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Development of a drug delivery platform using
multifunctional polymeric scaffold for scar
therapy
Vipul Agarwal, MApplSc
This thesis is presented for the degree of Doctor of Philosophy at The University of
Western Australia
School of Chemistry and Biochemistry
2015
Contents
Abbreviations .............................................................................................................. iv
Abstract ......................................................................................................................vii
Acknowledgement ...................................................................................................... ix
Statement of candidate contribution ........................................................................... xi
1 Introduction and literature review........................................................................ 1
1.1
Tissue engineering.............................................................................................. 2
1.2
Cellular scaffolds................................................................................................ 3
1.3
Physical features of the cell microenvironment ................................................. 4
1.4
Design concepts and strategies ........................................................................... 5
1.4.1
Physical properties ...................................................................................... 6
1.4.2
Mechanical properties ................................................................................. 7
1.4.3
Surface properties ........................................................................................ 8
1.5
Skin, injury and wound healing.......................................................................... 9
1.5.1
Haemostasis and inflammation ................................................................. 11
1.5.2
Reepithelialisation ..................................................................................... 13
1.5.3
Formation of granulation tissue ................................................................. 16
1.6
Role of TGFβ in fibrosis and scarring .............................................................. 17
1.7
Mechanisms of activation of extracellular TGFβ ............................................. 19
1.8
Exogenous mannose-6-phosphate act as an antagonist of TGFβ1 activation ... 20
1.9
Structural requirements for mannose-6-phosphate recognition to M6P/IGFII
receptor ...................................................................................................................... 22
1.10 Mannose-6-phosphate analogues ..................................................................... 23
1.11 Scarring ............................................................................................................ 27
1.12 Current treatments for skin wound healing ...................................................... 28
1.13 Cell based therapies .......................................................................................... 29
1.14 Matrix based therapies ..................................................................................... 30
i
1.14.1
Amniotic membrane ................................................................................ 30
1.14.2
Human cadaver allografts and xenografts ............................................... 30
1.14.3
Tissue engineered skin ............................................................................ 32
1.15 Nanofibrous scaffolds for skin tissue engineering ........................................... 34
1.16 Skin regeneration using electrospun scaffolds ................................................. 37
1.17 Electrospun hybrid materials ............................................................................ 41
1.18 Summary .......................................................................................................... 44
1.19 Hypotheses and Aims ....................................................................................... 46
2 Introduction to the series of papers .................................................................... 48
2.1 Development of a universal multifunctional scaffold ......................................... 48
2.2 Evaluation of mannose-6-phosphate analogues as potential anti-scarring agents ...
.......................................................................................................................... 51
2.3 Delivery of the lipophilic mannose-6-phosphate analogue PXS64 using an
electrospun PGMA scaffold ...................................................................................... 53
3 Series of papers ..................................................................................................... 56
A Functional Reactive Polymer Nanofiber Matrix .................................................... 57
Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor
Inhibitors by Intracellular Delivery ........................................................................... 62
Inhibiting the activation of transforming growth factor-β using a polymeric
nanofiber scaffold ...................................................................................................... 67
4 Conclusions and future work ............................................................................... 71
4.1
Tissue engineering of a nanoscaffold ............................................................... 71
4.2
Scar therapy and mannose-6-phosphate analogues .......................................... 73
4.3
Scaffold based delivery of analogue 2 ............................................................. 74
4.4
Future recommendations .................................................................................. 75
4.5
Final remarks .................................................................................................... 77
A Supporting information for papers .................................................................... 79
Supporting Information for ‘A Functional Reactive Polymer Nanofiber Matrix’ .... 80
ii
Supporting information for ‘Enhancing the Efficacy of Cation-Independent Mannose
6-Phosphate Receptor Inhibitors by Intracellular Delivery’ ...................................... 87
Supporting Information for ‘Inhibiting the activation of transforming growth factorβ using a polymeric nanofiber scaffold’ .................................................................... 98
B Published work not directly included in the thesis ......................................... 103
References................................................................................................................ 147
iii
Abbreviations
ANOVA
Analysis of variance
bFGF
Basic fibroblasts growth factor
CDMPR
Cation dependent mannose-6-phosphate receptor
CIMPR
Cation independent mannose-6-phosphate receptor
DBM
Demineralized bone matrix
DCFH-DA
2’, 7’-dichlorodihydrofluorescein diacetate
ECM
Extracellular matrix
EGF
Epidermal growth factor
ES
Electrospun
FTIR
Fourier transform infra-red
GAG
Glycosaminoglycan
HCAS
Human cadaver allograft skin
HDF
Human dermal skin fibroblasts
HPLC
High pressure liquid chromatography
HSF
Human scar fibroblasts
LAP
Latency associated peptide
LCST
Lower critical solution temperature
LTGFβ
Latent transforming growth factor β
M6P
Mannose-6-phosphate
M6P/IGF II
Mannose-6-phosphate/Insulin-like growth factor II
MMP
Matrix metalloproteinase
iv
MRI
Magnetic resonance imaging
NMR
Nuclear magnetic resonance
NP
Nanoparticle
Pa
Pascal
PBS
Phosphate buffer saline
PCL
Poly(ε-caprolactone)
Pd
Palladium
PEG
Poly(ethylene glycol)
PGA
Poly(glycolic acid)
PGMA
Poly(glycidyl methacrylate)
PLA
Poly(lactic acid)
PLGA
Poly(lactic-co-glycolic acid)
PNIPAM
Poly (N-isopropyl acrylamide)
PU
Polyurethane
ROS
Reactive oxygen species
RT-qPCR
Real time quantitative-polymerase chain reaction
SEM
Scanning electron microscopy
SQUID
Superconducting quantum interference device
TE
Tissue engineering
TEM
Transmission electron microscopy
TGFβ
Transforming growth factor β
TSP
Thrombospondin
UCNP
Upconverting nanoparticles
v
VEGF
Vascular endothelial growth factor
vi
Abstract
Tissue engineering is a multidisciplinary approach and has been used to promote
tissue regeneration, wound repair and to enhance drug delivery. In the case of burn
injury, despite the advances in treatment leading to reduced mortality, the problem
of permanent, disfiguring scar formation following healing is far from being
resolved.
