. Minimizing the
amount of virus spike used while still being able to reach the desired LRV target
is another approach.
Virus retention has been observed to decline with fouling for a variety of filters
[25,51,55,57,58]. If the virus spike required to achieve the target LRV causes excessive
fouling, alternative validation methods may be used to determine LRV at higher
throughput values [55]. These alternative methods may be used for validation after
consulting with the appropriate regulatory agencies.
4.6.7
Implementation
Once the virus clearance step has been optimized and virus validation studies
completed, an implementation strategy is required for robust process operation.
After determining the filter capacity (L/m2) required for a process during process
simulation/scale-up and virus validation studies, the filter area required for processing a given batch volume can be calculated. Various filter configurations are made
available by manufacturers to facilitate large-scale implementation.
Normal flow virus filters are operated either in constant pressure mode or in
constant flow mode. Typical factors to be considered during large-scale virus filtration
system design include the following:
Minimum and maximum batch volume.
Minimum and maximum flow rate; it is important to consider
flow rates during pre- and post-use water flush and for post-use
system cleaning/sanitization.
Maximum operating pressure and differential pressures across
the prefilter and the virus filter.
If in-line dilution is needed to maintain constant feed concentration, appropriate dilution and mixing hardware.
Minimum and maximum concentration and appropriate
instrumentation to span the range.
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Appropriate hardware and connections to enable the filters and
the system to be steamed, autoclaved or chemically sanitized.
Filter housing configuration – individual filters in parallel or a
multifilter housing.
System hold-up volume versus validated post-process buffer
rinse.
Pre- and post-use integrity tests.
In the case of normal flow virus clearance filters, pressure vessels are typically
employed for constant pressure operation. However, when very large process
volumes are involved, it may be easier to use a pump with a pressure feedback
loop to carry out constant pressure filtration.
A typical sequence of operations in a virus filtration process includes the following
steps, very similar to the protocol used for ultrafiltration process scale-down testing.
Filter installation and flushing – typically used to reduce
extractables. Follow manufacturer’s directions and validation
package.
Measurement of NWP – Confirm filter is within established
ranges.
Sterilization/sanitization – This step must integrate to the
downstream processing philosophy of the user. Some virus
filters are available pre-sterilized. Consider the Manufacturer’s
guidelines for autoclaving/SIP treatment of the virus filter.
Pre-use integrity testing.
Buffer pre-conditioning.
Processing and product recovery.
Post-production integrity testing.
To ensure that virus clearance is consistent with manufacturer’s claims and results
obtained during virus validation studies, filter integrity should be checked both preand post-use. To facilitate this, filter manufacturers have developed a variety of
destructive and nondestructive physical integrity tests that are related to virus
retention, which were discussed in Section 4.2.5. Ultimately, the objectives of
properly designed physical integrity testing are threefold:
to confirm that the virus removal filter is properly installed;
assurance that the filter is free from gross defects and damage;
confirmation that the filter removes viruses consistent
with both manufacturers’ specifications and end-user virus
validation studies.
Filter manufacturers should be able to provide evidence that integrity test methods
and acceptance criteria correlate to retention of viruses in the targeted size range
under standard conditions.
The complexity of integrity testing virus filters should not be overlooked when
selecting the virus removal filter for manufacturing. Key integrity test considerations
include performance, safety, logistics, validation, and regulatory support [32].
4.7 Membrane Adsorbers
Currently available integrity tests for virus removal filters can generally be classified
into three categories [50,59]. The first category is a particle challenge test, the second
type is a gas–liquid porosimetry test and the third type is a liquid–liquid porosimetry
test. A more detailed summary of the various tests along with troubleshooting
techniques can be found in the PDA TR41 [32].
While only nondestructive tests can be used pre-use, either type of test can be used
for post-use testing. Nondestructive tests are either gas–liquid or liquid–liquid
porosimetry tests. In general, a gas–liquid porosimetry test such as diffusion test
or pressure hold test is recommended to complement liquid–liquid porosimetry test
or particle challenge test to check for gross defects in the system.
Pre-use integrity testing can be performed either before or after sterilization/
sanitization. Post-sterilization integrity tests are particularly useful since they ensure
that the filters are not damaged during the sterilization process. However, in an
aseptic process, one must maintain system sterility during filter wetting and integrity
testing steps.
Prior to protein processing, a buffer flush is generally recommended in order to
displace WFI with the appropriate buffer. The buffer flush can be carried out using
the conditions that are employed during protein filtration (same DP, TMP or filtrate
flux). About 10 L of buffer per m2 of filter area is a reasonable volume of buffer.
After the buffer flush, the system is ready for protein processing. The protein
product should be processed using the process conditions and operating window
established during the scale-down optimization studies and virus validation studies.
In the case of normal flow filters, protein recovery may be enhanced with a buffer
rinse. The buffer rinse can be carried out using the conditions that are employed
during protein filtration (same DP or TMP or filtrate flux). Flush volume depends on
the upstream volume of the system and desired protein yield. About 10 L/m2 is a
reasonable flush volume for a well-engineered system.
4.7
Membrane Adsorbers
Historically, membrane devices have purified products in fluid streams by size-based
filtration. More recently, the use of membranes functionalized with specific ligands
has started to gain widespread use for the adsorptive purification of biotherapeutics.
While bead-based chromatography is widely employed and effective, membrane
chromatography has been heralded as a technology potentially suited for large-scale
applications due to its ability to integrate capture and purification steps for processing
large amounts of product in relatively short times [60].
In bead-based chromatography, most of the available surface area for adsorption is
internal to the bead. Consequently, the separation process is inherently slow since the
rate of mass transport is controlled by pore diffusion. To minimize this diffusional
resistance and concomitantly maximize dynamic binding capacity, small diameter
beads can be employed. However, the use of small diameter beads comes at the price
of increased column pressure drop. Consequently, the optimization of preparative
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chromatographic separations often involves a compromise between efficiency/
dynamic capacity (small beads favored) and column pressure drop (large beads
favored).
In contrast, membrane-based chromatographic systems (also called membrane
adsorbers), have the ligands attached directly to the convective membrane pores,
thereby eliminating the effects of internal pore diffusion on mass transport.
Additionally, the use of microporous membrane substrates with a tight membrane
pore size distribution coupled with effective flow distributors can minimize axial
dispersion and provide uniform utilization of all active sites [61,62]. Consequently,
mass transfer rates of membrane adsorber media may be an order of magnitude
greater than that of standard bead-based chromatography media [62], allowing for
both high efficiency and high-flux separations. Since single or even-stacked membranes are very thin compared to columns packed with bead-based media, reduced
pressure drops are found along the chromatographic bed, thus allowing increased
flow rates and productivities. The necessary binding capacity is reached by using
membranes of sufficient internal surface area, yielding device configurations of very
large diameter to height ratios (d/h) [60].
Properly designed membrane adsorbers have chromatographic efficiencies that
are 10–100 better than standard preparative bead-based resins. Consequently, to
achieve the same level of separation on a membrane adsorber, a bed height 10-fold
less can be utilized. Bed lengths of 1–5 mm are standard for membrane adsorbers,
compared to bed heights of 10–30 cm for bead-based systems. Due to the extreme
column aspect ratios required for large-volume membrane adsorbers, device design
is critical. To maintain the inherent performance advantages associated with membrane adsorbers, proper inlet and outlet distributors are required to efficiently and
effectively utilize the available membrane volume.
4.7.1
Membrane Chemistries
Membrane adsorbers are commercially available in a variety of chemistries ranging
from standard ion-exchange chemistries (strong and weak anion and cation exchangers) to hydrophobic interaction chemistries, reversed phase chemistries, and
affinity chemistries [61]. Membrane adsorber devices are traditionally available as
single sheets, stacked disks, radial flow systems, pleated devices, or hollow fibers [62].
Device sizes range from small-volume devices containing <1 mL for scouting
experiments and method development to devices containing several liters of media
for large-scale preparative separations.
Table 4.7 summarizes the currently commercially available membrane adsorber
products. Sartorius’s Sartobind family of single-use disposable membrane adsorbers comprises 15 layers of a nominal 3mm cellulosic membrane [23]. Small-scale
devices designed for method development employ a syringe filter format with a bed
volume of approximately 2.1 mL. Larger manufacturing scale devices employ a
radial flow cartridge with membrane volumes ranging from 7 mL up to 540 mL.
Sartorius also has a family of multiple-use membrane chromatography radial flow
Tab. 4.7
Commercially available membrane adsorbers.
Company
Membrane
Available
chemistries
Small-scale
device
Large-scale
device
Sartorius AG
Nominal 3 mm
(www.sartorius.com)
cellulosic
Pall Corporation
(www.pall.com)
Ion-exchange
15-layer stack 15-layer
affinity activated
(2.1 mL)
radial flow
specialized
(7–2100 mL)
Nominal 0.8 nm
Ion-exchange
16-layer stack 16-layer pleated
polyethersulfone
(0.35 mL)
capsule
(5–5000 mL)
units with device volumes ranging from several milliliter to 2100 mL. These devices
can also be run in series or parallel to achieve the necessary performance. Pall’s
Mustang family of disposable membrane adsorbers comprises 16 layers of a
nominal 0.8mm polyethersulfone membrane [22]. For method development, small
16-layer coins (0.35 mL bed volume) contained in a stainless-steel housing are
available. Manufacturing scale devices comprise a 16-layer pleated capsule with
membrane volumes ranging from 5 mL to 5000 mL. To process even larger
volumes of fluid, these devices can also be configured in series. Although Sartorius
AG and Pall Corporation are currently the only two companies with commercially
available membrane adsorbers, due to the relative infancy of membrane adsorber
technology, we can expect new competitors and new technologies to continuously
emerge.
4.7.2
Current Applications
Early work in the area of membrane chromatography was primarily focused on the
bind–elute purification of various proteins [62–66]. However, recent advances in
the theoretical and experimental understanding of the performance of membrane
chromatography has helped to focus the current applications of membrane
adsorbers to primarily two areas – flow-through applications aimed at removing
trace impurities and bind–elute purification of extremely large molecules essentially excluded from the internal pore structure of commercially available beadbased resins. Examples of flow-through applications include polishing applications
aimed at removing low levels of nucleic acid, host cell protein, endotoxin, virus, and
product aggregates and pre-capture applications aimed at removing impurities that
either foul the Protein A column or create product stability issues upon product
elution. Examples of large-molecule bind–elute purification include gene therapy
vectors such as viral vectors and plasmids, as well as extremely large protein
molecules.
4.7.2.1 Flow-Through Polishing
Due to the lower binding capacity of membrane adsorbers compared to bead-based
resins, universal adoption of membrane adsorber technology has been slow, even
147
148
though the high flux advantages provided by membrane adsorbers would lead to
higher productivity [67]. For trace-impurity removal applications, adsorptive
capacity is not a significant limitation. Rather, the adsorptive binding capacity
of bead-based columns used in this application are typically three to four orders of
magnitude larger than required since the columns are normally sized to achieve a
desired flow rate (i.e., the process is throughput limited). Since membrane-based
systems have a distinct flow rate advantage and sufficient capacity for binding trace
levels of impurities and contaminants, membrane adsorbers are ideally suited for
this application.
Flow-through anion-exchange membrane adsorbers have proven to be an extremely powerful polishing step designed to remove trace impurities from monoclonal
antibodies solutions. Under neutral pH conditions, many viruses, nucleic acids,
endotoxins, and host cell proteins are negatively charged and will bind strongly to an
anion-exchange membrane. In contrast, many monoclonal antibodies are typically
basic in nature (pI value >7), and, as such, will not interact with an anion-exchange
membrane adsorber.
This difference in charge has been successfully exploited to effect the separation of
various trace impurities from monoclonal antibodies [68–70]. Reported LRVs for
endotoxin (4 LRV), nucleic acids (6 LRV), and several mammalian viruses (2–6 LRV)
are consistent with those obtained with standard bead-based resins. To highlight the
performance advantages of membrane adsorbers, Phillips et al. has shown that for
efficiently designed membrane adsorbers, 6 LRV clearance of virus can be achieved at
residence times less than 0.4 seconds [69]. Residence times on the order of several
minutes are typically required to achieve this degree of separation with bead-based
chromatographic systems.
Zhou et al. [70] have conducted an economic analysis and concluded that singleuse disposable membrane adsorbers can be an economically viable alternative to
standard bead-based separations. Although the cost of single-use membrane adsorber media is typically higher than bead-based media that is reused for hundreds of
cycles, the use of membrane adsorbers often result in lower hardware costs,
significantly lower buffer usage, and savings in validation studies (column packing,
column reuse). Primarily due to savings in buffer, Zhou et al. conclude that
membrane adsorber technology is economically preferred.
4.7.2.2 Flow-Through Precapture
Protein A column chromatography is routinely used for the capture and purification
of monoclonal antibodies from cell culture harvest streams. Product elution from
Protein A columns is generally done at low-pH conditions where the possibility of
coeluted impurities precipitating is real. The consequences of this precipitation may
include clogging of downstream sterilizing-grade filters and fouling of the Protein A
column, both of which are detrimental to the robustness and process economics of
the downstream purification.
Adsorptive-based removal of these impurities prior to Protein A column chromatography has been shown to minimize the possibility of precipitation during product
elution. Shukla et al. [71] have exploited the adsorptive properties of depth filters and
anion-exchange resins prior to Protein A capture chromatography to minimize the
extent of precipitation. Lepore et al. [72] have shown that the use of anion-exchange
membrane adsorbers also work well for this application, with significantly higher
loadings compared to standard bead-based resins. Lepore et al. also conducted an
economic analysis indicating that the use of membrane adsorbers is economically
advantageous.
4.7.2.3 Large Molecule Bind–Elute Purification
For very large molecules, bead-based chromatography resins have been shown to
display very low dynamic capacities that decrease significantly with increasing linear
velocity [73]. The low dynamic binding capacities are mostly attributable to the
inaccessibility of the resin pores to the larger molecules. The dynamic capacity for
membrane adsorbers has been shown to be both significantly larger than highly
porous bead-based resins and essentially independent of flow velocity. These
qualities make membrane adsorbers ideally suited for the purification of very large
molecules.
The purification of plasmids and viral vectors for gene therapy applications are
uniquely suited to take advantage of the properties of membrane adsorbers. Pora [74]
describes the use of a Q membrane adsorber for the capture and purification of
plasmids. Using a 260 mL membrane adsorber, approximately 71 L of clarified lysate
containing 1.5 g pDNA was processed. The step yield was approximately 95 % and the
total cycle time was 24 minutes. The use of bead-based resins most likely would have
required significantly higher bed volumes and much higher cycle times to effect this
separation. Han et al. have successfully used various ion-exchange membranes for the
adsorption of Aedes aegypti densonucleosis virus (a mosquito specific parvovirus with
pI around 5.6) [75]). They conclude that membrane adsorbers may be ideally suited for
virus capture since nearly all of the ligands are available for interaction.
4.7.3
Future Trends
Membrane adsorbers will continue to find niche markets that are capable of
exploiting the inherent advantages of membrane chromatography compared to bead
chromatography – namely high efficiency due to minimal mass transfer effects (fast
separations), large external surface areas (high-binding capacities for very large
molecules currently excluded from the porous structure of bead-based resins), and
pre-packed disposable devices (ease-of-use considerations). Anion-exchange membrane adsorbers have been successfully employed for the efficient removal of several
impurities from monoclonal antibody streams. Other applications that require the
removal of trace level of impurities from a product stream can obviously exploit the
advantages of membrane chromatography. To effect these separations, however, may
require the development of novel membrane chemistries capable of achieving the
desired selectivity between the impurity and product molecule. It is quite likely that
new membrane adsorber chemistries will be developed as these new polishing
applications are identified.
149
150
Currently, one of the most visible trends in biotechnology manufacturing is
disposability. As entire downstream purification trains begin to migrate toward
complete disposability, membrane chromatography begins to play a more important
role. Although complete disposability may never be a reality on the manufacturing
scale, the flexibility and rapid change-out capabilities associated with disposable
manufacturing may be ideally suited for early phase manufacturing of clinical
product and contract manufacturing organizations. Additionally, disposable manufacturing may be a requirement in the area of personalized medicines where
minuscule batches are manufactured for the treatment of a single patient. In these
instances, membrane chromatography would play a critical role in the purification of
these products.
Finally, a focus on membrane adsorber device configuration could have an
important effect on the future of membrane chromatography. Historically, membrane adsorbers were incorporated into traditional filter configurations. Although
this was fine for polishing applications, the flow distribution and hold-up volumes
were not properly designed to make membrane adsorbers competitive in the bind–
elute purification markets. Improvements in these areas could significantly increase
the potential markets available for membrane chromatography.
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5
Membrane Applications in Red and White Biotechnology
Stephan Lütz, Nagaraj Rao
5.1
Introduction
In living systems, membranes and membrane processes play a crucial role in the
survival, growth, metabolism, defense, and reproduction processes. As our understanding of the role played by membranes gradually increases, the development of
new application areas becomes a logical consequence. Several of these newer areas
have been covered in other chapters of this book. In this chapter, we will focus on the
application of membranes in red and white biotechnology. ‘‘Red’’ biotechnology
refers to medical applications of biotechnology, starting from diagnostics and ending
with therapy. It also covers the biotechnological production of pharmaceuticals and
diagnostics. ‘‘White’’ biotechnology, by definition, is the application of nature’s toolset to industrial production.
Membranes are made up of natural materials (such as tissues) or synthetic
materials (such as certain polymers). They are permeable to certain substrates in
solution. The movement of molecules across a membrane is regulated in both
directions, giving the membrane its unique properties and wide applicability. In
Figure 5.1, commonly used filtration processes in biotechnology and the retention
properties of membranes are shown.
In reverse osmosis, particles, macromolecules, and low molecular mass compounds such as salts and sugars are separated from a solvent, usually water. The feed
solution often has high osmotic pressure, and this must be overcome by the
hydrostatic pressure applied as the driving force. Thus, microfiltration, ultrafiltration, nanofiltration, and reverse osmosis differ from each other in the size of the
particles being separated.
Interestingly, membrane processes are finding application in biotechnology at
practically every stage of production [1–3]. This includes sterile dosage of substrates
into the bioreactor, bubble-free aeration, and retention of biocatalysts using a
variety of techniques, as well as product concentration and recovery based on
different unit operations. These membrane processes are shown schematically in
Figure 5.2.
ISBN: 978-3-527-31480-5
156
5 Membrane Applications in Red and White Biotechnology
Fig. 5.1 Types of filtration and retention properties of membranes.
5.2
Types of Membrane Processes in Red and White Biotechnology
5.2.1
Bubble-Free Aeration
In the case of biotransformations where aeration is required for the supply of oxygen,
gas sparging is the preferred method because of the large mass-transfer rates and
operational simplicity. However, gas sparging can damage cultured animal cells,
since they are more sensitive to shear stress caused by vigorous mixing and gas
sparging. In such cases, bubble-free aeration is achieved by the use of silicon
membrane tubing, which allows the gas to diffuse into the medium without the
Fig. 5.2 Schematic representation of membrane processes
used in biotechnology.
5.2 Types of Membrane Processes in Red and White Biotechnology
formation of bubbles. Another advantage of such tubing is the fact that its permeability for carbon dioxide is greater than that for oxygen, because of which there is no
accumulation of carbon dioxide in the medium.
5.2.2
Filtration Processes
In the case of membrane processes using filtration techniques, the pore size of the
membranes plays a deciding factor in contributing to the efficiency of the process.
With the advent of membranes with fairly well-defined pore sizes, it is now possible to
carry out separation processes for the retention of biocatalysts, cofactors, salts, and
solvents.
The importance of biocatalyst retention on the economics of a biotransformation
can be judged from Figure 5.3. When the retention factor R is only 95 %, there is a
rapid decrease in the relative concentration of the biocatalyst over a certain number
of residence times. When the retention factor is greater than 99.99 %, there is no
significant loss in the biocatalyst concentration in the reactor, and a very large
number of residence times can be achieved in the continuous biotransformation
process.
5.2.3
Dialysis and Electrodialysis
In the case of dialysis, one or more solutes are transferred from one solution, called
the ‘‘feed,’’ to another solution, called the ‘‘dialysate,’’ through a membrane down
their concentration gradient. When pressure is employed besides the concentration
gradient for separation, the process is called pervaporation.
Fig. 5.3 Importance of high retention of biocatalyst on
membranes.
157
158
In electrodialysis, the separation of components of an ionic solution occurs in a cell
consisting of a series of anion- and cation-exchange membranes. These are arranged
in an alternate manner between an anode and a cathode to form individual electrodialysis cells. During the process of electrodialysis, there is an increase in the ion
concentration of one type in one type of compartment and is accompanied by a
simultaneous decrease in the concentration in the other type of compartment.
5.2.4
Adsorption of Microorganisms
In adsorption, material accumulates on the surface of a solid adsorbent having a
multiplicity of pores of different sizes. Adsorption is a common operation employed
for the purification of biological products. Various physical, chemical, or physicochemical forces may be involved in the adsorption phenomenon. Activated carbon
and ion-exchange resins may be used, depending on the nature of the material to be
adsorbed. In ion-exchange adsorption, the adsorbents have ionic groups with easily
dissociable counter-ions. In affinity adsorption, proteins may be absorbed biospecifically on the basis of their interaction with a complimentary tertiary structure.
Microorganisms and enzymes are also adsorbed onto inert surfaces in order to carry
out biotransformations.
5.3
Examples of Membrane Processes in Biotechnology
5.3.1
Bubble-Free Gassing
5.3.1.1 Hydrogen
Several naturally occurring enzymes use hydrogen as a substrate. For example, the
hydrogenase I obtained from the hyperthermophilic archaeon Pyrococcus furiosus
(PfH2ase), a microorganism found in the Volcano regions of Italy, can split hydrogen
heterolytically and catalyze regio- and enantioselective hydrogenation. The limiting
factors for this process are the solubility of hydrogen and the hydrogen transfer rate.
In nature, the solubility of a gas is increased by an enormous increase in the
membrane surface, as exemplified by the bronchi, bronchioles, and alveoli of the
lungs. Based on these principles, polymeric membranes are employed to introduce a
gas into a bioreactor. The polymeric membrane functions as a gas distributor and is
able to control the gas–liquid interface area and the mass-transfer coefficient
independently of each other [4]. The pressure can also be divided into two partial
pressures, one for the gas phase and one for the liquid phase.
For the regeneration of cofactors, sources of hydrogen that have been studied in
biotransformations include alcohols such as isopropanol, sugars such as glucose6-sulfate, formate, and molecular hydrogen. Continuous cofactor reduction using
biotransformations with molecular hydrogen as the source of the reducing agent is
5.3 Examples of Membrane Processes in Biotechnology
Fig. 5.4 Bubble-free H2-aeration in an enzyme membrane
reactor for continuous reduction of cofactor NADPþ.
possible when the biocatalyst accepts hydrogen as a substrate. The enzyme PfH2ase
has been used in batch and continuous experiments in order to reduce the cofactor
NADP+ to NADPH (Figure 5.4) [5,6].
Mertens et al. have described the potential applications of PfH2ase and similar
hydrogen-activating hydrogenases for the production of NADPH, biosensors, and
biofuel cells [7].
5.3.1.2 Oxygen
The bubble-free introduction of oxygen into a bioreactor is usually done with the help
of silicone tubings that pass through the bioreactor. Mammalian cells have been
immobilized on macroporous carriers made from small glass beads of up to 0.7 mm
diameter. These are then suspended and cultivated in a continuously operated
fluidized bed reactor. The introduction of oxygen occurs via a silicone tube that
passes through the reactor (Figure 5.5).
As a result, the entire length of the fluidized bed reactor is provided with oxygen,
irrespective of the axial position of the reactor. Depending on the oxygen transfer rate
and the oxygen consumption rate, the length of the tube is calculated. Efficient
mixing of the gas in the liquid helps in reducing the radial oxygen gradient.
Maximum cell densities of up to 3.3–50 107 cells/mL have been achieved using
this principle. Pharmaceutically important proteins, including antibodies and immunoglobulins, can be manufactured using this method [8,9].
The cultivation of CHO-2DS cell lines in a protein-free medium with bubble-free
aeration, accompanied by improved cell growth and better production of human
prothrombin, has been reported. [10]. Bubble-free aeration was found to be advantageous. BHK-21 cells have been fermented on a technical scale in protein-free culture
media in a membrane-airlift bioreactor [11]. The production of recombinant human
interleukin 2 was studied both with and without bubble-free aeration. By modifying
the SM1F7 culture medium by addition of the surfactant Pluronic F-68 (BASF), the
tolerance of the cells toward hydrodynamic stress was increased. The defoamer did
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Fig. 5.5 Bubble-free introduction of oxygen through a
silicone tube for cell culture.
not adversely affect the cultivation. The cultivation of the BHK-21 cells could be carried
out on a 1000-L scale without bubble-free aeration. This implies that a decision has to
be taken on a case-to-case basis if bubble-free aeration is a must or not. The human
tumor antigen and type I transmembrane protein Mucin MUC1 plays an important
role in the diagnosis of cancer cells. A continuous perfusion culture process by the
fermentation of CHO-K1 cells, using bubble-free oxygenation, has been developed [12]
. The optimum pO2 value was found to be at 40 % air saturation. The cell activity could
be maintained at a level above 85 %. A space–time yield (STY) of 100 mg/L/day could
be achieved for MUC1-IgG2a. The CHO-K1 cells could be transferred into the serumfree suspension culture with the help of the ProCHO4-CDM medium.
Studies on the control of bubble size, which has a direct bearing on the performance of mammalian cells, have been carried out by Nehring et al. [13]. Keeping the
bubble size small releases lesser energy when the bubbles are burst. Microbubbles
with 100–500 mm diameter can be generated in an aqueous media by using specially
designed hydrophilic materials such as porous ceramics. The ceramic gas bubbling
system consisted of a porous ceramic pipe filled with titanium dioxide and fixed on to
a steel structure. An early bubble-free reactor for the cultivation of mammalian cells
has been described by Schulze and Stahl [14]. This has been modified for use in the
cultivation of the phototropic microalgae Scenedesmus communis by Gorenflo et al.
Light energy was provided to the microorganisms by fitting the bioreactor with a new
system of irradiation of light. A cylinder made of stainless steel mesh was designed.
The pore size was twice the size of the cells (10 mm in diameter). As a result, the cells
were retained outside the cylinder and internal aeration was found to be optimal [15].
Bubble-free aeration often helps in stabilizing the biocatalyst. The extracellular
laccase of the white-rot fungus Pyconporus cinnabarinus (DSM 15225) was immobilized on DEAE-Sephadex after cultivation. The bubbling of oxygen through a silicone
tube in the enzyme membrane reactor used to carry out the biotransformation
stabilized the enzyme in continuously run experiments [16].
5.3.2
Membranes for Cell Retention
5.3.2.1 Higher Cells/Red Biotechnology
The fermentation of higher cells, such as hematopoietic stem cells, poses more
challenges than the cultivation of fungi and bacteria. This is because of the unique
requirements of the higher cells, including sensitivity to shear stress and the nature
of cell division.
In an interesting study exploiting the physical characteristics of cells and imitating
the naturally occurring hematopoietic microenvironment, Meissner et al. have
shown that it is possible to replicate bone marrow stromal cells by a three-dimensional cocultivation in porous microcarriers in a fixed-bed reactor [17]. A membraneseparated cocultivation method for the separation of hematopoietic and stromal cells
was employed, instead of a conventional coculture in microcarriers. During the
cocultivation, the two types of cells were separated with the help of a porous
membrane. The membrane pore size allowed the diffusion of the secreting stromal
growth factors as well as an intercellular contact through the membrane (Figure 5.6).
As a result, various degrees of expansion were observed for different colonyforming and burst-forming cells. For very early progenitor cells (CFU-GEMM) and
later progenitor cells (CFU-GM and BFU-E), expansion degrees of 4.2-fold, 7-fold,
and 1.8-fold were observed.
Using similar principles, cultivation of progenitor cells (colony-forming cells,
CFC, and cobblestone area forming cells, CAFC) was found to be more efficient. The
population of the CAFCs, for example, could be increased 39 times in the membrane
method described above, while only a sevenfold increase was observed in the
suspension cultures [18].
The production of IgG2a monoclonal antibodies by hybridoma cells has been
carried out using a perfusion culture system along with a ceramic membrane
module. In this so-called stirred ceramic membrane reactor system, consisting of
10 vertical, cylindrical tubes with an active surface area of 400 cm2, the volumetric
productivity increased sevenfold [19].
Fig. 5.6 Membrane-separated cocultivation of
hematopoietic and stromal cells.
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5.3.2.2 Whole-Cell Biotransformation
With the advent of simpler genetic engineering techniques, ‘‘tailor-made’’ microorganisms (so-called ‘‘designer bugs’’) are now being increasingly developed for
specific biotransformations. Ernst et al. have expressed a regio- and enantioselective
(R)-alcohol dehydrogenase from Lactobacillus brevis and a formate dehydrogenase
from Mycobacterium vaccae N10 into Escherichia coli. This functional overexpression
enabled the reduction of keto compounds into hydroxy compounds, for example, the
reduction of methyl acetoacetate to (R)-3-hydroxybutanoate [20]. The specific cell
productivity could be significantly improved by the introduction of a recombinant
cofactor regeneration system in the whole-cell biocatalyst. No byproduct formation
was observed, in addition.
Current trends indicate that such ‘‘designer bugs’’ will be increasingly used in
biotransformations.
The retention of microorganisms on membranes during biotransformation has
been utilized to manufacture chiral diol, (2R,5R)-hexanediol. In this process, resting
cells of Lactobacillus kefiri DSM 20587 were retained over an ultrafiltration membrane
with a molecular weight cut-off (MWCO) of 400 kDa. The reduction of the diol was
carried out at 30 8C under anaerobic conditions (Figure 5.7).
The selectivity for the diol was 78 % and the enantio- and diastereoselectivity were
found to be 99 %. A space–time yield of 64 g/L/day was achieved. Due to the high
selectivity of the process, downstream processing and product purification became
much simpler [21].
By using genetically modified E. coli coexpressing genes of LbDH and FDH from
M. vaccae, (R)-methyl-3-hydroxybutanoate with an enantiomeric excess of >99 % was
obtained, starting from methyl acetoacetate in this whole-cell biotransformation.
Reaction engineering enabled the lowest biocatalyst consumption in a continuous
reactor (0.9 gWCW/g). In comparative studies with enzyme-coupled systems, it was
concluded that reaction engineering techniques allow lower biocatalyst consumption
in whole-cell bioreductions. Whole-cell immobilization further reduced the biocatalyst consumption [22].
In an unusual application, whole cells of Pseudomonas putida type A1 were retained
in a hollow-fiber membrane reactor in order to degrade gas-phase toluene biocatalytically with a high oxygen availability. A porous, hydrophobic polyethylene
membrane functioned as a carrier for the active biofilm. Toluene and other
components diffused through the lumen of the membrane into the aqueous phase
and were biocatalytically oxidized in this phase containing nutrient and microorganism. Incoming toluene concentration between 86 and 97 % could be
degraded when the incoming air had a toluene concentration between 0.85 and
4.3 kg/m3/day [23].
5.3.3
Membranes for Enzyme Retention
The use of membranes for retention of enzymes in bioreactors forms a significant
part of red and white biotechnology. Since enzymes usually are large molecular
Fig. 5.7 Manufacture of (2R,5R)-hexanediol with cell retention.
weight compounds, with molecular weights ranging from 10 to 150 kDa or more,
membranes with defined pore sizes (or MWCO) find a wide range of applications for
enzyme retention during the biotechnological production of fine chemicals. A typical
example is the industrial production of amino acids by the acylase process using
enzyme membrane reactors by the German firm of Degussa. This process is in use
for over two decades now.
In addition to proteinogenic amino acids such as alanine, methionine (Figure 5.8),
valine, and tryptophan, nonproteinogenic amino acids such as O-benzylserine, norleucine, and nor-valine are manufactured on an industrial scale. Starting from the
racemic acetate mixture, only one enantiomer is hydrolyzed by the catalyst [24,25].
By establishing a direct phase contact in the pores of a hydrophobic, microporous
membrane, a hydrophobic product can be extracted from the stream coming out of a
bioreactor. Thus, an extraction step can be directly coupled to a biotransformation
step, thereby enabling the removal of poorly soluble products and unreacted educts
(Figure 5.9). Kruse et al. have used this method to couple an extraction step to an
enzyme membrane reactor with continuous cofactor regeneration [26].
A commercially available hydrophobic membrane (Celgard membrane in a LiquiCel lab module) functions as the extractor and extracts the hydrophobic product into
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Fig. 5.8 Racemic resolution of N-acetyl-methionine using acylase.
the organic phase. The regeneration of the hydrophilic cofactor takes place in the
aqueous phase unhindered. The residence time of the cofactor is decoupled from that
of the substrate. Several ketones have been reduced enzymatically, based on these
principles, to the corresponding, hydrophobic chiral alcohols. An NADþ-alcohol
dehydrogenase from Rhodococcus erythropolis was employed. There was an increase in
the total turnover number by a factor of 20–25.
In the above example, an extraction step has been used after a biotransformation
step to remove a hydrophobic product from the loop. In the case of substrates that are
poorly soluble in aqueous media, attempts have been made to introduce the so-called
Fig. 5.9 Coupling of a membrane-based extraction step to
a biotransformation process.
emulsion membrane reactors prior to the biotransformation step. By using a
hydrophilic ultrafiltration membrane, it has been shown that an emulsion of
2-octanone can be separated into its phases and the aqueous phase is saturated
with the organic substrate. This aqueous phase is fed into the enzyme membrane
reactor where the actual biotransformation takes place. The enantioselective reduction of 2-octanone to (S)-2-octanol is shown in Figure 5.10.
The biotransformation is catalyzed by a carbonyl reductase obtained from Candida
parapsilosis (CPCR) and is accompanied by continuous reduction of the cofactor
NADþ. By bringing the product stream in contact with the emulsion reactor, the
product is extracted. Production of (S)-octanol to the extent of 21.2 g/L/day and with
an enantiomeric excess of >99.5 % could be achieved over a period of four months
in a continuous experiment [27].
A variety of methods are now available to modify the carrier surfaces for immobilization of biocatalysts. Similarly, the surface of membranes can be chemically
modified in order to change its properties. Plasma-induced graft polymerization of
poly(a-allyl glycoside) has been used by Deng et al. to increase the biocompatibility and
hydrophilicity of a polypropylene hollow-fiber microfiltration membrane. A two-phase
membrane reactor using the lipase from Candida rugosa as the biocatalyst, immobilized on the membrane, was used for the hydrolysis of olive oil. Under optimized
reaction conditions, a volumetric reaction rate of 0.074 mm/L/h could be attained [28].
The importance of chiral amines as the building blocks to manufacture pharmaceuticals and agrochemicals is growing. Transaminases are being considered as
Fig. 5.10 Coupling of an emulsion membrane reactor with an enzyme
membrane reactor for biotransformation of poorly soluble substrates.
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potential biocatalysts for their synthesis. In one type of biotransformation, however,
the reaction is accompanied by the oxidation of one of the enantiomers to the
corresponding ketone, which causes strong product inhibition. To overcome this
problem, the enzyme membrane reactor has been coupled with a hollow-fiber
membrane contactor. The racemic substrates a-methyl benzylamine and 1-amino
tetraline yielded the products with enantiomeric purities of 98 and 95.5 %, respectively, when the (S)-specific v-transaminase from Vibria flusialis JS17 or Bacillus
thuringensis JS64 was used [29].
Similarly, hollow-fiber contactors have been used for the production of enantiomerically pure monoesters from meso diesters. A pig liver esterase was used, and the
polysulfone ultrafiltration hollow-fiber module had a MWCO value of 30 kDa. The
asymmetrical pores of the fibers served to immobilize the enzyme. The reaction and
the separation steps occur simultaneously in this module. The loss of enzyme activity
was found to be low [30].
Baccatin III is a precursor in the synthesis of the antitumor compound paclitaxel. It
has been continuously produced in an enzyme membrane reactor, starting from
10-deacetyl baccatin III by using an acyl transferase of plant origin, overexpressed in
E. coli. The substrate is obtained by methanolic extraction of the European yew
needles or cuts. The product was selectively separated from the biocatalyst in the
bioreactor and extracted in the solvent phase, made up of diisopropyl ether present in
an integrated module. The equilibrium of the reaction is shifted favorably by using
this method. A polysulfone membrane with a MWCO of 10 kDa was used. It was
observed that a reduction in the biotransformation temperature from 15 to 10 8C
improved the biocatalytic stability [31]. This example demonstrates once again the
importance of membrane processes in red biotechnology applications.
Miniaturization of the enzyme membrane reactor has led to the development of
micro enzyme membrane reactors with volumes of <200 mL. Such reactors function
as practical tools for education, research, and product development. The quantities of
biocatalyst and chemicals required for the experiments are significantly reduced [32].
5.3.4
Membranes for Cofactor Retention
In many cases, the catalytic activity of an enzyme comes into force only in the
presence of a cofactor, often called a coenzyme. While metallic cofactors are cheap to
obtain, cofactors such as NADþ and NADPþ are extremely expensive. For this reason,
cofactor regeneration plays a crucial role in the economics of a biotransformation.
Wichmann and Vasic-Racki have reviewed the advantages and disadvantages of the
different strategies of cofactor regeneration [33]. Closely related to cofactor regeneration is the retention of the cofactor in the biocatalytic reactor, so that it is not
subjected to the unit operations that the product stream has to undergo subsequently.
Cofactor retention by the use of membranes has been studied. Thus, for example,
in the continuous reduction of trimethyl pyruvate to L-tert-leucine, catalyzed by the
enzyme leucine dehydrogenase, the total turnover number (ttn) of the cofactor
NADH is dependent on the type of membrane employed.
Fig. 5.11 Dependence of total turnover number of cofactor
NADH on type of retention membrane.
When an ultrafiltration membrane was used in an enzyme membrane reactor, the
ttn was found to be 2325 (Figure 5.11). However, when a nanofiltration membrane
was used, the ttn increased to 7920.
Rosmarinicacid,anesterbetweencaffeicacidandachiraldihydroxyphenyllactic acid
(‘‘hydroxy-DOPA’’), has been investigated intensively for its anti-inflammatory properties. Rosmarinicacid has been producedby planttissue culture in significant quantities.
Biotransformation reactions also allow the synthesis of the naturally occurring product.
Thus, 3,4-dihydroxyphenyl pyruvic acid has been reduced to (S)-3,4-(dihydroxyphenyl)
lactic acid using a lactate dehydrogenase and a formate dehydrogenase, the latter for
continuous cofactor regeneration in a manner similar to the one described above.
The (S)-3,4-(dihydroxyphenyl) lactic acid has been converted to rosmarinic acid in a
second biotransformation using caffeoyl coenzyme A transferase [34,35].
The continuous regeneration of the cofactor NADPH during the Baeyer–Villiger
oxidation of 4-methyl cyclohexanone to 5-methyloxepan-2-one has been described by
Rissom et al. [36,37]. The product could also be synthesized in a repetitive batch mode
(Figure 5.12).
The chiral lactone had an enantiomeric purity of >99 % [38].
5.3.5
Application of Dialysis and Electrodialysis in Biotransformations
White biotechnology offers an interesting alternative to the synthesis of pyruvates.
Pyruvates are chemically synthesized at high reaction temperatures by oxidation
reactions using heavy metal catalysts, starting from tartaric acid, propylene glycol, or
lactic acid. In the case of biotransformations, a high extracellular pyruvate concentration greater than 500 mmol/L causes an inhibition of the microbial pyruvate
synthesis. Therefore, the product has to be separated, as soon as it is formed, in an
integrated process. This has been achieved by the integration of an electrodialysis
step by Zelic et al. [39]. The fermentation using the designer bug E. coli YYC2O2
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Fig. 5.12 Monooxygenase catalyzed Baeyer–Villiger oxidation.
Idh2A:Kan yields pyruvate in a concentration of more than 900 mmol/L under
optimized conditions.
Amino acids and peptides are interesting product groups for investigations using
electrodialysis, because of their charged nature at different pH values except the
isoelectric point. By combining electrodialysis with ultrafiltration, acid and base
bioactive peptides have been separated simultaneously by Poulin et al. [40]. Peptides
from a b-lactoglobulin (b-lg) hydrolysate obtained from whey were separated by using
an ultrafiltration membrane stacked in an electrodialysis cell. The process, called
electrodialysis with ultrafiltration (EDUF), showed a high selectivity. Starting from 40
peptides present in the raw hydrolysate, only 13 were recovered in the separated
adjacent solutions. In addition, among these 13 migrating peptides, three acidic–
anionic peptides migrated only into one compartment, while three basic–cationic
peptides migrated only into the other compartment, independent of the pH of the
hydrolysate. The ACE-inhibitory peptide b-Ig 142–148 showed the highest migration
value of 10.75 %. In addition to cationic and anionic membranes, a cellulose ester UF
membrane with a MWCO of 20 kDa was used in the above experiments.
5.3.6
Application of Pervaporation and Stripping in Biotransformations
Substrate-coupled cofactor regeneration (e.g., the oxidation of 2-propanol to acetone)
is interesting for NADPH-dependent enzymes as no efficient enzyme-coupled
system is yet available. Although higher conversions can be achieved by increasing
the 2-propanol concentration, this also leads to the formation of more coproduct
acetone. In such situations, it is essential to remove the coproduct as much as
possible, preferably in situ, in order to achieve higher conversions. During the
bioconversion of 5-oxo-hexanoic acid ethyl ester to the (S)-hydroxy ester, 2-propanol
was reduced to acetone and was removed by employing stripping or pervaporation.
The biotransformation was catalyzed by the NADPH-dependent enzyme carbonyl
reductase from C. parapsilosis (Figure 5.13).
Although both methods gave similar results with regard to product formation,
pervaporation offers several advantages over stripping on a technical scale with
regard to foaming and emission values [41].
The alcohol dehydrogenase from L. brevis (LbADH) is dependent on the cofactor
NADPH for its catalytic activity. A two-phase bioreactor has been used to carry out the
reduction of tert-butyl-6-chloro-3,5-dioxo-hexanoate (CDHE) to tert-butyl-(S)-4chloro-5-hydroxy-3-oxo-hexanoate (CHOE). During the biotransformation, a byproduct, tert-butyl-(4-oxo-4,5-dihydrofuran-2-yl)-acetate is formed due to the elimination
of HCl in aqueous media. In order to avoid this, the CDHE is dissolved in methyl-tertbutyl ether (MTBE) and the enzyme and 2-propanol are dissolved in the buffer. The
biphase emulsion reactor used gave the best results for the LbADH catalyzed
reduction of CDHE with respect to selectivity, substrate concentration, ttn, and
enzyme consumption, as well as downstream processing. Efficient separation of
acetone also improved the space–time yield [42].
Fig. 5.13 Pervaporation for acetone removal in a
biotransformation process.
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170
5.3.7
Nanofiltration and Ultrafiltration in Biotechnology
Several of the examples described in the previous sections employ either nanofiltration or ultrafiltration either during the biotransformation step or for subsequent
downstream processing.
The protective action of neuraminic acid-containing glycoconjugates in human milk
has been well documented. This has led to the synthesis of various molecules based on
neuraminic acid for possible therapeutic applications. The performance of a diafiltration membrane process for desalting of the permeate stream during the enzymatic
synthesis of CMP-Neu5Ac and GDP-Man has been studied. The nucleotide sugars
could be retained by a nanofiltration membrane in a cross-flow module. Only salts
were allowed to pass through. The purification steps yielded CMP-Neu5Ac in >95 %
purity and >90 % yield. GDP-Man gave values of 95 and 88 %, respectively [43,44].
In another well-established application of red biotechnology, the industrial production of recombinant tissue-type plasminogen activator (rt-PA) and other proteins,
starting from mammalian cells, employs tangential flow filtration using UF membranes for product isolation. Up to 5000 L/h of the medium can be processed with
protein yields generally >99 %. Cell vitality and cell density were maintained during
the process. Since the UF membrane had a pore size of only 0.2 mm, the filtrate could
be collected directly into containers after passing through a sterile filter. The latter
could also be steamed in place. Scaling-up of the process followed a linear pattern [45].
Recombinant HIV-1 transcription–transactivator protein (Tat-protein) from a
bacterial lysate could be separated easily by using a nylon membrane, modified
chemically with an acyl anhydride and subsequent covalent coupling of avidin and
biotin. In this affinity microfiltration technique, the product mTat primarily contained monomeric forms of the oligopeptidase sequence. While the membrane
process led to a fourfold improvement in protein recovery, nonspecific protein
adsorption was also observed [46].
5.3.8
Bioelectrochemical Applications
Electrochemical oxidation, ultrafiltration, extraction, and distillation have been
combined in a process to carry out the oxidative separation of racemic 1-phenyl1,2-ethane diol. The cofactor NAD+ was regenerated by anodic oxidation of NADH
using the mediator 2,20 -azinobis(3-ethyl-benzothiazolin-6-sulfonate (ABTS). An electrochemical, extractive enzyme membrane reactor was developed for this purpose [47].
In further studies, the cofactors NADþ and NADPþ have been electrochemically
reduced by an indirect method, using (pentamethylcyclopentadienyl-2,20 -bipyriddine
aqua)-rhodium(III) as the mediator. It was observed that for the reduction reaction,
adsorption of the mediator on the fixed-bed graphite cathode was essential. Stronger
adsorption led to lower mediator consumption. Using a coated titanium mesh as
the anode, values of ttn of up to 400 were achieved for the mediator. In order to hinder
the reoxidation of the product, a cation exchange membrane was placed between the
electrodes (Figure 5.14).
Fig. 5.14 Indirect electrochemical cofactor reduction.
Selectivity and conversion values achieved were practically quantitative. The
space–time yield was between 500 and 1000 g/L/day [48].
The first asymmetric electroenzymatic oxidation of thioanisole to (R)-methylphenyl sulfoxide with an enantiomeric purity greater than 98.5 % has been carried out by
Luetz et al. with in situ electrochemical generation of hydrogen peroxide [49]. The
hydrogen peroxide functioned as the oxidizing agent for the enzyme chloroperoxidase from Caldariomyces fumago. A dialysis tube membrane separates the counter
electrode (a platinum wire) from the working electrode (a carbon felt). The divided
cell (Figure 5.15) gave productivity values as high as 30 g/L/day when the experiments
were carried out on a 300 mL scale.
Fig. 5.15 Divided cell for electrochemical hydrogen
peroxide generation and biotransformation.
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In situ regeneration of cofactors such as NADH using electroenzymatic pathways
offer attractive, low-priced alternatives to the regeneration methods described in
earlier sections of this chapter. Considerable efforts are under way to achieve this. The
in situ electroenzymatic regeneration of NADH, during the formation of lactate from
pyruvate, has been studied in fixed-bed membrane reactors. The lipoamide dehydrogenase enzyme and the electron mediator methyl viologen were used. Porous
graphite electrodes were coated with a cation exchange membrane. Substrate,
product, enzyme lactate dehydrogenase, and coenzyme were present in the same
solution. Yield of lactate improved considerably as a result of the electroenzymatic
regeneration of NADH [50].
5.4
Summary
From the wide variety of applications and principles described above, it is evident that
membranes have a major role to play in industrial red and white biotechnology. This is
further fuelled by the fact that properties of membranes can be modified in a better
fashion now. In addition, a variety of polymeric materials are available for these
applications. Mechanical and chemical stability can be improved to suit individual
process requirements. The range of membrane-based applications in red and white
biotechnology will continue to show an upward trend.
5.5
Acknowledgment
The authors are grateful to Prof. Dr C. Wandrey for ongoing support and discussions.
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175
6
Membranes in Controlled Release
Nicholas A. Peppas, Kristy M. Wood, J. Brock Thomas
6.1
Introduction
Membranes have become increasingly important for use in separation processes
including microfiltration, ultrafiltration, gas permeation, pervaporation, dialysis, and
reverse osmosis. In the biomedical industry, synthetic polymer membranes are used
in a variety of applications including hemodialysis, blood oxygenators, and controlled
release of therapeutics [1,2]. Rose and Nelson [3] were responsible for the first
introduction of membranes for controlling drug release using osmotic pumps.
Folkman and Long [4] developed silicon rubber controlled release delivery systems
for anesthetics and cardiovascular treatments. With the creation of ALZA, membrane-based controlled release drug delivery systems (CR-DDS) became a commercial realization, creating both academic and industrial interest that energized the
field.
The OROS oral drug delivery technology was originally introduced to the market by
ALZA in 1973 as a gastrointestinal transport system. The OROS systems utilize
osmotic pressure [5] in combination with membranes that control water permeation,
drug release, or pushing force. OCUSERT was approved a year later for the control of
elevated intraocular pressure. This CR-DDS utilizes a thin multilayer membrane
device to control the release of the therapeutic pilocarpine at constant rate over a 7-day
period. ALZA further refined their use of membranes by developing transdermal
drug delivery system of which the first was approved by the FDA in 1981 for the
prevention of nausea and vomiting associated with motion sickness. These drug
delivery systems (DDS) employ a backing layer, a drug containing reservoir, a ratecontrolling membrane, and an adhesive layer that is capable of controlling rate of
release for periods of 1 day up to 1 week.
Membranes are utilized in CR-DDS to control the rate of permeation of a
therapeutic effectively allowing the rate of release of the drug to be finely tuned
according to the parameters associated with the polymer membrane [6]. The
permeation of the drug species through the membrane has been described by
two theories: solution-diffusion model and pore-flow model. The solution-diffusion
ISBN: 978-3-527-31480-5
176
6 Membranes in Controlled Release
model is the most commonly accepted theory to describe the phenomena of
CR-DDS employing polymer membranes. This model describes the drug molecules dissolution in the membrane and diffusion through the membrane due to a
concentration gradient. The polymer parameters responsible for controlling the
drug diffusion coefficient are degree of crystallinity and size of crystallites, degree
of crosslinking, degree of swelling, and molecular weight of the polymer. By
tailoring the properties of the membrane at the molecular lever, one can effectively
design materials that possess specific release kinetics [6,7]. The kinetics can be
designed so as to provide a sustained release of drug delivery in the desired
application of the membrane.
This review evaluates the release kinetics of the drug species through membranes
and describes how the various parameters associated with the polymer membrane
affect this release. The traditional types of membrane-controlled drug delivery
systems – diffusion-controlled, osmotically controlled, swelling-controlled, and
chemically controlled [8] – and current applications of membranes in CR-DDS will
also be discussed.
6.2
Controlled Release Kinetics
Dissolving or dispersing an active pharmaceutical ingredient (API), whether it is a
small molecule, peptide, protein, or other bioactive agent, in a membrane is a method
in which to control the release. Membranes can also act as rate-controlling layers
covering reservoirs of therapeutic agents. During the development of these CR-DDS, it
is necessary to develop and use mathematical models to describe the release kinetics.
Although the models are based upon diffusion equations, they are often referred to
kinetic models or expressions due to their time-dependent behavior of drug
release.
CR-DDS have been classified according to the physical mechanism that acts to
control the release of the therapeutic. Based on the mechanism of transport,
membrane-based DDS can be classified as diffusion-controlled, swelling-controlled,
osmotically controlled, and chemically controlled systems. The mathematical models
developed to describe the release kinetics [9] are used to (a) predict the rate of drug
release from and the transport of drug through the membrane and (b) describe the
mechanisms by which drug is transported. Mechanisms of the diffusion phenomena
in these CR-DDS are accurately elucidated with expressions that relate a mathematical model and the physical parameters that describe the membrane being used for
controlled release.
6.2.1
Diffusion in Membrane-Controlled Release
Two forms of Fick’s law of diffusion are often employed to describe drug release. The
Equations (1) and (2) are for one-dimensional diffusion where ci and ji are the
6.2 Controlled Release Kinetics
concentration and mass flux of drug i, respectively; x and t are position and time of
release; and Dip is the drug diffusion coefficient through the membrane:
ji ¼ Dip
dci
;
dx
ð1Þ
@ci
@ 2 ci
¼ Dip
:
@t
@x 2
ð2Þ
Certain assumptions are made to describe the release kinetics using Equations (1)
and (2), and these include (a) that the geometry of interest is in the form of a thin,
planar system, (b) that the diffusion coefficient is independent of drug concentration,
and (c) that ji is the drug flux with respect to the mass average velocity v.
Solutions to the Fickian diffusion equations are obtained when sufficient
knowledge of the initial and boundary conditions pertaining to the controlled
release environment are provided [9]. These solutions allow for the determination of concentration profiles from the normalized drug concentration, (c/co),
versus dimensionless position, (x/d), as a function of dimensionless Fourier
time, (Dipt/d2), where co and d are the initial concentration and slab
thickness, respectively. One can also evaluate the drug release rate, (dMt/Adt),
by differentiating the above equations with respect to position and then evaluating the derivative at the membrane–dissolution medium interface (Equation (3).
dMt
@ci
:
¼ Dip
Adt
@x x¼interface
ð3Þ
Finally, the total amount of drug release per cross-sectional area, (Mt/A), is given by
integrating the above expression over the experimental release time.
Mt
¼
A
ðt
dMt
dt ¼
0 Adt
ðt D
0
@ci
@x
:
ð4Þ
x¼interface
Solute transport through membranes has been extensively studied in recent
years. Much of the work, which characterized the polymer structure and the
diffusing solutes, has led to theories used to describe solute transport. In the area
of diffusion of drugs and proteins, am Ende et al. [10] studied the solute transport
through ionic hydrogels as a function of mesh size and environmental conditions
(i.e., pH and ionic strength) and determined that each factor plays a very important
role in solute transport. They also concluded that hydrogels may be tailor-made
for a release of a specific drug, protein, or peptide. Other studies showing the effect
of pH on drug transport from ionized hydrogels were done by Brannon-Peppas
and Peppas [11]. They found that pH-dependent hydrogels could be prepared to
exhibit zero-order or near zero-order release; this behavior could be attributed to
the effect of the pH on the relaxation, swelling, and release mechanism of the
hydrogel.
177
178
In biological systems, another important factor is polymer–solute interactions
when both the polymer and solute are ionized. Collins and Ramirez [12] studied these
interactions and determined that polymer–solute interactions decrease the transport
of the solute. Gudeman and Peppas [13] also studied these interactions using wellcharacterized interpenetrating networks of PVA and PAA by varying the content of
PAA (the ionic component) in the membrane.
Some of the early work in solute transport through hydrogels was done by Renkin
[14] who studied solute diffusion through porous cellulose membranes based on
Fick’s law:
dN
dc
¼ DA
:
dt
dx
ð5Þ
Here dN/dt is the solute diffusion rate, D is the solute diffusion coefficient in the
solvent, A is the apparent diffusion area, and dc/dx is the concentration gradient
across the membrane. Renkin [14] performed diffusion experiments and measured the rate of diffusion for a variety of solutes through inert, porous membranes. He compared his experimental results to predictions based on the theory
proposed by Pappenheimer and collaborators [15] and found that they were in
close agreement. He concluded that solute diffusion was a function of pore and
solute size.
Yasuda et al. [16,17] studied the relationship between salt rejection and water flux of
nonionic [16] and ionic [17] membranes using reverse osmosis. In the nonionic study
[16], they prepared a variety of membranes and monitored the water flux and salt
rejection until equilibrium was reached. They found that the size and solubility of
ions in the membrane were responsible for transport depletion, related water
permeation, v, and salt rejection, Rs, in the following manner:
Rs ¼
P2 RT
vþ
ðD p DšÞ
P1 y 1
1
:
ð6Þ
Here, P1 is water permeability, P2 is the salt permeability, y1 is the molar volume of
water, and ðD p Dsˇ Þ is the effective pressure. The difference in this study and the
ionic study is that ionic polymers were used in the study and that there is a different
relationship between Rs and K1. Equation (7) gives the relationship for ionic
membranes.
K1 ¼ A expfBRS g
ð7Þ
Here, A and B are constants, and the equation is independent of ionic charge and
nature or morphology of the membrane. Yasuda et al. [16] found that the principle
behind the salt rejection of ionic membranes is the difference in the transport
volumes for mobile co-ions and water. The repulsive forces between a fixed ion and a
mobile co-ion decreased the transport volume, creating a transport depletion of salt
flux relative to water transport.
6.2 Controlled Release Kinetics
Quinn and Anderson [18] studied hydrodynamic equations governing transport in
microporous systems (r 1 mm) accounting for Brownian motion and steric
restrictions. They showed that a one-dimension diffusion–convection analysis could
be used for such systems and developed a series of equations to account for the effect
of the pore wall on the solute–solvent drag.
Peppas and Reinhart [19] developed a theory based upon the free volume theory for
a three-component system (water, solute, and polymer). It predicted the dependence
of the solute diffusion coefficient on solute size, mesh size, degree of swelling, and
other structural characteristics of the hydrogels. Equation (8) is the Peppas–Reinhart
equation:
DSM
MC MC
k2 rs2
¼ k1
exp
:
DSW
Qm1
Mn MC
ð8Þ
Here, DSM and DSW are the diffusion coefficients of the solutes in the membrane and
water, respectively. The ratio of the two is known as the normalized diffusion
coefficient. Also MC is the molecular weight between crosslinks, k1 and k2 are
structural parameters of the polymer–water system, MC is the critical molecular
weight between crosslinks at which diffusion could not occur, Mn is the molecular
weight of the polymer before crosslinking, rs is the Stokes hydrodynamic radius of the
solute, and Qm is the degree of swelling of the membrane. A very important criterion
for this theory is that the diffusion occurs through highly swollen membranes. Using
well-characterized, amorphous PVA membranes [20], they were able to validate their
theory using experimental data.
Prausnitz and collaborators [21] used Monte Carlo simulations to develop a
modified size exclusion theory based on statistical distribution of chains in the
network. However, the theory does not consider ionic interactions or the effects of
side groups on the structure, but focuses on chains in a large region of space. The
intention of their theory is to provide a general understanding of partially ionized
polyelectrolytes where other theories may be built upon it.
The fields of membrane-based controlled release, which are three-dimensional
crosslinked polymer networks capable of imbibing significant amounts of water, have
been widely used in such applications because of their biocompatibility with the
human body and because many hydrogels exhibit characteristics similar to natural
tissue.
Selection of membranes used in such processes depends on the characteristics of
the polymer and the solute. Membranes have several important characteristics that
play an important role in solute diffusion. These include ionization of the
membrane, degree of swelling, and specific mesh or pore size. Membranes have
functional groups along the polymer chain that react to the external environment
(e.g., temperature, ionic strength, and pH of the swelling agent). The response of
the hydrogel may be a reaction or increase in the mesh size of the hydrogel. For this
work, the increase in mesh size is studied as a function of pH. The degree of
swelling is also a very important parameter because it describes the amount of
water that is contained within the hydrogel at equilibrium and is a function of the
179
180
network structure, crosslinking ratio, hydrophilicity, and ionization of the functional groups. It is calculated from swelling studies and can be used to determine
the molecular weight between crosslinks and the mesh size of the hydrogels. The
mesh or pore size is the space available for transport. Generally, the mesh size or
pores can lead to classification of hydrogels as nonporous, microporous, or
macroporous. Nonporous hydrogels have a mesh size between 20 and 100 A in
diameter. Microporous hydrogels have a pore size between 0.01 and 0.1 mm, where
transport occurs by a combination of diffusion and convection. Macroporous
hydrogels have pore size greater than 0.1 mm; transport may occur by convection.
These pore sizes determine the size of the solute permitted to diffuse through the
membrane.
The characteristics of the solute are as important as those of the membrane. The
size, shape, and ionization of the solute affect its diffusion through the membrane. In
the case of ionization, if the membrane and the solute are ionized, interactions may
occur that may hinder or assist in the diffusional process, depending on the charges
on the membrane and solute. If the charges are the same, the membrane repels the
charges of the solute and does not hinder, and in some case assist, transport. If
the charges are opposite each other, interactions between the membrane and the
solute may take place, hindering transport.
6.2.2
Physical Parameters of Controlling Release
Membranes are prepared from uncrosslinked or crosslinked, hydrophobic or hydrophilic polymers that can be moderately or completely swollen in water or in buffered
or physiological solution. These networks may be produced by a chemical reaction of
monomers or by physical entanglements of chains. They may contain functional
groups such as amino, hydroxyl, or carboxyl groups along the polymer chain.
Incorporation of such functional groups increases the hydrophilicity of the membrane leading to a hydrogel. Drug release from the polymer membranes can be
controlled and tailored by adjusting several different physical parameters of the
polymer including, but not limited to, molecular weight, crystallinity, crosslinking,
porosity, swelling, and branching.
6.3
Membranes and Solute Transport
6.3.1
Characterization of Membranes
In the presence of water, the functional groups become ionized and cause the
hydrogel to swell due to charge repulsion. This swelling process may be described in
terms of the equilibrium swelling ratio, Q, which increases as the swelling of the
network increases.
6.3 Membranes and Solute Transport
Fig. 6.1 Equilibrium degree of swelling of ionic hydrogels
as a function of solvent pH.
The swelling behavior of the hydrogel is affected by the pH of the swelling solution
as depicted in Figure 6.1. Upon ionization, the charges repel each other causing the
network to swell. In ionic hydrogels, swelling occurs when there is an increase in pH
from a value below to a value above the pKa of the network. For cationic hydrogels,
swelling occurs when there is a decrease in pH from a value above the pKa to a value
below the pKa of the network (Figure 6.2).
In addition to pH, the concentration of the functional groups on the network and
the ionic strength of the swelling agent affect the swelling of hydrogels. By increasing
the concentration of the functional groups, the swelling of the network increases.
Increasing the ionic strength of the swelling agent decreases swelling in the network.
As the ionic strength increases, the ions of the swelling agent counterbalance the
mutual repulsion of the functional groups in the membranes. This behavior was
observed by Khare and Peppas [22] in studying poly(2-hydroxyethyl methacrylate-comethacrylic acid) and poly(2-hydroxyethyl methacrylate-co-acrylic acid) using
dynamic and equilibrium swelling studies. Chou and associates [23] and Kuo and
associates [24] also observed this behavior in poly(2-hydroxyethyl methacrylate-comethacrylic acid), N,N-dimethyl aminoethyl methacrylate, and poly(hydroxyl ethyl
methacrylate). Through swelling studies similar to the ones used by Khare and
Peppas [22] and Chou et al. [23], Hariharan and Peppas [25] were able to develop a
Fig. 6.2 The effect of changing the pH of the swelling agent
on ionic hydrogel membranes.
181
182
model showing the effects of pH, ionic strength of the solution, and concentration of
ionizable groups on polymer.
The ionic nature of hydrophilic networks makes them ideal for uses in separation
and controlled release applications. In separation processes, the external environment of the hydrogel can be adjusted to produce the desired solute diffusion. In
controlled release systems, drugs can be delivered to a specific site in the body where
the pH is at a certain value or range of values (e.g., in the gastrointestinal tract). The
hydrogel would act as a carrier and release the drug at a desired rate once it has reach
the site of release. The ionic nature, as well as other structural parameters, of the
hydrogel has been studied to prepare hydrogels that can be tailor-made for a specific
application.
6.3.2
Solute Transport in Network Membranes
6.3.2.1 Structural Parameters of Membranes
Swelling studies can be used to determine the structure of membranes for controlled
release. The membranes are prepared and their polymer volume fraction in the
relaxed state is calculated using Equation (9). After each membrane has swollen to
equilibrium, the polymer volume fraction of the swollen polymer is calculated using
Equation (10).
y2;r ¼
Vd
;
Vr
ð9Þ
y2;s ¼
Vd
:
Vs
ð10Þ
Here, Vd, Vr, and Vs are the volumes of the polymer sample in the dry, relaxed, and
swollen states, respectively, and y2,r and y2,s are the polymer volume fractions of the
relaxed and swollen polymer gel, respectively. The volumes are calculated using
Equations (11) through (12), which utilize the weights of the dry polymer, Wd, the
relaxed polymer, Wr, and the swollen polymer, Ws, in air and heptane.
Vd ¼
Wa;d Wh;d
;
rh
ð11Þ
Vr ¼
Wa;r Wh;r
;
rh
ð12Þ
Vs ¼
Wa;s Wh;s
:
rh
ð13Þ
6.3 Membranes and Solute Transport
Here rh is the density of heptane. The densities of the swollen and dry polymer can
also be calculated using Equations (14) and (15).
rswollen ¼
rdry ¼
Wa;s
;
Vs
Wa;d
:
Vd
ð14Þ
ð15Þ
The equilibrium swelling ratio of a membrane can be affected by the ionic strength,
temperature, and pH of the swelling agent. Data taken from swelling studies can be
used to calculate the equilibrium swelling ratio:
Q¼
1
:
y2;s
ð16Þ
6.3.2.2 Determination of Molecular Pore Sizes
There are several methods to accurately determine the pore size of membranes.
While mercury porosimetry and BETmethods continue to be the two most important
methods to analyze macroporous membranes, molecular analysis can be used for
nonporous membranes. For example, in such membranes one can first determine
the average molecular weights between crosslinks, MC , in the case of crosslinked
membranes or the average molecular weights between entanglements in uncrosslinked membranes.
The molecular weights between crosslinks, MC , are calculated from the swelling
data using Equation (17) as discussed by Peppas and Merrill [26].
ðy=v1 Þ½lnð1 y2;s Þ þ y2;s þ xy22;s 1
2
:
¼
Mc;e Mn y2;r ½ðy2;s =y2;r Þ1=3 ð1=2Þðy2;s =y2;r Þ
ð17Þ
Here, Mn is the number-average molecular weight of the polymer before crosslinking, y is the specific volume of the polymer, V1 is the molar volume of the solvent,
y2,r is the volume fraction of the polymer in the relaxed state, y2,s is the volume
fraction of polymer in the swollen state, and x is the interaction parameter of the
polymer–solvent system in water.
A theoretical MC value can be calculated from knowledge of the nominal crosslinking ratio X as follows:
MC;t ¼
Mr
:
2X
Here, Mr is the molecular weight of the repeating unit of the polymer.
ð18Þ
183
184
Brannon-Peppas [27] developed equations to calculateMC for cationic and anionic
membranes. Starting from the Peppas–Merrill equation, she accounted for ionization of the polymers.
2 h i h
i
V1
Ka
y2;s 2
¼ lnð1 y2;s Þ þ y2;s þ x1 y22;s
pH
4I 10
þ Ka
y
"
#
V1
2MC
y2;s 1=3 1 y2;s
þ
1
y2;r
:
yMC
Mn
y2;r
2 y2;r
ð19Þ
Here, I is the ionic strength, pH is the pH of the buffer solution in which the
membrane was swollen, and Ka is the dissociation constant.
The membrane mesh size, j, defines the linear distance between consecutive
crosslinks. Indirectly, it indicates the diffusional space available for solute transport
and can be calculated using Equation (20).
1=3
j ¼ y2;s
1=2
2MC
Cn ð
Þ
1:
Mr
ð20Þ
Here, Cn is the Flory characteristic ratio, and l is the carbon–carbon bond length.
The crosslinking density of the membranes is calculated using Equation (21).
rx ¼
1
yMC
ð21Þ
6.4
Applications in Drug Delivery
Hydrophilic membranes are predominantly hydrogel membranes that swell in water
or biological fluids without dissolving. The swelling characteristics are the result of
crosslinks (tie-points or junctions), permanent entanglements, ionic interactions, or
microcrystalline regions incorporating various chains. In the last twenty-five years,
hydrogel membranes have been researched as prime materials for pharmaceutical
applications, predominantly as carriers for delivery of drugs, peptides, or proteins.
They have been used to regulate drug release in reservoir-type, controlled release
systems or swellable systems [28,29].
Hydrogel membranes are characterized as neutral, anionic, or cationic networks.
Their swelling behavior is governed by a delicate balance between the thermodynamic polymer–water Gibbs free energy of mixing and the Gibbs free energy
associated with the elastic nature of the polymer network. In ionic hydrogels, the
swelling is governed not only by the thermodynamic mixing and elastic–retractive
forces but also by the ionic interactions between charged polymer chains and free
ions. The over swelling behavior and the associated drug release kinetics are affected
by the osmotic force that develops as the charged groups on the polymer chains are
neutralized by mobile counterions. Electrostatic repulsion is also produced between
6.4 Applications in Drug Delivery
fixed charges and mobile ions inside the gel, affecting the over swelling of the ionic
gel. The equilibrium swelling ratios of ionic hydrogels are often an order of
magnitude higher than those of neutral gels because of the presence of intermolecular interactions including coulombic, hydrogen-bonding, and polar forces. This
means that drug or protein transport in ionic gels may be significantly faster than in
neutral gels [30].
Drug, peptide, or protein release through equilibrium-swollen membranes has
been analyzed using classical Fickian diffusion theories. Drug diffusion is expressed
in terms of hydrodynamic theories that consider the frictional characteristics of
spherical solute drugs as they diffuse through the mesh of a gel. Derivation of drug
diffusion equations through the mesh is based on the Fickian equation of drug flux
with additional terms for convection, wall partitioning, and restrictions due to the
tortuosity of the diffusional drug path. Various theories for drug transport in gels
have been based on the Eyring theory of rate processes; they utilize a free volume
approach to describe the probability that a drug molecule will pass across the available
mesh area. A successful model to describe drug transport in neutral hydrogels was
derived by Peppas and Reinhart [19]. This model relates the normalized diffusion
coefficient to the drug size, the equilibrium degree of swelling, and the gel mesh size.
Swollen membranes include a wide range of new materials that have been
considered for such applications. For example, chitosan is an extremely promising
carrier for drug delivery and has been studied either in the pure form or as a
copolymer with other important polymers, for example, poly(ethylene glycol) (PEG).
For example, Saito et al. [31] studied graft copolymers of the above for drug delivery
applications. PEG continues to be an important carrier for drug delivery mostly in the
form of homopolymeric or copolymeric hydrogels. Remuñán-López and Bodmeier
[32] studied the process conditions for the formation of chitosan-based carriers and
examined various applications of such systems either in colonic delivery or as
mucoadhesives. Again, blends or copolymers with PEG or larger molecular weight
poly(ethylene oxide) (PEO) were found to exhibit improved release properties.
Recently there has been increased research in the preparation and characterization
of membranes responsive to changing environmental conditions. Some of the
environmental conditions that can affect hydrogel swelling include pH, ionic
strength, temperature, and drug concentration.
Many properties contribute to the swelling of ionic hydrogel membranes. An
increase in the ionic content of the gel increases the hydrophilicity leading to faster
swelling and a higher equilibrium degree of swelling. Gutowska et al. [33] prepared
and tested a series of extremely versatile pH-sensitive gels that can be used for a wide
range of pharmaceutical applications. Such hydrogels were shown to be the prime
candidates for release of proteins, among other applications, including insulin and
calcitonin. For example, environmentally sensitive hydrogels have been studied as
possible controlled insulin release.
Recent advances in the development of ionic membranes have concentrated on
several aspects of their synthesis, characterization, and behavior. Major questions
that have been addressed in recent work include (i) synthetic methods of preparation
of hydrophilic polymers with desirable functional groups, (ii) synthetic methods of
185
186
preparation of multifunctional or multiarm structures including branched or grafted
copolymers and star polymers, (iii) understanding of the criticality and the swelling–
syneresis behavior of novel anionic or cationic polymers, (iv) synthesis and characterization of biomimetic hydrogels, (v) understanding of the relaxational behavior
during dynamic swelling, and (vi) modeling of any associated drug dissolution [7].
Hydrogel membranes may undergo volume-phase transition with a change in the
temperature of the environmental conditions. The reversible volume change at the
transition depends on the degree of ionization and the components of the polymer
chains. This type of behavior is related to polymer phase separation as the temperature is raised to a critical value known as the lower critical miscibility temperature or
lower critical solution temperature (LCST). Networks showing lower critical miscibility temperature tend to shrink or collapse as the temperature is increased above
the LCST.
Recently, in our laboratory, we composed hydrogel membranes of lightly crosslinked N-isopropyl acrylamide (NIPAAm) and methacrylic acid (MAA), which were
synthesized and characterized for their sensitivity to external conditions and their
ability to control release of two antithrombotic agents, heparin and streptokinase.
PNIPAAm is noted for its sharp change in swelling behavior across the lower critical
solubility temperatures of the polymer, while PMAA shows pH-sensitive swelling
due to ionization of the pendant carboxylic groups in the polymer. Hydrogel
copolymers of NIPAAm and MAA with appropriate composition were designed
to sense small changes in blood stream pH and temperature to deliver antithrombotic
agents, such as streptokinase or heparin, to the site of a blood clot. Experiments were
performed to show that hydrogels with certain compositions could show both
temperature- and pH-sensitivity.
Numerous other temperature-sensitive membranes have been used for drug
delivery including responsive, microporous hydroxypropyl cellulose gels modified
so as to exhibit LCST close to room temperature and copolymers based on NIPAAm,
butyl methacrylate, and acrylic acid (AA) with distinct LCST points [34,35].
Star polymers have been synthesized and characterized in the past twenty years.
Star and dendritic polymers may be used in a variety of pharmaceutical applications.
They can act as drug delivery membranes. These star chains could carry enzymes on
the short arms, while long arms with a nonreactive functional group could provide
protection for these enzymes. An alternative possibility would be to use star polymers
with long arms carrying a cell recognition moiety to adhere to a specific site with
shorter arms carrying a cytotoxic compound. In both these examples, the longer and
shorter arms carry two different reactable groups to permit coupling of different
molecules to the arms [28].
The main components of a star polymer are the core or foundation site from which
the diverging branches of the star structure start. One or more branching arms
emanate from this core site, each one incorporating a further branch point. Finally, a
terminal functionality is observed for each of the branches, usually having a reactivity
that allows it to further react in the dendritic structure.
Star-shaped polymers of PEO have been prepared by anionic polymerization on a
‘‘core’’ of crosslinked divinyl benzene and are presently studied for drug delivery
6.4 Applications in Drug Delivery
applications, particularly as hydrogels. These polymeric systems are particularly
promising because they can serve as micro- or nanoparticulate carriers for drug
delivery system development. The use of functionalities at the end of the star arms
allows immobilization of a wide range of biologically active agents. Such agents can
be antibodies, antigens, and polysaccharides forming interesting conjugates for
medical applications. Reaction of PEG at the end of the arms can lead to improved
‘‘stealth’’ particles that can be used in the body avoiding the reticulo-endothelial
system. Various enzymes can be immobilized. Thus, heparinase can be immobilized
to give nanoparticle star polymers, which can be used for postoperative blood
purification and removal of heparin. Fibrinoyltic enzymes such as streptokinase
and urokinase can be immobilized and used for lysis of thrombi in the blood. Urease
can be immobilized to produce microparticles, which will be useful in blood
purification by hemoperfusion. The advantages of such enzyme-immobilized star
polymer particles are enhanced stability, ease in separation and reuse, ability to
prepare enzyme free products, retention of a controlled enzyme microenvironment,
low or no immunogenic response, and low cost–high purity products.
A wide range of hydrophobic membranes has been studied in drug delivery
applications. They include the widely used ethylene vinyl acetate copolymer series
that has been used in many successful products, such as the early transdermal and
ocular therapy systems. Other hydrophobic polymers include polyethylene, polypropylene, various types of hydrophobic cellulose derivatives such as ethyl cellulose,
among others.
Recent work in the field is concerned with the formulation and evaluation of
delivery devices based on highly crosslinked acrylates. These materials, prepared in
loaded form by an extremely rapid, solvent-free photopolymerization process, are
multiacrylate-acrylic acid copolymers having nominal crosslinking ratios ranging
from 60 to 100 %. Release rates are therefore substantially lower than those exhibited
by typical hydrogel formulations. By manipulation of the structure of the PEGcontaining crosslinking monomer and of the AA feed composition ratio in the comonomer–solute mixture, systems possessing a wide variety of delivery characteristics are attainable. Additionally, the presence of ionogenic AA moieties along the
polymer backbone allows for environmental sensitivity and complexation-mediated
release under appropriate conditions.
Biodegradable membranes, by definition, change their chemical and potentially
physical form upon contact with the biological environment. There are two distinct
stages to the biodegradation process, especially in bulk degradation. The first stage is
restricted to the random cleavage of molecular linkages. The resulting decrease in
molecular weight produces some change in mechanical properties and morphology,
but no weight loss [36]. The second stage involves a measurable weight loss in
addition to chain cleavage. It begins when the molecular weight of the polymer has
decreased to the point that chain scission produces oligomers that are small enough
to solubilize and diffuse out from the network. These oligomers are necessarily
released into the adjacent tissue and therefore should be biocompatible.
The predominant means by which polymers degrade is by hydrolytic degradation,
enzymatic degradation, or a combination of the two. The degree to which each type of
187
188
degradation contributes to the degradation of a specific polymer may vary during the
degradation process and is dependent upon many factors, especially the environment
in which the degradation is taking place. The degradation is on a biological time scale
of hours to months, not an environmental time scale of years to decades. The
products produced from this degradation will be resorbed by the body or completely
eliminated by normal biological pathways.
Polymers that degrade by hydrolysis include poly(lactic acid), poly(glycolic acid),
poly(caprolactone), polyanhydride, poly(ortho ester), and polycyanoacrylate. Many
poly(amino acids) such as poly(L-lysine), poly(L-arginine), poly(L-aspartic acid),
poly(L-glutamic acid), and poly[N-(2-hydroxyethyl)-L-glutamine] are enzymatically
degradable [37].
After degradation has occurred, it is essential that the products of the degradation
be resorbed or eliminated from the body. For formulations that have been administered orally, the elimination of the polymers follow natural processes, but biodegradable polymers used parenterally or in other invasive techniques require more analysis
and design to ensure fully biodegradable formulations. In order for this elimination
to occur, the degradation byproducts, often small molecular weight polymer chains,
must first be solubilized before they can be eliminated. For the more water-soluble
polymer fragments, this is a fairly straightforward process. Other materials must first
be ionized before solubilization can occur.
There are three main classifications of degradable linkages that apply to biodegradble polymers as a whole. The category depends upon the location of the
biodegradable linkage: (i) the polymer backbone, (ii) at crosslinks, and (iii) at pendant
chains. Biodegradable formulations may include one or more of the mechanisms
mentioned. Furthermore, the biodegradation means described earlier (i.e., hydrolysis or enzymatic degradation) can be used in any of the types. Degradation of the
backbone and/or crosslinking agent usually is designed to lead to solubilization of
the material. Release of a pendant group, however, may not lead to solubilization of
the polymer matrix as the main, and potentially crosslinked, chain network will still
be in place (Mathiowitz et al., 1999).
From the previous analysis it is clear that membranes continue to have a prominent
position in the field of controlled release.
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5 Langer, R. and Peppas, N. A. (2003)
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Peppas, N. A. and Meadows, D. L.
(1983) Macromolecular structure and
solute diffusion in membranes: an
overview of recent theories. Journal of
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The role of crosslinks, entanglements
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Narasimhan, B., Mallapragada, S. K.
and Peppas, N. A. (1999) Release
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Siepmann, J. and Peppas, N. A. (2001)
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Peppas, N. A. (1995) Factors
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Brannon-Peppas, L. and Peppas, N. A.
(1988) Structural analysis of charged
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Collins, M. C. and Ramirez, W. F.
(1979) Mass transport through
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Gudeman, L. F. and Peppas, N. A.
(1995) pH-Sensitive membranes from
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Renkin, E. M. (1954) Filtration,
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Pappenheimer, J. R. (1953) Passage of
molecules through capillary walls.
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Yasuda, H. and Lamaze, C. E. (1971)
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Schindler, A. (1971) Salt rejection by
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Anderson, J. L. and Quinn, J. A.
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Peppas, N. A. and Reinhart, C. T.
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7
Drug Delivery Through Skin: Overcoming the Ultimate
Biological Membrane
7.1
Introduction
The prolonged constant drug level in the body is the ultimate goal of every drug
administration system. In fact, the constant drug level in the blood and the bypassing
of the hepatic ‘‘first-pass’’ metabolism are challenges for every delivery system.
Figure 7.1 presents the time variation of the drug concentration in the blood. The
traditional medical forms (tablets, injection solutions, etc.) provide delivery with
peaks, often above the required dose that may even be toxic for the patient. The
intravenous infusion can achieve constant drug delivery but requires hospitalization
and supervision of the patients. As a result the cost of the treatment is high and the
patient compliance is low.
Historically, the skin has been viewed as an impermeable barrier that protects the
body from external factors. Nevertheless, the introduction of drugs through it via
creams and gels has been used for a long time. In the past years, the skin has been
considered as a port for topical or continuous systematic administration of drugs. In
fact, for drugs with short half-lives, the transdermal drug delivery (TDD) provides a
continuous administration, rather similar to that provided by an intravenous infusion.
However, unlike the latter, the TDD is noninvasive and no hospitalization is required.
The drug level in the blood stays within the required acceptable limits (Figure 7.1).
The TDD via a patch (Figure 7.2) can provide continuous drug release through
intact skin into the blood stream for a long time and the delivery can be simply
stopped by removing the patch. The contact with the skin is usually achieved via an
artificial membrane. As we will see later the membrane can have or not have a
regulatory role in the drug delivery.
In this chapter, we will first give a short description of the skin structure, which is
one of the most important biological membranes, and of the fundamentals of skin
permeation. Skin is a great protective barrier. However, some innovative methods
have been already used to enhance TDD. The most promising ones (and rather close
to broad commercialization) will be discussed more in detail.
ISBN: 978-3-527-31480-5
192
7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane
Fig. 7.1 Variation of the drug concentration in the blood
during drug delivery.
7.2
Human Skin – Fundamentals of Skin Permeation
7.2.1
Human Skin Structure
The skin is the largest organ of the human body. It covers approximately 2 m2 of
surface area and receives about one third of all the blood circulating through the body.
Besides the protection against entry of foreign agents, it also plays a role in the blood
regulation and acts as a body thermostat. It is not our intention to present the skin
structure and properties in great detail in this chapter. The reader can find such
information in other excellent dedicated literature [1–4]. We will however focus our
attention on those properties important to TDD.
The skin is a rather complicated multilayer organ. It can be basically described in
terms of two basic tissue layers: the dermis and the epidermis (see Figure 7.3).
Fig. 7.2 Schematic illustration of a transdermal drug
delivery patch in contact with the skin.
7.2 Human Skin – Fundamentals of Skin Permeation
Fig. 7.3 A cross section of the human skin.
(Source: Reprinted from Ref. [1], with permission from
Elsevier, 2001.)
The dermis (thickness of 100–200 mm) forms the bulk of the skin and consists of
connective tissue elements. It provides physiological support for the epidermis via
blood and lymphatic vessels as well as nerve endings [5]. The epidermis, the top layer of
the skin (thickness of 100–110 mm), is composed of epithelial cells held together by
highly convoluted interlocking bridges. These bridges are responsible for the skin
integrity. The epidermis comprises several physiologically active tissues and a
physiologically inactive top layer: the stratum corneum (SC, 10 mm thick) that is
exposed to the external environment (see Figure 7.3). The drugs can potentially pass
through the skin either via the intact SC and/or via the hair follicles and sweat ducts
(see Figure 7.3). In fact, the average human skin contains 40–70 hair follicles and 200–
250 sweat ducts per square centimeters of area. However, both these appendages
occupy only 0.1 % of the skin; therefore, the SC is the main barrier to drug transport.