Scarring is the result of an imbalance of fibroblast activity resulting in an excess of
architecturally disorganised extracellular matrix protein deposition and is
predominantly mediated by the TGFβ pathway.
Whilst there are a number of promising therapeutic targets identified through our
increasing understanding of scar formation, there has been limited success in the
clinical translation. This can largely be attributed to difficulties with delivery,
stability and efficacy of treatments tested to date.
In this thesis the problems of drug delivery and stability have been addressed using a
combinatorial approach. First, a scaffold was developed to provide a platform for
drug delivery and stability. A novel multifunctional polyglycidyl methacrylate
(PGMA) scaffold was developed using electrospinning. The multifunctionality of
this polymeric scaffold was demonstrated for various applications. These include
surface functionalization of electrospun PGMA with poly (N-isopropyl acrylamide)
for stimuli response surfaces and development of multifunctional nanocomposites
with (NaGdF4:Yb, Er); Pd and Fe3O4 nanoparticles for upconversion fluorescence
imaging, sensing, and magneto-responsive properties.
This was followed by
exploration of modified analogues of a potential therapeutic target as anti-scarring
agents. Mannose-6-phosphate (M6P) has been shown to ameliorate scarring by
inhibiting the activation of TGFβ1, a necessary step required for its receptor
recognition and function. However, therapeutic delivery of M6P to a wound is
currently ineffective due to the low stability of M6P and limited ability to maintain
M6P concentration at the site of injury. Therefore whilst targeting TGFβ activity
through M6P has significant potential in burn therapy, stability and delivery issues
must be addressed before this can become a therapeutic reality. Two M6P analogues,
vii
PXS25 (analogue 1) and PXS64 (analogue 2), have been developed in collaboration
with Pharmaxis Ltd., to overcome the metabolic vulnerability of M6P whilst
retaining the receptor recognition function. Initial work was carried out to
investigate the biocompatibility and cytotoxicity capabilities of both analogues
compared to M6P in human dermal skin fibroblasts. Subsequently they were
investigated for their proficiency in regulating the expression of the critical fibrotic
marker, Collagen I. Analogue 2 was shown to significantly inhibit TGFβ1 mediated
up-regulation of collagen I gene expression. However, the lipophilic analogue 2 has
limited bioavailability. The final chapter addresses this specific problem by
incorporating analogue 2 into the multifunctional scaffold and testing the efficacy of
the combined drug/scaffold therapy on human dermal skin fibroblasts.
The tissue engineering approach presented herein demonstrated the potential of
combinatorial scaffold mediated drug delivery method to progress some of the
existing therapeutic targets into clinical therapies. The future work using porcine
wound healing model is a necessary extension to establish the efficacy and potential
of
this
combinatorial
approach
in
viii
vivo
before
its
clinical
translation.
Acknowledgement
I would like to begin by thanking my supervisors for their consistent support
throughout my candidature. I am really grateful to Prof Swaminathan Iyer for
offering me PhD scholarship and introduce me to a completely new stream of
research, to keep me on track and motivate me especially during the difficult times
and moulding me into a better researcher; Prof Fiona Wood for support and
guidance and brining clinical prospective into my work; Dr Mark Fear for bringing
biological insight, enabling me to resolve complex experimental problems and
improving my overall understanding. In addition, I would extend my gratitude
towards Prof Iyer and Dr Fear for going above and beyond to help me with my grant
application, you faith and support only project me to improve and become an
accomplished scientist. I would also like to thank my colleagues Dr Cameron Evans,
Dominic Ho, Diwei Ho, Dr Faizah Yasin, Alaa Munshi, Ruhani Singh, Ivan Lozic,
and Michael Bradshaw, Dr Rahi Varsani, Dr Tristan Clemons and Michael
Challenor for making me welcome and accepting me within the group with all the
humour and pranks and also to support me where possible in the lab. I would like to
acknowledge Dr Tristan Clemons, Dr Nicole Smith, Sumi Shrestha, Callum
Ormonde, Diwei Ho and Dr Marck Norret to proof reading this thesis. I would like
to thank the staff at CMCA especially Lyn Kirilak and Prof Paul Rigby for their
extended support with microscopy instrumentation and experiments. I would like to
also thank Pharmaxis Ltd and the team for inviting me into their lab and welcoming
me within their group, and also to consistently support me during my candidature.