7.2.2
Stratum Corneum – Main Drug Barrier
The drug transport through the skin evolves in a series of steps in sequence. The drug is
first absorbed in the SC, diffuses through it and through the lower layers of epidermis,
and then passes through the dermis and finally into the blood circulation. The rate of
penetration through the SC controls the drug delivery because the drug transport
through the deeper layers as well as through the vessel walls happens at a higher rate.
The control of delivery by the SC is due to its unique structure. The human SC has
about 18–21 cell layers of flattened, mainly dead keratinized, metabolically inactive
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194
cells (corneocytes). The individual corneocytes are 20–40 mm in diameter, in contrast
to 6–8 mm for the basal cell and may differ in size or packing depending on the body
site or their location within the SC [2]. These cells are formed and continuously
replenished by the slow upward migration of cells produced at the lower layers. The
SC contains only 20 % wt H2O in contrast to normal 70 % in the lower physiologically active layers.
The SC is composed of closely packed cells. The intercellular spaces are narrow and
occupied by other substances. In fact, the SC has been described as a wall with bricks
and mortar [1,6,7] (Figure 7.4). The corneocytes of hydrated keratin are the ‘‘bricks’’
impended in the ‘‘mortar,’’ consisting of lipid bilayers of ceramides, fatty acids,
cholesterol, and cholesterol esters. This arrangement creates a tortuous path that most
of the drugs should follow in order to pass through the SC. As we will see later, some of
the techniques aim to increase the drug transport through the skin by either increasing
the intercellular space or even completely disrupting the structure of the SC.
As it was mentioned earlier, the SC is partially hydrated (20 % wt H2O). When it
is in contact with water, it absorbs water slowly and when is fully hydrated, it absorbs
five to six times its weight of water. The water diffusion coefficient (Dw) through the
SC is comparable or slightly smaller than that of the biological membrane and in the
order of 1010 cm2/s. It is important to note that the thickness of the SC and the Dw
differs in different body sites (Table 7.1, adapted from [8]).
For example, the SC of the palm is much thicker than in other sites but also highly
water permeable (see Table 7.1). In any case, the Dw through the SC is three to four
Fig. 7.4 Drug permeation through SC.
(Source: Reprinted from Ref. [1], with permission from
Elsevier, 2001.)
Water diffusivity through the SC at different sites.
Tab. 7.1
Skin site
Thickness (mm)
Dw (cm2/s, 1010)
15.0
10.5
13.0
49.0
400.0
6.0
3.5
12.9
32.3
535.0
Abdomen
Back
Forehead
Back of hand
Palm
orders of magnitude lower than in the much thicker (200 mm) dermis
(Dw,dermis ¼ 2.106 cm2/s).
7.2.3
Drug Transport Through the Skin
7.2.3.1 Passive Diffusion
General – Transport Mechanism The drug delivery through the skin due to the drug
concentration difference between the patch reservoir (CD,res) and the skin (CD,skin) is
called drug passive diffusion and can be described by Fick’s law:
JDPD ¼
kD DD ðCD;res CD;skin Þ
;
‘
ð1Þ
where JDPD is the steady state drug flux through the skin, kD is the drug partition
coefficient in the skin, DD is the diffusion coefficient of the drug through the skin,
and ‘ is the skin thickness. When the CD,res >> CD,skin, then Equation 1 can be
written as
JDPD ¼
kD DD CD;res
¼ KDPD CD;res ;
‘
ð2Þ
where KDPD is the passive drug permeability coefficient through the skin.
The drug passive diffusion increases by using the maximum drug amount that can
be dissolved in the drug reservoir (the maximum solubility of the drug in the
reservoir), CD,S:
PD
JD;max
¼ KDPD CD;S :
ð3Þ
Drugs with high partition into the skin and high diffusion coefficient are good
candidates for a TDD system. In fact, Potts and Guy [9] suggested an empirical
relation to predict the drug permeability through the skin:
log½KDPD ðcm=hÞ ¼ 2:7 þ 0:71logkD;oct 0:0061ðMWÞD :
ð4Þ
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196
The kD,oct is the partition coefficient of the drug in octanol/water and can be
used as an indication of the drug partition into the skin. Because the drug has to go
through lipophilic and hydrophilic layers, it should be soluble satisfactorily in
both oil and water. One should, however, note that drugs with high octanol/water
partition usually have low solubility. Therefore, for each drug an optimum in
partition and solubility should be achieved. In addition, the lower the molecular
weight of the drug, (MW)D, the higher is its diffusivity through the skin. Equation
4 is often used to predict skin permeability, and the reader can find other proposed
models and empirical relations elsewhere [9].
Currently, the transdermal delivery mostly involves potent drugs of low
molecular weight (clodinine, estradiol, fentanyl, nicotine, nitroglycerin, testosterone, scopolamine, and others) that are active at low blood concentrations
(order of ng/mL or less) [10]. In order to expand to include a broader range of
drugs, various methods have been developed and will be described later in the
chapter.
Measurement of Drug Passive Diffusion Most of the studies for a TDD system start
with an in vitro evaluation of the delivery. Skin (full skin or only SC) is isolated from
animal and/or human sources and used in laboratory experiments. Figure 7.5
presents a piece of human skin obtained from a patient who had undergone a
cosmetic surgery.
The isolation and treatment of the skin is an important and labor intensive
procedure requiring a series of steps [11]. The skin is cleaned with ethanol and
dermatomed from the rest of the tissue (see Figure 7.5). Depending on the planned
experiments, the SC can be separated from the epidermis and then tried in N2
atmosphere. When access to normal skin samples is not possible, artificial skin
substitutes often used in burn surgery and the treatment of chronic wounds [12] may
be used for in vitro tests.
Fig. 7.5 Photo of human skin sample.
Fig. 7.6 The Franz diffusion cell.
The in vitro experiments are performed using various types of devices. One of
the most common devices is the ‘‘Franz’’ diffusion cell (Figure 7.6). The drug
permeates from the donor compartment through the skin and is detected in the
receptor compartment. The temperature is kept constant at 37 8C with circulating
water.
Figure 7.7 presents a typical example of the passive diffusion of timolol (TM)
[13] through pig SC. TM is a nonselective beta-adrenergic blocking agent that is
used in the management of hypertension, angina pectoris, myocardial infraction,
Fig. 7.7 Passive diffusion profile of TM through pig SC.
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198
Tab. 7.2
Some examples of passive transdermal drug delivery.
Drug – permeant
Skin type
References
Amino acids – dipeptides
Dextran – morphine
Dihydroergotamine
Thyrotropin-releasing hormone
Tetrapeptide
Piroxicam
Nalbuphine
Fentanyl
Diclofenac
Nicotine
Theophylline
Naltrexone
Timolol
Porcine skin
Human skin – in vivo
Rabbit skin
Mouse skin
Human skin
Human SC
Human SC
Rat skin
Human skin
Porcine skin
Human and porcine skin
[15]
[16]
[17]
[18]
[19]
[20]
[21]
[22]
[23,24]
[25,26]
[27]
[28]
[29–32]
and glaucoma. It undergoes extensive first-pass hepatic metabolism, and its
elimination half-life is 2–2.6 h [14].
Table 7.2 presents a selection of scientific literature concerning in vitro
passive diffusion transdermal delivery [15–32]. Transdermal products and short
patent overview will be presented later in the chapter. It is important to note
that if the results of the vitro study are promising, then in vivo study follows. For
this, often animals (mice, hairless rats, guinea pigs) and human volunteers are
used. Detailed discussion about these experiments is beyond the scope of this
chapter.
7.2.3.2 Iontophoresis
General – Transport Mechanism The concept of the TDD via a patch is widely known
and used in daily life. Nevertheless, the skin is a significant drug barrier and for
most drugs the skin permeability is low. The delivery can be assisted by electrical
energy. Iontophoresis implies the use of small amounts of physiologically acceptable electric current to drive charged drug molecules into the body [5]. The
iontophoresis device comprises two patches containing the two electrodes – one
the anode and the other the cathode and the power supply (Figure 7.8). The drug
formulation (drug dissolved in either liquid or gel reservoir) is placed in patch –
electrode having the same charge as the drug (in Figure 7.8, at the anode). The other
electrode/patch contains only reference electrolyte or gel. The two patches are
placed on the skin and connected to the power supply. Figure 7.9 presents an
iontophoresis patch in a more detailed manner. The drug is driven into the skin by
electrostatic repulsion. In addition, a bulk fluid flow or volume flow occurs in the
same direction as the flow of the counterions. This phenomenon, which accompanies iontophoresis, is called electroosmosis.
Fig. 7.8 Iontophoresis principle.
The steady state flux of a charged drug during iontophoresis comprises three
parts: the flux due to passive diffusion (JDPD ), the flux due to electromigration (JDEM ),
and the flux due to electroosmosis (JDEO ):
JDtotal ¼ JDPD þ JDEM þ JDEO :
ð5Þ
In most of the iontophoretic drug delivery systems, JDPD is very low and often
negligible. Nevertheless, when it can be measured, it is used to obtain the passive
drug permeability (KDPD , Equation 2). The electromigration can be described by the
equation [33]:
JDEM ¼
iD
;
zD AF
Fig. 7.9 Schematic illustration of an iontophoretic patch.
ð6Þ
199
200
where, iD is the drug ionic current flow, zD is the charge of the drug, A is the surface
area, and F is the Faraday constant. The drug current flow is related to the applied
current, I, via the equation
iD ¼ tD I;
ð7Þ
where tD is the transport number of the drug and represents the fraction of the total
current transported by the drug. The tD shows the importance of the presence of
competitive ions in the drug for the drug delivery – the higher the tD, the higher the
drug delivery efficiency. By combination of Equations 6 and 7, we get
JDEM ¼
tD I
tD I
¼
:
zD AF zD F A
ð8Þ
The ratio, I/A, is the current density. Based on Equation 8, the drug transport
should increase proportionally to the applied current density. Later, we will see that
this does not always hold.
The electroosmotic flux, JDEO , is the bulk drug flow occurring when a voltage
difference is applied across the charged skin. Rein [34] has proven 80 years ago the
electroosmosis flow in human skin, and recent studies confirmed the phenomenon
in skin from various sources [35–38]. The electroosmosis occurs always in the same
direction as the flow of the counterion and may assist or hinder the drug transport.
The electroosmosis increases in importance as the size of the drug ion increases.
For small ions, the drug flux increases mostly because of electromigration. For
bigger drugs, such as peptides and proteins, the electroosmosis might be the
dominant transport mechanism. An excellent recent review [39] describes some
examples using hairless mouse and human skin.
Factors Affecting Iontophoretic Drug Delivery Figure 7.10 shows a continuous flow
through transport cell, proposed by Van der Geest et al. [40] to measure the
iontophoretic drug transport. The drug is placed in the donor compartment (in
Figure 7.10, at the anodal chamber), and the electric current is applied via the two
electrodes connected to the power supply. A few minutes (2–3 min) after the current
application, the skins polarizes and its electrical resistance drops sharply [31]. The
drug permeates through the skin and is collected by the flow through solution
simulating the blood. The reference electrode compartment (in Figure 7.10, at the
cathodal chamber) contains only reference electrolyte.
Figure 7.11 presents a comparison of the TM delivery between passive diffusion
and iontophoresis through pig SC. The iontophoresis TM permeability is about four
times higher than the passive diffusion [31]. During the first few minutes of the
current application, the electrical resistance of the SC drops sharply from about
16 kV cm2 to about 3 kV cm2 and stays at the lower levels for several hours of
iontophoresis.
Fig. 7.10 The continuous flow through transport cell.
(Source: Adapted from Ref. [13].)
Drug Concentration In passive diffusion, besides the selection of the suitable
drug (based on its MW and partition into the skin), it is very important to achieve high
drug loading in the patch to maximize the drug delivery. In iontophoresis, the drug
concentration still has a great influence on the delivery. Usually, if the delivery is
performed under constant current density, the drug delivery increases as the drug
concentration in the patch is high. However, often the relationship is not linear. The
drug transport number seems to be reaching a plateau at high concentrations. An
Fig. 7.11 Typical result of the TM transport through pig SC
under passive diffusion and iontophoresis.
201
202
interesting finding is that in some cases the delivery is independent of the drug
concentration in the patch or even decreases at high concentrations. These discrepancies are often related to the competition of the drug with other species of the
background electrolyte for the current or to the accumulation of the drug to the skin
(‘‘skin fouling’’). The reader can find several examples concerning specific drugs
elsewhere [33].
Electric Current Based on Equation 8, the iontophoretic drug flux is directly
proportional to the applied current density. In fact, this has often been found in
the literature. In some cases, however, the iontophoretic transport can reach a
plateau suggesting again the saturation phenomena. For the same drug, the flux
level can be influenced by the drug formulation due to competition phenomena.
High amount of background electrolyte can result in lower drug efficiency. Often in
the literature, especially during in vitro experiments, high current densities are
applied. Nowadays, most scientists agree that 0.5 mA/cm2 is the maximum acceptable
current density producing minimal skin irritation and/or damage [13,40,41]. It is
always very important to consider this fact when designing a TDD system. High
currents can often achieve high drug delivery. However, the level of the current, the
time of application, and the potential health risks should always be taken into
consideration.
It is important to note that recent studies indicated that constant conductance
alternating current (CCAC) iontophoresis can enhance drug delivery as much as direct
current (DC). In fact, the results of CCAC delivery seem to show less inter- and
intrasample variability than that of the DC iontophoresis [42–44]. It seems that the
CCAC iontophoresis can provide more controlled TDD.
In iontophoresis besides the amount of the applied current, the type of
electrodes too has an important role. The conventional electrodes used are
classified as inert (metals such as stainless steel, platinum, carbon, or aluminum)
or reversible (Ag/AgCl, see Figure 7.10). The inert electrodes do not take part in the
electrochemical reaction, but they cause electrolysis of water leading to pH shifts
and consequently to skin irritation and perhaps to variations in drug delivery and
stability [5]. The Ag/AgCl electrodes can avoid these problems but participate in
the electrochemical reactions (therefore they are called ‘‘active’’ electrodes). The
use of Ag anode electrode requires the presence of Cl for the electrochemical
reaction:
Anode : AgðsÞ þ ClðaqÞ ! AgClðsÞ þ e ; E ¼ þ0:22V;
Cathode : AgClðsÞ þ e ! AgðsÞ þ ClðaqÞ ; E ¼ þ0:22V:
In most cases NaCl is used to provide the Cl. The presence of Naþ can have a
significant impact on the drug delivery because it competes with the positively
charged drug for the electrical current. In the AgCl cathode, Cl ions are released.
To have electroneutrality, cations, which also compete for the current, should be
released from the skin. As a result, the drug delivery efficiency can be very low.
The Ag/AgCl electrodes can also cause precipitation of peptides and protein
drugs. In recent work, Stamatialis et al. [45] found precipitation of salmon
Fig. 7.12 Loss of sCT due to the electrical current
application using pure Ag/AgCl electrodes and Ag/AgCl
agarose bridged electrodes.
calcitonin (sCT, a drug used in the therapy of hypercalcemia, postmenopausal
osteoporosis, and the treatment of Paget’s disease of bone) in contact with the Ag
electrode. To avoid the direct contact, the electrodes were separated from the sCT
solution by agarose salt bridges. Figure 7.12 shows that the agarose bridges avoid
sCT losses.
The pH The pH of the drug formulation can have a significant effect on the
iontophoretic drug delivery. One should select the pH when the drug is highly
charged to enhance the electromigration. Often, however, a compromise should be
achieved between the pH of the drug and the drug stability and solubility, and skin
irritation. For example, recent studies have shown that the degradation of sCT is
very low at low pH [45–47]. In fact the lowest sCT degradation is found at pH 3–3.3.
At such low pH, however, the risk of skin irritation and/or skin damage is high.
Therefore, as a compromise, minimum pH 4 was selected to achieve both low drug
degradation and skin irritation [45]. Besides the skin irritation, high amounts of Hþ
and OH should be avoided to improve the drug current efficiency. Finally, the pH
of the formulation can also influence the electroosmosis by changing the skin
charge.
All the above-discussed parameters affecting iontophoresis delivery have
been investigated extensively in the literature. Table 7.3 presents a selection of
scientific papers concerning delivery through various skin types in vitro and in vivo
[11,29,31,32,37,40,43,48–66].
203
204
Tab. 7.3
Some examples of iontophoresis transdermal drug delivery.
Drug – permeant
Skin type – condition
References
Timolol maleate
Salmon calcitonin
Timolol maleate
Apomorphine
Morphine HCl
Various dipeptides
Triptorelin
Amitriptyline HCl
Tetraethylammonium chloride
Nicotine
Defibrase
Tacrine HCl
Botox
Ketamine
Catechins from tea
Amino acids
Heparin
Amikacin
Dopamine agonist 5-OH-DPAT
Vapreotide acetate
Nalbuphine, sebacoyl
Antihistamines
Porcine SC – in vitro
Rat skin – in vivo
Human SC – in vitro
Human skin/SC – in vitro
Porcine skin
Cadaver human skin
Human epidermis
Human skin
Human epidermis
Rat skin
Porcine skin
Porcine skin – in vitro
Rat skin
Rabbit skin
Human SC/skin – in vitro
Porcine skin – in vitro
Human skin – in vitro
[31,32]
[48]
[29]
[11,40,49,50]
[51]
[52]
[37]
[53]
[43]
[54]
[55]
[56]
[57]
[58]
[59]
[60]
[61]
[62]
[63]
[64]
[65]
[66]
7.2.3.3 Electroporation
General – Transport Mechanism Electroporation is the structural perturbation of
lipid bilayer membranes caused by the application of high voltage pulses. For
TDD, high voltage pulses (100–1000 V) for short time period (100 ms–1 s) are
used [67]. In contrast to iontophoresis that acts primarily on the drug (by
electromigration and electroosmosis, causing structural changes in the skin
polarization, decrease of electrical resistance) as secondary effects [41]), electroporation acts directly on the skin, causing changes in the permeability of the tissue.
This involves the creation of ‘‘pores’’ (or aqueous pathways) for the drug transport
through the SC that are rather small (<10 nm), cover small surface area [68,69],
and are generally short lived (ms to s) [70–72]. The parameters affecting the
electroporation TDD are
Drug properties: charge/MW/hydrophilicity
Drug formulation: pH/competitive ions/viscosity
Electrical: Pulse waveform/voltage/number/length.
The effects of the drug properties and drug formulation on the delivery are similar
to those of passive diffusion and iontophoresis (and have been discussed earlier).
Here we will focus our attention on the electrical properties.
Pulse Specifications Two types of pulse waveforms are applied: the decaying and
square wave pulses. The exponential decaying pulses depend on the skin resistance and the electroporation system (electrode, conducting medium). This can be
a drawback for clinical application because the reproducibility of the delivery
would not be always possible. The square pulses, however, remain constant
during application and therefore, their effect and drug delivery could be controlled
better.
The drug delivery is generally enhanced when the number or the duration of the
pulses increases [73–78]. The enhancement is steeper at low voltages than at high
voltages.
Electrode Material and Design In most cases inert Pt electrodes are used because
the Ag/AgCl cannot be used for the application of high voltages for a long time. The
design of the electrode is of great importance, too. The parallel plate electrodes,
which are the simplest configurations to create a uniform field [79,80], can cause
nerve stimulation and/or skin irritation and burning. The meander electrode
design localizes the field on the superficial skin layer [80] and seems to be a better
option.
Safety Issues Jadoul et al. [41] have recently reviewed the effects of electroporation
and iontophoresis on the SC. Some changes of the SC, which are partly reversible,
occur for both methods:
Disorganization of the lipid bilayers of SC
Increase of skin hydration
Decrease of skin resistance; a larger decrease observed in
electroporation.
In contrast to iontophoresis that is generally considered a safe method, the safety
issues of electroporation are still under debate. It is rather difficult to define the
conditions that will be acceptable for clinical use. Many parameters such as pulse
voltage, pulse length, pulse frequency, duration of treatment, electrode size,
electrode design, and skin site are important [67,73]. The sensation and pain
thresholds exhibit a rather complex dependence on the current [68]. Tests on
human volunteers indicate that both sensation and pain depend on the volunteer
and besides physiological, often psychological issues are important. Generally, the
levels of sensation increase at high current/charge, pulse rate, and pulse length. The
quality of sensation is determined by the frequency, ranging from warmth to
tingling to vibration. Quite often muscle stimulation and blocking might occur.
205
206
Definitely, in the development of a drug delivery system, a balance between effective
drug delivery and safety should be reached following carefully the safety guidelines
and standards of responsible organizations [68].
The in vitro electroporation experiments can be performed using devices similar
to those used in iontophoresis (Figure 7.10). In the scientific literature, several
important studies show the significant effect of electroporation on the transdermal drug delivery (see Table 7.4 [30,36,78,81–94]). Usually, two different experimental protocols are used: intermittent application of more ‘‘short’’ (<10 ms)
high voltage pulses and less number (<15) or ‘‘long’’ (100–500 ms) medium
voltage pulses (50–250 V) [73,74]. Figure 7.13 shows a typical example of a TM
permeation profile under passive diffusion and electroporation. In this case,
passive diffusion was followed for 2 h, then 10 pulses of 400 V during 10 ms were
applied, and the diffusion after electroporation was evaluated for 6 h. Owing to
pulses, pores are created into the skin that cause significant increase in the drug
transport.
The combination of electroporation with iontophoresis was found to be very
effective to improve transdermal delivery of big drug molecules [36,94,96].
In this case, after the pulse application and the creation of the pores on the
skin, constant current iontophoresis is applied. Besides, electroporation
has been used in combination with ultrasound [97] and chemical enhancers
[98–100]. More details about these techniques will be described later in this
chapter.
Tab. 7.4
Some examples of electroporation transdermal drug delivery.
Drug/penetrant
Skin type
References
Insulin
Salicilic acid
Methotrexate
Piroxicam
Carboxyfluorescein
Timolol
Atenolol
Human parathyroid hormone
Nalbuphine
Buprenorphine
Sodium nonivamide acetate
Fentanyl
Buprenorphine
Physostigmine
Terazosin HCl
FITC-dextran
Sulforhodamine
Salmon calcitonin
–—
Rat skin
Pig skin
Porcine skin
Porcine skin
Human SC
Human SC
Porcine skin
Mouse skin
Human skin
Human skin
Human epidermis
Human epidermis
Human skin
Rat skin
Rat skin
Human skin
Human epidermis
[81]
[82]
[83]
[84]
[84]
[30,36]
[36]
[85]
[78]
[86]
[87]
[88]
[89]
[90]
[91]
[92]
[93]
[94]
Fig. 7.13 TM flux from the carbopol gel 1 % (w/w) through
pig SC under the electroporation protocol (passive
diffusion – pulse application – postelectroporation
diffusion) [95].
7.2.3.4 Other Methods
Skin Penetration Enhancers The TDD can be improved using skin penetration
enhancers (otherwise called sorption promoters or accelerants) [101]. These
compounds penetrate into the skin, interact with skin components, and reversibly decrease the skin resistance to transport. The penetration enhancers should
[102]
Be nontoxic, nonirritating, and nonallergenic
Have no pharmaceutical activity on the body
Act fast in a predictable and reproducible way
Not cause irreversible changes in the skin
Not cause loss of body components
Be compatible with the drug and other components (buffer, etc.)
Be cosmetically acceptable.
An effective penetration enhancer of all the above requirements should improve
the drug passive skin permeability (KDPD , see Equation 2) that it may increase the DD or
improve the partitioning of the drug into the skin (kD). In a recent excellent review,
Williams and Barry [101] summarized the potential mechanism of action of the
enhancers:
Denaturation or modification of intercellular keratin of SC
Effects on the dermosomes that are responsible for the
cohesion between corneocytes
207
208
Modification of the intercellular lipid domains and
therefore reduction of the resistance of bilayer lipids
Alteration of the solvent nature of SC and therefore
increase of drug partition and/or drug diffusivity.
In the literature, one can find several methods to prepare drug formulations with
penetration enhancers. Table 7.5 presents a short overview of the mostly used
enhancers. Some details about their action are included, too [103–128].
Ultrasound-Assisted TDD – Sonophoresis Ultrasound is defined as sound having a
frequency beyond 18 kHz and has already been used in medicine, for example, in
physical therapy, in dentistry or for diagnostic purposes. There are basically three
types of conditions [129]:
High frequency (3–10 MHz) – for diagnostic purposes
Medium frequency (0.7–3.0 MHz) – for therapeutic purposes
(physical therapy)
Low frequency (18–100 kHz) – for lithotripsy, liposuction,
cancer therapy, dentistry, transdermal delivery, and others.
The ultrasound energy penetrates the body tissue and is absorbed by the tissue.
The effects on the tissue can be
Thermal: Owing to ultrasound the temperature of the tissue
increases. The rise varies with the ultrasound intensity and
exposure time.
Cavitational: Owing to ultrasound-induced pressure, gaseous
cavities are formed in the tissue. The cavities can collapse
causing significant changes in the surrounding tissue.
Acoustic streaming: As a result of the sound waves, a onedimensional flow current that affects the surrounding tissue
develops.
The ultrasound can increase the TDD – a method called sonophoresis – from a
small percentage to several orders of magnitude. The mechanism of the improved
TDD has been studied extensively in the past 20–25 years. Two excellent recent
reviews discuss in detail the developments in this field [129,130]. The most dominant
mechanism suggests that the ultrasound interacts with the intercellular lipids of the
SC that is somewhat similar to the action of the penetration enhancers. Mitragotri
et al. [130,131] have evaluated the effects of the various ultrasound-related phenomena to the skin structure and drug delivery. The low-frequency ultrasound seems to be
more effective. For example, an ultrasound of 20 kHz of intensity 225 mV/cm2
applied for 100 ms every second can increase the drug delivery up to 1000-fold higher
than a therapeutic ultrasound [131]. Mitragotri et al. [130,131] suggested that the
cavitation effects cause disorder to the lipids of SC, increasing the water transport
through it. The drugs use these aqueous channels to penetrate at high rates. The
Result
Effective at low concentration
(0.1–0.5 %), interact with
lipid domains of SC
Pyrrolidones (N-methylPartition well in the SC and
2-pyrrolidone (NMP) or
alter the solvent nature
2-pyrrolidone (2P)) [111,112]
of the skin
Fatty acids [113–115]
Interact and modify the lipid
Oleic acid [116,117]
domains of SC
Alcohols, fatty
Increase drug solubility in the patch,
alcohols, glycol [118–121]
increase drug partition into the skin,
solvent drug dragging
Anionic, cationic: interact with
Surfactants anionic, cationic,
intercellular keratin and swell the SC
nonionic (sodium lauryl
sulfate – Tween 80) [122–125] Nonanionic: only minor interaction
Modify solvent nature of SC,
Essential oils, terpenes,
improving drug partition
terpenoids (menthol,
Increase drug diffusivity
eucalyptus) [126,127]
Prepare vehicles to
Phospholipids [128]
carry drug through the skin
Occlude skin surface, increase
tissue hydration
Dimethyacetamide (DMAC),
dimethyl formamide (DMF)
[106–108]
Azone [109,110]
Sulfoxides [103–105]
Problems
Erythema and damage of SC at high
concentrations. Bad odor in the mouth
DMF can cause irreversible
damage to the skin
No
Anionic, cationic: irritants
potentially damaging to the skin
No
No
Promote both hydrophilic
and hydrophobic drugs
Liposomes fuse with SC
lipids and collapse
increasing the drug
partition to SC
No
No
Enhances skin transport of steroids, Not clear
antibiotics, and antiviral agents
(hydrophilic and hydrophobic)
Cause short-lived erythema
and toxic reactions
Action
Increase SC humidity, swelling
High DD, KD for hydrophilic
of corneocytes
Change intercellular keratin conformation,
distort packing geometry of lipid domains
Change intercellular
keratin conformation
Water
Skin penetration enhancers for TDD delivery.
Enhancer
Tab. 7.5
210
thermal effects induced by the ultrasound seem to have a relatively lower impact on
the drug delivery. Merino et al. [132] showed that about 25 % of the enhancement of
the mannitol delivery can be attributed to the increased temperature. Most of the
enhancement is due to cavitation effects.
In low-frequency sonophoresis, the drug enhancement is determined by four
main parameters: frequency, intensity, duty cycle, and application time. The intensity,
duty cycle, and time can be combined to a single parameter, the total energy delivered
to the skin:
Energy ¼ intensity time:
ð9Þ
Generally, at a certain frequency the drug enhancement increases beyond a certain
energy dose. At each frequency, there is a threshold intensity below which no delivery
enhancement occurs [130,133]. Beyond this threshold, the delivery increases strongly
until a second threshold is reached: the decoupling intensity. Beyond the decoupling
intensity, no further drug delivery enhancement occurs. The first threshold is really
low at low ultrasound frequency probably due to the cavitations that are more
pronounced at low frequency. For example, for porcine skin, the threshold is
0.11 W/cm2 at 19 kHz and 2 W/cm2 at 93.4 kHz [130]. It is, however, important
to note that at very low frequency, the drug enhancement is indeed strong but seems
to be localized to certain areas [133]. The effect seems to be more homogenous
beyond an optimum frequency of around 60 kHz [130].
Finally, sonophoresis has already been used successfully in combination with skin
penetration enhancers to deliver mannitol [134] and iontophoresis to deliver heparin
[135]. Table 7.6 presents more examples of sonophoresis drug delivery literature
[131,132,136–149].
Microprojection/Microneedle Patch The microprojection or microneedle patch is a
recent development for the TDD. It consists of a microneedle array that is applied to
Tab. 7.6
Some examples of sonophoresis transdermal drug
delivery.
Delivered drug
References
Insulin
Mannitol
Glucose
Heparin
Inulin
Morphine
Caffeine
Lidocaine
Ketoprofen
Salicylic acid/corticosterone
[136–141]
[132,142,143]
[132,143]
[144]
[143]
[145]
[145,146]
[147]
[148]
[131,149]
Fig. 7.14 Micromachined hollow microneedles:
(a) SEM picture of a 350 mm high microneedle and
(b) array of needles with a pitch of 555 mm.
(Source: Reprinted from Ref. [154], 2003IEEE.)
the skin. The needles penetrate the skin and create superficial pores through which
the drug can be delivered [150]. Such patch has been used in conjunction with passive
diffusion and iontophoresis to deliver therapeutical amounts of antisence oligodeoxynucleoside (ODN) to guinea pigs in vivo [150]. Martanto et al. [151] have also
delivered insulin successfully to diabetic hairless rats. Alternatively, the needles have
been coated with a model protein antigen ovalbumin (OVA). After the needle
penetration through the skin, the OVA was delivered to hairless guinea pigs in vivo
[152]. Besides, Chabri et al. [153] manufactured silicon-based arrays and delivered
genes to human breast skin in vitro. Gardeniers et al. [154] have recently developed a
method to fabricate hollow microneedle arrays in silicon (see Figure 7.14). The patch
containing these needles has improved the diclofenac delivery by a factor of 750 with
respect to passive diffusion.
Recently, Park et al. [155] prepared microneedle patches of biocompatible and
biodegradable polymers such as polyglycolic acid (PGA) and their copolymers. The
first results showed that calcein and bovine serum albumin (BSA) can be delivered
effectively through human cadaver epidermis [155]. The polymeric needles
provide additional safety in comparison to silicon or stainless steel ones. If the
needles break into the skin, they will safely degrade in time without risks to the
patients.
Magnetophoresis Magnetophoresis is an approach to enhance the delivery of a
diamagnetic drug across skin using a magnetic field. The concept seems to be
attractive, but it has not been exploited yet. The first results using benzoic acid
showed that the delivery can be enhanced with respect to passive diffusion. The drug
flux increases with the applied field strength [156].
211
212
7.3
Transdermal Drug Delivery System – Structure/Design
In the past two to three decades the TDD systems have become part of our daily life.
The first ones developed and the most popular ones are still based on passive drug
delivery. In the past few years, however, the iontophoresis TDD systems are slowly
being introduced. Yet, these systems have not found broad application outside the
hospital though the prospects are positive.
In this section, we will separate the TDD systems into two main categories:
The passive systems: traditional delivery due to drug
concentration difference between the patch and the skin
The active systems: where an extra driving force such as electric
current (iontophoresis/electroporation), ultrasound
(sonophoresis), or others is used.
The materials and basic technologies for the construction of the patches are the
same for both systems. Therefore, it will be described only once for the passive
systems. Of course, the active systems require extra components to provide the
driving forces (electrical current, ultrasound, and others). These elements will be
described separately.
7.3.1
Passive TDD Systems
7.3.1.1 Types
The technologies developed to provide controlled passive drug delivery can be
classified into two main categories:
1. Membrane or reservoir systems: In this case, the drug is incorporated into a reservoir (liquid or gel) placed between a drug
impermeable layer and a membrane (see Figure 7.2). The
device also includes an adhesive layer on the external surface of
the membrane to achieve a firm contact with the skin. The drug
release from this system can be controlled by varying the
reservoir composition and the drug permeability through the
membrane (by tailoring the material, porosity or thickness)
and/or through the adhesive. Several successful commercial
TDD systems are based on this design (see Table 7.7, adapted
from [157,158]).
2. Matrix systems: In this case, the drug is incorporated (dissolved
and/or distributed) into a polymer matrix (see Figure 7.15).
There is no membrane and the adhesive layer is added when
the matrix itself is not adhesive. Table 7.7 presents some
examples of systems of this type, too.
7.3 Transdermal Drug Delivery System – Structure/Design
Tab. 7.7
Some commercial passive TDD systems.
Trade name
Company
Type
Drug
Action
Nitroderm
Nitrodur
Deponit
Nitrodisc
Frandol- Tape
Alza/Ciba
Key/Schering
Lohman/Schartz
G.D. Searle
Nitto Electric Ind.
Reservoir
Matrix
Matrix
Matrix
Matrix
Antianginal
Antianginal
Antianginal
Antianginal
Antianginal
Catapres
Duragesic
Kimete Patch
Transderm – Scop
Estraderm
Minitran
Alza/Boehringer Ing.
Alza/Ivers/Jansen
Myun Moon
Alza/Ciba
Alza/Ciba
3M
Reservoir
Reservoir
Reservoir
Reservoir
Reservoir
Reservoir
Niconil
Nicoderm
Nicotrol
Elan
Alza
Cygnus
–—
–—
–—
Nitroglycerin
Nitroglycerin
Nitroglycerin
Nitroglycerin
Isosorbide
dinitrate
Clonidine
Fentanyl
Scopolamine
Scoparamine
Estradiol
Glyceryl
trinitrate
Nicotine
Nicotine
Nicotine
Antihypertensive
Narcotic analgesic
Antimotion sickness
Antimotion sickness
Hormonal
Antianginal
Antinicotinic
Antinicotinic
Antinicotinic
7.3.1.2 Materials
The major parts of TDD systems are composed of polymers. The drug impermeable
layer, the drug reservoir, the pressure adhesive layer and the artificial membrane are
all prepared from polymers. The range of the polymers used is really broad; natural
polymers (gelatin, starch, etc.), semisynthetic (hydroxyl propyl cellulose, nitrocellulose, cellulosic), synthetic (polysiloxane, polybutadiene, polyisoprene, silicone
rubber, polyesters, polyurethane, polyethylenevinylacetate polyacrylamide, polyvinylalcohol, polysulfone, polymethylmethacrylate, etc.). The reader can find excellent
overview of the polymers used elsewhere [157,159,160].
In the membrane (or reservoir) system, the membrane is generally the part in
direct contact with the skin and acts as the interface between the drug reservoir and
Fig. 7.15 Matrix TDD system.
213
214
the skin to give optimal control for the transdermal drug delivery. The membrane
should have the following requirements:
It should be made of biocompatible material to avoid skin
irritation.
It should control the drug delivery (the permeability of the drug
should be lower through the membrane than through the skin).
In this case the transdermal bioavailability of the drug becomes
independent of any possible intra- and/or interpatient
variability in skin permeability.
The drug adsorption on it should be low.
The issue of the drug-controlled delivery will be addressed in more detail in the
following section.
7.3.1.3 Skin or Device-Controlled Delivery
The issue of skin or device-controlled delivery has been discussed in the scientific
community for a long time. In the transdermal drug delivery by a patch (see
Figure 7.2), the total permeability (KD,total) of the drug through the membrane
and the skin is given by
1
1
1
¼
þ
;
KD;total KD;mem KD;skin
ð10Þ
where KD,memb and KD,skin represent the permeability of the drug through the
membrane and the skin, respectively. Depending on the ratio of KD,memb and KD,skin,
the delivery may be primarily skin-rate controlled or primarily membrane-rate
controlled. When the ratio KD,memb/KD,skin is less than 0.2, the delivery is considered
to be membrane controlled. When the ratio KD,memb/KD,skin is larger than 5, it is
considered to be skin-rate controlled. If the ratio KD,memb/KD,skin is in between 0.2 and
5 the systematic dosage received is controlled by both the skin and the membrane.
As we have seen earlier, the passive drug transport can have great intra- and
interpatient variability. Moreover, the delivery of some drugs should have a very welldefined therapeutic window (i.e., where a dose above a certain limit is ineffective or
even toxic to the patient). To ensure that the drug delivery through the patient skin is
always the same, the device should have a major control on the delivery. Based on this
fact, a lot of scientists have worked on patches where the drug delivery was designed
to be much lower than through the skin. However, for most drugs the safety issues are
not that critical. Moreover, the passive TDD is so low that even great variability
between patients cannot cause safety problems. In these cases, in order to deliver as
much drug as possible, the device should not impose any restriction and/or control
on the drug delivery. It should only be used as a storage of the drug. Most of the
modern TDD systems have no controlling membranes.
Often, devices with quite different release kinetics when in contact with buffer
solutions may perform quite similarly when in contact with skin in vitro and in vivo. In
these cases, the skin controls the delivery, and possible variability in the actual delivery
arises solely from variations in patients’ skin permeability. It is important to note that
often in the literature for TDD, the term ‘‘rate controlling’’ is unspecified, if not
misleading. A characteristic example has been presented in the work of Hadgraft et al.
[161]. The authors studied the delivery of nitroglycerin from four different commercial
‘‘rate controlling’’ TDD systems having different drug loading and surface areas:
Transderm – Nitro: drug ¼ 10 mg, area ¼ 20 cm2
Nitrodur II: drug ¼ 80 mg, area ¼ 20 cm2
Deponit: drug ¼ 32 mg, area ¼ 32 cm2
Minitran: drug ¼ 36 mg, area ¼ 13.3 cm2.
All these systems were designed and confirmed to deliver [161] in vitro 10 mg in
24 h. However, the differences in their specifications do not allow the estimation of
their control in drug delivery. Guy and Hadgraft revisited this study [162] and
calculated the fractional rate control of the device by the equation:
Mdevice ðfractional control by deviceÞ ¼
Qtotal
;
Qdevice
ð11Þ
where Qtotal and Qdevice represent the amount of the drug transported in a given period
of time under steady state conditions through the combination of device and skin and
the device alone (in mg/cm2), respectively. The fractional control by the skin is then
given by MS ¼ 1 Mdevice. Thus, Mdevice ¼ 1 implies that the delivery is controlled
entirely by the device (often the membrane); however, when Mdevice < 1, the skin is
contributing to the control process. Guy and Hadgraft found [162] that only Deponit
plays a significant role in the control of nitroglycerin delivery (Mdevice ¼ 0.87). For the
others, the results were mixed: Transderm – Nitro/Mdevice ¼ 0.45; Nitrodur
II/Mdevice ¼ 0.13 and Minitran/Mdevice ¼ 0.28. One of the key parameters to be
matched in TDD is the amount of drug absorbed by the skin on time. Moreover, for
safety reasons, the drug loading should be as close as possible to the amount absorbed
by the skin.
7.3.1.4 Commercialization – Patents
In the field of passive transdermal delivery, several companies have been active
in the past 15–20 years. Since the beginning of the 1990s, Alza, Merck, and
Ciba/Novartis are the ‘‘big players.’’ In recent years, 3M presents a dynamic
profile in the area as well, and it has marketed a series of products used in
transdermal patches such as membranes, impermeable backing layers, adhesives,
and so on.
In the patent literature, one can find hundreds of inventions concerning passive
TDD. In these patents, the application of a rate controlling dense or microporous
membrane for the drug delivery plays a rather important role. The claims mainly
address the use of specific membranes. These membranes either exclusively control
the drug delivery or partially control the drug transport together with other
components of the invention, such as the adhesive layer. The membranes are either
commercial or laboratory-made. They cover a big list, including hydrophobic and
215
216
hydrophilic materials. They may be symmetric or asymmetric, and their pore size
may vary from 1 nm to 100 mm but preferably lower than 5–10 nm. Table 7.8
presents a selection of recent patents for passive TDD systems.
7.3.2
Active TDD Systems
7.3.2.1 Types
Iontophoresis and electroporation are the most widely used active systems. In the
past decades, they have been broadly tested for drug delivery in the clinic. In
iontophoresis, a broad range of drugs has been tested in vitro, and in vivo such as
steroids, lidocaine, nonsteroidal inflammatory drugs (NSAIDs), histamine, antihistamines, antibiotics, and others [5]. In electroporation, mostly the big drug molecules
have been delivered, such as heparin, calcitonin, luteinizing hormone–releasing
hormone (LHRH), and others [5].
7.3.2.2 Materials – Devices
The basic elements of the active patches are common to those of the passive patches.
Polymers are also used for the drug impermeable layers and the skin contacting
membrane. Buffer solutions and/or gels are used for the drug formulation, and so on.
Of course, in iontophoresis and electroporation, specific attention is given to the
construction of the patches due to the presence of the electrodes for the current
application. Often to avoid direct contact with the drug, the electrodes are placed in a
separate compartment created by a membrane. The separate compartment is filled
with conductive gel for the current conduction. The basic drug formulation composition varies between manufacturers, but typically is a highly conductive gel, too. In
iontophoresis, mostly Ag/AgCl electrodes are used. The electroporation electrodes
are mostly manufactured from inert material (Pt) and have a meander structure for
better voltage distribution and patient safety. Figure 7.16 presents some examples of
commercially available iontophoresis electrodes.
In almost all patches – electrodes, an artificial, nonrate-limiting membrane, is used
in contact with the skin. The drug delivery is regulated via the external driving forces,
namely the current (iontophoresis), the voltage pulses (electroporation), the ultrasound (sonophoresis) and so on. In these cases, inter- and intrapatient variability in
drug – skin permeability is much lower than in the passive systems.
In iontophoresis, the electrodes are connected to a relatively small (‘‘walkman’’ or
‘‘discman’’ size) power supply. In the United States, several companies manufacture
such devices. Typical examples of such the devices are Dupel (Empi, St Paul, MN),
Phoresor II (Iomed, Salt Lake city, UT), Iontophor II (Dynatronics, Salt Lake city, UT),
and others [5]. Figure 7.17 presents an example.
The electroporation power supplies have been primarily manufactured for
genetic manipulation of living cells and have been adapted for TDD. These
devices are much bigger in size than the iontophoresis devices because of the
need of the capacitor. A partial list of manufacturers in the United States includes
Genetronics Inc. (San Diego, CA), CytoPulse Sci. (Columbia, MD), Bio-Rad
Tab. 7.8
Selection of patents on passive TDD.
Patent
Inventor/owner
Short description
DE 19738643/1999
US 4951657/1990
Mueller –
Lohman
Pfister et al.
US 5869089/1999
Risheng Wu
US 5585111/1996
Peterson
US 4913905/1990
Fankhauser –
Ciba
US 5904930/1999
Fisher et al.
US 5284660/1994
Lee et al.
US 5462745/1995
Enscore et al.
CZ 287678/2001
Bracher et al.
WO 0130316/2001
Dreyer – 3M
WO 0003698/1999
Unihart Corp
US 5613958/1997
Kochinke et al.
EP 0916339/1989
Pharmacia et al.
WO 0023644/2006
US 078601/2006
US 121102/2006
3M
Noven Pharma
C.C. Ming
WO 058287/2005
B. Stefan et al.
Regulation of scopolamine delivery by
polyethylenevinylacetate (PEVAc) membrane
Dense membrane of polydimethyl siloxane
(PDMS) and polyurethane (PU) for TDD
Nuclear track microporous membrane made
from alpha particles or modified EVAC for TDD
Dense membrane (Cotran 9702, PEVAc, etc.) or
microporous (Cotran 9710) controls TDD
Membrane (dense or microporous) controls TDD.
Material could be hydrophobic and hydrophilic,
and the membrane could be symmetric
or asymmetric
Delivery of tamoxifen. Microporous membrane
(polypropylene, PEVAc, and silicone) controls
the drug delivery to a varying degree.
When skin penetration enhancer is applied,
the membrane controls TDD
Membrane (of hydrophilic or semihydrophilic
polymer) impermeable to drug in a dry state
and permeable in a hydrated or wet state.
The patch, wetted manually by sweat or
just prior to use
Drug membrane flux no greater than the drug
skin flux (dense Cotran 9702, low-density
polyethylene membrane). Rate control is
achieved by the thickness of both
membranes – adhesive
Testosterone delivery. The reservoir is separated
from the skin by the rate-controlling
adhesive layer
Drug in adhesive reservoir layer the thickness
of which controls the drug delivery
Concentration of solubilizers, penetration
activators, or providing a membrane
control the delivery
An adhesive layer containing plasticizer
type enhancer (10–40 %). Membrane
(optionally used) having pore size between
about 0.01 and about 3.0 mm and being
nonrate-controlling membrane
Adhesive and/or preferably nonporous
membrane controls nicotine delivery
Reservoir TDD system
Transdermal system for delivery of estradiol
Transdermal system for delivery of estrogen
and/or progestin
Transdermal system for delivery of hormones
217
218
Fig. 7.16 Iontophoresis electrodes from Iomed.
(Source: Iomed, Salt lake City, UT, USA, 2005, printed with
permission.)
(Richmond, CA), and others [5]. Figure 7.18 presents an example of power unit for
iontophoresis and electroporation (printed with permission from Moor Instruments Ltd, Devon, UK).
7.3.2.3 Commercialization – Patents
Commercialization of iontophoresis/electroporation patches for self-use by the
patient is still not achieved. The drug delivery is still rather complex for the patient
and the cost of the technology high. Besides, regulatory approval of such systems
seems to be more difficult than the passive systems. The manufacturers should
present to the regulatory authorities, besides information about the drug, a detailed
description of the electronics of the patch and the equipment, as well as of the drug
delivery protocols (e.g., current density or voltage, time of application, etc.). In
principle, the patches – electrodes can be filled with the drug formulation of choice
and can be broadly used. The drug delivery protocols, however, should be drug specific.
Iontophoresis is already approved in the United States, for the topical delivery of
lidocaine and epinephrine for local analgesia (Iomed, USA). Devices available on the
market for delivery of local anesthetics and corticosteroids include Phoresor II
Fig. 7.17 Phoresor II auto iontophoresis device
from Iomed.
(Source: Iomed, Salt lake City, UT, USA, 2005,
printed with permission.)
Fig. 7.18 Iontophoresis and electroporation
devices from Moor Instruments Ltd.
(Source: Moor Instruments Ltd, Devon, UK,
printed with permission.)
(Iomed), Empi Dupel (Empi, USA), Life-Tech Iontophor (Life-Tech, USA), and Henley
Intl Dynaphor (Henley Intl, USA). In addition, devices for iontophoresis of pilocaprine
are on the market, including the CF Indicator (Scandipharm, USA), among others.
Several companies have been trying to commercialize miniature patch systems. A
partial list includes the Alza Corp. (USA), Becton Dickinson (USA), Fournier (France),
Hisamitsu Pharm. (Japan), and Cygnus (USA). In addition, companies such as Thera
tech and Genetronics have developed electrotherapeutic devices, electrodes, and so on
applied in transdermal delivery systems with a combination of various driving forces
(such as iontophoresis þ electroporation, or iontophoresis þ sonophoresis, or iontophoresis þ magnetophoresis/magnetoporation).
Table 7.9 presents a selected list of patents describing active TDD systems. The
iontophoresis and electroporation systems seem to dominate the field. In most of the
patents the control of the drug delivery is left to the applied driving forces. The role of
the membrane is rather limited to either a holder or support of the drug reservoir
and/or to prevent contact of the drug solution and the electrode. In these cases, its
porosity and pore size are just enough to avoid uncontrolled leakage of the drug to the
skin surface and skin fluids to pass to the drug reservoir. However, there are
inventions that provide rate control by the membrane to avoid:
(i) Dependence of the drug transported by iontophoresis
through the skin on the skin electrical resistance (therefore
the transdermal drug bioavailability becomes dependent on
any possible intra- and/or interpatient variability in skin
permeability). In patent US 5167616, the iontophoretic flux of
metoclopramide coming from the individual’s back through skin
could be 50–150 % higher than through other skin sites
because of lower electrical resistance of the skin at the back site.
(ii) There are safety features to prevent excessive drug
electrotransport if the patch is applied on damaged skin or on
body surface that has somehow been compromised.
(iii) If the permeability of the drug through the membrane material
is much higher than the permeability of the drug through skin,
the accumulation of a large quantity of the drug on the skin
surface could cause skin irritation.
In EP 0847775A1/Mori-Hisamitsu Pharm., and US 6167301/Flower et al.,
the above-mentioned disadvantages have been extensively addressed, and an
219
220
iontophoresis system that measures the resistance of the skin and in response to
adjusts the output voltage (drug delivery) is proposed. In several other patents the
membrane-controlled drug delivery has also been proposed (see Table 7.9). Finally
some patents describe the combination of driving forces, for example, in patent WO
0035533/Hofmann – Genetronics, the combination of iontophoresis and electroporation for the delivery of drugs and genes is described. In US 5947921/Johnson et al., the
application of chemical and/or physical enhancers and ultrasound for transdermal
drug delivery is proposed, and an increase of permeability of magnetic particles is
claimed if a magnetic field is applied in combination with ultrasound.
Tab. 7.9
Selection of patents on active TDD.
Patent
Inventor/owner
Short description
US 5667487/1997
Henley
US 5362308/1994
Chien et al.
EP 0748636A2/1996
Yanai et al. – TakedaHisamitsu Pharm
WO 0066216/2000
Kontturi et al.
EP 0813887A2/1997
Higo-Hisamitsu
Pharm
EP 931 564/1999
Theeuwes et al.
US 5445607/1995
Venkateshwaran
et al. – Thera tech
Application of ultrasound + iontophoresis.
Drug in a porous material: polyethylene,
paper, cotton silicone, PTFE, ceramic
A membrane separates two gel layers. The
top gel layer contains ion exchange resins
and contacts the electrode. The bottom layer
contains the drug (proteins and peptides
including calcitonin)
Iontophoresis delivery of protein and
peptides. Low protein adsorptive membrane
(hydrophilic or hydrophilized hydrophobic
material) acts as a support for the drug
(dry or in solution). An ion exchange
membrane protects the drug from the
electrode. The applied current could
be constant or pulsed
Iontophoresis delivery. An ion exchange
membrane is used to separate the
electrode from the drug solution-gel and a
microfiltration membrane is placed
in contact with skin
Current (constant or pulsed) delivers protein
and peptides through animal skin. A patch
contains an electrode–conductive gel
reservoir and a separate drug loaded
membrane
No drug flux is obtained when current is
off and no linear increase of drug-flux
when the current is on
Skin doesn’t control the drug delivery due
to application of a penetration enhancer.
The proposed rate controlling membrane
should be electrically sensitive. No drug
permeation during passive diffusion but
linear increase of drug permeability
with the current is obtained
7.4 Conclusions and Outlook 221
7.4
Conclusions and Outlook
The passive TDD technology is well established and has ‘‘won the hearts and minds’’
of the patients. The application and the market of the passive patches are expected to
grow further in the next years. New products that are easier to use with better quality
materials will be developed. The main expansion of TDD is expected to come from
the active systems. Using active systems such as those based on iontophoresis and
electroporation, the transdermal delivery of several more drugs with better patient
compliance (reduced side effects, etc.) will become possible. Slowly these emerging
technologies will overcome the technical difficulties and find broader applications.
One issue that is always in the spotlight is the need for rather bulky, heavy,
and complicated power supplies and other components. However, the latest
Fig. 7.19 Iontophoresis patches containing (a) both electrodes
on the same patch from Iomed (Source: Iomed, Salt lake City,
UT, USA, 2005, printed with permission.) and (b) both
electrodes and built-in battery, from Travanti Pharma Inc.
(Source: Travanti Pharma Inc., St Paul, MN, USA, 2005, printed
with permission.)
222
developments in miniaturization of systems might give a significant help on this
issue. The miniaturization of the components and the development of suitable
microcomputer for regulation and control of delivery will make these technologies
safer and therefore available outside the hospital, too. Figure 7.19 presents some
example of patches containing both electrodes and/or the battery built in making
their application much easier.
Currently, the majority of the active systems are more expensive than conventional
drug delivery systems. However, as the technology solves the technical problems,
they will become cheaper. Nevertheless, even if the costs are relatively high the patient
benefits from these technologies may be able to justify the extra costs.
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8
and Biohybrid Organ Technology
Thomas Groth, Zhen-Mei Liu
8.1
Introduction
8.1.1
Application of Membranes in Blood Detoxification
Membranes have found a multitude of applications in biotechnology and medicine
where controlled exchange of solutes between different phases or compartments is
required [1–4]. Particularly the application of membranes for blood detoxification
(hemodialysis) has been well established since several decades [1,2]. Nowadays,
membranes for hemodialysis represent the largest market segment in the field of
membrane production [3]. Applications of membranes for blood detoxification such
as hemodialysis to treat kidney failure are linked to transport of solutes to remove
(toxic) metabolites, salts, and water from the organism. On the other hand, the
undesired loss of substances, such as proteins, must be prevented [2,4]. The transport
process is basically controlled by the molecular cut-off of membranes not only to
remove toxins and excess of salts and water but also to avoid any passing of essential
components from blood, such as serum albumin and other proteins [4,5]. Elimination of substances from human blood can be also accomplished by specific or
selective adsorption of substances onto the surface of membranes [6]. This has been
achieved by the exploitation of the affinity of molecules to membrane bulk materials
(e.g., by hydrophobic interaction) [6,7] or to specific surface-attached functions and
ligands [8,9]. Quite novel developments also aim to use imprinting technologies
specifically to adsorb molecules from liquids, which may be interesting for applications in blood detoxification as well [3,10,11]. However, it must be emphasized here
that in the majority of biotechnological and biomedical applications of membranes,
nonspecific adsorption and adhesion processes have to be minimized [1,4,5]. This
is necessary to prevent biofouling, which may cause blocking of membrane pores
by adsorbed proteins and attached cells from the surrounding liquids [1].
A further requirement is the prevention of the undesired activation of coagulation
and inflammation during blood contacting applications, which represents
ISBN: 978-3-527-31480-5
228
8 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology
life-threatening risks for the patients [2,4,5]. The control over adsorption and
adhesion processes is basically obtained by using hydrophilic polymers for membrane formation, such as cellulose and its derivatives [12], blending or grafting
of hydrophobic polymers with hydrophilic polymers, for example, poly(vinylpyrollidone) [13,14], using copolymers consisting of hydrophobic and hydrophilic
co-monomers [15,16], and coating or grafting of membrane surfaces by
hydrophilic macromolecules such as poly(ethylene glycol) or phosphatidylcholine
polymers [17,18]. Also, the binding of bioactive molecules like heparin has been
used to avoid surface-induced adsorption and activation phenomena at least in
blood-contacting applications [19,20]. Overall most techniques to prepare membranes for blood detoxification and oxygenation aim to reduce efficiently the
adsorption of proteins and thus adhesion of cells from blood.
8.1.2
Requirements to Support Adhesion and Function of Cells
By contrast, applications in tissue engineering and biohybrid organ technology
require that tissue cells adhere on material surfaces [21]. Adhesion of cells is a
prerequisite for a multitude of cellular functions such as movement, growth,
differentiation, and survival [22]. Cell adhesion is not a simple on–off event. It
has been shown that the strength of adhesion controls the degree of cell spreading
and hence the shape of cells on a substratum [23]. The shape of cells, however, seems
to be a regulator of cellular functions [23,24]. For example, it has been shown that
spreading of cells from connective tissue such as fibroblasts promotes their growth
and function in terms of extracellular matrix (ECM) synthesis [25,26]. Figure 8.1
shows as an example the effect of surface chemistry and wettability on the shape and
functional activity of human dermal fibroblasts. It is visible that a methyl-terminated,
nonwettable surface inhibits adhesion and spreading of cells (Figure 8.1a). As a
consequence, the synthesis of a fibronectin matrix is greatly reduced (Figure 8.1c).
In contrast, an amine-terminated, wettable surface promotes spreading of cells
(Figure 8.1b) and fibronectin matrix synthesis (Figure 8.1d). However, epithelial cells
such as hepatocytes tend to dedifferentiate if they spread too strongly on a substratum
while they maintain their functional activity when attached moderately on a substratum having a round morphology [27,28]. Figures 8.2 and 8.3 show as an example
human hepatoblastoma cells – a hepatocyte cell line cultured on a highly adhesive
support (Figure 8.2a) in comparison to a less adhesive substratum (Figure 8.2b) on
which cell spreading is greatly reduced. The copolymer in Figure 8.2a possesses
primary amine functions and is called aminoethylmethacrylate (AEMA), while
the copolymer in Figure 8.2b contains N-vinylpyrollidone (NVP) as co-monomer
and is named NVP 20 (see also Section 2.2). Please also note that cells make stronger
intercellular contacts on the less adhesive membrane, which contains NVP 20. If
the functional activity is investigated as shown in Figure 8.3, it is visible that
the detoxification capacity of cells, which is measured by the conversion of
7-ethoxycoumarin (ECOD), is lower on the highly adhesive substratum (AEMA)
in comparison to the less adhesive NVP 20. From the literature it is further known
8.1 Introduction 229
Fig. 8.1 Effect of surface chemistry on shape,
adhesion formation and functional activity of
dermal fibroblasts. Cells on methylterminated self assembled monolayers (SAM,
a and c) have reduced spreading and lack
focal adhesions (a), which is corroborated by
reduced functional activity of cells indicated
by the absence of fibronectin matrix
formation (c). Only spots of fibronectin have
been deposited on the surface. In contrast,
amine-terminated SAM (b and d) promote
adhesion and spreading of cells with
formation of well-developed focal adhesions,
which are visible as white streaks in the cell
periphery (b). This is accompanied by the
formation of an abundant extracellular matrix
of fibronectin (c). Cells and matrix were
stained with fluorescence-labeled antibodies
against vinculin (a and c) and fibronectin (b
and d) and visualized by fluorescence
microscopy. Bars in upper panel have a length
of 100 mm and in the lower panel of 50 mm.
that a complete failure of cellular attachment particularly for epithelial cells results in
the onset of apoptosis, a mechanism by which cells undergo controlled cell death [29].
It has been shown that cell adhesion is linked to the presence and conformation of
specific attachment proteins on material surfaces [30]. Adhesive proteins such as
fibrinogen, fibronectin, vitronectin, and von Willebrand factor are normal components of blood plasma and may adsorb on biomaterial surfaces [31]. The ECM, which
surrounds cells in tissues, is composed of structural proteins, like collages, adhesive
230
Fig. 8.2 Comparison of C3A hepatoblastoma cell
attachment on polymer membranes composed of highly
adhesive amine group containing copolymer AEMA (a) and
a hydrophilic low adhesive copolymer NVP 20 (b) by
staining cells for vinculin. Note the large cell size on AEMA
and the reduced spreading on NVP 20. Size of scale bar is
50 mm.
proteins, and glycosaminoglycans [32]. These proteins provide attachment to cellular
receptors mainly integrins and deliver signals important for surivival, growth, and
differentiation [33,34]. The physicochemical properties of the material surface such as
surface energy (wettability) and electrical surface potential (zeta potential), which are
Fig. 8.3 Comparison of metabolic function of
hepatoblastoma cells in terms of P450 activity indicated by
the ability of cells to convert 7-ethoxycoumarin. It is shown
that cells cultured on the less adhesive substratum NVP 20
have a significantly higher activity than cells on more
adhesive substratum AEMA.
8.2 Application of Membranes in Tissue Engineering
dependent on the chemical composition have an impact on the adsorption of proteins
from surrounding liquids, such as blood or tissue fluids [35,36]. Accordingly, cell
growth and function were found to be strongly related to the wettability of materials
[37–39]. Thepresenceofpolarorchargedfunctionalgroups,suchasaminoorcarboxylic
groups, has been identified as promoting, while apolar groups like methyl may inhibit
cellular attachment, growth, and function [40,41]. Hence, the tailored synthesis of
copolymers or the modification of biomaterial surfaces to obtain specific quantities of
functional groups can be used as a tool to optimize the biocompatibility of materials.
It is also known that cell substratum interactions depend on the underlying
topography [42]. Even quite small topographical changes have been shown to cause
certain cellular responses [43]. For example, structures in the range of 100 nm are
sufficient to induce guidance of cells [44,45]. There is also evidence for the influence of
surface porosity on cell surface interactions [46–48]. It was demonstrated that changes
in surface topography created by variation of pore size in the range of 0.1–0.8 mm
supported corneal epithelia outgrowth, in contrast to the inhibitory effect of pores in
the range of 0.9 mm or greater [49]. Earlier studies have also indicated that skin
outgrowth can be promoted on membranes with pores in the range of 0.025–1.2 mm,
while tissue growth was inhibited on surfaces having pores larger than 3–8 mm [46].
The overall chemistry and topography of material surfaces can be used to control the
adhesion and function of cells. This is particularly important when low-cost applications are envisaged, which shall allow a later biomedical application of the biomaterial.
8.2
8.2.1
Introduction to Tissue Engineering and Membrane Applications
Tissue engineering has been defined by Langer and Vacanti as a novel discipline
aiming at the restoration of tissue structure and function [50]. This shall be achieved
by the cooperation of three components. First, some type of scaffold that shall guide
tissue regeneration gives mechanical support to cells. The scaffold is merely a
degradable material, which shall dissolve in a timescale necessary for cells to replace
it by their own extracellular matrix components (e.g., newly formed bone). The
scaffold possesses typically certain porosity to allow the colonization with cells,
neovascularization, and transport of oxygen, nutrients, and waste products [51].
Second, the scaffold may be loaded or covered with bioactive molecules, which
stimulate anchoring, growth, and differentiation of cells [52]. However, this is not an
absolute requirement. The material or its surface may have some inherent bioactivity,
which stimulates cells (e.g., calcium phosphates for osteoblasts) [53]. Third, the
scaffold must be colonized by cells, which can occur in vitro or in vivo (in situ). The
cells are preferentially autologous, obtained by biopsies and seeded on the scaffold.
Another possibility is that cells colonize the scaffold by migration and ingrowth from
surrounding body fluids or tissues [54]. In most tissue engineering applications,
231
232
Fig. 8.4 Scanning electron micrograph
showing the establishment of a multilayer of
keratinoctyes on a polymer membrane.
degradable scaffold materials are an absolute requirement. However, there are also
applications when nondegradable materials are possible. This holds in cases when
guidance of tissue regeneration is required as it is known for bone regeneration (see
Section 2.3) or specific arrays of cells, such as for an artificial retina that must be
maintained for a long time in a specific spatial arrangement [55]. A possible
disadvantage of using nondegradable materials is that a second surgical intervention
may be needed to remove the material after healing.
Membranes can be prepared from a variety of materials including degradable and
nondegradable polymers [1–4]. (Flat) membranes are interesting in applications,
when two-dimensional tissues shall be replaced. This holds for regeneration of
epithelia, which represent single or multilayers of cells covering the outer and inner
surface of our body. Figure 8.4 shows as an example of such a structure a multilayer of
keratinocytes cultured on a polymer membrane. Membranes may also provide
guidance for cell growth in certain applications or can be used in a tubular setting
either made of hollow fibers or (rolled) flat membranes to provide guidance for
regeneration of bone, nerves, blood vessels, and other tubular tissues. Table 8.1 gives
a general overview about applications of membranes in tissue engineering. The
following sections will highlight some established or potential applications of
membranes in the field of tissue engineering.
8.2.2
Tissue Engineering of Skin
Treatment of deep-skin defects such as burns or chronic skin ulcers in elderly or
diabetic patients represents a serious problem in medical care. Medical therapies aim
to restore the normal skin structure, which consists of an underlying dermal part with
fibroblast and other types of cells, and also blood vessels embedded in extracellular
matrix material overlaid by the epidermis as a nonvascularized multilayered epithelium [71,72]. This includes the restoration of the dermal part as a prerequisite for
proper regeneration of the epidermis [72]. Application of autologous split skin grafts
is still the best medical standard but not feasible if the wound size is too large or the
constitution of the patient does not allow a further skin injury at the donor site [73,74].
First attempts to repair deep-skin defects by tissue engineering were directed to the
regeneration of the epidermis. Rheinwald and Green developed more than 30 years
Tab. 8.1
General overview on membrane applications in tissue engineering.
Membrane sources
Applications
Reference
Poly[(ethylalanato)1.4(imidazolyl)0.6
phosphazene] film
Polylactide film
Polylactide membrane
Polyurethane membrane
Polylactides and polyurethanes
membrane mixture
Expanded polytetrafluoroethylene
Collagen membrane
Collagen membrane
Collagen membrane
Polyglycolide membrane
Chitosan membrane
Heparin–chitosan complex membrane
Genipin-crosslinked gelatine membrane
Peripheral nerve repair
[56]
Reduction of posterior adhesion
Diaphyseal bone regeneration
Potential for ventricles
Potential for skin repair
Potential for pancreas
Skin repair
[57]
[58]
[59]
[60]
[61]
[62]
Periondontal tissue regeneration
Skin repair
Periodontal tissue regeneration
Repair injured urinary tract
Bone regeneration
Skin repair
Skin repair
Skin repair
[63]
[64]
[65]
[66]
[67]
[68]
[69]
[70]
ago a technique to prepare cultured epidermal autografts (CEA) in vitro, which are
derived from skin biopsies of the patient [75]. The cells grow in vitro up to confluence
as a multilayered keratinocyte sheet, which has to be separated from the culture flask
by enzymatic treatment. However, the CEA are quite fragile because they lack any
mechanical support. Procedures for their production are quite labor consuming and
require at least three weeks. Unfortunately, the healing of these autografts is often
worse than that of split skin grafts. Even months after transplantation, the epidermis
can still get lost due to blistering [72,74].
Ideally, wound dressings, which are used for temporary coverage of skin injuries
have to adhere to the wound site and must be porous enough to allow oxygenation
and removal of wound exudates. Furthermore, they must prevent dehydration and
infection of the wound [74,76]. It has been recognized that polymer membranes
meet the majority of these requirements. If keratinocytes could be cultured on
membranes they would combine the properties of a wound dressing and at the
same time being a carrier for cell transplantation. So far, a limited number of
approaches have been put forward to apply membranes in tissue engineering of
epidermis. One of the examples is the application of a polyurethane membrane
(HydroDerm, HD Innovative Technologies, Ltd), which has been developed
purposely as a wound dressing [77]. Polyurethanes have been known for their
good biocompatibility in blood-contacting applications. Also, their mechanical
properties can be varied depending on their chemical composition from stiff to
elastic materials [78]. Therefore, they are applicable as wound dressings [77]. In the
study described here, keratinocytes were seeded in vitro on the polyurethane
membrane. It was observed that cell growth was delayed when compared to
standard tissue culture polystyrene. However, under in vitro culture conditions,
the formation of a multilayered epidermis was observed, indicating that the
233
234
membrane supports the differentiation of cells [77,79]. If the membrane was placed
upside-down at early pre-confluent stages of cell growth on the wound of an animal
model, then a partial reformation of an epidermis was obtained. To achieve this,
cells must be able to transmigrate from the membrane surface to the wound bed to
form a basal layer of keratinocytes [80]. The delayed growth of keratinocytes might
be attributed to the chemical composition of the polyurethane membrane, which
was not especially adopted to promote keratinocyte attachment and growth.
To adjust the chemical composition of membranes to the requirements of specific
cells in different applications, we have developed a number of membrane types,
which are based on copolymers of acrylonitrile [81–83]. The co-monomers applied
here were acrylonitrile, hydrophilizing, nonionic N-vinylpyrollidone (NVP 5, 20, and
30 mol%), aminoethylmethacrylate or aminopropylmethacrylate (APMA, both about
1 mol%), and sodium methallylsulfonate (NaMAS, about 2 mol%). Copolymers
were compared with the poly(acrylonitrile) (PAN) homopolymer. Membranes were
prepared by phase inversion and possessed porosities in the ultrafiltration range.
Interestingly, membranes prepared from copolymers AEMA or APMA promoted
growth of keratinocytes in vitro [83]. Figure 8.5 shows a comparison of scanning
Fig. 8.5 Comparison of cultures of keratinocytes on PAN
(a and c) and APMA (b and d) with scanning electron
microscopy after 7 days of culture with low (a and b) and
high (c and d) magnification. Note that the aminefunctionalized copolymer APMA supports rapid monolayer
formation (b) with close cell–cell-contacts (d) compared
to PAN where no monolayer of cells was established (a and c).
electron micrographs of keratinocytes cultures on PAN (a and c) and APMA (b and d).
It is visible that cells on APMA cover the entire membrane surface and have close
intercellular contacts. In contrast, keratinocytes cultured on PAN did not reach a
confluent monolayer. Also, intercellular contacts between the cells are less well
expressed. Moreover, it was possible to show in organotypic coculture models that
keratinocytes placed upside-down on a dermal equivalent (composed of a collagen gel
with embedded fibroblasts) were able to transmigrate from the membrane into the
artificial wound bed and started to grow and differentiate there, which means to build
up a neo-epidermis [82]. Figure 8.6 shows a cross section through membranes with
neo-epidermis and dermal equivalent stained with hematoxilin-eosin. It is shown
that the APMA membrane (a small and c high magnification) promotes the formation of a neoepidermis in contrast to a hydrophilic NVP membrane (c small and d
high magnification). To explore further the possibility that polymer membranes with
amine groups may support growth and differentiation of keratinocytes, we developed
a blend made of poly(ether imide) (PEI) and poly(benzimidazole) and prepared
membranes by phase inversion method [83]. PEI has the advantage of good
membrane formation properties and high-thermal stability [84], which makes this
polymer interesting for biomedical applications. It was found in this study, that the
membrane could support the growth of keratinocytes and may be used as a temporary
carrier of keratinocytes as well. During our studies we also discovered the possibility
Fig. 8.6 Organotypic culture of keratinocytes
immobilized on APMA (a and c) and NVP 5
(c and d) membranes placed upside-down
on a collagen gel with embedded fibroblasts
for 14 days. Constructs were fixed and
cryosections were stained with hematoxylin/
eosin. It is visible that APMA supports the
formation of a neoepidermis (a and c)
indicated by the thick blue layer beneath the
membrane compared to NVP 5 (b and d).
(a and b) Magnification 10, (c and d)
magnification 20.
235
236
to functionalize PEI by a simple wet chemical procedure using different diamines or
polyamines [85]. It was observed that covalent binding of poly(ethylene imine)
(PETIM) of lower molecular weight may provide advantages to keratinocyte culture,
such as rapid establishment of a pre-confluent cell layer on the surface and subsequent
release of the cell layer from the membrane support [86]. While the first organotypic
culture models have shown advantageous effect of the polymer compositions and also
first animal experiments have been carried out, clinical studies are still necessary.
Beside synthetic polymers, natural polymers or combinations with polyesters have
also been used to prepare membranes for the transplantation of keratinocytes. The
advantage of using natural polymers or polyesters, such as poly(e-caprolactone)
(PCL), is that materials are degradable, which avoids the necessity to remove the
material at later stages of medical treatment. Hyaluronic acid (HA) is a major
component of the extracellular matrix in connective tissues and plays a promoting
role in wound healing [87]. However, HA is water soluble and forms hydrogels, which
have poor mechanical properties. Therefore, HA has been esterified with benzyl
groups and crosslinked, which decreases its solubility in water and improves its
mechanical stability [88]. This material has been used to prepare different kinds of
scaffold materials for engineering of skin, cartilage, and other tissues. Thereby,
esterified hyaluronic acid (HYAFF) has been used as a delivery system for keratinocytes. The commercial product is produced by Fidia Advanced Biopolymers and
denoted as HYAFF-11 or Laserskin. It is a membrane that additionally contains a
regular pattern of large pores generated by a laser. Keratinocytes are seeded on both
sides of Laserskin and may migrate through the pores to cover the underlying dermal
structures of the skin. Laserskin is applied mainly for the treatment of nonhealing
skin ulcers with good success [89]. Also, membranes made of blends from collagen
and PCL are used for tissue engineering of epidermis and dermis. One of the examples
is a membrane based on a collagen scaffold, which is prepared by a freeze drying
process and subsequent impregnation with PCL solution [90]. The resulting scaffold
represents a microporous membrane with large pores in the range of about 40 mm. It
was shown that both dermal fibroblasts and keratinocytes were able to grow on this
scaffold. A specific coculture model was developed where the membrane separates the
fibroblast building the dermal component and keratinocytes making the epidermis
[90,91]. Such a construct can be used in the treatment of deep-skin defects when not
only epidermis but also the dermal part of the skin must be replaced. Recently, pure
PCL membranes have also been suggested as a tool for tissue engineering of skin [92].
These membranes were prepared by solution casting, subsequent biaxial stretching,
and laser modification to obtain large pores for transmigration of cells. First experimental work with cell cultures could show that these membranes can be applied as
culture vehicle for both fibroblasts and keratinocytes [92,93].
8.2.3
Tissue Engineering of Bone
Bone (also called ‘‘osseous tissue’’) serves multiple functions in the body, including
support of body structure, protecting internal organs, acting as major calcium
reservoir and facilitating movement (in conjunction with muscles). One of the most
common bone diseases is bone fracture. Autologous bone grafting remains the gold
standard for the treatment of bone defects. The main advantages of autologous bone
grafting are its viability, immunocompetence, osteogenic activity, and high incorporation rate. However, donor site morbidity, long operative time and bleeding, risk
of infection, rapid graft resorption, and limited availability are common restrictions
of this procedure [94]. The applications of polymers for bone tissue repair and
reinforcement, for regeneration of cartilage associated with bones, for helping in
partial replacement of bones by metallic parts, and as carriers of antibiotics to the
infected bone tissues have led to the development of bone tissue engineering [95]. By
the implantation of dental and orthopedic devices and bone-substitute materials,
bone tissue engineering has been practiced since a long time in the repair of bone
fractures [96].
In 1956, Bassett et al. first used a physical barrier to prevent the migration of
unprofitable cells to the wounded region and to enable the proliferation of useful
cells [97]. Since 1960s, polymeric membranes have been used in experimental
animals to facilitate healing of bone defects [98]. In 1982, Nyman et al. attempted
the first utilization of guided tissue regeneration (GTR) for the periodontal surgery
[99]. The basis of GTR technique depends on the use of a mechanical barrier placed
between the root surface and the full-thickness periodontal flap. After the placement of a Millipore filter membrane, a histological analysis demonstrated new
cementum formation and the insertion of Sharpey’s fibers on the previously
diseased root surface. Dahlin et al. [100] originally applied this technique to bone
regeneration in a bone loss area to establish the concept of ‘‘osteopromotion.’’
Based on Dahlin’s technique, Buser et al. [101] proposed the term of ‘‘guided bone
regeneration’’ (GBR), which aims to promote bone augmentation by a barrier
membrane. GBR membranes also have an important function that encourages
bone growth. GBR membranes are occasionally utilized with dental implant [102]
or bone-grafting materials [103]; a typical membrane set-up for bone defects is
illustrated in Figure 8.7.
A number of membrane forms and materials have been applied in animals and
humans by GTR or GBR technique to repair bone diseases. The below table is a list of
membranes adopted in bone tissue engineering (Table 8.2).
As an effective barrier in GTR or GBR, utilized membranes can be divided into two
categories: resorbable and nonresorbable. The most popular nonresorbable membrane utilized for GTR or GBR is extended polytetrafluoroethylene (ePTFE) [105],
which is a biological inert material and can be safely applied clinically. Many studies
have demonstrated successful bone regeneration by using ePTFE alone [123,124] or
in combination with filling material for bone augmentation and regeneration [125].
In a bilateral mandibular defect, the use of high-density PTFE membranes to
facilitate guided bone regeneration in the rat was examined [126]. On both the
medial and lateral aspects of the mandible, the experimental side was covered with
PTFE membrane, while the opposite side served as a control. At 10 weeks postoperatively, complete ossification was noted on the PTFE treated side, while the
control sides displayed little osseous regeneration.
237
238
Fig. 8.7 Schematic presentation of
membranes set-ups in a critical-size
segmental bone defects in the sheep tibia.
A. single external microporous membrane;
B. two microporous membranes placed
externally and internally in the defect as a
‘‘tube-in-tube’’ construct. The microporous
membrane used here is poly(L/DL-lactide)
80/20 %. M: membrane; OD: osteoperiosteal
defect in the diaphysis of tibiae which was
then stabilized with a bilateral AO external
fixator. Revised from Gugala, Z. and
Gogolewski S. (2002) Injury, International
Journal of Care Injured, 33, S-B-71–6.
Tab. 8.2
Overview on use of membranes in bone tissue engineering.
Year
Scientists
Membrane
Reference
1960
1961
1982
1984
1987
1987
1987
1988
1992
1992
1993
1995
1995
1999
1999
2000
2004
2005
2005
Lindhorne
Goldhaber
Nyman et al.
Nyman et al.
Yaffe et al.
Pitaru et al.
Blumenthal et al.
Magnusson
Gogolewski et al.
Wanzhang et al.
Vuddhakanok et al.
Jansen et al.
Nishimura et al.
Ishikawa et al.
Veronese et al.
Wang et al.
Lee et al.
Desai et al.
Liao et al.
[98]
[104]
[99]
[105]
[106]
[107]
[108]
[109]
[110]
[111]
[112]
[113]
[114]
[115]
[116]
[117]
[118]
[119]
[120]
2005
Fujihara et al.
2006
Kuo et al.
Polyethylene tube
Millipore filter
Millipore filter
Gore-tex membrane (ePTFE)
Collagen membrane
Collagen membrane
Collagen membrane
Poly-L-lactic acid
Polyurethane
Calcium alginate film
DL-PLGA membrane
PLA/hydroxyapatite membrane
Cementum-impregnated gelatine membrane
Alginate membrane
Polyphosphazene membrane
Polyesterurethane membrane
Polycarbonate membrane
Alumina membrane
Hydroxyapatite/collagen/PLGA
composite membrane
Polycarprolactone/CaCO3
composite nanofiber
Chitosan membrane
[121]
[122]
Despite the successful applications of nonresorbable membranes in bone tissue
engineering, they must be removed by a second operation. This additional surgical
trauma is a negative factor both to the patient and to the newly formed tissue. Hence,
resorbable membranes have been developed for bone tissue engineering to avoid a
second surgical intervention, and have become widely used in conjunction with GTR
and GBR [127,128]. There are two types of resorbable membranes, namely synthetic
polymers and natural biomaterials. Collagen membranes have been introduced as
resorbable material in 1987 [106]. Since that time numerous applications of collagen
membranes in GTR have been developed. For example, Taguchi et al. [127] used a
collagen membrane for guided bone regeneration in rat maxillae. Standardized
artificial bone defects (1.0 1.0 2.5 mm3) were covered with a collagen membrane (Bio-Gide, composed of porcine type I and type III collagen fibers). After four
weeks it was visible that membrane-associated and cavity-derived bone had completely filled the defect. The Bio-Gide membrane appeared to have so-called osteoconductivity, resulting in a well-augmented alveolar ridge. While collagen membranes
have excellent cell affinity and biocompatibility, they are normally weak in strength
and difficult to manipulate. Therefore, polylactides or copolymers of glycolide and
lactide been found widespread applications for bone fixation [94,109,112]. However,
their acid degradation products may lead to an inflammatory response. Moreover, the
cell affinity of these polymers is rather low if compared to a collagen membrane.
Hence, to meet the demands of resorbable membranes in GTR or GBR both in
mechanical property and biocompatibility, composite membranes were developed.
For example, a three-layered degradable composite membrane was prepared,
which contained one layer of 8 % nano-carbonated hydroxyapatite/collagen/poly
(lactic-co-glycolic acid) (nCHAC/PLGA), one layer of pure PLGA membrane, and
one layer of 4 % nCHAC/PLGA [121]. Results revealed that the addition of nCHAC
increased the biocompatibility and osteoconductivity, while the composite membrane
had sufficient mechanical strength. It was concluded that polymer-nanoparticle
composite membranes are promising materials for bone tissue engineering application due to their superior mechanical properties, improved durability, and surface
bioactivity [129].
For membranes used in bone tissue engineering, the healing efficiency depends
on the surface chemical and physical properties of membranes. Usually the chemical
properties are related to surface chemical groups that control surface free energy,
surface charge, and hence the ability to support attachment and function of cells (see
Section 2.2). Gogolewski et al. [130] tested the effects of polylactide membranes on the
healing process of rabbits’ diaphyseal segmental defects. The polylactide membranes
were made either of poly(L/D-lactide) or poly(L/DL-lactide). After one year, they found
in most cases that both membranes led to complete bone regeneration in the defects.
Obviously, the differences in quantities of stereo isomers of lactides did not have a
significant effect on osteoblasts. They further investigated the effects of plasma
treatment of polylactide membranes on cellular response of osteoblasts. The membranes were subjected to low-temperature oxygen, ammonia, or sulfur dioxidehydrogen plasma treatments. The results revealed that for all treated membranes, the
attachment and growth of osteoblasts were greater in comparison to untreated
239
240
polylactide. The treatment with ammonia plasma, which generates different kinds of
amine and amide groups on the surface, was found to be most efficacious [131].
On the other hand, surface properties of membranes are also related to the surface
morphology such as texture of surface, presence of pores, pore structure, pore size,
and pore distributions. The effect of surface topography on osteoblast cultures was
tested using different polyesterurethane membranes [117]. The membranes prepared had either a fairly dense and smooth structure (membrane A), or a rough and
sunken surface supported by a porous sublayer (membrane B), or a porous surface
with particles that were supposed to be generated by the polymer crystallization
(membrane C). Rat osteoblasts culture experiments revealed that on membrane C
there was a greater number of cells with better spreading and a flattened appearance,
which may be due to the presence of particles on the membrane surface. Degradation
experiments further revealed that the degradation of membranes was obviously
affected by the membrane’s surface morphology; rough surface could decrease the
degradation rate, especially for those with particles on the surface. Due to its relatively
fast degradation during early stages of healing, membrane A was found to be not
suitable for the repair of bone defects [117]. For polycarbonate membranes with
different pore sizes (pore diameter from 0.2 to 8.0 mm), on the other hand, the
response of MG63 osteoblast-like cells was quite different. For membranes with pore
a diameter of 0.2–1.0 mm, cells adhered and spreaded easily. When pore size
increased, cells became spherical in shape, which was unfavorable for their functional activity [118]. The overall membrane porosity seems to be also a useful tool to
manipulate osteoblast adhesion, proliferation, and function.