Because of the multidisciplinary aspect to my project, I had the privilege to work
with some really good people including Prof Charlie Bond, Dr Bernard Callus,
Megan Finch, Benjamin Gully, Dr Foteini Hassiotou, Dr Ben Corry, Dr Natalie
Smith, Dulharie Wijeratne and Dr Lindsay Byrne, without whom this thesis would
not have been possible. For the motivation, persistent support and having faith in me
to pursue my research ambitions, I would like to thank Dr Karen Stack, Dr Stephen
Newbery, Dr Manab Sharma, Sinu Sharma and Murray Frith. For funding, I would
like to thank Australian Research Council (ARC), National Health & Medical
ix
Research Council (NHMRC), Australian Nanotechnology Network (ANN),
Australian Academy of Science and The University of Western Australia.
Finally, I would like to thank my family, my girlfriend and my friends for their faith
and selfless support enabling me to pursue my goals.
x
Statement of candidate contribution
This thesis contained published work and work prepared for publication, some of
which has been co-authored. The bibliographical details of the work and author
contributions are outlined below.
Refereed journal articles included in the series of papers
1.
Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D., Gong, P., Ye, W.,
Singh, R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre, T., Wood, F.M.,
Fear, M., Lorenser, D., Sampson, D. D., Zdyrko, B., Smith, N.M., Luzinov, I., Iyer,
K.S., A Functional Reactive Polymer Nanofiber Matrix, RSC Advances (Submitted)
VA and DH developed the initial concept of PGMA electrospinning. VA
collaborated with WY and DH to synthesize upconverting nanoparticles, AM to
synthesize palladium nanoparticles and acquired TEM images on nanoparticles, RH
to synthesize magnetite nanoparticles and acquired TEM images on fibers, PG to
perform NIR Room Temperature Emission Spectroscopy, FMDY to perform
hydrogen gas sensing analysis and RCW to perform squid analysis. YG, BZ, IL
provided the PGMA; remaining authors supervised the work. Contribution by VA:
75%
2.
Agarwal, V., Toshniwal, P., Smith, N. E., Smith, N. M., Li, B., Clemons, T.
D., Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B., and Iyer, K. S.,
Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor
Inhibitors by Intracellular Delivery, Angewandte Chemie International Edition
(Submitted)
xi
PT and TDC carried out protein expression studies. NES and BC performed docking
experiments; rest all the authors supervised the work. Contribution by VA: 90%
3.
Agarwal, V., Wood, F. M., Fear, M., and Iyer, K. S., Inhibiting the activation
of transforming growth factor-β using a polymeric nanofiber scaffold, Nanoscale
(submitted)
FMW, MF and KSI supervised the work. Contribution by VA: 90%
Refereed journal articles included in the appendix
1.
Agarwal, V., Tjandra, E.S., Iyer, K. S., Humfrey, B., Fear, M., Wood, F. W.,
Dunlop, S. and Raston, C. L., Evaluating the effects of nacre on human skin and scar
cells in culture, Toxicology Research, 3, 223-227 (2014)
EST performed the viability and reactive oxygen species assay on HaCaTs; BH
provided the pearl shells; remaining authors supervised the work. Contribution by
VA: 90%
2.
Eroglu, E., Chen, X., Bradshaw, M., Agarwal, V. , Zou, J., Stewart, S.G.,
Duan, X., Lamb, R.N., Smith, S.M., Raston, C. and Iyer, K. S., Biogenic production
of palladium nanocrystals using microalgae and their immobilization on chitosan
nanofibers for catalytic applications, RSC Advances, 3, 1009-1012 (2013)
VA
performed
and
optimised
protocol
for
electrospinning,
performed
characterisation and contributed to the manuscript. Contribution by VA: 25%
3.
Eroglu, E., Agarwal, V., Bradshaw, M., Chen, X., Smith, S. M., Raston, C.
and Iyer, K. S., Nitrate removal from liquid effluents using microalgae immobilized
on chitosan nanofiber mats, Green chemistry, 14 (10), 2682-2685 (2012)
VA
performed
and
optimised
protocol
for
electrospinning,
performed
characterisation and contributed to the manuscript. Contribution by VA: 40%
xii
Conference presentations
1.
Agarwal, V., Schilter, H., Ho, D., Guo, L., Hassiotou, F., Jarolimek, W.,
Wood, F. M., Fear, M., Iyer, K.S., An Innovative Tissue Engineering Approach
Towards Scar Therapy, Fifth International NanoMedicine Conference, Sydney,
Australia, 30 June- 2 July 2014 (Oral presentation)
2.
Agarwal, V., Ho, D., Schilter, H., Guo, L., Hassiotou, F., Jarolimek, W.,
Wood, F. M., Fear, M., Iyer, K.S., Novel Scaffold Approach Towards Scar Therapy,
International Conference on Nanotechnology and Nanoscience (ICONN2014),
Adelaide, Australia 2014 (Poster presentation)
3.
Agarwal, V., Schilter, H., Guo, L., Jarolimek, W., Wood, F. M., Fear, M.,
Iyer, K.S., Scarless wound healing: A new approach, Sixth International Conference
on Advanced Materials and Nanotechnology (AMN-6), Auckland, New Zealand, 1115 February 2013 (Poster presentation)
4.
Agarwal, V., Wood, F. M., Fear, M., Iyer, K.S., 6. Development of scarless
wound healing platforms, School of Chemistry and Biochemistry Research Forum,
University of Western Australia,
Perth, Australia, 2 November 2012 (Poster
presentation award)
5.