Until now the effect of membranes in bone regeneration is not completely understood. They may depend on the location of bone defects, their size, the vascularization,
and also the type of animal. One model, namely the rabbit radius model, has suggested
that membranes have a number of functions namely to concentrate and store the
osteogenic components (growth factors) released from the bone ends and bone
marrow in the ‘‘medullary cavity’’ formed by the membrane (i), to protect the site
against the massive ingrowth of soft fibrous tissues into the defects (ii), and to serve as
a guiding scaffold migrating osteogenic cells into the defect (iii) [132,133].
8.2.4
Further Tissue Engineering Applications of Membranes
Besides the abundant research on membranes for tissue engineering of skin and
bone, a broad range of other applications of membranes have also been envisaged.
Table 8.3 gives an overview on membrane applications besides skin and bone tissue
engineering.
For example, membranes were also tested for regeneration of components of the
nervous system [134,135]. A typical example is the gelatine composite membrane
prepared from tricalcium phosphate and glutaraldehyde-crosslinked gelatine [136].
Several neurotrophic factors such as nerve growth factor (NGF) and brain-derived
neurotrophic factor (BDNF) were covalently immobilized using carbodiimide. The
effects of these membranes on the regeneration of the sciatic nerve in rats across a
Tab. 8.3
Further applications of membranes in tissue engineering.
Scientists
Membranes sources
Application
Chen et al.
Qin et al.
Gelatine-tricalcium phosphate
Polylactide/polytrimethylene
carbonate
N,O-Carboxymethyl chitosan
Crosslinked hyaluronate/collagen
Polyaniline/gelatine
Nerve guidance channel
[134]
Preventing postoperative adhesion [138]
Zhou et al.
Shih et al.
Li et al.
Hsiue et al. Gelatine
Riboldi et al. Polyesterurethane
Vernon et al. Collagen
Zhu et al.
Collagen-g-poly
(DL-lactide-co-glycolide)
Preventing postoperative adhesion
Preventing peridural adhesion
Potential for cardiac or
neuronal tissue constructs
Neural retinal transplantation
Potential for skeletal muscle
tissue engineering
Cell alignment or cell perfusion
Esophagus tissue engineering
Reference
[142]
[143]
[147]
[148]
[149]
[150,151]
[152,153]
critical size defect of length 10 mm were investigated. The gastrocnemic muscles
begin to degenerate when the sciatic nerve is severed. Because they regain their mass
proportional to the amount of re-innervation [137], the weight ratio of reinnervated
gastrocnemic muscle has been used as a means to evaluate peripheral nerve repair.
The results showed that the average size of newly grown nerve axons was the greatest
for membranes immobilized with NGF, while the BDNF-immobilized membrane
had a much higher gastrocnemic muscle weight ratio than the blank membrane.
Although the effects of different growth factors immobilized onto gelatin membranes seemed to be quite complex, which could not be simply illustrated by the
sciatic function index, these types of membranes may provide a solution for
peripheral nerve repair.
Postoperative adhesion formation (e.g., the formation of adhesions between gut
and the surrounding tissue) is a serious surgical problem, which may lead to
complications for the patient such as infection and chronic pain [138–140]. Polymeric
membrane covers maybe a useful method to prevent adhesions after visceral surgery
[141–143]. Using a blend membrane prepared from polylactide and poly(trimethylene carbonate) (PLA/PTMC), intestine studies were conducted on the ascending
colon in rabbits’ abdomen, which was covered by the blend membrane to investigate
their effects on adhesion prevention. Histological observations showed that the PLA/
PTMC blend membrane could act as a physical barrier before their degradation. The
distribution of adhesion scores in the intestinal models revealed that PLA/PTMC
blend membranes were able to prevent postoperative adhesion formations [138].
A novel research direction is the combination of electrical conductive polymers
such as polypyrrole or polyaniline with biodegradable materials [144–147]. Zhang
et al. [145] prepared a blend membrane from polypyrrole and polylactide and tested its
in vivo tissue response by subcutaneous implantation in rats. Both alkaline
phosphatase activity and acid phosphatase activity were used to evaluate the inflammatory response to the implanted membranes. The alkaline phosphatase activity was
similar to that of the pure PLA membranes, indicating that the addition of polypyrrole
241
242
did not intensify the inflammatory response. Acid phosphatase activity for materials
containing polypyrrole was similar or even lower than that of the membrane without
polypyrrole. The histological findings were in agreement with the enzymatic results,
indicating that the polypyrrole and PLA composite membranes provoked no abnormal tissue response [145]. Cultures of fibroblasts on the composite membranes were
conducted based on the hypothesis that the composites can be used to modulate the
cellular by direct electrical stimulation [146]. Membranes experienced mediumrange electric current stimulations, which promoted cell growth in comparison to
nonstimulated controls. Hence, these composite membranes may be a potential
substratum for the arrangement of electrical responsive cells or can be used as
a scaffold for electrically stimulated tissue regeneration. Another inherently conductive polymer, polyaniline was blended with gelatine. The blend was used for
electrospinning of nanofibers to prepare a microporous membranous scaffold [147].
Cultures of H9c2 rat cardiac myoblast cells showed that the scaffold supported
attachment, migration, and proliferation of cells. This implies the possible application of this membrane-like material based on nanofibers for tissue engineering of
functional cardiac, cardiovascular, and neuronal tissue constructs.
A further example of application of membranes in tissue engineering is the
applications of a gelatine sandwich-like membrane as implant in neural retinal
transplantation [148]. Using two gelatine membranes sandwiched by a retina cell
sheet, the retina-gelatine membrane construct was implanted into rabbits. Histological results revealed that the newly developed implant survived in a proper
orientation of the retinal layer, which indicates the potential of this approach for
tissue engineering of retina. Membranes were also tested for tissue engineering of
skeletal muscle [149]. Collagen membranes with different morphologies were tested
for their ability to control muscle cell alignment [150,151]. A poly(L-lactideco-caprolactone) membrane was also used for porcine esophageal cells cultures;
the mitochondrial activity assay, and cell morphology indicated that this kind of
membrane may be potential scaffold materials for esophagus repair [152,153].
Overall, there is a multitude of established and future applications of membranes
in the emerging field of tissue engineering.
8.3
Membranes in Biohybrid Organ Technology
8.3.1
Organ Failure and Biohybrid Organ Technology
The term biohybrid (artificial) organ was first introduced by Chick et al. in 1975 to
denote a device containing living cells separated from the body by a synthetic material
to substitute the function of an organ or specific tissues [154]. Today both terms
bioartificial organ and biohybrid organ are used. Biohybrid organs can resemble an
implantable device for the treatment of chronic diseases, such as secretion of insulin
in diabetes [155], human growth factor in dwarfism [156], and Parkinson’s disease
8.3 Membranes in Biohybrid Organ Technology
[157,158]. These applications usually require that cells are located in a compartment
separate from the surrounding tissue to maintain the structural integrity of the
immobilized cell mass. Moreover, these cells are typically allogenic or xenogenic in
origin, which requires their isolation from the immune system of the recipient.
Membranes have been identified as useful tools providing protection from host
immune responses but also exchange of substances and oxygenation of immobilized
cells [159].
The requirements to replace a specific tissue or organ cover a large range, which is
determined by the specific activity (e.g., secretion of hormone) per cell and quantity of
the agent required by the body [159]. Because of the relatively low quantities of
hormones needed by the body, a low amount of cells of about 106–107 is necessary for
replacement of most endocrine functions. This corresponds to a cell volume of about
10 mL. In contrast, implantation of Langerhans islets cells to treat diabetes requires
already 109 cells, which matches to about 1 ml volume. Full replacement of liver
function however needs between 1010 and 1011 cells, which is more than 100 mL cell
suspension [159]. It is quite clear that replacement of endocrine functions can be still
achieved by an implantable device while liver or kidney replacement therapies
require extracorporal machines.
Different configurations of implantable devices have been developed, such as
microcapsules, hollow fiber, or planar diffusion chambers with flat membranes [159].
A majority of developments aim to encapsulate the cells into hollow spheres, which
can be implanted, for example, into the peritoneal cavity. Microcapsules are advantageous regarding the exchange of solutes due to their large surface to volume ratio.
However, their application can be problematic if the device must be later retrieved. To
overcome this limitation hollow fiber membranes have been also used to encapsulate
cells with endocrine functions. The hollow fibers were introduced into blood vessels
or in the lumbal region of the backbone, from where they can be also removed. While
the current review is focused on application of membranes in biohybrid organs for
blood detoxification, others have reviewed application of cells in implantable devices
for the replacement of endocrine functions in more detail [159,160]. Table 8.4 gives an
Survey on implantable biohybrid organs with type of organ, type of membrane and
membrane geometry.
Tab. 8.4
Type of organ
Type of membrane
Heart
Liver
Liver
Pancreas
Pancreas
Pancreas
Polyurethane
Alginate/poly-L-lysine/alginate
Regenerated cellulose
Alginate
Alginate
Alginate-cellulose sulfate complexed
with poly(methylene-co-guanidine)
Polyacrylonitrile-sodium
methallylsulfonate copolymer
Alginate-polylysine
Pancreas
Parathyroid
Membrane
geometry
Reference
Flat membrane
Microcapsules
Hollow fiber
Hollow fiber
Microcapsules
Microcapsules
[59]
[161]
[162]
[163]
[164]
[165]
Hollow fiber
[166]
Microcapsules
[167]
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244
overview on implantable biohybrid organs with type of organ, geometry of the device
and type of membrane.
The separation of toxins from blood is crucial for the treatment of end-stage renal
disease (ESRD), acute renal failure, and also fulminate liver failure. The current
majority of extracorporal support and replacement therapies of these organs is based
on physical processes that comprise dialysis, filtration, and adsorption [2,4–7].
However, kidney and liver represent organs with multiple functions. Kidney has
basically a filtration function, but also a re-adsorptive and endocrine function that
cannot be replaced by artificial organ technology [168]. Likewise, the numerous tasks
of the liver, such as conversion of toxins generated by the metabolism, removal of
xenobiotics, protein synthesis, and so on, cannot be substituted by hemofiltration or
adsorption technologies alone [169,170]. Hence, it is not surprising that the conventional organ replacement therapies are associated with an increased morbidity
and mortality, which is known from long-term treatment with hemodialysis in ESRD
and particularly in acute renal or in fulminant hepatic failure [171–173]. To date,
organ transplantation seems to be the most successful therapy for acute and chronic
failure of internal organs like heart, kidney, liver, lung, and so on. However, there is an
increasing gap between the number of donor organs available and the patients on the
waiting list. Because of the shortage in donor organs, the advent of biohybrid organs
may be a solution of the described problem because they may support failing organs,
which lead to organ regeneration in acute liver and renal failure or may bridge the
time to find a suitable transplant for the patient.
Kidney and liver mainly consist of connective tissue, endothelial, and epithelial
cells. Epithelial cells that represent the functional units, such as proximal tubule cells
in the kidney or hepatocytes in the liver, have a polar organization where the
cytoplasm is separated into an apical and a basolateral region. Main morphological
characteristics of epithelial cells are the presence of tight junctions and microvilli on
the apical cell membrane, which are both related to transepithelial transport
processes [174]. Tight junctions separate the basal and apical region controlling
the exchange of solutes along or against concentration gradients. Microvilli on the
apical cell surface increase the surface area for the uptake of substances and are an
important feature of many epithelia as well. Typical epithelia have a planar structure
with an underlying complex of basal membranes composed of different extracellular
matrix proteins, such as collagen IV, fibronectin, laminin, and so on [175]. Figure 8.8
shows the typical morphology of kidney epithelial cells with presence of a tight
junction between cells and microvilli on the apical region obtained by transmission
electron microscopy (TEM).
Biohybrid liver and kidney resemble a combination of immobilized epithelial cells
on a suitable type of carrier, which is in many cases a polymer membrane in a
bioreactor (see, e.g., [176]). Organ cells in biohybrid organs have to make an intimate
contact with the surface of the membrane but also to develop close cell–cellconnections, which is a prerequisite for their survival and high functional activity
(see also Section 1.2). On the other hand the blood, which has to be detoxified will
contact the other side of the membrane and may not become activated by the
synthetic material. Finally as a third important requirement, the transport properties
Fig. 8.8 Transmission electron micrograph of
kidney epithelial cells with typical
ultrastructural features such as microvilli on
the apical cell surface (arrows) and tight
junctions (asterisk) between cells.
of the membrane must be adjusted to allow a sufficient exchange of oxygen, small
molecular weight solutes, such as toxins, smaller proteins, but must also protect the
immobilized organ cells from the immune system of the recipient [159,177,178].
Figure 8.9 shows a simplified set-up of a biohybrid organ to emphasize the complex
requirements to membranes in this application.
8.3.2
Biohybrid Liver
Acute (or fulminant) hepatic failure after intoxication or viral infections is characterized by defective blood protein and clotting factor synthesis, gluconeogenesis,
urogenesis with impairment of plasma detoxification, neurological complications
(often associated with cerebral edema), and finally multiorgan failure [179].
Current therapies based on a multitude of measures including hemodialysis or
Fig. 8.9 Schematic set-up of biohybrid organs for blood
detoxification highlighting the crucial role of membranes as
support for epithelial cells, as contacting surface for blood
components and barrier controlling the exchange of
solutes between the different compartments.
245
246
hemofiltration have a very poor outcome with survival rates as low as 20 %. It has been
stated that orthopic liver transplantation is the best medical therapy with a survival
rate higher than 90 % after one year [179]. In light of the fact that access to donor
organs is limited for a specific patient due to not only availability of organs but also
immunological mismatch between donor and recipient, development of biohybrid
liver support systems bears great hope to bridge the time to find a suitable transplant
or to support even regeneration of the damaged liver [179]. Biohybrid liver support
systems replace the function of the failing liver through the actions of hepatocytes
cultured in a bioreactor module. Different concepts of biohybrid liver devices have
been realized, which have been reviewed in more detail recently [176,181]. Among
them membrane-based systems may provide certain advantages due to their ability to
separate donor hepatocytes from the immune system of the recipient (i), to provide a
substratum for cell attachment (ii), and the large surface-to-volume ratio allowing a
compact bioreactor design (iii). Critical issues, however, are the ability of membranes
to provide adequate attachment for hepatocytes and their ability for bidirectional mass
transfer. The latter requirement must be fulfilled to provide nutrients to sustain cell
viability and export of certain cellular products. Moreover, it is known that many toxins
in blood are albumin bound and contact hepatocytes directly after passage through the
fenestrated endothelium in the liver [176].
At present the majority of membranes used in biohybrid liver support systems
have been those developed for other biomedical applications such as hemodialysers
or oxygenators [170,182]. These membranes were selected because of their relative
good biocompatibility in contact with whole blood or plasma. The molecular cut-off of
these membranes is chosen to prevent exposure of cells in the bioreactor to
components of the immune system, such as immunoglobulins (about 160 kDa)
or factors of the complement system (>200 kDa), but alternatively permeation of
serum albumin (about 70 kDa) shall be still possible. Therefore, some groups have
selected membranes with a molecular cut-off of around 100 kDa [176]. Conversely,
also membranes with a lower cut-off were used to prevent the entrance of cellular
proteins from donor cells into the blood stream of the patient, which is important in
cases when porcine or cancer hepatocytes are applied in the bioreactor [180,183,184].
Few groups have selected membranes with microfiltration properties showing that
immunoisolation of porcine cells is not an absolute requirement [185]. Table 8.5 gives
an overview on biohybrid liver support systems, which underwent clinical testing and
application.
It must be noted that the majority of membranes used in biohybrid liver systems
have not been optimized regarding the promotion of attachment and function of
hepatocytes. The majority of current biohybrid liver devices are only used for
treatment sessions of 6–8 hours, each patient receiving multiple sessions. The main
reason for this limitation is that the primary hepatocytes have unstable function, and
lose viability and essential enzyme activities with time in culture [186]. Particularly,
quite hydrophilic membranes based on cellulose and its derivatives have shown to be
less suitable for primary hepatocytes [182]. Membranes with limited compatibility for
hepatocytes can be coated with collagen or fibronectin, which improved the functional activity and survival of hepatocytes [187]. The development of membranes with
Tab. 8.5 Overview on biohybrid liver systems using hollow fiber membranes undergoing clinical
testing.
Name of devices/
company
Type of cells
Type of
membrane
Cut-off/
pore size
Reference
ELAD Amphioxus
Cell Technology
HepatAssist
Circe Biomedical
MELS/Charité
Humboldt Univ
C3A hepatoblastoima
cells
Porcine hepatocytes
Cellulose acetate
70 kDa
[192]
Polysulfone
0.2 mm
[193]
Porcine or human
hepatocytes
100 kDa
oxygenation
80 kDa
[188]
BLSS/Excorp
Biomedical Inc.
Porcine hepatocytes
Combination of
polyamid,
polypropylene,
polysulfon
Polysulfone
100 kDa
[191]
polymer compositions, which specifically support the attachment, function, and
viability of cultured hepatocytes, would significantly improve the length of treatment
time possible for patients with acute liver failure. Catapano and co-workers have
shown that cell adhesion and urea synthesis increased with increasing wettability and
surface roughness of polypropylene membranes [188,189]. The chemical composition of the membranes also markedly influences the properties of the cultured cells,
and in general, adhesion and liver function is better maintained on moderate
hydrophilic than on hydrophobic membranes [28,190]. A few groups have also
tested other polymer membranes than those developed for hemodialysis or blood
oxygenation successfully, such as polyetherketone coated with proteins as a substratum for hepatocytes [191]. We have developed a range of copolymers based on
acrylonitrile, which have been already introduced in Section 2.2 and investigated the
adhesion, growth, and functional activity of C3A hepatoblastoma cells as well. This
cell line has been used also in clinical studies with the extracorporeal liver assist
device (ELAD) mentioned above [184]. We found that particularly hydrophilic
copolymers, which contained 20 mol% N-vinylpyrollidone (NVP 20), allowed moderate attachment of these cells and supported cell–cell adhesions, which seems to
stimulate the functional activity of cells [192]. During these studies we also explored
the impact of pore size in the ultrafiltration range on C3A cell attachment, growth,
and function. We used membranes based on NVP 20 and found an inverse relationship between initial cell attachment, which was higher on membranes with larger
pores, while cell growth was enhanced on membranes with smaller pores [193].
Figure 8.10 shows examples of this work as a comparison of membrane surface
morphology between 6 and 12 nm (a and d) pore size visualized by scanning electron
microscopy (SEM). It is obvious that membranes with 12 nm pores possess a higher
surface roughness. Studies on cell adhesion with staining of viable cells show that the
size of aggregates was bigger on membranes with 12 nm (e) than on 6 nm pores (b).
Moreover, it was also observed that intercellular connections visualized by immunostaining of E-Cadherin and confocal laser scanning microscopy were more
expressed on membranes with 12 nm (f) than 6 nm (c) pores, which indicates a
247
248
Fig. 8.10 Comparison of C3A
hepatoblastoma cell growth and morphology
on membranes with different pore size made
of NVP 20. Upper panel shows membranes
with a mean pore diameter of 6 nm (a),
vital staining of cells after 4 days of culture (b)
and expression of intercellular junctions
(E-Cadherin) visualized by
immunofluorescence microscopy (c).
Lower panel shows membranes with mean
pore diameter of 12 nm (d) and vital
staining of cells (e) and expression of
intercellular junctions. Note the larger size
of aggregates and the more abundant
expression of E-Cadherin (arrows) on 12 nm
membranes. Scale bars (a and d) 0.2 mm,
(b and e) 100 mm, and (c and f)
20 mm.
better function of these cells. However, the functional activity measured by the
activity of liver-specific P450 enzymes was maximal at intermediate pore size [193] as
shown in Figure 8.11. The blood compatibility of these copolymers is also of interest
for the potential application in biohybrid liver systems. It was found in parallel
investigations that introduction of N-vinylpyrollidone improves the hemocompatibility of membranes considerably, when compared to polyacrylonitrile as standard
membrane for hemodialysiers [194]. In further studies with primary hepatocytes, we
also explored the effect of other co-monomers on survival and functional activity of
cells. Here, a polyacrylonitrile membrane containing acrylamido-2-methyl-propansulfonic acid as co-monomer was found to be the most compatible option for
maintaining primary hepatocyte morphology and function. It maintained viable,
functional cells for at least 16 days in culture [195]. Overall, tailoring membrane
properties regarding chemical composition and porosity has a great potential to
improve further the functional activity and to prolong the survival of hepatocytes in
the biohybrid liver systems.
Fig. 8.11 Comparison of growth of C3A hepatoblastoma
cells in dependence on mean pore diameter of NVP 20
membranes. Please, note that initial number of cells after
two days is higher on membranes with larger than on
smaller pores. However, cell numbers are higher after 7
days on membranes with smaller pores indicating better
growth on smoother surfaces.
8.3.3
Biohybrid Kidney
Despite the progress in medical technology, patients with acute kidney failure still have
a poor survival prognosis [196]. Also, patients having end-stage chronic failure treated
with hemodialysis suffer from an increased morbidity and have also a limited life
expectancy [197,198]. Reasons for the increased morbidity and mortality of these
patients are numerous uremic toxins that accumulate in the body [199]. Kidney
transplantation is currently the best therapy, but the limited number of donor organs
is the cause for long waiting lists of patients. Since the beginning of the twentieth
century researchers worked on a device to replace the function of the kidney. In fact, the
kidney was the first organ whose function was substituted by an artificial organ – the
hemodialysis [200]. However, hemodialysis therapy is not a long-term solution since
about 50 % of end-stage renal failure patients develop dialysis-related amyloidosis [201]
and other serious health problems. The main reason for this situation is that hemodialysis can only partly replace the filtration function of the kidney glomerulus but does
not substitute many other renal functions, such as transport, metabolic, endocrine, and
immune activities [201]. A biohybrid kidney must overcome these drawbacks combining the previously established functions of artificial kidney, such as filtration and
adsorption, with the activity of cells derived from the kidney tubule system [202].
In 1987 Aebischer et al. suggested for the first time to combine hollow fiber
membranes with kidney tubular cells to allow for the transport of substances
across the cell-attached membranes [203]. This concept was named bioartificial
kidney [204]. The principal set-up of a biohybrid kidney has been realized by Humes
et al. in 1999 [205], which is shown in Figure 8.12. It basically consists of a commercial
249
250
Fig. 8.12 Principal set-up of a biohybrid kidney as
implemented by Humes and coworkers that combines a
synthetic hemofilter, which generates an ultrafiltrate in
series with a tissue-engineered proximal tubule device for
metabolic, endocrine and immunological functions. This
figure is a reproduction from [214].
hemofiltration device, which generates an ultrafiltrate like the kidney glomerulus.
This primary urine in analogy to the human kidney is transported to a proximal
tubule device (RAD), which consists of a hollow fiber bioreactor containing kidney
epithelial cells. These cells provide the same active transport processes as in the renal
tubule along with a multitude of cellular metabolic reactions. They produce some
reabsorbate, which flows across the membrane and mixes with the blood of the
patient. It has been anticipated by the group of Humes through preclinical and
clinical studies that the biohybrid kidney improves survival in acute renal and also
multiorgan failure due to sepsis [206,207].
Until now polysulfone (PSU) hollow fiber membrane modules have been used as
bioreactor for the biohybrid kidney device [208,209] although it has never been
explored in detail which type of commercial membrane is the most biocompatible for
kidney epithelial cells. Saito et al. compared the biocompatibility of cellulose acetate
and polysulfone for a porcine kidney epithelial cell line (LLC-PK) in a recent paper
and observed that polysulfone had superior properties [210]. In the same study the
authors stated that human kidney epithelial cell did not attach and grow that well on
polysulfone as they did on a polyimide membrane. Obviously, because of reduced
ability of polysulfone to support attachment, growth, and function of kidney epithelial
cells, the group of Humes used ProNectin-L to coat the membrane surface. ProNectin-L is a synthetic protein sharing the cellular attachment domain of laminin,
which is a component of the kidney tubular basement membrane [201]. Also, other
authors have used extracellular matrix proteins to improve the biocompatibility of
commercial membrane materials for kidney epithelial cells [210,211].
However, coating with extracellular matrix proteins is an additional step during the
preparation of the bioreactor and increases the risk of infections. Therefore, we have
looked for the ability of different polymers to support adhesion, growth, and function
of kidney epithelial cells [212,213]. Among them we used polyacrylonitrile and its
copolymers introduced in Sections 2.2 and 3.2 and compared them with PSU and a
commercial Millipore filter, which represents a blend of cellulose acetate and
cellulose nitrate. Investigations with flat membranes were performed with MDCK
cells, which represent a canine kidney epithelial cell line. The cell proliferation assays
revealed a significantly reduced cell attachment and proliferation on AEMA, NaMAS,
and PSU in comparison to the other membranes. The morphology of cells on the
various flat or hollow fiber membranes was studied with SEM, while the ultra
structure of cells and their contact with membranes were studied with TEM. It was
observed that confluent layers of cells existed on most of the membranes except on
PSU and NaMAS where the cells were not uniformly distributed. TEM of MDCK cells
cultured for 8 days showed typical morphological features of functional active
epithelial cells, such as tight junctions and microvilli on most of the membranes
except on PSU and NaMAS, where the cells formed much less microvilli, which
indicates a reduced functional activity of these cells. Figure 8.13 shows a comparison
Fig. 8.13 Scanning electron micrographs of
kidney epithelial cells grown on the inner
surface of hollow fibers made of polysulfone
(a) and NVP 5 (b). Note that on polysulfone
only single cells are visible, while on NVP 5
after 7 days a confluent layer of cells has been
established.
251
252
Fig. 8.14 Transmission electron micrographs
of kidney epithelial cells cultured between
PSU and NVP 5 membranes. It is visible that
cells make very close contact to the NVP 5
membrane and develop numerous microvilli
as feature of functional active epithelial
cells. For details look at [213].
of cell growth of kidney epithelial cells in contact to PSU and NVP 5 membranes. It is
visible that MDCK cells grow better and make up a confluent layer on NVP 5 (b) in
contrast to PSU (a), where poor attachment and growth was observed on the inner
surface of a hollow fiber membrane. On NVP 5 cells also make very close contact with
the membrane surface and develop typical features of epithelial cells, such as tight
junctions and microvilli as shown in Figure 8.14. The functionality of kidney
epithelial cells was estimated by transepithelial resistance (TR) measurements
and impedance analysis that are validated methods to estimate the properties of
epithelial sheets, such as tight junction formation and active transport. Values for
functional epithelia are normally around 2000 V cm2, which is reached on standard
Millicells (MC) membranes [212]. Figure 8.15 shows the results of TR measurements. It is visible that not only materials with a lower content in NVP but also AEMA
0.1 mol % provided conditions where the TR values were comparable to that of
normal epithelia. In contrast, membranes with acidic sulfonate groups such as
AMPS and also PSU to some extent do not support the functional activity of kidney
Fig. 8.15 Transepithelial resistance of MDCK cells cultured
7 days on different supports (MC ¼ Millicell – reference
material). It is shown that cells on NVP 5 and NVP 20 have
resistance values comparable to the reference material
Millipore, which indicates a high functional activity of cells
on these membranes.
8.4 Summary and Conclusions 253
epithelial cells. The same investigations were also carried out with LLC-PK cells and
yielded comparable results regarding growth, morphology, and functional behavior
on the different membrane materials [214]. In summary of the investigations, it can
be stated that copolymers with low quantities of basic functionalities such as NVP 5
and AEMA 0.1 were the most suitable supports, while copolymers with acidic
groups such as AMPS, polysulfone, and NaMAS did not support kidney epithelial
cells. This is in line with the finding of others that coating with extracellular matrix
proteins is necessary to improve the cytocompatibility of polysulfone for epithelial
cells [206–211].
8.4
Summary and Conclusions
Membranes have found a multitude of applications in medicine. This holds not only
for conventional artificial organ therapy for blood detoxification and oxygenation
such as hemodialysis, hemofiltration, and cardiopulmonary bypass. Membranes
have been also implemented in the field tissue engineering basically as guiding
structures providing support for cell attachment and also helping to accumulate
growth factors and cytokines in the wounded area to promote healing of defects. For
helping cell attachment, growth, and functionality, the chemical composition of the
membrane surface as well as their surface topography play a crucial role. Membranes
have been also used in tissue engineering applications simply as a scaffold to place
materials for augmentation of tissues such as bone. While the transport properties of
membranes in tissue engineering applications have not been paid specific attention
so far, this is the case in biohybrid organ technology. Here membranes have several
functions particularly when in intimate contact with blood for detoxification purposes. Hence, membrane materials must be blood compatible on one hand side and
cell compatible on the other. This cannot be achieved so easily because blood contact
requires a noninteracting surface while cell contact needs a surface that is able to
support adsorption of proteins and adhesion of cells. A tailored design of copolymer
composition [192,195,213,214], the surface modification of one side of the membrane with specific ligands [194] or the application of specific membrane formation
techniques resulting in bilayer membranes from different polymers [215] can be a
solution to the problem. The permeability or pore size of the membrane may also
have several effects such as protection of xenogenic or allogenic cells from the
immune system of the recipient, control of exchange of solutes important for the
metabolic activity of cells, and also creating a specific topography that has an
influence on the behavior of immobilized organ cells. So far, a large number of
attempts have been put forward to produce implantable devices basically for the
replacement of endocrine functions. On the other hand, support of blood detoxification has been achieved by extracorporal biohybrid organs to replace liver and kidney
function particularly in acute failure of these organs. Overall, it may be concluded that
membranes represent an essential component in many tissue engineering applications and biohybrid organ technology. However, it may be stated for certain
254
applications that the full potential of membranes has not yet been exploited due to the
restriction to standard polymers and other reasons.
8.5
Acknowledgments
The authors gratefully acknowledges the cooperation with Wolfgang Albrecht,
Gregor Boese, Nathalie Faucheux, Günter Malsch, Dieter Paul, Klaus Richau,
Michael Schossig, and Barbara Seifert from GKSS Research Centre, George Altankov
and Natalia Krasteva from Bulgarian Academy of Sciences, Ulrich Gross and
Frederique Fey-Lamprecht from Free University Berlin, Volkmar Thom from Sartorius AG, Mathias Ulbricht from University Essen-Duisburg, Hans-Jürgen Stark and
Norbert Fusenig from German Cancer Research Centre Heidelberg. I am very
thankful to Mrs. Ruth Hesse for her excellent technical assistance in a number of
experimental studies mentioned in this review. Also, the financial support of some of
our own studies reviewed herein by BMBF (grant FKZ 01GN0119) and the Commission of European Union (BE 97-4326) is gratefully acknowledged.
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9
Membranes in Bioartificial Pancreas – An Overview
of the Development of a Bioartificial Pancreas,
as a Treatment of Insulin-Dependent Diabetes Mellitus1
Ana Isabel Silva, António Norton de Matos, I. Gabrielle M. Brons,
Marı́lia Clemente Velez Mateus
9.1
Introduction
9.1.1
Diabetes and Its Treatment
Diabetes mellitus is an autoimmune chronic disease characterized by high blood
glucose levels and long-term complications including micro- and macroangiopathic
lesions leading to retinopathy, neuropathy, and nephropathy. The insulin secretion is
either impaired or entirely destroyed in the pancreas of a diabetic patient, as a result of
which the blood glucose concentration climbs to levels far exceeding its normal
range. The cells that produce insulin, b cells, are located in structured cell clusters
named islets of Langerhans, which make up for 1–2 % of the total mass of the
pancreas [1].
There are two types of primary diabetes (type I and type II), differing in etiology,
pathophysiology, and treatment. Type I occurs when the pancreas ceases to produce
insulin due to a massive immunological attack on the islets by the patient’s own
immune system. This causes the destruction of all b cells in the islets in the
pancreas; thus, type I diabetes patients depend upon exogenous insulin therapy for
life. Type II diabetes is a different, milder type of diabetes, which occurs when the
pancreas does not produce enough insulin or when the insulin produced cannot be
used by the body due to resistance to the action of insulin. Only a small portion of
type II patients need insulin to control hyperglycemia; usually the ingestion of oral
hypoglycemic drugs and an appropriate diet are enough to stabilize blood glucose
levels [1].
The earliest known record of diabetes is mentioned in a third Dynasty Egyptian
papyrus (1500 B.C.) [2]. However, it was not until the early nineteenth century that
1
This is a preprint of an article published in ‘‘An
Overview on the Development of a Bio-artificial
Pancreas as a Treament of Insulin Dependent
Diabetes Mellitus’’, Silva A. I. et al. Medicinal
Research Reviews, Vol. 26, No. 2, 181–222, 2006
Copyright# 2005 Wiley Periodicals, Inc. (http://
www3.interscience.wiley.com./cgi-bin/jtoc/
34185/)
ISBN: 978-3-527-31480-5
9 Membranes in Bioartificial Pancreas
264
Tab. 9.1
Main events in the history of diabetes.
Year
Event
1500 B.C.
1869
1889
First known record of diabetesa
Paul Langerhans discovers the cellular structure of the pancreasb
Josef von Mering and Oskar Minkowski found diabetes could
be induced in dogs by removing their pancreasesc
Frederick Banting and Charles Best extract insulin from the
pancreases of dogsd
Two major types of diabetes, insulin dependent and insulin
independent are recognized
First pancreas transplant performede
First islet allograft transplant for diabetes is performede
First living donor pancreas transplant in the worlde
Human monocomponent insulin is launched, the world’s first insulin
preparation identical to human insulin. It is extracted from the
pancreas of pigs and converted to human insulinf
First islet human allograft transplantation using a corticosteroid-free
immunosuppressive regimeng
1921
1959
1966
1974
1979
1982
2000
a
Ebers Papyrus.
Berlin Pathological Institute, Germany.
c
University of Strassburg, Germany.
d
University of Toronto, Canada.
e
Fairview-University Medical Center, University of Minnesota, USA.
f
Novo Nordisk A/S.
g
University of Alberta, Edmonton, Alberta, Canada.
b
Source: Adapted from Refs. [2] and [3].
the cause of diabetes mellitus became clearer. Table 9.1 presents some important
events in the history of the treatment of diabetes.
Currently the more common treatment for insulin-dependent diabetes mellitus
(IDDM) is lifelong insulin therapy, which requires daily injections and frequent
monitoring of blood sugar concentration. An alternative to insulin therapy is
pancreas or islet transplantation. The first pancreas transplantation reported was
limited by poor graft survival and high morbidity. In 1978, new surgical techniques
together with new immunosuppressive regimens became available, greatly improving the outcome of pancreas transplantation [3].
Transplantation of isolated pancreatic islets into the liver by intraportal infusion
offers normalization of glucose metabolism similar to that after whole pancreas
transplantation, requiring less risky invasive surgery [3]. In the year 2001, an islet
transplantation group based in Edmonton, Canada, published a revolutionary paper
on human islet allotransplantation describing the use of a different immunosuppressive regimen together with improved islet isolation methods. The follow-up of
the transplanted patients [4,5] showed such good islet function over time that this
method has become the preferred islet transplantation method for IDDM patients.
Still, the patients are subjected to lifelong immunosuppressive drugs and the therapy
is conditioned by the scarcity of human donors.
9.1 Bioartificial Pancreas
9.1.2
The Bioartificial Organ Concept
A large number of people all over the world suffer from hormone deficiencies.
Currently, many of these hormone deficiency diseases can be controlled by the
patients with regular intakes of the lacking hormone. However, oral replacement
therapy may only delay the onset of complications of the disease, since large
variations still occur in the levels of metabolites.
Currently, the only definite cure for these diseases is partial or whole organ
transplantation, although the risks involved in transplantation are high. The surgical
risk can be reduced drastically, if instead of a whole organ, a specific cell transplant
can be performed. Transplantation of allogeneic organs or cellular tissue from one
individual to another or one species to another as in xenotransplantation will always
stimulate strong responses from the recipient’s immune system, which means that
the recipients will have to take immunomodulating drugs for the rest of their lives in
order to avoid organ or cell rejection. In the long run, these drugs have significant
toxicity that may give rise to other complications, for example, an increased risk of
infection and cancer.
Over the last 30 years, researchers have been trying to develop devices to prevent
the need for immunosuppressing drug regimes. These so-called bioartificial organs
generally contain cells or cell clusters within a synthetic biocompatible semipermeable membrane, which separates the foreign tissue from the host’s immune system
[6]. They can be inserted into the blood supply or implanted somewhere in the body
with diffusion potential and are meant to fully mimic the behavior and function of a
healthy organ. These devices are eliminating the need for immunosuppressing drugs
by protecting the cells from the onslaught of immune cells or antibodies, which are
not able to cross the synthetic biocompatible semipermeable membrane due to their
higher molecular weight. Ideally, the protected cells or cells clusters are kept viable
and functional by having access to nutrients and oxygen and stimulatory agents,
which can easily diffuse across the membrane and produce the lacking hormones or
factors on a minute-to-minute basis or when needed and/or for long periods of time.
The bioartificial organ implant procedure is generally simpler and less risky than
whole or partial organ transplantation. Biocompatible synthetic membranes developed for these devices have also other applications, like dialysis, in the development
of a bioartificial kidney device [7,8].
Many parameters can be varied in the bioartificial organ conception, namely, the
biocompatible polymer used and its intrinsic characteristics, the cell immobilization
or suspension media, the existence or not of coimmobilized molecules or cells, the
number of devices used and the implantation site, just to mention a few. From the
number of adjustable parameters, the complexity of developing a functional prototype of such a device is easily perceived.
This report attempts to present some of the more recent developments in the
conception of a bioartificial pancreas (BAP) device meant to mimic the insulin
secretion of the natural organ.
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9.2
Bioartificial Pancreas
Several different models of bioartificial organs have already been studied like the
liver, pancreas, and kidney [7–12]. The membrane characteristics and geometry vary,
depending on the specific function of the organ the bioartificial device should mimic.
In any model there are some requirements the device must obey to be accepted as an
alternative therapy: biocompatibility and good diffusional properties of the membrane, easy retrievability or biodegradability of the device, and maintenance of cell
viability and functionality for long periods of time (when cells are immobilized in the
device).
A BAP containing viable and functional islets of Langerhans protected by a
biocompatible semipermeable membrane is, in theory, a much better solution than
currently available IDDM therapies because it would provide a tighter blood glucose
control without the need for immunosuppressive drugs. It also makes possible the
use of islets from nonhuman donors thereby reducing the problem of scarcity of
human pancreata for isolation of islets for transplantation.
The physiological feedback inherent in pancreatic islets allows for normoglycemia
to be maintained at all times, even during high glucose stimulatory events; therefore,
the transport of stimulatory substances to b cells and in response the release of
insulin from them is fast and efficient. Consequently, a successful BAP model must
provide fast stimulation and response time to secretagogues, which translate into a
fast and efficient diffusion potential and must not allow insulin to stay in the vicinity
of the cells, thus damaging the physiological feedback system.
The material used in a successful BAP model must also be fully biocompatible and
should not trigger the immune response system by activating inflammatory agents or
stimulating fibrosis around the device.
Researchers all over the world have studied different membranes and membrane
geometries, as well as different immobilization and coimmobilization methods.
Devices are usually classified in two groups: intravascular and extravascular
devices. Intravascular devices are implanted as an artery-to-vein (AV) shunt in the
recipient’s body and are usually made of tubular hollow fiber membranes [13,14].
Extravascular devices can be further classified into two different categories, macroand microcapsular devices. Macrocapsular devices can have three main geometries:
flat sheet [15,16], sealed hollow fibers [17,18], and macrospheres [19]. These devices
have not only been mainly implanted in the peritoneal cavity, but have also been
implanted in other regions, for example, the subcutaneous tissue site. Microcapsular
devices consist of islets suspended in a polymeric gel material (e.g., alginate), gelled
and encapsulated in a biocompatible synthetic membrane, to produce microspheres
[20]. They are usually implanted in the peritoneal cavity, the kidney capsule, or
subcutaneous site. The working principle of all of them is the same: to allow
allogeneic or xenogeneic islet transplantation to succeed, by separating the islets
from the recipient’s own body cells with an artificial biocompatible semipermeable
membrane.
9.2.1
Immunoprotection and Biocompatibility of Implanted Devices
Immunoprotection in bioartificial devices concern the separation of transplanted
tissue from the host by a membrane or film forming a barrier against the passage of
immunologically active solutes and cells, thus preventing immune rejection of the
transplant. The pore size of the semipermeable membranes must ideally prevent the
passage of cells as well as the diffusion of proteins involved in the humoral
component of immune rejection, while allowing nutrients and metabolites to diffuse
[21].
Since the cut-off point of synthetic semipermeable membranes is an average value,
there are often a small number of larger pores that allow the transport of larger
molecules than suggested by the nominal cut-off definition. Therefore, complex
issues of rate of transport, threshold concentration of critical proteins, adsorption and
denaturation in contact with polymeric materials, and interactions between proteins
involved in immune reactions in an environment where their relative concentrations
may be far from normal may all have some impact in immunoprotection [21].
The most important aspects of any graft success are the nature of the membrane
and its biocompatibility with the body system. Biocompatibility involves a wide range
of requirements. A biocompatible material is defined as any substance (other than a
drug), synthetic or natural, that can be used as a system or part of a system that treats,
augments, or replaces any tissue, organ, or function of the body without causing a
reaction against it [22]. With intravascular implants, clot formation is the most
serious issue because it could threaten the patient if thrombi are shed into the
bloodstream. With extravascular implants, the most serious threat is the development of fibrotic tissue encapsulation over the membrane or encapsulant surface.