Agarwal, V., Wood, F. M., Fear, M., Iyer, K.S., Development of Hybrid
Hydrogel for Scarless Wound Healing, First International Conference on BioNano
Innovation (ICBNI 2012), Brisbane, Australia, 18-20 July 2012 (Oral presentation)
6.
Agarwal, V., Wood, F. M., Fear, M., Iyer, K.S., Regulating the Migratory
Behaviour of Fibroblasts on a Hydrogel Scaffolds, International Conference on
xiii
Nanotechnology and Nanoscience (ICONN2012), Perth, Australia, 5-9 February
2012 (Oral presentation)
7.
Agarwal, V., Eroglu, E., W., Wood, F. M., Fear, M., Iyer, K.S., In vitro
evaluation of electrospun pluronic F-127 dimethacrylate copolymer towards wound
healing in burn injuries, Australian Nanotechnology Network Early Career
Symposium, Sydney, Australia, 21–22 November 2011 (Oral presentation)
Vipul Agarwal
Prof Swaminathan Iyer
Candidate
Co-ordinating Supervisor
xiv
Chapter 1
Introduction and Literature Review
Each year 100 million people develop scars in the developed world alone,1 with
causes ranging from elective surgery to severe trauma. Scars result not only in
aesthetic deficits but also long-term functional and psychological problems.2
According to WHO estimates, 6 million people suffer from burn injury each year of
which over 300,000 succumb to their injuries.3 Further, chronic skin ulcers
contribute an additional 6 million patients annually.4 In attempts to address the
functional and aesthetic deficits, the field of tissue engineering has seen a surge of
interest and applications in recent years. However, current clinical treatment of
scarring still centres on a surgical approach. Small molecules such as mannose-6phosphate (and many other biological factors) have been shown to have significant
potential to promote healing and reduce scarring both in elective surgery and after
injury. However, the potential of these factors to deliver improvements in the clinic
has been significantly hampered by issues of stability and delivery. This thesis will
describe the preparation of a multifunctional universal nanofibrous scaffold with the
potential to be used as a delivery vehicle for wound repair. Subsequent work will
assess the potential of more stable mannose-6-phosphate analogues which may
overcome some of the limitations of the endogenous ligand that has previously been
trialled. Finally, a combinatorial approach will be investigated using the
scaffold/analogue combined.
In this chapter the literature detailing the application of tissue engineering and the
key considerations for tissue engineering approaches will be discussed. The
pathophysiology of skin injury and wound healing will also be reviewed. Finally, the
cell therapy and tissue engineering approaches to skin repair will be discussed and
the concept of multifunctional scaffolds, electrospinning and their potential for drug
delivery and enhanced healing will be introduced.
1
1.1 Tissue engineering
Tissue engineering (TE) has evolved as an interdisciplinary science combining
principles from nature and integrating them with material and engineering science
with the goal of developing functional substitutes for damaged tissues and organs. 5
The materials serve as temporary scaffolds and promote the organisation of cells to
form a functional tissue with the cells providing a source for repopulation.6 The
general strategy behind TE is to seed cells within a scaffold that mimics the
architecture of the replacement tissue while providing environmental cues that
promote tissue regeneration. Tissue engineered skin equivalents have been in
clinical use since 1997.7
In tissue engineering, a scaffold is usually required to provide a platform or niche
that promotes the desired behaviour of cells (for example proliferation of epidermal
cells to promote wound coverage).8 The scaffolds are intended to replicate or
enhance the natural ECM environment in order to retain cell viability and control
cellular behaviour. For example, demineralized bone matrix (DBM, bone from
which mineral and cells have been removed, leaving only proteinaceous material)
have been implanted in the muscle to initiate bone formation in the neighbouring
muscle tissue. This led to the commercial production of recombinant DBM from
cadavers for implantation in bone defects.7
There are three important criteria to be considered for the development of tissue
engineering scaffolds: i) the 3-dimensional micro-structure of the scaffold such as
inter-connectivity of the network pores and subsequent porosity allowing the
exchange of gases, nutrients and waste and facilitating cell adhesion, spreading and
tissue formation, ii) mechanical parameters catering to specific tissue type such as
scaffold morphology like linearity, plasticity, flexibility or anisotropy, and iii)
successful delivery of cells, drugs, growth factors and/or cytokines.8
2
1.2 Cellular scaffolds
Cellular scaffolds have been developed using natural and synthetic materials and
their composites, harnessing the specific advantages of each component and
amalgamating them to achieve the desired application. Usually, natural materials are
based on purified ECM components (eg. collagen, gelatin, laminin, hyaluronic
acid)9-11, often containing multiple components to create a composite substrate
(Matrigel®, Integra®).12 Other natural materials are also commonly used, mainly
derived from plant or animal sources including silk, agarose and chitosan.13 The
main advantages associated with the use of natural materials are their biological
activity and biocompatibility. In contrast, synthetic materials are employed to
overcome the shortfalls experienced in the use of natural materials, for example
manufacturing and process variability and an inability to control their physicochemical properties.8
Single component scaffolds lack the complexity and often functionality of
biomaterial scaffolds, such as mechanical properties, electrical activity or cues for
cell-matrix interactions14.8 Composite materials are therefore generally preferred.