One of the major reasons for BAP graft failure is the development of fibrotic and
inflammatory responses to the chemical surface composition of the device. In
addition to the effects of chemical surface composition, the implant shape, morphology, and detailed microscopic topography at the material–tissue interface can
influence a foreign body response. For example, interaction of macrophages with the
surface is a crucial aspect of biocompatibility because these cells can induce
proliferation and activation of other cell types [1].
The foreign body reaction to implanted materials, resulting in fibrotic overgrowth,
is a complex process of sequential events: an acute inflammatory response leading to
chronic inflammatory responses and development of granulation tissue that ultimately leads to the formation of a fibrotic tissue capsule composed largely of collagen,
macrophages, fibroblasts, and few, if any, capillaries. The presence of thick fibrotic
tissue surrounding the membrane capsule prevents proper nutrient and oxygen
diffusion, quickly leading to massive loss of islet viability and functionality.
The intensity of this fibrotic response can vary greatly from individual to individual, even within the same species. Some authors [23] concluded that the extrapolation to humans of findings in rodents is very difficult and, in the particular case of islet
transplantation, even high animal models like pigs can differ significantly from that
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of human recipients. Although pigs are metabolically similar to humans, their
fibrotic response is potentially different.
Aside from the foreign body response induced by an inadequate capsule material, a
far more dangerous reaction can occur when allo- or xenogeneic tissue or molecules
are exposed or released from the capsular environment into the adjacent tissue and
initiate recognition of the foreignness of the tissue causing immune rejection.
The blockage of immune molecules is a complicated task. Although it is relatively
easy to prevent the passage of cytotoxic cells, macrophages, and other larger cellular
immune molecules including high molecular weight antibodies through the semipermeable membrane, a potentially more serious problem is blockage of humoral
immune components such as low molecular weight cytokines as well as tissue antigens
secreted by both the cells inside the membrane and the cells outside the chamber.
Antibody binding to a cellular transplant, by itself, usually does not cause a
cytotoxic reaction but it is the binding of the complement components that initiates
the cytotoxic events [24]. C1q is the key molecule that must be retained by the
immunoisolation membrane. This is the largest of the complement components and
the first that binds to two or more exposed antibody Fc domains. In the absence of
immune cells, binding of antibody alone to a cell surface antigen without complement activation cannot lethally damage the cell and is of no consequence unless the
epitope is from a site on the cell (e.g., receptor) important for cell function [1]. In order
for this cytotoxic process to occur, both C1q (and the other smaller components of the
complement system) and the antibody must cross the membrane. This can be
minimized or prevented, theoretically, if membrane pore size is controlled.
9.2.2
Vascular Devices
In intravascular devices, islets are encapsulated within one or several hollow
biocompatible membrane tubes or fibers and then implanted into the blood system
of the recipient’s body (Figure 9.1) [14,25–30]. One of the most relevant in vivo studies
was performed with a device that consisted of two concentric tubular membranes,
with one end anastomosed to an artery and the other to a vein. The blood flowed
through the lumen of the inner membrane, and the islets were immobilized in the
annular space between the two membranes [14].
The tubular design was one of the first designs ever experimented with and
published about. Earlier devices were transplanted in vivo into dogs [25,26], or ex vivo
in rats [27–29] and monkeys [27,30]. However, although ex vivo devices maintained
recipient normoglycemia for several months, the in vivo grafts never lasted more than
a few days (2 days in dogs). This was due to several different problems, some
identified at the time, like the presence of blood coagulation in the fibers (thrombosis), others remain the subject of investigations.
The use of intravascular devices presents three great advantages when compared to
extravascular ones: First, theislets embedded inside the chamber are near to the flowing
blood in the hollow fiber and have the most direct access to nutrients and oxygen;
secondly, the constantly flowing blood stream allows immediate recognition of
Fig. 9.1 Schematic representation of the device used by
Ref. [14].
Source: Adapted from Monaco et al. [14], with permission
from Lippincott, Williams & Wilkins, copyright 1991.
molecular changes in stimulatory agents like glucose and consequently immediate
release of insulin; thirdly, a higher diffusion potential is facilitated by the blood flow
allowing the glucose and insulin to be carried away faster. Thus, the islets are properly
fed, maintain their viability and functionality, and the delay in the response to a glucose
peak is reduced. However, all these advantages depend on the size of the hollow fiber
and the size of the surrounding chamber containing the islets. Diffusion problems may
be encountered related to thewidthof thechamber: toonarrow wouldmean insufficient
tissue is grafted, and too wide would mean islets suffer from hypoxia and insufficient
nutrition, followed by central necrosis and loss of functionality and viability. Nonetheless, this device design is probably the one that mimics more closely the physiological environment of the islets. Moreover, cell re-seeding or complete device removal
is easier to perform in case of functional failure. Other major problems are associated
with this kind of device [1]: some risk is involved in the surgical implantation procedure,
but a high risk is presented by the development of thrombosis, especially near the
anastomosed ends. Also, the blood flow is slowed down by the curvature of the device,
resulting in turbulence and damage to thrombocytes followed by clot formation.
9.2.2.1 Biocompatible Materials in Vascular Devices
Several different polymeric hollow fiber membrane materials have been proposed
as biocompatible materials for islet encapsulation. Zekorn et al.[31] studied
performance and biocompatibility of 50 kDa molecular weight cut-off (MWCO)
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membranes made of different membrane materials: Thomapor (polyamide), Cuprophane HDF (cellulose), and Amicon XM-50 (polyvinyl chloride acrylic copolymer).
The membrane devices with seeded rat islets were implanted in the right lobe of the
liver of streptozotocin-induced (STZ-induced) diabetic Lewis/Han rats. From these
experiments the authors determined that islets enclosed in the selected membranes
could not survive in a sufficient number for more than a few days.
There was evidence that this was the result of protein coating of the membrane
during the culture period, not a problem of islet adaptation to the hollow fibers
culture conditions. Instead of being released, insulin was nonspecifically adsorbed to
the membrane surface. Only when all available adsorption sites were filled did
insulin begin to be released into the lumen [31]. Protein precoating of the Amicon
XM-50 and PEEK (polyetheretherketone hollow fibers) membranes was then performed by preincubation of the membranes in medium supplemented with 10 %
FCS (Fetal Calf Serum)[32]. It improved islet function and cell viability. However, the
improvement of diffusional properties of the membranes by protein coating might
not have been the only cause for device-increased performance.
A very important aspect of membrane effectiveness in a bioartificial organ is the
ability to prevent toxic agents like lymphocytes and cytokines from passing through or
adhering to the membrane. Diffusion of toxic cytokines into the inner space causes the
destruction of the islets, and cellular adherence to the outer surface of the membrane
causes fibrosis and loss of cell viability due to impaired diffusion of nutrients and
oxygen. Some tests performed with the 50 kDa MWCO PEEK hollow fibers showed
that encapsulation in this particular membrane effectively protected the islets from the
toxic effect of Interleukine-1 (IL-1), one of the cytokines thought to play a role in the
immune destruction of islets [33,34]. Since the previous work showed improved
insulin diffusion with protein precoated membranes, the PEEK was also preincubated
with medium supplemented with 10 % FCS. After 2 days of in vitro culturing
encapsulated islets in IL-1 supplemented medium, the islets were still able to respond
to a glucose stimulus. They secreted insulin in amounts similar to those of control
islets (free floating, without IL-1 supplement) and higher insulin levels compared with
those of free islets cultured under the same conditions in IL-1 supplemented medium.
Amicon XM-50 membranes were still the ones preferentially used in vascular and
extravascular device experiments over the last decade [14,31,32,35–40], especially in
the most significant set of experiments in vivo with an intravascular prototype,
conducted by Monaco and his co-workers [14]. Experiments were performed with
dogs, using a prototype device consisting of a single-coiled large-bore Amicon XM-50
membrane, contained in a disk-shaped acrylic housing. The device was AV-anastomosed. To prevent aggregation, which leads to loss of function, the islets were
suspended in the annular membrane compartment, which was filled with sodium
alginate (SA) hydrogel.
The authors demonstrated the long-term patency of this nonthrombogenic membrane device by implanting unseeded devices and performing the follow up for up to
20 months. The membranes could withstand arterial pressure for long periods of
time without losing physical integrity. Nonetheless, 10 out of 22 devices implanted
eventually occluded.
These in vivo results represented the best results ever obtained for allo- or
xenogeneic implants: The longest graft survival times of 284 and 80 days, respectively,
without administration of insulin were presented in this article. The most important
cause for graft failure in these experiments by Monaco and co-workers was the
apparent loss of islet function and/or viability. Fibrotic tissue formation was not
observed on the blood stream side of the semipermeable membrane. Infection was an
important cause for graft failure, but the main problem seemed to be an inadequate
insulin secretion and a delayed response time to glucose stimulation. This poor
performance may have been due to an insufficient number of islets implanted, which
was confirmed by results from a double-device implantation experiment.
Maki et al. [40] devised three other prototypes of larger pore-sized membrane
material (MWCO of 70 kDa) and geometry similar to those already described. Totally
pancreatectomized dogs were used as models of diabetes, and xenogeneic porcine
islets were seeded into these vascularized BAP models, without administration of
immunosuppressive drugs. The prototypes differed from each other in length (30
and 45 cm) and volume (7.1–15 mL).
The overall reduction of exogenous insulin requirements for the prototypes varied
between 54 and 74 %. In seven of nine dogs, administration of exogenous insulin
between 4 and 16 U/day was needed to keep the glucose levels below 11.1 mmol/L.
Two animals maintained glucose control without additional insulin therapy for more
than 8 months. A device removed and dissected on day 271 showed acute vascular
occlusion of the device, which the authors attributed to hypercoagulability, but there
was no evidence of lymphocyte or inflammatory cell infiltration. The improvement in
long-term survival of discordant xenogeneic encapsulated islet grafts and in glycemic
control with a single implanted device was more significant when larger prototypes
were used, but not necessarily seeded with a larger number of islets.
Searching for membranes adequate for intravascular use, with better diffusional
properties and which induce less host sensitization and consequent immune
responses, is the main current research subject [41–44]. Recently, hydroxymetilated
polysulfone (PSU) was tested in vitro and used in the development of a prototype for a
vascular device [45].
Nonmodified PSU had already been used for xenotransplantation of porcine islets
into CD1 mice [17]. Graft survival time was very limited, 13 days. Cytotoxic anti-pig
islet cell antibodies were present in the recipients after 8 days of coil implantation.
The potential of PSU for BAP was demonstrated. The smooth surface of the
membrane allowed little or no fibrosis to appear.
Lembert et al. [45] used a smaller cut-off, 50 kDa, and performed surface modifications to improve the permeability characteristics of the biomaterial [45]. PSU is
hydrophobic and adsorbs large amounts of proteins, particularly insulin. Rendering
it hydrophilic would improve protein diffusion. To reduce diffusion distance between
the islets and the bloodstream, the islets were encapsulated in narrow capillaries,
which were then coiled to form the inner surface of an artificial vessel (Figure 9.2).
The in vitro testing showed that, over a 3-day period, this material allowed the lag
time for insulin to be reduced to 5 min and the maximal release time to be reduced to
30–45 min, which is similar to that of free-floating islets. The reduced kinetics of
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Fig. 9.2 Vascular prosthesis: (a) Diffusion chamber
(b) Compartment with coiled capillary;
Immobilization
matrix, * islet, –—Immunoprotective membrane,
!Blood flow.
Source: Modified from Lembert et al. [115].
glucose-induced insulin occurred only when PSU was chemically modified by
substitution with hydroxymethyl groups, at a degree of substitution near two. The
prototype is still waiting to be tested in allogeneic islet transplantation in pancreatectomized pigs.
Some authors also tried to ascertain the biocompatibility of other materials by
in vitro [46] and in vivo [15,37,47] experiments using extravascular devices.
9.2.3
Extravascular Devices
Extravascular BAP models present some advantages like a smaller surgical implantation risk, easier retrievability (for macrocapsules), possibility of device re-seeding,
and size flexibility. However, they also have some major drawbacks, the most
important of which is the diffusional hindrance to the transport of nutrients,
hormones such as insulin, oxygen, and metabolites. Other problems also limit
the success of these BAPs, like the fibrotic response and cellular adherence to the
outer wall of the membrane, loss of physical integrity of the capsule, especially in the
case of microcapsules, which in turn make total device retrievability difficult.
9.2.3.1 Implantation Sites
Extravascular devices can be classified by the size of the device into two categories:
macrocapsular and microcapsular devices. Several implantation sites have been
proposed and studied for the different types of capsules: epiploic chambers in the
great omentum, area around the splenic vein, external wall of the stomach, back wall
of abdominal muscles and peritoneal cavity [23], renal cavity, kidney capsule,
subcutaneous sites [48], and so on.
Icard and co-workers [23] studied the best implantation site for an extravascular
hollow fiber device in pig recipients. XM-50 hollow fiber macrocapsules with 100-mm
wall thickness were filled with human islets and implanted in several sites of STZinduced diabetic pigs. The preferential implantation site (lower fibrotic overgrowth)
that emerged from this study was the peritoneal cavity. This became the implantation
site chosen by most other authors to perform in vivo studies with extravascular hollow
fiber BAP even when using other animal models.
The subcutaneous implantation site was first used by Lacy et al. as an attempt to
minimize fibrotic response and cellular adherence by the host and also to determine
if this was a better nutrient diffusion site compared with the intraperitoneal site [36].
Others [48,49] evaluated device performance in the subcutaneous site, and it has been
used especially when prevascularization is performed, prior to seeding of the device.
Prevascularization is promoted in an attempt to minimize diffusional limitations and
mimic the pancreatic beta cells, where the cells are directly in contact with the
capillaries present within the islets. Compared to intraperitoneal implants subcutaneous implants involve even less surgical risk, are easier to seed and re-seed, or
remove if failure occurs.
9.2.3.2 Macrocapsules
Macrocapsules can have, essentially, two different geometries: tubular and planar
sheet. The tubular membranes are usually made up of a hollow fiber membrane, with
the islets immobilized inside with both ends sealed. These systems sometimes have
groups of individual tubes put together, called modules, to reduce device size and
increase the number of islets seeded into it, maintaining the transfer area. Some
authors who performed studies with intravascular devices also attempted to develop
hollow fibers extravascular devices, using the same membrane material [39,50]. The
overall graft survival time under similar experimental conditions but with extravascular instead of intravascular devices was shorter, as expected, since there are
additional nutrient, oxygen, and metabolite diffusion problems that condition the
islets survival when inside the capsule.
Flat sheet membranes are essentially made up as large circular or rectangular
sealed chambers with the islets immobilized inside. Due to their size and geometry
these devices are generally implanted in the peritoneal cavity. This configuration is
especially used for in vivo experiments with vascularized devices, which are preferentially implanted in the subcutaneous space. Because it is difficult to maintain
membrane planarity, membrane breakage frequently occurs due to the large size of
the device [51], rapidly leading to device failure. This geometry is, however, very
useful when performing islet cells viability studies [52] or when trying to assess
biocompatibility, diffusional, and immunogenic properties of a particular membrane
[15,46]. Insulin and glucose transport have been mathematically modeled for this
device geometry [53,54].
Biocompatible Materials in Macrocapsular Devices The biocompatibility of several
different membrane materials and their ability to work as immunoprotective barriers
in xenotransplantation of islets were established with in vivo experiments [15,37,47].
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Prevost et al. [47] evaluated the AN69 Hydrogel biocompatibility by transplanting
and comparing the efficacy of free-versus-encapsulated islets. AN69 membranes of
80 kDa MWCO, common for renal dialysis, are composed of 69 % acrylonitrile and
31% sodium methallyl sulfonate.
Four different groups of normal or STZ-induced diabetic Lewis rats were implanted
with either 6000 free Lewis rat islets or 9000 Lewis rat islets seeded in an AN69 hollow
fiber membrane. The diabetic animals transplanted with free islets remained normoglycemic for only 3 weeks, whereas those transplanted with encapsulated islets
remained normoglycemic for 54 and 70 days, respectively, until the removal of the
membrane devices, after which the animals became hyperglycemic again. The
biocompatibility of AN69 proved to be satisfactory, since host reaction to the syngeneic
implant was reduced to a thin layer of fibroblasts. The implantation area in these
animals remained intact: No tissue necrosis or cellular inflammation was observed.
Nonetheless, neither glycemia nor diuresis completely returned to normal level, which
implied the need for optimization of the diffusion potential.
Lanza et al. [37] had performed extravascular xenoimplants with capillary hollow
fibers (XM-50) in STZ-induced diabetic rats, to determine the protection against
rejection, in the absence of immunosuppressive or anti-inflammatory drugs. To
minimize fibrosis formation, membranes with smooth external skin were used,
which did not result in a fibrotic reaction, after more than 150 days postimplantation.
However, these results can hardly be extrapolated since the immune system differs
greatly between different species [23].
After conducting studies using intravascular devices, Amicon XM-50 was also tested
in the macrocapsular configuration. Allogeneic in vivo transplantation studies in dogs,
made diabetic by total pancreatectomy, were conducted to determine overall graft
survival, glucose control and adequate insulin release, as well as immune responses
induced by the device [38]. After transplantation into the peritoneal cavity of 155–248
chambers, each embedded with 2–4 105 islet equivalents, three of the six dogs
regained complete endocrine pancreatic function for 51 to more than 82 days
(experiments were still ongoing at the time of publication). These results showed
that this device was capable of responding to glycemic stress and could control blood
glucose concentrations in large diabetic patients, albeit for relatively short periods of
time (compared with 284 days in [14]). The difference in graft survival time was
thought to relate to the change from an intravascular to an extravascular BAP, which
affects nutrients, oxygen, and insulin availability and diffusion rates. Still, the authors
proved it was possible to achieve exogenous insulin independence for prolonged
periods of time without the need of any immunosuppressive or anti-inflammatory
regimen.
A planar diffusion chamber was made of Nuclepore membranes with a pore size of
0.1 and up to 0.6 mm to assess the membrane biocompatibility and the influence of
pore size on permeability properties [15]. Diabetic Wistar rats were implanted with
transgenic mouse insulinoma cell clusters (MIN6 cells) trapped in a mixed matrix
(1 % (w/v) agarose, containing 0.005 % (w/v) HEMA(2-hydroxy ethyl methacrylate)
copolymer and 0.15 mg/mL collagen gel) inside the Nuclepore membrane chambers,
which were sandwiched and edge-shielded with silicone.
Only 0.1 mm pore-sized Nuclepore membranes completely reversed diabetes
and normalized glycemic levels for up to 20 weeks, suggesting that a device
with these characteristics might provide a good option to improve diabetes
treatment.
Later, these Nuclepore membrane devices with the same pore-size range were
further used to perform in vivo xenotransplantation of encapsulated MIN6 cells into
STZ-induced diabetic Wistar rats [55]. Both devices with 0.1 or 0.2 mm pore-sized
membranes, transplanted intraperitoneally, reversed the diabetic state of the
recipients for at least 3 months. Outside this membrane pore-size range, diabetes
was only partially reversed and recipient hyperglycemia returned to preimplantation levels in less than 3 weeks. The immunoisolating system with the semipermeable membrane/mixed matrix reduced the IgG and complement permeation into
the membrane. This immunoprotective effect was best when 0.1 mm pore-sized
membrane was used.
These reports explain the classic compromise needed for workable devices in
any BAP: Small pore sizes prevent nutrients and insulin from diffusing in and out
of the inner space of the membrane device, but larger pore sizes allow immunoglobulins (IgGs) and other cytotoxic factors to enter this space and destroy the
islets.
Diffusional and permeability properties of membrane materials can also change
during the lifetime of a BAP device implanted into an animal. In order to clarify
one possible mechanism, protein coating of the membrane was performed.
The flat BAP with the AN69 membrane sheet (area of 700 mm2, MWCO of
65 kDa circular sandwich prototype) was coated with proteins before in vitro tissue
culture in an attempt to mimic an in vivo peritoneal environment [46]. Its
permeability to glucose and insulin was then assessed in culture, and the results
compared with those obtained previously with similar membrane devices
implanted in rats and recovered 1 or 7 days afterward.
The implanted noncoated AN69 membrane presented no tissue necrosis or
cellular inflammation after 1 or 7 days and maintained a well-preserved porous
structure (no cellular adhesion or fibrin deposits on the surface). The amount of
proteins adsorbed into the implanted and retrieved membranes were much higher
than the amount adsorbed into the in vitro protein-coated of membrane over the
same time period. Clearly, different mechanisms were involved in the adsorption
of proteins into the membrane in vivo and in vitro, depending, respectively, on the
protein composition of the peritoneal fluid and the culture medium. Even so, a
significant reduction in glucose and insulin diffusion occurred in both precoated
and noncoated implanted membranes. Both membranes showed a 50 % reduction
of glucose and insulin permeability, indicating that protein coating is a likely
reason for the permeability change in vivo.
This is in disagreement with the study by Zekorn et al. [31,32], which showed an
increase in insulin permeability when XM-50 and PEEK hollow fiber membranes
were subjected to protein coating. The membranes used in each study had quite
different physicochemical properties, which is probably the basis for this discrepancy in results.
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Biocompatibility of Inner Immobilization Matrix Aggregation of islets within the
membrane capsule results in impairment or loss of islet function within 1–2 weeks
due to hypoxia and central necrosis [36]. To prevent aggregation and preserve islet
function, the combined use of gel entrapment matrices together with immunoprotective barriers has been studied.
A mesh-reinforced polyvinyl alcohol hydrogel tubular (MRPT) device was used for
in vitro studies of biocompatibility of material and long-term maintenance of islet
function [56]. Agarose-dispersed islets seeded in the MRPT device showed an even
dispersion of the intact islets with an acceptable time response to glucose stimulation.
Other substances, including collagen, hyaluronic acid, and SA were studied to
prevent islet aggregation [57]. Islets were immobilized in collagen, type IA, 0.5 %
hyaluronic acid, 1 % SA gel, and 5 or 7 % agarose, before being seeded into the MRPT.
Of all substances tested, only the alginate gel allowed islets to stay intact and evenly
dispersed for 3 weeks. The differences in the gelling procedure for each substance
could have influenced this outcome, especially the consistence of the gel and gelling
time. Agarose was gelled by immediate cooling, minimizing the chance for islets to
aggregate.
Polymeric mesh structures were also used to assess the efficacy of coimmobilization of molecules or cells as an immunoprotection strategy. Pollok et al., for instance,
used a polyglycolic acid (PGA) polymer mesh structure as a scaffold and performed
in vivo allo- and xenotransplantation studies with islets coencapsulated with autologous chondrocyte within these scaffolds [58]. Chondrocytes were used to determine
if they could act as an immunoprotection barrier.
Islets from Lewis rats were seeded in a PGA sheet scaffold (1 1 0.06 cm) and
then encapsulated in a monolayer of chondrocytes. After 3 days in static culture, the
capsules were implanted subcutaneously into nude mice. The encapsulated islets
remained viable and functioned in vivo for up to 2 weeks and recently for up to
5 weeks. The glucose/insulin feedback mechanism remained intact [59,60]. This
procedure would need to be repeated in a diabetic animal model.
Influence of Membrane Surface Structure on Biocompatibility Lacy et al. [36] studied
the influence of the outer surface (roughness) of the encapsulating membrane with
respect to the fibrotic and immune response using subcutaneous and intraperitoneal
xenoimplantation of encapsulated islets. Individual rat islets were first encapsulated
in poly-L-lysine (PLL) coated alginate microspheres. To prevent islet aggregation islets
were embedded in a SA hydrogel. The encapsulated islets were then immobilized
inside two different types of hollow fibers: one fiber type had a totally fenestrated
outer wall and the other had a smooth outer wall.
This immobilization process did not interfere with the transport of insulin or
nutrients, as was proven by experimental controls. Biocompatibility of both fibers was
similar and acceptable (only a thin layer of collagen and fibroblast were present
surrounding each membrane). Better immunoprotection and control of glycemic
levels were obtained with the smooth fiber containing a lower number of islets in the
subcutaneous position while the fenestrated fiber yielded the best results in the
intraperitoneal position. The authors attributed these differences to a greater
bioburden of cells throughout the walls of the fenestrated fibers and to a lower oxygen
tension in the subcutaneous than in the intraperitoneal implantation site. Besides,
further experiments [35,38] showed host fibrotic response, and BAP failure could be
minimized if the outer surface of the membranes was smooth and continuous.
In another xenogeneic experiment, macroencapsulated dog islets, immobilized
with SA gel inside 50–80 kDa MWCO acrylic tubular membranes, were transplanted
into the peritoneal cavity of diabetic BB/Wor rats [37]. Eight of the ten rat recipients,
xenotransplanted with 1.2104 to 2.5104 encapsulated islet equivalents, showed
normal glucose levels for at least a month, whereas two animals sustained normoglycemia for 2 and 8 months. Tubular membrane chambers with smooth external
skin reduced host inflammatory pericapsular responses in these rats. The external
membrane surfaces were generally free of fibrosis and host cell adherence. However,
many capsules retrieved from animals at 1 and 2 months post-transplantation were
broken or bend. This implied the need for improvement in membrane length, wall
thickness, and/or diameter. The overall data indicated that the implants were capable
of responding to glycemic stress by correcting hyperglycemia within 1 h. Tolerance to
intravenous injection of glucose (0.5 g/kg body weight) was significantly improved
but not normalized.
More recently, a group of researchers described their work on the development of a
prototype called the Islet Sheet [51]. The Islet Sheet is a multilayered construction of
alginate: the islets, inserted in a thin flat sheet without polymer reinforcement, are
sandwiched on both sides by an acellular immunoprotective alginate layer. The
overall thickness of the Islet Sheet has to be the smallest possible (the typical device
measures approximately 4 cm 8 cm 250 mm). Like all devices implanted in the
peritoneal cavity, it not only relies on passive diffusion for transport of nutrients,
oxygen, insulin, and other metabolites, but also has to assure protection of the islets
to prevent host immune response; so a compromise must be reached. As is the case
with hollow fiber devices [35,37], the outer membrane wall has to be smooth and
continuous; any irregularity in the texture of the surface can lead to fibroblast
attachment and trigger fibrosis.
Allogeneic islets encapsulated in the Islet Sheet and implanted in the omentum of
a pancreatectomized beagle dog maintained fasting euglycemia for 84 days [51]. Still,
a strong fibrotic capsule (mainly collagen) developed around the macrocapsule. This
caused the loss of sheet planarity, further stimulating fibrotic response.
Neovascularization of Macrocapsules In order to improve islet nutrient supply and
oxygenation as well as insulin secretion, a new concept, that of vascularized devices,
was developed during the last decade. To build such a device researchers produced an
artificial implantation site, a scaffold, for the islets. The scaffold is usually implanted a
few days or weeks before seeding it with the islets to promote vascularization or
neovascularization (development of new capillary vessels around and inside the
device) and minimize damage to the islets due to prolonged hypoxia, especially
immediately after transplantation.
One group of researchers determined the performance of an artificial
islet implantation site made of a polytetrafluoroethylene (PTFE) solid support
277
278
sheet [61]. Human islets were encapsulated in a PTFE scaffold and transplanted
into the epidermal fat pad of diabetic nude mice. After 6 months, recovery of the
implanted devices showed that more than 50 % of the cells stained positive for
insulin. Close to the membranes were numerous small blood vessels, suggesting
the device had induced vascularization. There was no foreign body response
against the scaffold membrane.
The effect of protein coating as promoter of device vascularization was explored
with the initially developed PTFE device [49]. The scaffolds were coated with acidic
fibroblast growth factor (a-FGF) and implanted in rat recipients 4 weeks prior to
isogeneic islet seeding. Only the solid PTFE supports that were coated with a-FGF
were found covered and penetrated by a large number of blood vessels. The
successful islet isografts allowed the diabetic Albino Oxford (AO) rat recipients to
remain normoglycemic for the duration of the experiment (6 months). Large
numbers of viable islets were found in the solid supports retrieved from the
intraperitoneal cavity. The PTFE scaffold allowed successful transplantation of islets
to the peritoneal cavity, transforming it into a more efficient transplantation site.
Other authors also attempted to induce vascularization of BAP devices, using two
different membrane constructions and comparing their performance at different
implantation sites (renal subcapsule, intraperitoneal cavity, subcutaneous site) [48].
Thin fibrils of polyvinyl alcohol (PVA) or PGA were woven into soft, pliable, mat-like
structures to produce flat sheets where free or encapsulated allogeneic islets were
seeded.
Free islets implanted under the renal capsule performed better. The best results
with encapsulated islets were obtained with PGA polymer sheets implanted subcutaneously (four out of five mice maintained normoglycemia for 3 months after
implantation). Successfully grafted devices showed many intact islets and were
neovascularized. The authors could not determine which specific contributions of
the PGA material caused this graft success. But the bad results obtained with PVA
polymer were probably a consequence of PVA being more loosely woven than PGA,
thus causing islet leakage from the membrane.
The efficacy of prevascularization of other planar diffusion chambers was assessed
by performing xenotransplantation experiments of macroencapsulated rat islets into
the subcutaneous site of STZ-induced diabetic Swiss Webster Nude mice [62].
The 20 mL macrocapsules used (Thera Cyte, from Baxter) were made of three
different layers of membrane laminates: an inner cell-retentive 0.45 mm membrane,
laminated to a 5 mm pore diameter vascularizable PTFE membrane, with an outer
polyester mesh providing support. Islets in culture medium were seeded into
capsules previously implanted subcutaneously (2 weeks earlier); their performance
was compared with implants without prevascularization.
Prevascularization did not improve the outcome of the implants; the time needed
to achieve normoglycemia was similar, between prevascularized or non-prevascularized macrocapsules.
Evidence of AN69 hollow fiber membrane neovascularization potential was also
reported [18]. Intraperitoneal xenotransplantation of porcine islets into diabetic
NOD (Non Obese Diabetic) mice was originally performed, in this study, to
determine the ability of the devices as immunoprotection barriers. Seeded capsules
(40 cm-long fibers, 500 mm i.d., 80 mm wall-thickness and 80 kDa MWCO)
reversed the diabetic state in only 18 out of 54 recipient NOD mice for 6 weeks.
This showed that the fibers provided at least transient protection from xenograft
rejection to porcine islets.
Capsules retrieved from successfully implanted mice showed functional islets
26 days after implantation. Two out of four fibers retrieved from mice that remained
hyperglycemic contained islets showing a response to glucose. These fibers showed
perifiber inflammatory tissue (fibrotic reaction and cellular adherence). Additionally, neovascularization of the perifiber tissue was observed in more than 20 % of
the random sections. This neovascularization process may have improved the
supply of nutrients and oxygen to islets. Nonetheless, the efficiency of the
encapsulated islets transplant was limited by the percentage of successful recipients (only 37 %) and by the duration of graft function (6 weeks). Although it was
not proven, the formation of the perifiber inflammatory tissue might have been
associated with the promotion of neovascularization, having a beneficial effect on
the viability of entrapped islets.
Smooth surface regenerated cellulose hollow fibers (Thomapor) have also been
reported to spontaneously promote vascularized connective tissue overgrowth,
when implanted intraperitoneally in rodents [63]. Nevertheless, only grafts without
surrounding connective tissue had a stimulatory response to high levels of glucose.
In an attempt to improve oxygen delivery to the islets encapsulated in agarose-PSSa,
a study was undertaken where a subcutaneous site was coimplanted with basic
fibroblast growth factor (b-FGF) impregnated microspheres, as shown in [48,49].
The b-FGF-impregnated microspheres were implanted into C57BL/6 mice 2 weeks
before implantation of the macrocapsular device [64]. The device implanted was made
of a mixture of agarose crosslinked with polystyrene sulfonic acid (PSSa) with a tubular
shape. This type of hydrogel is more usually found in experiments with microcapsular
devices (vide Section Agarose-PSSa Microcapsules). The performance of devices coimplanted with or without b-FGF-impregnated microspheres was assessed. Intraperitoneal glucose tests were performed 1 and 2 months after transplantation. Vascularization assessment around the implantation site was histologically performed with
subcutaneous tissue samples. In two recipients, graft survival was evaluated after
device retrievement at 1 and 2 months post-transplantation.
All the recipients of the subcutaneous xenografts returned to normoglycemia. The
average normoglycemia period was 68.4 25.6 days, in contrast with the control
groups without b-FGF-impregnated microspheres, which maintained normoglycemia
only for 9.5 5.1 days. Glucose tolerance tests at 1 and 2 months with b-FGFþ mice
showed significant improvement but were still not as good as for control mice. b-FGFimpregnated gelatine microspheres, unlike nonimpregnated ones, induced significant
subcutaneous neovascularization in diabetic mouse 14 days after injection.
Viability assessment in recipients of the b-FGFþ group revealed the islets within the
graft were viable and intact and no fibrotic overgrowth or pericapsular cellular infiltrate
were present around the surface of the graft. In the b-FGF– and control groups,
however, remnant islets in the removed graft were found to be necrotic, although no
279
280
significant fibrotic overgrowth was observed, indicating islet malnutrition as the
probable cause of necrosis.
When recipients returned to normoglycemic state, they were subjected to
an intraperitoneal injection glucose tolerance test. Their insulin secretion levels
were found to be improved but reduced and returning more slowly to basal
levels than in normal animals, which pointed to plasma insulin responses being
insufficient.
The cylindrical device presents many technical advantages: it is injectable, retrievable, durable, difficult to mechanically disrupt, and suitable for subcutaneous
implantation. The additional use of b-FGF-impregnated gelatine hydrogel or gelatine
microspheres strongly induced neovascularization and tissue granulation around the
implantation or injection site. However, biocompatibility of this agarose-PSSa mixed
gel-based BAP device was attenuated over time, which eventually led to graft failure.
Also, if a tighter homeostatic control of glucose metabolism could be achieved by this
xenotransplantation technique, it could be possible to achieve long-term reversal of
IDDM in human patients.
9.2.3.3 Microcapsular Devices
Aside from macrocapsular devices, researchers have also developed a microspherebased BAP. In this case, the cells are immobilized inside microspheres of a
biocompatible material usually alginate or agarose gel, coated or not with a semipermeable membrane, and implanted in the patient’s body. This kind of device was
the most extensively used in the development of a BAP mainly because of the
simplicity of its construction and its flexibility for modifying key components. This
construction flexibility allowed researchers to play with key parameters like wall
thickness, pore size, or membrane compactness to overcome significant nutrient and
oxygen diffusion limitations seen in all extravascular implantation sites. Nonetheless, the geometry presented technical problems, of which the major was in
many cases the difficulty to retrieve the microspheres from the patient’s body,
especially if fibrotic reaction or cellular adherence occurred. Several problems were
addressed initially and others were evidenced when microcapsules were used in vivo,
like the development of a strong fibrotic response, difficulties in microcapsule
construction, and ideal polymer composition.
Poly-L-Lysine (PLL) and Alginate-PLL-Alginate (APA) Microcapsules Sun et al. were
the first to report the use of microcapsules as a BAP [65] and later optimized prototype
characteristics. Briefly, the microcapsules were produced in the following manner:
Islets were suspended in a SA solution, and droplets of the suspension were
produced by syringe pump extrusion with a droplet-forming apparatus and dropped
into a calcium chloride solution, for bead gelling to occur. The alginate beads were
then incubated in a PLL solution, washed, and incubated with sodium citrate to
liquefy the alginate gel inside the capsule. Several alterations to this method were
made in follow-up experiments, namely the droplet generation process, alginate
purity, and its concentration and incubation times. Consequently, capsule properties
changed which, as will be shown, had a crucial effect on graft survival. A special
modification with respect to bead size allowed obtaining microcapsules smaller in
diameter than the previous ones (250–350 mm), most of them containing only one
islet (Table 9.2).
Many studies were performed where chemical composition of capsules [66,67],
purity [68,69], bead size [70,71] and shape, integrity, and mechanical properties
[72–74] of APA microcapsules were examined and related to the ability to induce or
not an immune response, when implanted in an animal recipient.
Other modifications included the use of fetal rat islets [75] because their isolation
technique is relatively simple, requiring no Ficoll separation or laborious handpicking. The results proved that these islets were well suited for this kind of device
construction. Sun and co-workers also used Wistar rats as donors and Balb/c mice
as recipients (concordant xenografts) and a similar microencapsulation method
[20] but with smaller capsules, which more than tripled the duration of normoglycemia following transplantation when compared to the already good results
obtained with fetal rat islets. Since smaller size beads allowed easier access to
oxygen and nutrients, the islet viability was increased, which, in turn, promoted
prolonged graft survival. The eventual graft failure that was observed for some
recipients may have occurred for several reasons: insufficient number of islets
transplanted, presence of cellular overgrowth, or capsules clumping together
(which can lead to islet central necrosis). All of these aspects were absent in
recipients that remained normoglycemic and emphasized the need to clarify the
influence of capsule characteristics on graft survival including mechanical structure, size, and alginate composition.
Zhou et al. tried this microcapsular prototype in a large animal model with pig
islets into diabetic monkey recipients. Porcine islets were microencapsulated in
alginate-PLL microspheres and transplanted into spontaneously diabetic Cynomologus monkeys. The islets were isolated and purified by means of two different
processes, which yielded two islet preparations with different degrees of purity, 75–
90 % and more than 95 %, respectively. The microcapsules used were similar to
those of [20]. The general condition of all recipients improved substantially [76].
Sun et al. reported the first ever long-term discordant xenograft function without
any immunosuppression in a large animal model, resulting in normal glycemic
control [77]. Several months after transplantation, the reappearance of hyperglycemia could be reversed after a second implant with a similar number of encapsulated islets. In two recipients that never became insulin independent, graft
failure was attributed to damage caused to islets by contaminating exocrine tissue
(these recipients were transplanted with lower purity islet preparations).
The long-term survival of the microencapsulated islet grafts in these experiments
could be attributed to two major factors: the strength of the capsular membrane and
the purity of the porcine islet tissue. The purity of the alginate solution was also very
important with respect to the biocompatibility of the capsules. This was investigated later by other authors, using immunogenic assays.
Effect of Alginate Chemical Composition Alginate gels exist in varied forms and
compositions and usually their exact and complete composition is not known, which
281
M.c.s had smooth
surfaces and Ø
between 400
and 600 mm
STZ-induced diabetic
Balb/c mice, i.p.
transplant, seven
with free islets
and nine with m.i.s
Balb/c mice, i.p.
transplant, four
with free and
10 with m.i.s
800–1000 free or
encapsulated
Wistar Rat fetal
pancreatic
islets/recipient
900–1000 Wistar
rat free islets
or m.i.s/recipient
M.c.s with Ø
between 250
and 350 mm
N/A
STZ-induced Wistar
Lewis Rats, i.p.
transplant, five
with m.i.s
Wistar rat
islets (number
of islets N/A)
Microcapsule
parameters
Recipient/
implantation site
Experiments with APA microcapsular devices.
Number of islets
transplanted
Tab. 9.2
Until
hyperglycemia
returned
Until
hyperglycemia
returned
APA m.c.s
2–3 wks, in vivo
PLL: 37 kDa MW
APA m.c.s: PLL
with 17–22 kDa MW
15 wks, in vitro
Experiment
duration time
APA m.c.s
Encapsulation
method/material
End of experiment,
recurrence of
hyperglycemia
Recurrence of
hyperglycemia,
in vivo
End of experiment,
recurrence of
hyperglycemia
Loss of
functionality,
in vitro
Device
removal reasons
M.i.s morphologically and
functionally intact
throughout in vitro culture
studies (15 wks)
Five recipients of m.i.s
maintained normoglycemia
for 3 wks
Reversal of mice diabetic state
for between 54 and 171 days,
without immunosuppression
Capsules removed from two
recipients (144 days
postimplantation) were
essentially free of fibrosis
and cell overgrowth and
had minimal or no
structural damage
Xenografts functioned
for periods
ranging from
172 to 308 days
Recovery of approximately
80 % of the capsules
implanted
About 70 % of recovered
capsules free of cell
overgrowth and with
viable islets
Major
outcomes/results
[20]
[75]
[65]
Ref.
Mongrel dog islets
(number of
islets N/A)
One, two, or three
transplants of
3104 – 7104
porcine free
islets or m.i.s/
recipient
One or two
injections of
3104 – 7104
porcine m.i.s/
recipient
Spontaneously
diabetic dogs
with IDDM
transplanted into
the abdominal
cavity, three with
free islets, seven
with m.i.s, four
reimplanted
with m.i.s
Four spontaneously
diabetic
cynomolgus
monkeys, i.p.
transplant
with porcine m.i.s
Spontaneously diabetic
cynomolgus monkeys
i.p. transplant, two
with free islets and
nine with m.i.s
Same as
Ref. [65],
with high-G
alginate
Same as
Ref. [20]
Same as
Ref. [20]
High-G
alginate-PLL
m.c.s
APA m.c.s
APA m.c.s
Graft survival
occurred when
recipient
remained
euglycemic
w/o any
exogenous
insulin
Until
hyperglycemia
returned
Until
hyperglycemia
returned
Until
hyperglycemia
returned
Recurrence of
hyperglycemia
Recurrence of
hyperglycemia
Not referred
Normoglycemia restored
from 120 to 803 days, in
seven of nine recipients
No antiporcine islet
antibodies detected in the
recipients, even after return
to hyperglycemia
Recovered m.c.s free of cell
overgrowth, physically intact
and not clumped
Insulin secretion of recovered
islets was similar to that of
freshly isolated ones
Graft survival occurred in all
animals transplanted with
m.i.s for 63–172 days
(CysA administered
to all recipients)
Lower exogenous insulin
requirements to maintain
euglycemia, following graft
failure before transplantation,
suggesting that islet survival
occurred several days beyond
those reported
Complete restoration of
normoglycemia for
106–390 days,
w/o exogenous insulin or
immunosuppression
[78]
[77]
[76]
Recipient/
implantation site
In vitro assay
Normoglycemic
AO rats,
i.p. implant
Five to six Balb/c
mice and five
to six Wistar
rats, i.p. transplant
In vitro assay
None
ICCs from porcine
fetuses (number
of ICCs N/A)
(Continued )
Number of islets
transplanted
Tab. 9.2
Average alginate
bead Ø: 0.7 mm;
one or two
ICCs/capsule
Same as in
Ref. [69]
M.c.s Ø between
500 and
1800 mm
Microcapsule
parameters
APA high-G and
high-M, empty
with fetal porcine
ICCs or with
Dynabeads
(Dynal, Oslo,
Norway) covered
with BSA
PLL Bariumalginate m.c.s
with Manugel
and Keltone LV
APA m.c.s,
external alginate
layer made of
Manugel or
Keltone LV
Encapsulation
method/material
Different
types of m.c.s
implanted
twice, with a
2-month
interval
1 month
In vitro assay
Experiment
duration time
End of the
experiment
End of the
experiment
In vitro assay
Device
removal reasons
Keltone LV: lowest overall
percentage of protruding
islets; islet-containing
defective capsules were
lowest with 800-mm Ø m.c.s.