For example, bone is made up of collagen and ceramic like hydroxyapatite, based on
which polymer-ceramic composites have been widely used in bone tissue
engineering15-19.8 Hydrogels have also been explored as they can imitate the crosslinked architecture of ECM through their cross-linked network of monomers,
oligomers or polymers.20 However, the major limitation in their use is their lack of
mechanical strength necessary for tissue engineering applications. 8 Gelatin
methacrylate hydrogels for example are known to promote cell adhesion and
spreading but lack the required mechanical strength.10 Carbon nanotubes have been
used to reinforce the mechanical properties of gelatin methacrylate by significantly
enhancing the compressive modulus, while retaining the porosity and cell
adhesiveness21.8
In designing a scaffold, one of the most desirable features is biodegradability; they
must gradually degrade over time and simultaneously get substituted by the naturally
deposited ECM and newly formed tissue.8 Consequently, linear aliphatic polyesters
such as poly(lactic acid), poly(glycolic acid) and poly(ε-caprolactone) have been
3
extensively investigated owing to their biodegradability, achieved through the
hydrolysis of their ester bonds and the ability to control their degradation rate.8
Scaffold structure is another important feature that needs to be considered during
fabrication of cellular scaffolds. Initially the emphasis was on macroporous
structures to facilitate mass transfer of vital molecules, developed using
microspheres, salt leaches or gas foams.8, 22 Such micron-scale scaffolds however do
not recapitulate the nano-scale dimensions of the ECM.23 In order to generate such
nanometer-scale dimensions techniques such as electrospinning,24 molecular selfassembly,25 soft lithography,26 and phase separation27 have been employed.8
1.3 Physical features of the cell microenvironment
The important physical properties to consider towards the cell’s microenvironment
are substrate mechanics and surface topography.8 Mechanical properties of the tissue
are dependent on their anatomical location. For example, the elastic modulus of the
soft tissue of the brain (0.5 kPa) is low compared to intermediate muscles and skin
(~ 10 kPa), and hard precalcified bone (>30 kPa).8, 28 One of the major limitations in
recreating the native cellular environment is to reproduce the intricate physical
features found in tissues which are viscoelastic with non-linear, anisotropic and
heterogeneous mechanical properties. It is an important consideration while
designing scaffolds because the optimal scaffold should completely replicate the
replacement tissue.8 For example, the biological nanostructure of the heart
influences its biochemical, electrical and mechanical functions and its ECM
comprised of dense elastic fibrillar collagen, elastin bundles and proteoglycans.5, 29 It
promotes mechanical coupling between cardiomyocytes resulting in aligned and
anisotropic cell bundles promoting not only intercellular interactions but also with
neighbouring capillaries and nerves.5 This bundled elongated cellular structure is
critical for the function of the cardiac syncytium, enabling the muscles with
exclusive rhythmic contraction and mechanical and electrical properties to facilitate
the blood pumping function of the heart.5 Cardiomyocytes have been shown to adopt
random morphology on flat surfaces as they lose their elongated morphology post
4
isolation. This is most likely because they lose the cues required to instruct their
formation and interactions with their environment.5 Strategies adopted to address
this problem of cell alignment in engineered cardiac tissue include applying
mechanical stretching,30 interstitial fluid flow, electrical stimulation31 and
microcontact printing,32 with limited success.5 Alternatively, a model cadiac tissue
was developed by controlling the nanotopography of the scaffold that mimics the
ECM matrix of the native myocardial tissue.5, 33 Polyethylene glycol hydrogel was
fabricated to yield grooved arrays with ridge widths ranging from 15 to 800 nm
which when cultured with cardiac cells resulted in their self-assembly aligned to the
direction of the scaffold.5 It resulted in elongation of cadiac cells to form
anisotrophic cell array which could result in rhythmic contraction necessary for
normal heart function.33 Despite the progress made, the challenge of patterning such
nano-arrays within 3D scaffolds towards anisotropically aligned tissues still
remains.5
1.4 Design concepts and strategies
Three-dimensional (3D) scaffold constructs have more demanding requirements for
efficient cell motility compared to the 2-D monolayer cultures to achieve tissue
uniformity and to avoid heterogeneous tissue growth.34 The supply of oxygen and
other nutrients along with waste removal and cell motility reply on mass transfer
properties of the scaffold.35 Considering the primary mechanism behind mass
transport is diffusion, scaffolds need to be designed after carefully considering its
diffusion characteristics.36 Cells in vivo reside within 100 µm of a capillary for
efficient nutrient supply.35 Neotissue growth on engineering scaffolds has been
reported to be preferentially limited to the peripheral regions (100-200 µm) of the
scaffold because of the limited oxygen supply.35, 37, 38 This is still a major limitation
in the engineering of thick 3D tissues because it limits the cellular distribution and
bioavailability of nutrients especially in the interiors of the scaffold as the cells
preferentially localise to the periphery.
35
Therefore scaffolds for tissues with low
metabolic activity have been easier to synthesize. For example, cartilage, an
5
avascular tissue known to have low metabolic activity, has been successfully
fabricated into a thick 3D construct.39 Similarly, other structural tissues known for
their mechanical function such as blood vessels, heart valves, ligaments, tendons and
skin may have lower demand for high oxygen and nutrient transport and therefore be
easier to replicate.35, 40
1.4.1 Physical properties
One of the main features required facilitating nutrient and oxygen transport is the
pore size of the engineered construct. Pore size is critical as it influences tissue
ingrowth and cellular adhesion especially in the internal surface area of the matrix.
For example, small pores are impeded by the growing tissue preventing tissue
ingrowth and ECM production. Alternatively, big pores lack cellular recognition of
the surface topography preventing neovascularisation.35 Although the pore size is
susceptible to change in an in vivo environment,35, 39 it has been demonstrated that
optimal size varies with differing cell types and their respective architecture. For
example, where 5 µm is optimal for neovascularisation, fibroblasts ingrowth
demands 5-15 µm whereas hepatocytes require 20 µm, keratinocytes need 20-125
µm, and 100-700 µm size pores are required for bone regeneration.35, 39, 41, 42 Ma et
al. studied the effects of pore size in a 3-D polyethylene terephthalate nonwoven
fibrous matrix on long-term tissue development of human trophoblast ED27 cells in
terms of cell morphology and spatial organisation.43 They showed that human
trophoblast ED27 cells with average diameter of 14 µm were not able to bridge the
gap between fibres with a pore diameter of 20 µm. Instead they formed large
aggregates and started to differentiate in larger pores. On the contrary, cells rapidly
bridged the gap on the matrix with low porosity, a pore diameter of around 15 µm
and promoted cell spreading and proliferation. Further, the cell differentiation was
inhibited, highlighting the importance of pore size on cellular behaviour and
response, in terms of cell morphology, differentiation and proliferation, towards
scaffold design features. In yet another similar study, canine microvascular epithelial
cells formed a thin endothelial lining only when grown on scaffolds with average
pore size of 90 µm, whereas vascular smooth muscle cells preferred the pore size of
6
107 µm in order to form uniform tissue.35 Dermal fibroblasts, on the other hand, did
not respond to any pore size.35, 44 Similarly, higher pore size greater than 250 µm has
been shown to promote angiogenesis with faster cellular ingrowth when compared to
smaller pore sizes.35 Despite the elaborate evidence on the impact of pore size and
their distribution on angiogenesis, inflammation and tissue infiltration in vivo, the
optimal pore size and range is still fairly unpredictable, especially in the context of
regenerating complex tissues with multiple different cell types45, 46.35
In addition to pore size, cell transport mechanisms like diffusion, attachment and
migration is dependent on total porosity, pore interconnectivity and scaffold surface
area.35, 41 In order to achieve uniform cellular penetration and ingrowth, scaffolds are
required to be highly porous with open interconnected geometry and a large surface
area: volume ratio for example, scaffolds with over 90% porosity have been shown
to have optimal surface area for cellular-matrix interactions warranted for efficient
diffusion.35, 41, 47, 48 Highly porous scaffolds, however, lack the mechanical integrity
required for tissue regeneration.36, 49, 50 Therefore, a fine balance between porosity
and mechanical strength need to be established for the development of an optimal
replacement scaffold.35
1.4.2 Mechanical properties
Studies determining the mechanical properties of various tissue types revealed that
most tissues are heterogeneous, viscoelastic, nonlinear and anisotropic materials51.35
Scaffolds are required to be mechanically strong to carry out its function while
withstanding the hydrostatic or pulsatile pressures encountered in vivo.35, 41, 42 The
bulk properties of the constituents have been used to predict the mechanical
properties of the designed scaffolds. In addition, there has also been a direct
correlation between scaffold architecture and structural features such as pore
distribution, fibre diameter and orientation and mechanical strength of the
scaffold52.35 This is of particular interest in the case of scaffolds with high total
porosity and low material content, for example in hydrogels and electrospun
scaffolds. It has also been demonstrated that the cellular response, such as cell
contractility, motility, adhesion, spreading and differentiation, can be influenced by
7
the physical properties of the scaffold like softness and hardness53-55.35 Engler et al.
reported that matrix properties such as softness and elasticity has
a profound
influence on stem cell lineage and their commitment to a particular phenotype.55 It
was shown that human mesenchymal stem cells display a more neurogenic
phenotype on soft substrates (0.1-1 kPa), a myogenic phenotype on moderately hard
matrices (8-17 kPa) and an osteogenic phenotype on hard substrates (25-40 kPa).35
In another study, soft substrates with Young’s moduli less than 1-1.5 kPa have been
reported to enhance the differentiation of neural progenitor cells.56 This property of
the substrates can be exploited to achieve tailored cell proliferation and
differentiation56.35
1.4.3 Surface properties
Cellular adhesion and growth has been shown to be correlated to the surface
properties of scaffolds including topology and chemical characteristics.57,
58
Chemical properties correspond to the cellular interactions including adhesion and
protein interactions at the material surface. Tamada and Ikada investigated 13
different polymeric surfaces for fibroblast adhesion, growth and collagen synthesis,
with a range of different surface energies using goniometer (contact angle
measurements).59 It was concluded that proliferation was independent of surface
chemistry, whereas cell adhesion which relies on cell protein-substrate interactions
was correlated to surface wettability. A positive correlation was observed between
the surface charge and density of adsorbed proteins which corresponded to better
cellular adhesion.35, 60 Furthermore, cellular behaviour was altered corresponding to
the surface topography from angular edges, abrupt grooves or other surface
indentations to smooth surfaces.61-64 A number of studies have reported the
directional motility of cells along the fibres and ridges fabricated on the substrate
surface which is explained by a phenomenon called “contact guidance”.35, 65
As highlighted above, an ideal scaffold is an amalgamation of chemical, mechanical
and structural properties. It should be biodegradable but structurally strong, should
be porous but also flexible. Therefore, an ideal scaffold should be multifunctional,
incorporating cues for cellular processes like adhesion, extracellular matrix
8
production and migration, but also have potential applications such as drug delivery
and multimodal imaging capabilities.