Manugel: percentage of
incomplete encapsulation
and defective m.c.s lowest
for 500-mm Ø (apparently
due to swelling properties
of Keltone LV gel)
High-M m.c.s were found
more biocompatible
than high-G ones
Minor modifications in the
encapsulation procedure
lead to high variability of
m.c.s biocompatibility
and graft survival
Mice developed antibodies
against the high-M but not
the high-G material
Rats did not develop any
antibody responses to either
of the m.c.s
Major
outcomes/results
[80]
[67]
[66]
Ref.
1000–2000 Fisher
rat or 2000–5000
mongrel dog
m.i.s/mice
recipient
300000 mongrel
dog m.i.s/beagle
dog recipient
29 STZ-induced
diabetic Balb/c
mice and one
beagle dog,
islets i.p. implant
Microspheres of
250–300 mm
in diameter,
1–3 islet/m.c.
Prepurified
high-G and
high-M
alginate m.c.s
(no PLL used)
In vivo, graft
failure
determined
of the end of
the experiment
Islets cultured
for 4 days
in vitro
Persistence or
return to
hyperglycemic
state
Immune response proved
not only to be connected to
alginate type but also to
recipient species
Twenty-three out of 29 mice
transplanted with rat m.i.s
sustained normal plasma
glucose levels for
56–419 days.
Twenty-two out of 40 mice
transplanted with dog m.i.s
sustained normal plasma
glucose levels for
21–630 days
M.c.s recovered from mice
contained rat islets; a few
exhibited central damage,
indicating insufficient
amounts of nutrients and
oxygen in the peritoneal cavity
Mongrel dog m.i.s retrieved
from mice showed
well-granulated endocrine
cells (a few necrotic islets)
Fibrosis and cell adhesion
were observed when islets
protruded from the m.c.s
Beagle dog remained
euglycemic (after gradual
withdrawal of exogenous
insulin), for over 400 days
[68]
Recipient/
implantation site
STZ-induced
diabetic AO/G-rats,
i.p. transplant,
empty or isletcontaining m.c.s
AO/G rats, i.p.
transplant,
three with
empty m.c.s,
seven with
small m.c.s,
five with
large m.c.s
AO/G or Lewis
rats (number
of islets N/A)
Lewis rat islets
(number of
islets N/A)
(Continued )
Number of islets
transplanted
Tab. 9.2
M.c.s with two
different Ø
ranges: 400–500
(small) and
700–800 mm
(large)
M.c.s with
600–700 mm Ø
and 100 kDa
MWCO
Microcapsule
parameters
APA m.c.s
Purified and
nonpurified
SA (40%
guluronic
acid, 60%
mannuronic
acid contents)
Encapsulation
method/material
4 wks after
transplantation
or 2 wks after
recurrence of
hyperglycemia
Empty capsules
removed
1-month
postimplantation;
graft failure
determined or
the end of the
experiment
Experiment
duration time
Persistent or
recurrent
hyperglycemia
Persistence or
recurrence of
hyperglycemia
Device
removal reasons
Within 1 month of
postimplant, all empty
nonpurified alginate m.c.s
were overgrown with fibrotic
tissue and surrounded by
collagen deposits
Majority of purified alginate m.c.s
freely floated in abdominal
cavity, w/o adhesion to
abdominal organs
Ten out of 13 rats with allogeneic
m.i.s implants (purified
alginate) remained
normoglycemic for 6–20 wks
No relation was observed
between the duration of graft
function and the number of
capsules overgrown by
fibrotic tissue
Nonovergrown capsules had
nonviable islet remnants,
after graft failure
One in seven recipients
became normoglycemic
with small m.c.s (>5 %
vital b cells in native
pancreas) for 8 wks
Major
outcomes/results
[70]
[69]
Ref.
2000–3000
Sprague–Dawley
rat m.i.s/recipient
STZ-induced
diabetic Balb/c
mice, i.p.
transplant,
20 with m.i.s
Two populations of
m.c.s: standard
(Ø 350 mm,
method of
Ref. [82]) and
compact
(Ø 300 mm
modification to
method of
Ref. [82])
Adequacy of
encapsulation:
% of inadequate
m.c.s in total
number of
islet-containing
m.c.s
APA m.c.s
Until blood
glucose levels
rose above
200 mg/dL
for m.i.s
1, 3, and 6-month
after implantation,
for empty m.c.s
Blood glucose
levels rose
above
200 mg/dL
Majority of small m.c.s
adhered to the abdominal
organs (very low m.c.s
retrieval rate); extensive
fibrosis and cellular
infiltration present; islets
in m.c.s were necrotic and
contained no vital b-cells
All five recipients of large
m.c.s became
normoglycemic for 7–16 wks
Majority of large m.c.s were
free floating in the peritoneal
cavity and nonadherent to
abdominal organs
Recipients of compact m.c.s
maitained blood glucose
levels below 180 mg/dL
for 108 days; standard
m.c.s maintained these
same levels for
only 76 days.
Empty APA m.c.s were
recovered 6 months
postimplantation: 81%
compact ones intact, 38%
standard ones intact
Fibrotic neogrowth occurred
only when m.c.s were
broken or damaged
(islets exposed)
[73]
Recipient/
implantation site
Normal Lewis rats,
i.p. transplant,
four with
alginate beads,
14 with APA m.c.s
AO/G rats, i.p.
implant, five
with free islets
and nine with m.i.s
Adult porcine or
bovine calve
islets (number
of islets N/A)
2800–3000 free
Lewis rat islets
or 2500–2900
Lewis rat m.i.s/
recipient
(Continued )
Number of islets
transplanted
Tab. 9.2
Alginate beads
or APA m.c.s
APA m.c.s
Uniform smooth
spherical m.c.s, Ø
800–900 mm; 1–3
islets/m.c.
Encapsulation
method/material
Alginate beads:
Ø 300–600 mm,
1–2 IE/
microsphere
m.c.s: Ø
800–1200 mm,
66 kDa MWCO
(measured)
Microcapsule
parameters
5–6 wks
2 wks, 1, 2,
3, 7, and
9 months
Experiment
duration time
Relapse into
hyperglycemia
Removal of
device at
the end
of the
experiment
Device
removal reasons
Islets encapsulated in
uncoated alginate beads
nonviable for 2 wks or
1 month after implantation
in normal rats
Islets immobilized in m.c.s
survived transplantation
and had similar viability
at preimplantation and
removal days
Nonadherent and intact m.c.s
retrieved, with external
surfaces mostly free from
fibrosis or host cellular
adherence.
M.i.s removed at 9 months
had an average twofold
increased insulin secretion
with a static glucose challenge
Similarly good results were
obtained for the xenotransplantation of encapsulated
bovine islets into rats
Four out of nine recipients
of m.i.s remained
normoglycemic for
10–14 wks (three of them
showed b-cell regeneration
in native pancreases)
Major
outcomes/results
[83]
[74]
Ref.
Healthy B6 or Balb/c
mice or Sprague–
Dawley rats
Same as in
Ref. [73]
APA m.c.s
4 wks
Removal of
device at
the end of
the experiment
M.c.s retrieved from animals
that returned to hyperglycemic
state were overgrown with
fibrous tissue and contained
no viable cells
Islet recovery ratio was
significantly different for
allogeneic versus xenogeneic
and syngeneic versus
xenogeneic transplants:
60 % versus 0 % and
1.6 % versus 4 %;
neogrowth was very
different in both cases:
1.6 % versus 4 % and
2.8 % versus 4 %
B6 mice and Sprague–Dawley
rats are incompatible with
APA m.c.s, but Balb/c mice
are a good model for
transplantation of these m.c.s
[84]
Abrevitations: m.i.s, microencapsulated islets; m.c.s, microcapsules; wks, weeks; STZ, streptozotocin; Ø, diameter; i.p., intraperitoneal; N/A, not available; IDDM, insulindependent diabetes mellitus; CysA, cyclosporin; w/o, without; ICCs, islet-like cell clusters; IE, islet equivalents; NPCC, Neonatal porcine pancreatic cell clusters.
C57BL/6 (B6) mice
or Balb/c mice
or Sprague–
Dawley rats,
2000–3000
m.i.s/recipient
Removal of
device at
the end of
the experiment
290
means they may contain substances toxic to the islets or to the recipient. Pure
alginates are linear polysaccharides composed of homopolymeric regions of 1,4linked b-D-mannuronic (M) and 1,4-linked a-L-guluronic (G) acid, interspaced with
regions of mixed sequence [68]. But commercial alginates are usually crude products
extracted from algae and are contaminated with different inflammation provoking
components, of which only polyphenols and endotoxins have been identified [69].
The alginate composition was studied by Soon-Shiong et al. [78]. This was the first
report describing long-term successful reversal of spontaneous diabetes in a large
animal model transplanting allogeneic islets with this kind of device. The report
focused on the enhancement of microcapsule biocompatibility and mechanical
integrity by using high-G-alginate PLL microcapsules and introducing modifications
to the microencapsulation method. Longer periods of euglycemia were verified for
high-G PLL capsules.
In an attempt to determine if the development of fibrosis around the microcapsules was due to the chemical composition of the alginate used in the immobilization process, De Vos et al. studied the influence of the size of the microspheres
and their guluronic and mannuronic acid alginate content (G:M ratio) in microcapsule physical resilience and ability to completely envelop the islets, in vitro [66].
Encapsulation was performed as usual, followed by an additional external alginate
layer, of Manugel or Keltone LV; Manugel has higher G:M ratio than Keltone LV.
The percentage of incomplete encapsulation varied with capsule size and type of
alginate used. The authors concluded from this study that alginates with high G:M
ratio were better suited for encapsulation. Quantitatively, incomplete encapsulation
was the most important manifestation and was lower with Manugel than with
Keltone LV. Since Manugel has a higher G:M ratio than Keltone LV, it is associated
with less swelling in the absence of divalent ions, which diminishes the possibility of
protrusion (Figure 9.3).
De Vos et al. continued to investigate the effect of alginate composition (this time
with barium-alginate microcapsules) on the induction of fibrotic response induced by
different chemical composition alginates [67].
The fibrotic reaction was more severe against Manugel alginate (capsules over-
Fig. 9.3 Schematic presentation of the process of swelling
and shrinkage during the encapsulation procedure.
Swelling and shrinkage contribute to the occurrence of
incomplete encapsulation.
Source: Reproduced from De Vos et al. [66], with
permission from Lippincott, Williams & Wilkins, copyright
1996.)
grown and adherent to abdominal organs) than against Keltone LV alginate (capsules
free floating and without adherent cells). No fibrotic response was observed with both
kinds of alginates in the absence of the PLL membrane, indicating the problem
related to the binding of PLL to guluronic acid present in alginate, which was in
agreement with other authors (see the section ‘‘Assessment of fibrotic response’’,
Ref. [79]).
Kulseng et al. [80] tried to determine the influence of the alginate composition on
the ability to induce antibody responses or act as an adjuvant to antibody responses to
antigens leaked from the microcapsule. The authors compared high-G and high-M
alginates performance. Two different in vivo sets of experiments were performed: the
first, to determine the antigenicity of empty capsules; the second, to determine the
responses of encapsulated porcine islet-like cell clusters (ICCs), when transplanted
into mice and rats. They showed that high-M alginate had a higher immunestimulating activity than high-G alginate, both in terms of antibody and cytokine
production, thus supporting results by [81].
To determine the generation of antibodies against ICCs, rats received 50 free or
encapsulated ICCs, in high-G alginate-PLL capsules. Antibodies against ICCs were
detected 6 weeks after implantation of free or encapsulated islets, and there was an
increase in antibody levels 2 weeks after the second injection; after 8 weeks the
antibodies started to decrease. This result clearly showed that ICCs, either free or
encapsulated, always evoke a humoral immune response, which decreases with time.
Surprisingly the highest antibody response occurred in animals transplanted with
encapsulated islets. However, as was suggested before, this happened because
proteins released from the encapsulated ICCs can pass through the membrane
and bind themselves to the outer capsule surface, triggering complement activation
and local inflammation, thus negating the protective effect of the microcapsular
membrane.
By implanting into rats BSA covered Dynabeads (commercial beads), with or
without alginate coating, it was further shown that the alginate coating completely
prevented antibody response, where the absence of it lead to the generation of
antibodies against BSA.
Effect of Alginate Purification Cochrum et al. [68] also conducted in vitro and in vivo
xenotransplantation experiments with microcapsules, but this time high mannuronic and high guluronic acid alginate were purified (proteins, fucose, and polyphenols
were removed) before being used to produce PLL microspheres [68].
This study [68] reported prolonged reversal of diabetes in mice and dog by
transplantation of encapsulated xenogeneic islets. Furthermore, it showed that highly
purified mannuronic and guluronic calcium-alginates were stable in vivo and could
provide immunoisolation of islets, especially in conjunction with a small diameter
biocompatible membrane coating. The absence of fibrosis on implanted empty
spheres and absence of host mononuclear cell infiltration (except where islets were
exposed) was supportive of the efficacy of alginate purification in removing immune
response inducing contaminants. However, when the survival of the allotransplants
291
292
in rodents (56–419 days) was compared with that shown by Sun and co-workers [20]
(172–308 days) without purification, no significant improvement on graft survival
was observed, thereby questioning the need for this alginate purification.
The reduced size of the microcapsules used in this work led to protrusion of islets
from the capsules. This fact rendered the proper evaluation of the effect under study
impossible. Where total recovery of microspheres was possible, the PLL coatings
were generally free of fibrosis and host cells; however, if islet tissue was left exposed
on the microsphere surface, host mononuclear cell infiltration occurred, as was later
confirmed by [66]. The influence of capsule size on graft survival is mentioned below.
De Vos et al. also compared the performance of purified and nonpurified SA [69].
The capsules had a larger diameter and pore size than those of [68].
In order to test the effect of the immune response against allogeneic islet
transplants, isografts were performed. The mean graft survival and percentage of
capsules retrieved was very similar with both iso- and allografts, indicating that the
limited graft function was not caused by insufficient immunoprotection of the
allogeneic tissue.
The biocompatibility of the grafts were influenced both by the chemical composition of the alginate and by the individual imperfections of the microcapsules (similar
to the case of extravascular hollow fiber devices). Graft failure could have had several
causes, including cytokine release (induced by small fibrotic reactions toward
exposed islet tissue), insufficient number of islets (leading to graft exhaustion),
and problems of nutrient diffusion.
The main cause of limited duration of graft function, in this case, seems to be the
presence of the capsule itself combined with the choice of implantation site. The
lifespan of the islets is very much conditioned by the implantation site. Functional
microencapsulated islets remain free floating in the peritoneal cavity. They depend
completely on diffusional phenomena for supply of nutrients, growth factors, and
other substances essential for the homeostasis of islet cells. This fact necessarily led
to the idea of minimizing, as much as possible, the size of microcapsules, but without
compromising the integrity of islet encapsulation procedures. However, for longterm survival of the graft, further functional studies on the diffusion and its
limitations are needed.
Effect of Microcapsules Size/Mechanical Stability Chicheportiche and Reach studied, in vitro, the effect of PLL microcapsules size in insulin secretion of microencapsulated islets [71]. They used 47 kDa MW PLL polymer chains to produce two
populations of microcapsules: 350 mm (small) and 650 mm (large) average diameters, respectively. Static glucose challenges with evaluation of insulin concentrations were performed. Insulin secretion increased significantly with small
microcapsules, but insulin secretions of both groups of microcapsules were much
lower than that of free islets. These results clearly showed the significance of capsule
diameter in the diffusional process and, consequently, on the encapsulated islets’
survival. Islets encapsulated by small microcapsules respond faster to stimulation
than larger microcapsules, possibly due to the time taken for nutrients and
metabolites to reach the islet within the capsule, and/or accumulation of insulin in
the microcapsule. The driving force for diffusion is the difference in concentration
across the membrane and the surrounding milieu, therefore the exchange of
stimulatory agents and the ensuing secretory compounds will result in a longer
time lag for larger microcapsules.
De Vos et al. also studied the influence of the size of the microcapsules. They
hypothesized that a portion of microcapsules in any transplantation procedure
would always be overgrown with fibrotic tissue due to the unavoidable presence of
protruding islets. They used Sephadex beads (S-beads) to simulate islet encapsulation and studied their distribution and placement in PLL microcapsules, using a
common encapsulation protocol. S-beads immobilized in Ca-alginate beads (no
PLL outer capsule) were mostly located in the periphery of the latter, although some
were located centrally; a small number of S-beads even protruded from the
Ca-alginate beads. APA microcapsules with immobilized S-beads showed a higher
number of these beads of protruding from the capsule, than those with Ca-alginate
beads [66].
In another set of experiments [70], they used microcapsules of two different size
ranges and assessed the effect of islet protrusion from the microcapsules in graft
survival. With small sized microcapsules only one in seven recipients became
normoglycemic and remained so for the first 8 weeks after transplantation. Thus,
reducing the microcapsule diameter was associated with not only a substantial
reduction in volume but also a substantial increase in the percentage of inadequately encapsulated islets. Inadequate encapsulation is a major factor inducing
fibrotic overgrowth and subsequent necrosis of the islet cells, and may, therefore,
lead to graft failure. The inflammatory response developed seemed to be proportional to the number of inadequate microcapsules. The percentage of inadequate
capsules was higher for small-sized microcapsule (24 %) than for larger sized
microcapsule (6 %) preparations. In contrast, all five recipients became normoglycemic for 7–16 weeks after transplantation with islets encapsulated in large
capsules. The majority of capsules were found free floating in the peritoneal cavity.
Occasionally, some large capsules were found to be adhering to the omentum.
However, no islets were found in these attached large capsules. The islets in
nonadherent capsules were found to contain large necrotic zones with only a few
viable b cells left. The main factor that seemed to lead to graft failure with large
microcapsules was islet necrosis, brought about by insufficient nutrition and
oxygenation.
These results seemed to be in disagreement with other reports [20,75–77] using
small microcapsules (250–350 mm in diameter). Size of the microcapsules, however, is not the only factor influencing graft survival. Alginate purity, PLL MW cutoff, and especially the mechanical stability of the microcapsules are very important,
and may influence the success of the encapsulation procedure.
Mechanical stability of APA microcapsules was addressed by Hsu et al. [73], who
evaluated microcapsules membrane integrity and its influence on immunoprotection. Calcium concentration in the alginate gel was increased to prepare predictably
more compact APA microcapsules, and their performance was compared to those
293
294
of a standard APA microcapsule preparation method [82]. Although APA microcapsules protected the encapsulated islets from direct antibody attack, neither
could prevent donor antigens from diffusing out of the capsule and inducing host
immune responses: cytokines involved in tissue rejection are small enough in size,
can enter the microcapsules, and damage the islets. They can also induce fibroblast
proliferation, causing the disintegration of the microcapsule.
It was very likely that the membrane was rendered more resistant to biodegradation because of its more compact membrane, with less water molecules and
stronger and closer ionic interaction between interacting polymers, emphasizing
that better in vivo long-term glycemia control results were obtained with compact
APA microcapsules.
Lanza et al., who had performed several hollow fiber studies, also studied the
mechanical stability of standard APA and newly developed biodegradable microcapsules, varying alginate concentrations [74]. Calcium alginate microcapsules
coated with a 10 kDa MW PLL plus an external alginate layer were produced (800–
1200 mm in final diameter). The alginate concentration was varied (0.75–1.5 %) to
produce different sets of microspheres, with which mechanical stability and
permeability studies were performed. Good in vivo mechanical stability of the
microcapsules was observed only for alginate concentrations of 1.5 %. No capsule
breakdown was observed when 1.5 % alginate microcapsules were transplanted,
and virtually all of the microcapsules were recovered from the animals and had
remained intact for the whole duration of the study (1 year). IgG was detected on the
external surface of the capsules, but no appreciable amounts of IgG were found
inside. The resulting microcapsules allowed transplantation of discordant bovine
and porcine tissue into rats and excluded IgGs for over a year, while maintaining
islet viability similar to preimplantation levels.
One can conclude from these experiments that membrane and alginate gel
characteristics synergistically interact with each other and together determine the
durability, integrity, and biocompatibility of APA microcapsules.
Fibrotic Response to PLL-Microcapsules Islets encapsulated in APA-microcapsules
have also been used as a BAP for evaluation of the induction of a fibrotic response [83,84].
In the work of Fritschy et al., Lewis rat islets were encapsulated in APA microspheres
and implanted in STZ-induced diabetic AO rats. The capsules prepared were of 800–
900 mm diameter [83]. Although microencapsulated islet allografts were able to restore
normoglycemia, there was reduced insulin secretion. The microcapsules used were
relatively large, thereby delaying insulin and glucose diffusion kinetics. Insufficient
islet mass transplanted could also have contributed to graft failure, but the most
important factor was the occurrence of fibrotic overgrowth, further hindering diffusion
of nutrients and metabolites.
The intensity of fibrotic response can vary greatly from individual to individual, even
within the same species. Phylogenetic differences of islet donors and recipients
strongly influence the fibrotic overgrowth of microcapsules [84]. Hsu et al. used islets
from three different rat and mice donor species. They were microencapsulated in APA
microcapsules and transplanted intraperitoneally into healthy concordant and discordant mice or rats. Three parameters were assessed: islet recovery ratio (IRR),
neogrowth, and glucose-stimulated insulin secretion (GSIS).
With C57BL/6 mice as recipients, neogrowth in syngeneic islet transplantation was
significantly less than that of allo- or xenotransplantation. When these mice were used
as donors, however, the neogrowth score was less severe than that of syngeneic
transplantation and the IRR was significantly higher for allogeneic transplantation.
No islets were recovered from xenografts. These results showed that phylogenetic
disparity affects graft survival, but the authors were unable to clarify the immunological
processes involved.
Barium-Alginate Microcapsules Instead of using alginate coated with PLL membrane, several researchers [85–87] crosslinked alginate microdroplets with barium
ions and developed beads that are usually referred to as Ba-alginate microcapsules.
Their performance as immunoprotection barriers has been studied from then on
(Table 9.3).
In vitro Biocompatibility Assessment In vitro studies were performed to test the
properties of Ba-alginate microcapsules with pig islets [85]. Encapsulation of islets in
Ba-alginate was shown to be a simple and quick method of immunoprotection,
allowing islet functionality to be preserved, although insulin output was greatly
diminished when compared to that of free islets.
A similar in vitro study compared the performance of Ba-alginate capsules versus
thin layer (10 mm) alginate-coated islets [86]. The results also emphasized the
importance of having low microcapsule thickness. Still the effectiveness of such
thin capsules as immunoprotective barriers was not always good, especially due to
lower mechanical stability and deficient islet encapsulation.
Effect of Microcapsule Size Concerned with the microcapsule diffusional limitations, Petruzzo et al. developed a different encapsulation process, allowing the
production of Ba-alginate capsules of 200 mm average size, within the size range
of 50–500 mm, and pore size between 60 and 70 kDa [88].
The capsules retrieved from in vivo rat experiments showed well-preserved islets
and absence of fibrosis, if they were still from normoglycemic animals. Those
removed from animals that regressed to a hyperglycemic state were free of fibrosis
but contained unhealthy islets. The authors attributed this to damage to the islets
during the isolation and encapsulation procedures, because porcine islets are
extremely sensitive and prone to disintegration.
Assessment of Fibrotic Response Evaluation of fibrotic response to Ba-alginate
capsules was extensively performed in syngeneic [89], allogeneic [87,89,90], and
xenogeneic [79,91] transplantation experiments.
Allogeneic transplantation of Wistar rat islets into preimmunized and non-preimunnized STZ-induced Lewis rats, to evaluate recipient immunological response,
was performed [90]. There was tissue reaction by the host, which seemed to be
295
In vitro assays
In vitro assays
10 STZ-diabetic
rats (species
not mentioned)
i.p. transplant
50 rat or porcine
m.i.s (species
names not
mentioned)
50 Lewis rat
m.i.s for static
and 100 Lewis
rat m.i.s for
perfusion
assays
150–200 porcine
free islets or
m.i.s in vitro
5000–7500 m.i.s
in vivo (species
not mentioned)
Recipient/
implantation site
M.c.s Ø:
50–500 mm,
average
diameter
200 mm;
60–70 kDa
pore size
N/A
N/A
Microspheres
parameters
Experiments with Ba-alginate beads.
Number of islets
or islet cells
Tab. 9.3
Ba-alginate beads
Ba-alginate beads,
with or without
a 10 mm-thickness
alginate coating
Ba-alginate
beads
Encapsulation
method/material
In vitro: 135 min
glucose
perfusion
experiments
In vivo: 2,
4, 12, and
24 wks
120 min static
glucose
stimulation
Perfusion glucose
stimulation
for 60 min
120 min static
glucose
stimulation
7 or 14 days
in culture
Experiment
duration time
Until graft failure
(two consecutive
blood glucose
levels higher
than 10 mm)
–—
–—
Device
removal reasons
M.c.s recovered from normoglycemic
animals: well-preserved islets; no
fibrotic overgrowth; m.c.s recovered
from hyperglycemic animals: free of
fibrotic overgrowth, but contained
unhealthy islets
Rat m.i.s responded to glucose
static challenges at 7 and
14 days with a fivefold increase
in insulin secretion, substantially
lower than the insulin output
of free-floating islets
Porcine m.i.s had poorer insulin
secretion response to the glucose
challenge, which worsened
after 14 days
M.i.s and alginate-coated islets had
similar insulin response to
static glucose challenges.
M.i.s had a 2-min delay compared
to controls, in perfusion
experiments; alginate-coated
islets had no measurable delay
No observable difference of insulin
secretion of free islets versus m.i.s
in perifusion experiments, in vitro
In vivo xenografts functioned
for 17–180 days
Major outcomes/results
[88]
[86]
[85]
Ref.
STZ-induced
diabetic Lewis
rats i.p.
transplant,
seven immunized
and nine
nonimmunized
STZ-induced
diabetic Lewis
rats i.p.
transplant,
seven syngeneic,
10 allogeneic,
four with
empty capsules
6000 Wistar rat
free islets
and/or 3500
Wistar rat m.i.s
6000 Lewis
and Wistar
rat free islets
or 3500 Lewis
and Wistar
rat m.i.s
M.c.s Ø:
600–700 mm
m.c.s average Ø:
600 mm, 1 or
2 islets/m.c.
Ba-alginate
beads
Ba-alginate
beads
15 wks
1 month
Graft failure or
the end of
the experiment
Relapse into
hyperglycemia
or the end of
the experiment
Four out of nine nonimmunized
recipients of m.i.s remained
normoglycemic for 1 month;
the remaining five stayed
normoglycemic for 28 days,
or less
Five out of seven preimmunized
recipients were normoglycemic
on day 28; nonrelapsed into
hyperglycemia on days 7 and 10.
Intact tissue was present in devices
explanted from normoglycemic
animals
In failing animals, islets were destroyed
or badly damaged and fibrotic
overgrowth was increased
No differences between immunized
and nonimmunized recipients
Reversal of hyperglycemia in six of
seven animals with syngeneic
transplants for 15 wks
Two out of 10 recipients of allogeneic
transplants stayed normoglycemic
for more than 100 days; remaining
eight animals stayed normoglycemic
for 0–56 days
Negligible foreign body reaction
to empty m.c.s
Slightly higher reaction to syngeneic
m.i.s than empty m.c.s
Allogeneic normoglycemic transplanted
animals: m.i.s free from fibrotic
overgrowth, with viable islets;
relapsing hyperglycemic animals:
m.i.s with severe fibrosis,
nonviable islets present
[89]
[90]
Recipient/
implantation site
Four groups of
rats: 25 Balb/c
mice treated
or untreated
with HOE 077,
transplanted
with empty m.c.s
10 Balb/c mice
treated or
untreated with
HOE 077
transplanted
with m.i.s, in
the peritoneal
cavity or the liver
STZ-induced
diabetic NOD
mice (seven
with free islets,
12 with m.i.s)
and Balb/c mice
(three with free
islets, 11 with
m.i.s), i.p.
transplant
Porcine islets
(number of
islets N/A)
9000–10 000
NOD or
B6AF1
mice free IEs
or m.i.s
(Continued )
Number of islets
or islet cells
Tab. 9.3
49, 105, and
210 days,
for m.i.s
1 month for
empty m.c.s
Ba-alginate beads,
highly purified
high-G or
high-M alginate
M.c.s Ø:
900–1100 mm
in vitro: static 24 h
assays, on 2
consecutive days
In vivo: 4 wks
Ba-alginate
beads
M.c.s Ø:
250–450 mm,
Experiment
duration time
30–50 m.c.s
implanted
in the liver
Encapsulation
method/material
Microspheres
parameters
The end of the
experiment,
graft failure
The end of
the experiment
Device
removal reasons
Ba-alginate m.i.s allogratfs survived
for >340 days and xenografts
for >350 days
Retrieved m.c.s free of capsular
overgrowth and with viable islets
Retrieved m.i.s had a 5–10-fold increase
of glucose-stimulated insulin
secretion (perfusion)
23% encapsulated islet viability,
after 4 wks, in recipients treated
with HOE 077, thinner fibrotic
overgrowth
No difference detected between
the HOE 077 treated and untreated
groups after i.p. transplant
Administration of HOE 077 significantly
reduced fibrosis after intrahepatic
transplant
Percentage of retrieved empty m.c.s
lower for high-G than high-M
Ba-alginate
15 % encapsulated islet viability in
untreated recipients, 4 wks after
intrahepatic transplantation;
high fibrotic overgrowth
Major outcomes/results
[87]
[91]
Ref.
10 000 IEs of
free or
encapsulated
NPCCs
STZ-induced
diabetic B6AF1
mice, i.p.
transplant, 32
with m.i.s
Same as
in [87]
Ba-alginate beads,
highly purified
high-G or
high-M alginate
2, 6, and
20 wks
The end of
the experiment
Two-six out of 32 animals transplanted
with microencapsulated NPCCs had
normalized blood glucose levels
for 20 wks
M.c.s removed at 2, 6, and 20 wks still
had increased insulin secretory
responses to static glucose challenges
High m.c. recovery: 91, 85, and 93%
at 2, 6, and 20 wks, respectively
80 % of m.c.s free of overgrowth
at 2 and 20 wks
[79]
300
triggered by alloantigens that had been released out of the bead, suggesting the need
for a reduction in pore size. Also, in failing animals most of the islets were destroyed
or severely damaged and the fibrotic and inflammatory reaction toward beads was
markedly increased, when compared to transplanted animals still normoglycemic.
No differences in inflammatory response were observed for immunized and nonimmunized animals (immunization will be further discussed in Section 9.2.4).
The same group of researchers proceeded with syngeneic or allogeneic islets
transplants, for longer periods [89]. Syngeneic transplanted microbeads were mostly
recovered with little cellular or fibrotic overgrowth and had mostly viable cells. The
reaction to syngeneic implants was only slightly bigger to that of empty microcapsules. When allogeneic transplant was performed, only the animals that sustained
normoglycemia for more than 100 days had microcapsules free from fibrotic overgrowth and containing viable cells; those retrieved from recurring diabetic animals
had severe fibrotic overgrowth, nonviable islet remnants, and decreased number of
microcapsules.
Zhang et al. verified that a strong fibrotic response occurred in vivo to Ba-alginate
material and that the use of anti-fibrotic agent HOE 077 (developed as an anti-fibrotic
agent for human liver fibrosis therapy) could significantly reduce it if encapsulated
islets were xenotransplanted into adequate sites [91]. Empty capsules retrieved 4
weeks after implantation had a marked pericapsular infiltrate (PCI) surrounding the
capsules. Administration of HOE 077 after intrahepatic transplantation resulted in a
significant reduction of PCI formation. However, treatment with this substance after
intraperitoneal transplantation showed no inhibitory effect on fibrotic overgrowth.
HOE 077 exists in a much higher concentration in the liver and only inhibits
inactivation of cells present on this organ, which explains this outcome. This proved
that the administration of an anti-fibrotic drug could be a different but feasible
approach to improve islet viability and graft survival of microencapsulated islets.
Two highly purified alginates (high-G or high-M) were used to encapsulate mice
islets in Ba-alginate microcapsules and transplanted into the peritoneal cavity of NOD
and Balb/c mice. The ability of such a microcapsule as an immunoprotective barrier
against allorejection, was compared to crude SA beads coated with a PLL membrane,
namely in the prevention of a fibrotic response [87].
Overall, the Ba-alginate capsules made of highly purified alginate proved to be
stable and highly biocompatible and protect against allorejection and autoimmunity
in murine models, even without a PLL barrier. If such a barrier provides the same
kind of immune protection in larger animal models is still unknown.
Considering these positive results, these highly purified alginate capsules were
used to determine their immunoprotective ability for neonatal pancreatic cell clusters
(NPCC), when these are transplanted into STZ-induced diabetic B6AF1 mice [79].
Aside from graft survival or rejection times, the proliferation and maturation of
NPCCs in the Ba-alginate capsules were also evaluated.
In animals with successful grafts, maturation of the NPCCs significantly increased
with time. Thus, in agreement with a superior number of b-cells, the total insulin
content per islet equivalent also increased in the NPCCs: 4-fold, 12-fold, and 10-fold,
at 2, 6, and 20 weeks, respectively.
Recipient IgG was found throughout the NPCC-containing capsules at 6 and
20 weeks after transplantation, which confirms that IgG diffuses through the gel
without exerting any adverse effect on the encapsulated NPCCs.
The authors could not fully explain why these simple capsules worked so well for
xenografts, but they advanced two probable reasons for it: the avoidance of PLL, which
might work as an adjuvant of xenograft reaction and the use of highly purified
alginate, with very low endotoxin levels. This experiment showed that it is possible to
successfully transplant encapsulated NPCCs into diabetic mice and that they can
differentiate into b-cells in this environment, thus providing a large number of
insulin-producing cells for transplantation.
Comparison of APA With Ba-Alginate Although these materials have been extensively used to develop microcapsular BAPs, their comparison has been studied using
macrocapsular devices [19] to avoid interferences due to a nonsmooth outer membrane surface, the existence of protruding islets, or the clumping in vivo of the
microcapsules.
Trivedi et al. performed and compared two different islets macroencapsulation
protocols (protocol 1: islets are suspended in alginate, drops are gellified and coated
with PLL and finally covered with alginate, as already described; protocol 2: islets are
suspended in alginate and the droplets crosslinked with barium ions). Both kinds of
macrocapsules (35 islets/each; 2–3 mm in diameter) were studied in vitro and
in vivo by the syngeneic transplantation of encapsulated Lewis rat islets into the
peritoneal cavity.
Static glucose challenges performed with Ba-alginate macrocapsules, before graft
transplantation, showed these had significantly lower total insulin secretion in basal
media, compared to free islets, and no significant response on exposure to a high
glucose challenge.
STZ-induced diabetic Lewis rats syngeneically transplanted with protocol 1 or
protocol 2 macrobeads became normoglycemic significantly faster than animals with
free islets transplanted under the kidney capsule. Regarding protocol 1 graft
performance: after a meal challenge, no significant difference was observed in
the plasma glucose profiles between alginate-PLL macrobeads and kidney capsule
recipients. The plasma insulin levels for the kidney capsule group displayed a sharp
increase after a meal, unlike the macrobead group, for which plasma insulin
remained at basal level. Grafts using protocol 2 macrobeads showed divergent
behavior: after an overnight fast, animals with Ba-alginate macrobeads had significantly lower plasma glucose levels than kidney-capsule grafted animals. Also, their
plasma insulin levels increased significantly compared to basal insulin levels in
response to a meal at 100 and 120 min.
During the intravenous infusion tests, both rat groups, protocol 1 and protocol
2 macrobeads, had no measurable increases in plasma insulin, whereas in control
rats (normal rats or diabetic rats with free-islet grafts) there were clear insulin
responses.
Islet retrievability after 5 weeks of transplantation was larger for Ba-alginate
(80–90 % intact) than for APA macrocapsules (50% broken) indicating lesser
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mechanical stability of the materials in the latter. Retrieved capsules from protocol
1 had mostly viable islets but with some evidence of central necrosis, unlike those of
protocol 2 for which islets had a healthy appearance.
Several reasons were advanced to explain the poor insulin secretion dynamics
observed. It may, in part, be due to the implantation site, as was shown by other
researchers, although the authors thought that diffusion limitations were responsible for the absence of acute insulin responses. This directly relates to the large size
of the macrobeads, which unquestionably induced time lags in glucose and insulin
diffusion. The residual negative charge on alginate may also play a role in binding
the insulin to the gel, which would slow the dynamic response to glucose
stimulation.
Agarose Beads Without Immunoprotecting Membrane Since a synergistic effect
seemed to occur between the agarose gel, PLL membrane, and the immune response
reaction [67,79], instead of conjugating alginate bead formation and PLL membrane
or other polymeric membrane coating, some authors tried to use simple agarose
beads to microencapsulate islets and tested their performance in vivo.
The immunoprotective efficiency of agarose capsules was tested by Iwata et al. [92].
C57BL/6 mice islets were microencapsulated with agarose and implanted in STZBalb/c and NOD mice. Balb/c mice remained lifelong normoglycemic and four of
five NOD mice had functional graft survival for more than 80 days. It was therefore
determined that a capsule made of this material only can effectively immunoprotect
islets and prolong graft functioning in these recipient mice, without any
immunosuppression.
Comparative experimental microcapsular geometry studies with agarose gel
constructed BAP were also performed [93]. Three geometries were studied:
microbead, rod, and disc. The devices were in vitro and in vivo analyzed. In vitro
examination showed the islets survived well and released basal insulin regardless of
the geometry of the device used. In vivo, five of the eight rods and five of the six disc
recipients could not function for longer than 15 days, although low dosages of 15deoxyspergualin, an immunosuppressive drug, were used. This occurred even
though no difference in fibrotic response could be observed between the three kinds
of devices. The best graft survival was obtained with microcapsules, which was
attributed to their smaller diameters and lower diffusion times.
These same capsules were used for allotransplant of microencapsulated islets in
dogs [94]. Implantation of microencapsulated islets completely supplanted exogenous insulin therapy in three out of the five recipients for 28, 42, and 49 days. Good
fasting glucose control was achieved but response to an intravenous glucose tolerance
test remained abnormal. The histologic appearance of retrieved allografts revealed
viable islets and no evidence of adherence or infiltration of immunocytes or
inflammatory cells. Thus, although with limited success, they showed it is possible
to use these microcapsules to reverse IDDM in a larger animal model.
Agarose hydrogel microcapsules have also been used to relate graft survival time
with the amount of islets encapsulated. Ao et al. [95] compared the performance
of microcapsule devices, with different amounts of islets seeded, to the performance
of free islets (xenogeneic mongrel dog islets transplant into alloxan-induced
diabetic nude mice). Long-term graft survival occurred when 500 encapsulated islet
equivalents (IE) or more were implanted; best results were obtained with 1000 IE
(all mice transplanted had more than 119 days of graft survival). Two thousand free
islets implanted in the kidney capsule, or more than 4000 islets intraperitoneally,
were needed to reverse diabetes. These results have, however, a limited scope since
diabetic nude mice are immuno-incompetent animals.
The above-mentioned results are modest, when compared, for example, to those of
[68], which used similar experimental parameters but encapsulated islets in APA
microcapsules. These authors had better graft survival times, both with dog islet
allografts and xenografts, and the microcapsules were usually free of fibrotic overgrowth (except if islets were exposed or protruded from the capsule) (see Table 9.2).
Comparing both results, the number of islets transplanted seems to be crucial to
obtain long-term graft survival and fibrotic overgrowth appears to occur not due to the
presence or absence of the PLL membrane or to the alginate gel used but rather due to
an efficient coating of the islets, thus preventing direct exposure of islet tissue to the
host immune cells.
Agarose-PSSa Microcapsules A mixture of agarose crosslinked with polystyrene
sulfonic acid (PSSa) coated with a polyanionic carboxymethyl cellulose (CMC)
membrane material was also used for the microencapsulation of pancreatic islets or
islet cells. The first time this kind of microcapsule was used, a hamster-to-mouse
xenotransplantation was performed for material biocompatibility assessment [96].
The longest survival time achieved was 90 days, and three out of five NOD
recipients showed prolonged normoglycemia.
Biocompatibility experiments proceeded with xenotransplantation of cells from
a b-cell line, MIN6, to the subcutaneous space of STZ-induced diabetic Lewis rats
[97]; the MIN6 cell line seems to retain the characteristics of glucose metabolism
and GSIS similar to those of normal pancreatic islets cells.
Microspheres were prepared with agarose and PSSa as before to yield droplets of
300–500 mm in diameter, and these were covered with hexadimethrine bromide to
form a polyanionic complex, preventing PSSa from leaking, and further coated with
CMC. Two thousand microcapsules were formed, seeded with MIN6 cells and
implanted in the subcutaneous site for a week, recovered and, afterwards, subjected
to in vitro static glucose challenges.