1.5 Skin, injury and wound healing
Skin is the largest organ on the human body. Adult skin consists of two main tissue
layers: a thin top layer called epidermis which is mainly composed of keratinocytes
while the lower epidermal layer is made up of melanocytes, cells responsible for
skin pigmentation and an underlying supporting layer called dermis which is mainly
composed of collagen synthesising cells, fibroblasts (Figure 1).66, 67 The main role of
epidermis is protection against microbes and to help regulate body temperature.68
The layer sandwiching these two layers is a 20 nm thick membrane called basement
membrane. It is made up of hemidesmosomal structures and its main role is to
mechanically stabilize the interaction between the epidermis and dermis.67
9
Figure 1: This diagram of human skin shows the two main layers of skin — the upper epidermal barrier layer
and the lower, much thicker, dermis. Figure and caption taken from MacNeil 2007. 69
The main role of dermis which is 2-5 mm thick is to provide support to the
overlaying epidermis.67 It provides considerable tensile strength and elasticity to
skin mediated by the intertwined arrangement of collagen fibers, has specialized
components and structures.67 Collagenous architecture is composed of varying
amounts of interwoven elastin fibers, proteoglycans, fibronectin and other
components.67, 68 To avoid infection, injury to the skin demands rapid intervention
where the healing time is dependent on extent and depth of the injury. Epidermal
lesions mostly heal within a week.66 Wound healing is a dynamic and interactive
process involving synergistic interactions between blood cells, ECM, and
parenchymal cells.70 Wound healing is accomplished by three successive but
overlapping stages: inflammation, reepithelialisation and remodelling (Table 1).
Failure or delays in healing can lead to significant scarring and potentially fatal
sepsis and therefore interventions to promote healing are critical to reduce morbidity
and mortality associated with extensive skin injuries.
10
Table 1: Phases of wound healing, major types of cells involved in each phase, and selected specific events.
Taken from Falanga 2005.71
1.5.1 Haemostasis and inflammation
An insult to the skin causes the disruption of blood vessels and extravasation of
blood constituents. As the blood clots it re-establishes haemostasis and provides a
provisional ECM (Figure 2).70 As the platelets aggregate they not only facilitate the
formation of a hemostatic fibrin mesh but also secrete several mediators of wound
healing, such as platelet derived growth factor, a chemoattractant that recruits and
activate macrophages and fibroblasts.70 Coagulation of platelets generates numerous
vasoactive mediators and chemotactic factors and activates complementary
pathways by injured and activated parenchymal cells. This then results in the
recruitment of inflammatory leukocytes to the injury site.70
11
Figure 2: A Cutaneous Wound Three Days after Injury. Growth factors thought to be necessary for cell
movement into the wound are shown. TGF-β1, TGF-β2, and TGF-β3 denote transforming growth factor β1, β2,
and β3, respectively; TGF-α transforming growth factor α; FGF fibroblast growth factor; VEGF vascular
endothelial growth factor; PDGF, PDGF AB, and PDGF BB platelet-derived growth factor, platelet-derived
growth factor AB, and platelet-derived growth factor BB, respectively; IGF insulin-like growth factor; and KGF
keratinocyte growth factor. Reproduced with permission from A. J. Singer and R. A. F. Clark, N. Engl. J. Med.,
1999, 341, 738-746. Copyright Massachusetts Medical Society.
Next, neutrophils infiltrate the wounded area to initiate cleaning the injury site of
foreign particles and bacteria, which are then extruded with the eschar or
phagocytosed by macrophages.70 Chemoattractants such as fragments of ECM
proteins, transforming growth factor β (TGFβ) and monocyte chemoattractant
protein 1 (MCP1), attract monocytes to the injury site. These monocytes become
activated macrophages that release growth factors and additional cytokines to initiate
the formation of granulation tissue.70 Interaction of macrophages with the ECM
through their integrin receptors, stimulate them to phagocytose microorganisms and
fragments of ECM and simultaneously, stimulate monocytes to differentiate into
pro-inflammatory or reparative macrophages
72 70
.