Retrieved agarose-PSSa microbeads were intact, spherical, and smooth. Cells
were found viable inside the capsules, and adhesions or fibrotic overgrowth were
rarely found in the subcutaneous site. The in vitro static glucose challenges
showed that the encapsulated cells had a 2.7-fold increase of insulin secretion
relative to basal secretion levels, which confirmed cell functionality. The
studies evidenced the microcapsules biocompatibility and immunoprotective
properties.
Microspheres alone were rarely used for implantation in the subcutaneous site.
However, tubular capsules together with b-FGF impregnated microspheres, both
made of PSSa, were tested in [48], as was discussed in Section Neovascularization of
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Macrocapsules. Average graft survival time was 68.4 25.6 days, also due to
neovascularization induced by the presence of the b-FGF impregnated beads.
These results confirmed and improved the ones from Ref. [97].
The biocompatibility of the material produced by this crosslinking method was
thus confirmed, since no fibrotic overgrowth or pericapsular cellular infiltrate
were observed either with b-FGF+ and b-FGF– groups, even though islets retrieved
from the second group were found necrotic. Nevertheless, graft survival time was
reduced, when compared to similar xenograft experiments with Ba-alginate beads
[89].
Regarding the success in long-term reversal of diabetes with this type of BAP,
PSSa capsules were used in Hamster-to-rat xenotransplantation studies, this time
to determine the amount of islets needed to reverse diabetes in this model [98].
Long-term graft survival was only achieved in recipients with 3000 encapsulated
islets, the mean period of normoglycemia being 124 days. Encapsulated islets,
recovered from recipients when they returned to hyperglycemia, had mostly
retained their morphologic appearance but lost functionality. A fibrous layer
surrounded the surface of the microbeads and was probably the primary reason
for graft failure. In spite of this, the authors believed these microcapsules could be
used to reverse diabetes in higher mammals, with a more complex immune
system.
The ability of the agarose-PSSa microspheres to prevent porcine islet degradation had been determined in vitro [94]. The encapsulated porcine islets were kept
morphologically intact, viable, and functional during 7 days in culture, whereas the
nonencapsulated ones were rapidly disintegrated. An extensive study was then
conducted to examine the use of agarose-PSSa microcapsules in xenogeneic
transplantation of islets in higher animal models and determine if their immunoprotective ability is as good as that of smaller animals [99].
In the larger animal discordant xenotransplant, transplantation of pig islets into
pancreatectomized diabetic beagle dogs, two out of five animals became normoglycemic without exogenous insulin for 50 and 119 days [99]. Agarose-PSSa
microcapsules could protect the xenogeneic islets from humoral attack by the
host’s immune system. The limited success was attributed especially to the
difficulty of isolating and handling pig islets. Pancreatic pig-to-dog xenografts
with microencapsulated islets were not attempted again.
These results are slightly better than those of Ref. [14], without the problem of
device occlusion, which occurred with intravascular devices. Even better results
than those presented in Ref. [99] were obtained by Ref. [40], which achieved
xenograft survival for more than 8 months. However administration of a low
dosage of exogenous insulin was necessary, which might have prevented an earlier
recurrence of hyperglycemia.
Islets or Islet Cells Covered With a Siliceous Membrane Over the last 7 years,
silanization of islets (reaction of silanes with exposed hydroxides on the cell surface)
has been proposed as a method of microencapsulation. The potential for long-term
survival of viable and functional islet grafts was evaluated [100–104].
The permeability of two sol–gel synthesized silicate membranes was assessed [102]:
Si(OCH3)4 (tetramethoxysilane, TMOS) and (CH3O)3Si(CH2)3NH2 (3-aminopropyltrimethoxysilane, APTrMOS). Microcapsules were prepared in a similar manner as
APA capsules, with Ca- or Ba-alginate, although the siliceous substances were used
instead of PLL, and their permeabilities to glucose, myoglobin, ovalbumin, bovine
serum albulin (BSA), and g-globulin were determined.
BSA and g-globulin could adsorb onto the silicate derived from TMOS alone. Only
the membrane derived from APTrMOS allowed g-globulin (the biggest molecule used
in this study) to permeate, even with high concentrations of this siliceous substance.
Permeability assays were performed with Ca-and Ba-alginate capsules coated with a
mixture of silanes (best ratio APTrMOS/TMOS ¼ 2.40). For this ratio of silanes,
coated Ba-alginate beads had good permeability to g-globulin, but coated Ca-alginate
beads could reject g-globulin. This membrane was found to have a MWCO of
approximately 60 kDa. These changes in permeability could be explained by the
change of electric charge of substances used during capsule formation and electrostatic interaction of Ba and Ca ions with amino and carboxyl groups.
The in vitro insulin secretion indexes of encapsulated rat islets were shown to be
dependent on the molar ratios of the silanes [103]. For APTrMOS/TMOS ¼ 0.6, onefourth of the ratio for optimal permeability, the ability of islets to secrete insulin upon a
glucose stimulus was hampered, while for ratios of 1.2 and 2.4 stimulation was similar
to that of free islets.
In vivo xenotransplantation studies (encapsulated Wistar rat islets peritoneally
inserted into STZ-induced diabetic DDY mice; 1000 islets/recipient) demonstrated
marked cellular overgrowth just after the 21st days of transplant. That led to islet
necrosis and recurrence of hyperglycemia. Some mice (6/10) were able to maintain
normoglycemia levels up to 2–3 months after transplantation. The cause of overgrowth, although not due to the chemistry of the membrane, had not yet been defined
[104].
Desai et al. analyzed, in vitro, the behavior of islet cells when in contact with the
microfabricated immunoprotecting membrane, using two microfabricated half capsules bound together with silicon elastomer adhesive or when direct deposition of a
thin film of siliceous material on the cell’s membrane surface was performed [100].
Preliminary biocompatibility studies had already been performed with silicon membranes, which suggested that silicon substrates were well tolerated and nontoxic in
vitro and in vivo [24].
Using a microfabrication technology (combination of UV-lithography, silicon thinfilm deposition and selective etching of sacrificial oxide layers) a homogeneous
siliceous membrane was formed. This technology allows thickness and pore size
uniform distribution to be controlled, while maintaining original cell morphology. The
microfabricated biocapsule had 78 nm pore-sized membranes. Rat islets were placed
in the microfabricated capsules made of boron-doped silicon and polysilicon
(Figure 9.4), cultured, and their viability and functionality assessed. The insulin
secretion response obtained with encapsulated and free islets was similar, which
validated the pore density and length chosen when constructing the microcapsules,
and showed these biocapsules provide a stable functional environment for the islets.
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Fig. 9.4 Diagram of basic microfabricated
immunoisolation biocapsule concept.
Source: Reproduced from Desai et al. [24] with permission
from Elsevier, copyright 2000.
Insulin secretion levels still decreased after 1 month of culture, though. After 4 weeks
islets removed from the capsules showed very little central and no peripheral necrosis
and were attached to the inorganic substrate.
In vitro and in vivo assays were then performed with micromachined biocapsules of
18 or 66 nm pore-sized membranes and insulinoma cells (RIN and bTCC6F7) [24].
The in vitro assays showed both kinds of insulinoma cells had higher stimulatory
indices (stimulatory/basal insulin secretion) with 66 nm than with 18 nm membranes, even though the latter had higher immunoprotection abilities. In vivo,
microfabricated capsules with RIN and bTCC6F7 cells remained mechanically intact
until removal (8 days postimplantation). Both types of transplanted microencapsulated
cells remained viable and had higher stimulatory indices with 66 nm than 18 nm
pore-sized membranes in vivo, indicating that the pore size greatly affected the
secretory response of the cells. However, in vitro experiments with recovered capsules,
showed a sharp reduction in the stimulatory indices of both cell lines with 66 nm
biocapsules, but not with 18 nm biocapsules, when compared to those of nonimplanted microencapsulated islets. This suggests a good immunoisolation effect of the
18 nm pore-sized biocapsules, which was lacking with the 66 nm biocapsules.
Although no long-term in vivo experiments were performed so far with these
micromachined biocapsules, they remain a very attractive alternative, since microfabrication technology allows the precise control of pore size, configuration, and
distribution, and thus protect encapsulated islets from the host immunologic
response.
Recently, another research group produced the microcapsules by direct deposition
of a siliceous material, after suspending the islets in a filter holder and fluxing
them with air saturated by CH3SiH(OEt)2 and Si(OEt)4 (50/50 molar ratio). The
siliceous membrane formed had at least 0.1-mm thickness (BioSil membrane).
Capsule performance was evaluated in vitro and in vivo [101].
This encapsulation procedure allowed the islets to preserve their original
dimensions and viability (about 90 %). The membrane formed was continuous
and homogeneously distributed over the islets’ surface. Also, perifusion experiments indicated that islet function was not reduced by the encapsulation method.
The siliceous envelope did not reduce the insulin diffusion rate when compared to
free islets.
Islets from inbred Lewis rats were encapsulated with the siliceous membrane
and transplanted into Lewis or Sprague–Dawley (SD) rats, under the left kidney
capsule, and normoglycemia was restored. The implant promoted a faster return to
normal glucose levels and a prolonged effect in glycemic control than in animals
transplanted with free islets. Longer graft survival time was observed for encapsulated islets, both with syngeneic (6 out of 10 rats remained normoglycemic for at
least 70 days) and allogeneic transplants (six out of 10 rats remained normoglycemic for at least 55 days). Marked improvement was observed with encapsulated
allografts, both in terms of prolonged graft survival and faster decrease in glycemia
levels.
This clearly demonstrated that the siliceous membrane is well tolerated by
the living organism, allows proper feeding of islets, provides immunological
protection, and allows a fast response to hyperglycemia episodes, thus improving the long-term effect of the islets grafts. In view of these favorable results,
the testing of these microcapsules in a larger animal model with more complex
immunological defenses and more important diffusional limitations is
needed.
Multicomponent Microcapsules Containing Poly-Methylene-co-Guanidine At the
Vanderbilt University, Nashville, USA, Dr Taylor Wang and co-workers developed
a method for SA-CS/PMCG microcapsules fabrication in which a combination of
polyanions, SA, cellulose sulfate (CS), and the polycation poly(methylene-coguanidine) (PMCG) were used, together with a specific encapsulation technology,
resulting in polyelectrolyte complexation. The authors claim that capsule parameters, like size, wall thickness, mechanical strength, permeability, and surface
characteristics, can be controlled independently [105].
Animal trials were conducted after characterization of microcapsules (morphology, strength, molecular exclusion limits, and biocompatibility) [106] with a
rat-to-mouse intraperitoneal transplantation model. SD rat islets (1000/recipient) were encapsulated in 800 mm microcapsules (100 mm wall thickness;
230 kDa of exclusion limit) and transplanted into STZ-induced diabetic C57/
BL6 mice and into spontaneously diabetic mice. The authors followed up
transplantation in both systems, for about 6 and 3 months, respectively, before
recurrence of hyperglycemia. Microcapsules retrieved after 1 year from C57/BL6
transplanted animals were shown individualized and absent of fibrosis, and the
encapsulated islets still retained significant insulin response to glucose stimulus.
In contrast, microcapsules inserted in NOD mice were referred to as clumped and
surrounded by fibrosis; the failure possibly occurred due to immune or autoimmune attack.
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9.2.4
Influence of Recipients Sensitization to Donor Antigens in Graft Survival
Immunoprotecting barriers are far from being absolute. BAP devices may sensitize
diabetic recipient patients to donor antigen, because seeded islet tissue can still
trigger an immune response [39]. Xenogeneic humoral response can occur with
extravascular immunoisolation devices, even with a membrane that prevents passage
of large molecular weight substances. Due to the release of donor proteins by the
islets, which leak through the membrane, the host indirect immune response system
is activated. The number of islets implanted seems to condition this humoral
response, by causing an overload of the system with immune complexes. The islet
tissue mass needed to reverse diabetes may be significantly less when using other
types of BAP. Microcapsules, narrower-bore membrane chambers, or vascular
devices may evoke a smaller xenogeneic humoral response.
In order to verify the hypotheses, Lanza et al. tried a different BAP design [50].
Injectable membrane small capsules made from biodegradable materials
(MWCO < 150 kDa) seeded with bovine calve islets were introduced in STZinduced diabetic rats and mongrel dogs. Several weeks after postimplantation,
the injected capsules were recovered from the host animals. They still contained
multiple viable and functional islets, mostly free of fibrotic overgrowth, with
sufficient insulin response to high glucose levels (one to threefold and four to
sixfold increase above basal levels, in rats and dogs, respectively). The authors
successfully proved that discordant islet xenograft survival can be achieved in rodents
and dogs, without immunosuppression, with this encapsulation technology, thus
establishing its clinical utility.
9.2.5
Islet Oxygenation Studies
A crucial aspect in the prevention of islet loss of viability and functionality is to have
proper feeding and oxygenation conditions [107–112]. Schrezenmeir et al. performed extensive studies of islet oxygenation [109]. They used both piscine islet
organs, called Brockmann bodies (BBs) from Osphronemus gorami and Lewis rat
islets, together with a special recessed-type microelectrode to determine oxygenation conditions around and inside the islets in culture, correlating islet functionality with oxygen distribution throughout the islets. The islets were encapsulated in
smooth surface 12 kDa MWCO Thomapor hollow fibers (regenerated cellulose,
2.2 and 1.5 mm outer and inner diameter, respectively) [63].
Free and encapsulated rat islets were cultured in vitro for 34 days, and their
functionality determined for 5 and 34 days. The secretion of insulin by free islets
under stimulatory conditions at the beginning and end of the culturing period was
significantly higher than the basal values, and increased significantly during the
observation interval. The encapsulated islets also had increased insulin release under
stimulatory conditions, at the beginning, but considerable loss of viability was
observed at the end of the culture period. A typical finding in these cases was central
necrosis. This clearly demonstrated the burden introduced by additional unstirred
water layers through the membrane.
In vivo tests of graft survival and functionality proceeded with the implantation
of insert peritoneally encapsulated BBs and rat islets in adult male Wistar rats.
Histological findings confirmed that islet function was best preserved if the
immunoseparating membrane remained free from surrounding connective
tissue.
Finally, in vitro tests with individualized BBs were performed to determine oxygen
distribution profiles. BBs were cultured in vitro, in unstirred aerated medium (pO2
tension in medium was 140 mmHg), and the oxygen tensions were measured.
Oxygen tension inside the islet organs declined with organ depth from the pressure
of the surrounding medium following a sigmoidal curve. The decline began about
300 mm outside the organ and attained values close to 0 mmHg in the vicinity of the
organ center, at 300–400 mm from the organ’s surface. The oxygen tension fell to
30 % of the oxygen saturated surroundings at 64 mm inside the organ tissue with
basal glucose concentrations and at half this distance with stimulatory glucose
concentrations. Also, the depths for which the oxygen tension fell to 10 % of initial
oxygen value were 176 24 mm with basal glucose concentrations and
128 21 mm with stimulatory ones.
Given these results, three main aspects were made clear:
1. The diameter of unstirred water layers should be reduced to a
minimum, and the outer limit of the device should not exceed
300 mm; preventing fibrous tissue formation or inducing
vascularization was later proven to produce better results
[64];
2. Energy-consuming tissue layers should be reduced to
unavoidable dimensions;
3. Euglycemia improves the chances of survival and function
of the graft, by preventing unnecessary energy oxygen
expenditure (since O2 is needed for the degradation of glucose).
To clarify the relation between oxygen supply to the islets and their viability and
functionality and also to assess the critical tissue volumes and critical pO2 tensions for
intact islet function, other studies followed [113]. Three groups of nonencapsulated
BBs were established, identical in mass of viable cells: smaller organs (diameter
680 mm), larger organs (diameter 1120 mm), and larger organs segmented to
produce approximately three equal parts (each organ produced three segments of
diameter 805 mm). The organs were kept in culture for several days. On days 4/5
and 18/19, insulin secretion was tested for all the BBs at basal and stimulation
glucose levels. No significant difference was observed 4 days after isolation. Two
weeks later, however, the values of insulin secretion for large organs were substantially lower than those of small or segmented organs, both under basal and stimulated
conditions. This confirmed the previous hypotheses that increased energy expenditure and oxygen consumption for insulin secretion increase the decline of oxygen
tension to the center of the BB.
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Comparing the results of smaller islet organs with those of larger islet organs
showed pO2 tension had a tendency to be higher in the smaller ones at the same
defined distances, but the difference is not very significant. A shift of sigmoidal curve
to minor pO2 values occurred when higher glucose concentrations were used. In
larger organs central necrosis inevitably occurs, since pO2 values reach a point where
function is hindered. The calculated critical value for these organs is between 8 and
16 mmHg. This is lower than but close to the values reported for rat islets (12 or
20 mmHg [107,108]), which indicates that BBs are less sensitive to low pO2 than rat
islets or that the different adopted methodologies in each case may have contributed
to the discrepancy of results.
Aside from static culture of BBs for a prolonged period of time, freshly isolated BBs
were cultured also, for 14–16 h only, for pO2 levels measurement. The pO2 profiles
had a sigmoidal shape and a decline in oxygen tension beginning in the medium
surrounding the BB, as before. Oxygen tensions at 50 mm outside the organ were
consistently and significantly lower in stimulating than in basal level of glucose
concentrations.
Based on the data assessed, the authors assumed that, under normoglycemic
conditions and high oxygen tension, the outer rim of piscine islets that are sufficiently
supplied with oxygen is approximately 200 mm in depth. This distance will be
reduced if glucose concentration is elevated and oxygen tension in the surrounding
milieu is lower than the one used in this study. It also seems to be smaller in
mammalian islets.
Schrezenmeir et al. analyzed the effect of coimmobilizing the islets with hemoglobin to facilitate oxygen diffusion and obtain better islet survival, preventing
hypoxia and islet central necrosis [110]. Capillary 20 kDa MWCO Cuprophan Type
11/600 membranes were used to encapsulate the islets, suspended in sodium
alginate gel. The devices were implanted in Wistar rats and several in vitro and
in vivo experiments were performed, with and without the alginate matrix and/or
hemoglobin.
The best results in vitro were obtained with the islets suspended and hemoglobin
molecules immobilized in the alginate matrix. By the time the assay ended (34 days),
about 1/3 of the islets in the membrane were still 100 % viable (cells with intact
chromatine structure).
This membrane-gel matrix system was then tested in vivo. Very poor results
were obtained without hemoglobin, and the improvement with coimmobilized
hemoglobin molecules was not substantial (by the end of the assay only
approximately 2 % of the cells were intact). Even so, the authors concluded that
the permanent presence of hemoglobin might improve function and viability of
the immunoisolated islets. However, they could not distinguish to which hemoglobin properties this was due, oxygen donation or nitric oxide/oxygen radical
scavenging.
A recent study performed by this group, with piscine islets encapsulated in
microspheres, allowed to clarify the function of hemoglobin in the improvement
of oxygen distribution to the islets and their functionality [112]. The results obtained,
which will be discussed in further detail when microsphere devices are mentioned,
still did not completely clarify the mechanism responsible for longer graft survival,
when islets are coimmobilized with hemoglobin.
9.2.5.1 Oxygenation of Ba-Alginate Microencapsulated Islets
Schrezenmeir et al., who had already performed some studies with piscine islets, free
or encapsulated in macrocapsular extravascular devices, to determine the oxygen
distribution in the islet and its influence in graft survival, continued their studies
using Ba-alginate microcapsules.
They were the first to study the influence of convection and microencapsulation on
oxygen distribution in islet organs encapsulated in Ba-alginate microcapsules. The
oxygen measurement was performed by means of polarography using oxygensensitive microelectrodes [111].
When free BBs were cultured with and without convection, significant differences
were observed in pO2 values. With convection, the values outside and inside of BBs
were considerably higher. Without convection, the oxygen-depleted zone outside the
BBs was larger, the diameter of the oxygen-consuming rim was smaller, and the
oxygenation values in the BBs center were reduced to 6 % of the initial value. Better
results were obtained, in every case, with convection.
Microencapsulation eradicates the effect of direct convection on the BBs. The
authors tried to clarify if encapsulation exerts a similar effect as lack of convection and
the way in which it influences oxygen distribution. The oxygen profiles of nonencapsulated BBs without convection corresponded well to those of microencapsulated BBs; all electrode positions inside and outside the BBs showed significantly
lower pO2 values when compared to nonencapsulated BBs.
According to these results, the encapsulation has primarily an effect on the
extension of the oxygen-depleted zone outside the organ. Thinning down the alginate
layer, and thus achieving better oxygenation, could reduce the depth of this zone.
Compared to single mammalian islets, piscine islet organs have a much higher cell
mass, so a smaller decrease of pO2 in mammalian islets could be expected. However,
a similar cell mass to that of BBs can be attained when mammalian islets are packed
in a BAP. To avoid these problems, the authors suggested a packing mechanism
using alginate immobilization for increased spacing between islets or singlecell suspensions and choosing appropriate transplantation sites, with increased
convection.
They also showed, in an additional study, that islets from different animal models
can have different oxygen requirements and, therefore, different hypoxia-tolerance
capabilities [114]. For example, rat islet cells cultured in 38 mmHg tolerate hypoxia
better than porcine islets, and the neonatal form is more hypoxia tolerant than the
adult, which means the choice of the islet donor species can influence the outcome of
xenotransplantation.
As a continuation to another study already performed with hollow fibers, a study
was performed to clarify whether hemoglobin may improve oxygen supply within the
microcapsules by increasing its oxygen solubility [112]. In the previous study the
presence of coimmobilized hemoglobin definitely improved encapsulated islet
organs survival but the mechanism responsible for this was not made clear.
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In this study, the islets were cultured free or encapsulated, with or without
hemoglobin, and functional tests were performed, in vitro. Naked BBs had considerably higher pO2 outside and inside the islet tissue than BBs encapsulated
without any hemoglobin and these had higher pO2 at any electrode position than islet
organs encapsulated with hemoglobin. However, on the second and third days of
culture, BBs of all groups showed normal function with a significant increase of
insulin release as a response to a glucose stimulus, when compared to basal insulin
secretion levels. After 4 weeks both basal and stimulated insulin secretion had
decreased in all groups, particularly the stimulated secretion. Viability after 4 weeks
of culture was similar for naked BBs and BBs coencapsulated with hemoglobin
(approximately 65 %) but lower for BBs encapsulated without hemoglobin (approximately 55 %).
In the previously mentioned report, it had already been shown that microencapsulation reduced oxygen availability and tension in the islet tissue. In this study
this was confirmed. Also, the addition of hemoglobin to the capsule matrix
significantly improved the viability following long-term culture. The oxygenconsuming rim tended to be higher in BBs encapsulated with hemoglobin
compared to BBs encapsulated without hemoglobin. The oxygen plateau reflecting
central necrosis tended to be lower with hemoglobin in the matrix. The improved
survival of islet tissue in BBs encapsulated with hemoglobin does not seem to be a
result of increased oxygen supply, though. Other factors must be considered, since
the introduction of hemoglobin into the capsule tended to reduce the oxygen
tension at the organ surface and significantly decrease oxygen tension at the organ
center. Thus, the mechanism by which the beneficial effect of hemoglobin on islet
cell survival is mediated remains unclear, but it is definitely not due to improved
oxygen availability.
9.3
Final Comments
From this review of the developments toward a BAP device, major aspects are
evidenced:
i. None of the devices developed so far can bring fibrotic response
to acceptable negligible levels;
ii. Diffusional limitations imposed by the encapsulating material
and the capsule size are crucial for graft success.
The first aspect remains as one of the main reasons for graft failure and limited
graft survival. Regarding the second aspect, although some authors achieved good
results with microcapsules, graft survival time and response time to a glucose
stimulus still have to be improved. The use of new biomaterials and/or of coimmobilization of other molecules, cells or cell aggregates with the islets, are two possible
strategies to overcome this problem. Also, device design other than the ‘‘classical’’
macro- and microcapsular design has been attempted with some degree of success
and may eventually lead to better results in the future.
References
9.4
Acknowledgment
Ana Isabel Silva wants to acknowledge the PhD grant attributed to her by the
Portuguese Government through FCT Fundação para a Ciência a Tecnologia
SFRH/BD/10594/2002.
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normalization of diabetes by
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Aomatsu, Y., Nakajima, Y., Ohyama,
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M., Ko, S., Nagao, M., Tatekawa, Y.,
Sho, M., Ikeda, N., Kanokogi, H.,
100
101
102
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104
105
106
Kobayashi, T., Urizono, Y., Yamada,
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Kawakami, K. (2001) Newly
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Kawakami, K. (2002) Alginate/
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Yarmush, M. L. (1989) Effect of
oxygen on isolated pancreatic tissue.
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Hyder, A., Laue, C., Schrezenmeir, J.
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cells of different ages and species to
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Lembert, N., Petersen, P., Wesche, J.,
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Planck, H., Becker, H. D., Ammon,
H. P. (2001) In vitro test of new
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271–276.
321
Index
a
Abel, John 2
ACCUREL process 59, 82
active electrodes 202
active systems 212
active TDD Systems 216
acute phase reaction 37
adhesion 228, 241
advanced glycation end products 34
affinity chemistry 146
air diffusion tests 120
alginate 280, 300
–chemical composition 281
–effect of purification 291, 300
–guluronic and mannuronic acid
content 290
–inflammation induction 290
Althane 9
Althin Medical AB 9
ALZA 175
AM50-Bio 29
AM-SD18M 31
AN 69ST 13, 37
AN69 acrylonitrile and sodium
methallyl sulfonate copolymer 13,
27f, 30ff, 274
anaphylactoid reactions 26
angiotensin converting enzyme 27
anion-exchange membrane
adsorbers 149
APA (alginate-PLL-alginate)
overview 282
APMA 234f
apoptosis 31f
applications in drug delivery 184
artificial lungs 49
Arylane 16, 19
arylether 16
Asahi AM-UP-75 37
Asahi Medical 8, 11f, 16, 34
Asahi PAN 27
asymmetric membranes 97, 118
autoclaved assemblies 112
b
B. diminuta 105
B. revundimonas diminuta 101
back pressure 116
bacteriophage 120
BAP, (bioartificial Pancreas) 266
–classification 266
–implantation sites 266
Baxter Healthcare Corporation 9
Bemberg 4
benzyl-modified cellulose 11
biocompatibility 24, 267, 273, 276, 281, 290,
295, 303
–assessment 272, 295, 303
–fibrotic overgrowth 267, 279, 290f
–foreign body reaction 267
–macrocapsules 273f
–membrane surface structure 276
–purity of alginate solution 281
–requirements 267
–thrombosis 267, 268
bind-elute 147
binding capacity 147
bioartificial organs 264
–Immunoprotection 266
biodegradable membranes 187f
bioelectrochemistry 171
biohybrid organ 242ff, 249, 253
biological membrane 191
biotransformation 162
BK series U/P/F 14
blood 75
–specific features during application 75
blood plasma treatment 87
–medical indications 87
blood fractionation 69
ISBN: 978-3-527-31480-5
322
Index
blood oxygenation 49
bone 236ff
bradykinin 26ff, 37, 38
bubble point equation 112
bubble point 96, 101
bubble-free gassing 158
buffers 103
c
cake model 102
calcitonin 203
capacity 94, 104, 139
–trials 107
capillary membrane 84
–velocity profile 84
capillary membrane plasma separation
filter 86
capsules 110
cardiodiapulmonary bypass
(CPB) 64
cartridges 110
cassettes 116
categories for hemodialysis
membranes 22
caustic stability 98
cell 103
–culture media 107
–culture 103
–debris 103
–harvest 107
–lysate 103
cell activation 31
Cellophane 3
cellulose 5
cellulose acetate (CA) 9
cellulose acetate dialyzers 29
cellulose diacetate 5
cellulose triacetate 10
charge of the drug 200
chemical compatibility 104
chromatography columns 103
clarification 103
clearances 25
cofactor regeneration 166
collodium tubes 3
colony forming units 100
compatibility 104, 115
complement activation 29
complement factor C3b 29
complement factor D 35
composite membranes 95, 115
constant conductance alternating
current (CCAC) 202
contact phase activation 13
controlled release kinetics 176
convective clearance 23
convective treatments 9
Cordis Dow 9
CR-DDS (controlled release drug delivery
system) 175f
crossflow rate 127
crossflow (TFF) 116, 137
cross-laid knitted hollow fiber mats 63
cross-wound mat device 64
Cuoxam process 4
cuprammonium rayon 8
Cuprophan 4, 8, 27–29, 33
current density 200
cytokine elimination 36
cytokine generation 35
cytokine mRNA production 35
d
dead-end 137
DEAE cellulose 5
DEAE-cellulose (Hemophan) 29
DEAE-modified cellulose 10, 28
degradable polymers 236
degranulation of neutrophils 34
dense membranes 55, 57
dermal fibroblasts 229
dermis 192
description of diffusion 20
device controlled delivery 214
diabetes mellitus 263
–characterization 263
–history 264
–insulin-dependent diabetes
mellitus 263
–STZ-induced 270
–treatment 263
diafiltration 123, 125, 129
diafiltration optimization parameter 130
dialysis patients 1
dialysis 157, 168
dialyzer 2
–constructions 17, 20
–housing 18
–sales 21
DIAPES 16
diffusion
–coefficient of the drug 195
–membrane controlled release 176
–tests 112
diffusive flow 132
direct current (DC) 202
dirt holding capacity 102
double filtration first 74
Index
double filtration set up 74
drug ionic current flow 200
drug partition coefficient 195
dynamic binding 149
e
electrodialysis 157, 168
electromigration 199
electroosmosis 198
electroosmotic flux 200
electroporation 204
–electrode material and
design 205
–pulse specifications 205
–safety issues 205
elute purification 147
endogenous retroviruses 135
endogenous viral clearance 135
endotoxins 32
ENKA company 4
enzyme retention 163
epidermis 192
epithelial cells 252
erythropoietin (EPO) 13
ESRD patient population 2
EVAL 27, 33
Excebrane membrane 33f
extractibles 137
extravascular devices 272ff
f
factor D 35
factors H and B 30
fed batch 128
fermenter 102
fiber bundle 19
fibrotic response
–Ba-alginate microcapsules 295
–PLL microcapsules 294
Fick’s law 176, 177
filter selection 104
filtration mode 137
flat membrane filter 86
flow rate 104
flow-through 147, 148
flux excursion 126
FMC FX class series 19
formulation 121
fractionation membrane 81
Franz diffusion cell 197
freeze-thaw 141
Fresenius polysulfone 8, 15, 16, 28,
29, 33, 35
Fresenius 6, 16
FX-class of dialyzers 20
g
Gambro 16
gamma irradiation 105, 110
gas transfer 53
–carbondioxide 53
–oxygen 53
gas transfer rates 58
–microporous PP membrane 58
–PMP membrane 58
glyco-protein GMP-140 29
gold nanoparticles 120
good manufacturing practice 135
graft survival 270
–extravascular 274, 279, 281, 293, 300,
302f., 308
–hemoglobin 311, 312
–intravascular 270, 271
–oxygen limitations 309
–siliceous membrane 307
Graham, Thomas 2
guarding layer 97
guided bone generation 237, 239
guided tissue generation 237, 239
h
Haas, Georg 1ff.
Hagen–Poiseulle’s law 23
heat lung machine 65
Helixone 16
Helixone membranes 6
hemapheresis 70
hematocrit (Hct) 21
hemodialysis 1
Hemoflux 17
hemolytic dispositions 85
Hemophan 10, 28, 30, 35f.
hepatoblastoma 230, 249
hollow fiber 114
–devices 117
–dialyzers 17
Hospal (Gambro) 12
Hospal AN69 27
Hospal/Code 16
hourglass 95
housing 18
human skin structure 192
hydrogel 180–186
hydrophilic 105, 112
hydrophobic 105
hydrophobic interaction chemistry 146
hydrophobic microporous flat sheet
membranes 52
323
324
Index
i
m
IDDM (insulin dependent diabetes
mellitus) 264
IDEMSA, Spain 8
immune rejection 267, 294
–cytotoxic reaction 268
–humoral immune
response 268, 291, 308
inert electrodes 202
inner diameter 24f.
insulin secretion 271, 292, 301, 305,
308
–insulin release time 271, 292,
301
integrity test 98, 112, 132
–air flow integrity test 132
integrity 125, 144
interleukin-1 35
intermediate buffer exchange 121
ion-exchange 146
islets covered with a siliceous
membrane 304f.
–3-aminopropyltrimethoxysilane, APTrMOS 305
–tetramethoxysilane,
TMOS 305
islet-like cell clusters (ICCs) 291
islets of Langerhans 263, 276
–depth of oxygened outer
rim 310
–direct deposition of a siliceous
membrane 306
–gel entrapment matrices 276
–nutrient supply and
oxygenation 277, 293
–oxygenation studies 308
–preventing aggregation 276
macrocapsules 273, 301
–acrylonitrile 274
–alginate-PLL-alginate (APA) 280, 282
–AN69 274
–biocompatibility 274
–Amicon XM-50 274
–AN69 274
–poly-L-lysine (PLL) 276
macrovoid 115, 118
magnetophoresis 211
manufacturing of microporous
membranes 81
matrix systems 212
melt spin stretch process 55
Membrana GmbH 4, 8, 16
membrane adsorbers 146
membrane effectiveness 270
–citotoxicity testing 270
–diffusional properties 271
–effect of pore size 306
–insulin permeability 275
–insulin release time 271
–pore-sized 275
–thrombogenicity testing 270
membrane is asymmetric 13
membrane or reservoir systems 212
membrane oxygenators 50
membrane processes 155
membrane production 60
–melt spin stretch process 60
–temperature-induced phase separation
process 60
membrane selection 124
membrane sieving 123
membrane wall thickness 7
microcapsules 280, 292ff, 300, 302, 305,
307, 311
–agarose beads 302
–agarose-PSSa 303
–Barium-alginate 295f
–comparison of APA with
Ba-alginate 301
–effect of size 292, 295
–encapsulation process 303, 307
–fabrication 307
–fibrotic overgrowth 300
–fibrotic response 294f
–immune rejection 294
–mechanical stability 294
–microfabrication technology 305
–multicomponent, containing polymethylene co-guanidine 307
–oxygenation islets 311
k
Kalle company 3
kallikrein-kinin system 26
keratinocytes 232, 233, 234, 236
kidney 244, 245, 249
Kimal 16
kinin system 37
Kolff, Willem 3
l
linearly scaleable 124
liquid porosimetric 120
liver 243, 244, 245, 246
log reduction value 120
lontophoresis 198
Index
–poly-L-lysine (PLL) 280
–protruding islets 293
–swelling and shrinking
phenomena 290
–use of anti-fibrotic agent 300
microfiltration membranes 93
microporous capillary membranes 62
–blood inside 62
–blood outside 62
microporous membranes 55, 61, 83
–melt spin stretch process 83
–production 55
microporous PMP membrane 59
–dense outer skin 59
microporous PP membrane 56
microprojection/microneedle
patch 210
Minntech 16
Moiré structure 19
molecular pore sizes 183
molecular weight cutoff 22
monoclonal antibody 129
Murine Leukemia Virus 142
mycoplasma 94, 101
n
neonatal pancreatic cell clusters
(NPCC) 300
nervous system 240
NFF, see normal flow
Niels Alwall 3
Nikkiso 16
noncellular blood components 78
–size 78
normal flow (NFF) 118, 137
normal molecular weight cutoff 116
Nucleopore 274
o
OROS 175
osmotic pressure 122
oxidative stress 32
oxygen species production 32
oxygenation membranes 55
oxygenator 54, 61
–extraluminal flow 54
–membrane makeup 61
–operational modes 61
OXYPLUS 59
p
PAN DX 27, 37
PAN 27
PAN17DX 31
pancreas 243
pancreas, bioartificial 265
–exocrine tissue contamination 281
–extravascular devices 272
–graft survival 270
–hollow fibers 268
–implantation sites 272
–insulin adsorption 270, 271
–long-term discordant xenograft
function 281, 304
–oxygen distribution 308, 311
–vascularization 277
particle challenge tests 120
Parvoviruses 136, 142
passive diffusion 195
passive drug permeability coefficient 195
passive systems 212
passive TDD systems 212
patch 191
PCL (poly-e-caprolactone) 236
PEG-grafted cellulose 11
PEPA 16
PES, see polyethersulfone
PH effect 181f, 185
pilot-scale studies 141
plasma fractionation membranes 79f
–producers 80
–sieving coefficients 80
plasma kallikrein 38
plasma separation 75
–set up 74
plasma separation membranes 77, 79
Plasmaflo 79
Plasmaphan 79
plasmapheresis 69
–procedures 83
pleated cartridges 93, 97
Pmax 109
PMMA, see Polymethyl Methacrylate
polyacrylonitrile 5, 29, 33
polyacrylonitile membranes 13
polyacrylonitrile (PAN) 12, 234f, 251
polyamide (PA) 5, 17, 36
polyamide S 16
Polyamix 16
polyethersulfone (PES) 5, 95, 117
polyethylenimine (PEI) 13
Polyflux 17
polylactide 239
polymethyl methacrylate (PMMA) 5, 7, 14, 27,
30, 33–36
Polyphen 16
polysulfone (PSu) 15
polysulfone 28, 31, 36, 95, 250, 251, 253
325
326
Index
polysynthane 11
polyurethane 17
polyvinylpyrrolidone 5
poly-L-lysine (PLL) 276
–coimmobilizing the islets with
hemoglobin 310
–comparison of APA with
Ba-alginate 301
–effect of barium crosslinking 301
–effect of size 292
–encapsulation process 280
–geometry 273
–Mechanical stability 293, 301
–polyvinyl alcohol hydrogel
tubular (MRPT) 276
–sodium methallyl
sulfonate 274
–vascularization 277
pore blockage 102
pore-plugging 107
–model 141
post-harvest concentration 121
potting material 18
prefiltering 140
process control 134
process development 124, 139
process flux 115
processing plan 128
product recovery strategy 133
protein A column
chromatography 147f
PVDF (polyvinylidene fluoride) 117
r
reactive oxygen species 33
red biotechnology 155
regenerated cellulose 8, 117
retention 94, 104, 105, 115
–assurance 94
–confidence 100
retention assurance 101
retrovirus like particles (RVLPs) 135
retroviruses 136
reversed phase chemisty 146
Rhone Poulenc 12
robust 133
robustness 94, 115
rotating disc plasma separator 87
rotating drum 3f
rotating plasma separation device 86
s
Salmon calcitonin 203
Saxonia 16
scaffold 231
scalability 115
scale-up 131
–linear-scale 131
scaling 109
sieving coefficients 21f
simulation study 141
size exclusion 113
sizing filters 107
skin 191, 232
skin penetration enhancers 207
skin fouling 202
skin Controlled Delivery 214
small-scale testing 111
spacer yarns 19
spiral wound module 116
steaming 98
steam-in-place 112
sterile filtration 103
sterile microfiltration 93
sterility 145
–assurance 99f
sterilization 99, 110
–autoclave sterilization 110
–ethylene oxide (ETO) 26
–hydrophobic sterilization
membranes 102
–method 112
–pre-sterilized 99
–sterilizing assurance 100
–sterilizing grade 100
sterilizing grade 101, 105
sterilizing membranes 105
Stewart, Richard 4
stratum corneum 193
stretching process 61
structural parameters 182
surface modification 98
symmetric membranes 118
t
tangential flow 118
TAT formation 27
temperature effect 186
temperature-induced phase separation
(TIPS) 55
Terumo 11
Terumo polysulfone 34
TFF, see cross flow
Thalhimer, William 3
thermally induced phase separation 82
thrombin-anti thrombin III complex (TAT) 27
thrombogenicity 26
Index
ultrafiltrate (QF) 24
ultrafiltration 113, 121
–factor (UE) 21
–membranes 114f
–process 121, 123, 128
–theory 122
ultrasound-assisted TDD –
sonophoresis 208
–high frequency 208
–low frequency 208
–medium frequency 208
undulation 19
vascularization 273, 278, 303
–AN69 hollow fiber 278
–fibroblast growth factor 278f, 303
–hydroxymetilated polysulfone 271
–neovascularization of
macrocapsules 277
–polyglycolic acid (PGA) flat sheet 278
–polystyrene sulfonic acid (PSSa) 279
–polytetra fluoroethylene (PTFE)
scaffold 278
–regenerated cellulose hollow fibes 279
–subcutaneous implantation site 273,
303
virus capture 149
virus filters 143
–virus clearance filters 144
virus filtration 113, 137
–filtration process 137
–validation 142
virus removal 117, 136
–virus clearance filters 117
–virus removal membranes 117
virus spike 142
Visking Company 3
Vitabran E 16
vitamin E 12, 33, 34
vitamin E-modified cellulosic 11
vitamine C 34
Vmax 107ff, 141
v
w
vascular devices 268ff.
–Amicon XM-50 (polyvinyl
chloride acrylic
capolymer) 269
–Cuprophan HDF
(cellulose) 269
–hydroxymetilated
polysulfone 271
–polyethertherketone
(PEEK) 270
–Thomapor (polyamide) 269
wall thickness 9
water permeability 116
Weingand 3
wetting 112
white biotechnology 155
timolol 197
tissue 267
tissue engineering 227ff, 228, 231,
232, 236, 240, 241
TNFa (tumor nekrose factor alpha) 35
Toray Industries 14, 16
Toraysulfone 16
transdermal drug delivery (TDD) 191
transmembrane pressure (TMP) 21,
23
transplantation 264, 265
–allogeneic or xenogeneic cellular
tissue 265
–pancreatic islets 264
transport number 200
u
x
Xanthogenate 9
z
zeta potential 38
327

Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes

Transcription

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