This interaction also stimulates
both monocytes and macrophages to express colony-stimulating factor 1, a prosurvival cytokine; tumor necrosis factor α, an inflammatory cytokine; and platelet12
derived growth factor, an important chemoattractant and mitogen for fibroblasts.70
Other key cytokines secreted by monocytes and macrophages include transforming
growth factor α, interleukin-1, transforming growth factor β, and insulin like growth
factor 173.70 Macrophages are an integral part of the wound healing process critical
to wound repair, with macrophage deficient animals shown to have markedly
defective wound healing.74, 75
Figure 3: Schematic diagram of wound reepithelialization models. (a) Basal KCs at the leading edge of the
wound that are firmly attached to surrounding basal and suprabasal KCs actively migrate to close the wound in
the “sliding” model of reepithelialization. Arrow indicates direction of movement of basal KC. (b) With the
basal KCs firmly attached to the BMZ, suprabasal KCs roll onto the wound matrix in the “rolling” model of
reepithelialization. Arrow indicates movement of suprabasal KCs as they tumble over basal KCs. Arrowheads
indicate initial cut edge. Figure and caption taken from Usui 2005.76
1.5.2 Reepithelialisation
Two mechanisms behind reepithelialisation have been proposed in the literature: the
“rolling” model and the “sliding” model (Figure 3). The “rolling” model postulates
that basal keratinocytes remain strongly attached to the basement membrane, while
suprabasal keratinocytes at the wound margin are activated to roll over into the
wound site77, 78.76 The “sliding” mechanism on the other hand, postulates that basal
keratinocytes are the principal cells mediating the migration and wound closure.
Both basal and suprabasal keratinocytes remain strongly attached to the leading edge
and basal keratinocytes are passively dragged along as a sheet79, 80.76 A variant is
also proposed where a large population of suprabasal cells migrate out of the wound
(Figure 4).76
13
Figure 4: Schematic diagram of a new model of reepithelialization. Both basal and suprabasal KCs are
activated to respond to the complex microenvironment created by a wound. Suprabasal KCs undergo
dramatic changes in response to injury and are possibly the primary source of cells available for wound closure.
Figure and caption taken from Usui 2005.76
Reepithelialisation is initiated spontaneously after injury.70 Epidermal cells from the
skin appendages initiate the cleaning process by removing the clotted blood and
damaged stroma.70 Simultaneously, cells alter their phenotype which includes
retraction of intracellular tonofilaments;78 dissolution of most intercellular
desmosomes, important for physical connections between the cells; and formation of
peripheral cytoplasmic actin filaments, crucial for cell motility81, 82.70 Furthermore,
the loss of basement membrane and hemidesmosomal links disrupts the interactions
between juxtaposed epidermal and dermal cells, allowing the lateral movement of
epidermal cells.70 Epidermal cells interact with a variety of ECM proteins via their
integrin receptors. These include proteins such as fibronectin and vitronectin which
are interspersed with stromal collagen I at the wound margin and interwoven with
the fibrin clot in the wound space83-85.70 The epidermal cells start to migrate and
dissect the wound to separate desiccated eschar from viable cells.70 Collagenase as
well as the activation of plasmin by plasminogen activator is produced by migrating
epidermal cells in order to degrade the ECM required for their migration between
the collagenous dermis and the fibrin eschar86.70 Plasminogen activator is also
14
known to activate collagenase (matrix metalloproteinase I) which propels the
degradation of collagen and ECM proteins.
Figure 5: A Cutaneous Wound Five Days after Injury. Blood vessels are seen sprouting into the fibrin clot as
epidermal cells resurface the wound. Proteinases thought to be necessary for cell movement are shown. The
abbreviation u-PA denotes urokinase-type plasminogen activator; MMP-1, 2, 3, and 13 matrix
metalloproteinases 1, 2, 3, and 13 (collagenase 1, gelatinase A, stromelysin 1, and collagenase 3, respectively);
and t-PA tissue plasminogen activator. Reproduced with permission from A. J. Singer and R. A. F. Clark, N.
Engl. J. Med., 1999, 341, 738-746. Copyright Massachusetts Medical Society.
Epidermal cell proliferation begin few days after injury.70 Local release of growth
factors, such as epidermal growth factor, transforming growth factor α and
keratinocyte growth factor;87 and absence of the neighbouring cells on the wound
margin are thought to be the principal reasons behind this epidermal response.70 The
reepithelialisation results in the synthesis of basement membrane proteins laid down
in an orderly sequence from the margins of the wound inwards, mimicking the
original architecture in a zipper like fashion88.70 Epidermal cells return to their
normal phenotype, establishing juxtaposition to the re-established basement
membrane and underlying dermis once the epidermal barrier has been restored.70
This also acts as a trigger for next step of inflammation and matrix remodelling.
15
1.5.3 Formation of granulation tissue
Four days post injury, granulation tissue starts to invade the wound space (Figure 5).
Macrophages, fibroblasts and blood vessels simultaneously start to infiltrate into the
wound space.70 Angiogenesis and fibroplasia is constantly supplemented by the
growth factor secreted by macrophages; while fibroblasts lay down ECM necessary
to support cell ingrowth and blood vessels carry oxygen and nutrients crucial to
sustain cell metabolism.70 Fibroblasts, nurtured by growth factors, especially
platelet-derived growth factor and transforming growth factor β1 (TGFβ1), in
addition to the ECM proteins, begin to proliferate and start to express integrin
receptors for their migration into the wound.70 Platelet derived growth factor, along
with many other similar factors, have been shown to accelerate healing of chronic
pressure sores and diabetic ulcers when applied exogenously to these wounds. 89, 90
Newly synthesised provisional ECM promotes granulation tissue formation by
providing the conduit for cell migration88.70 It has been postulated that the
fibronectin architecture and fibronectin specific integrin receptors on fibroblasts
control the rate of granulation tissue formation.70 These fibroblasts are responsible
for the synthesis, deposition, and remodelling of the ECM. Cell migration into the
blood clot of crosslinked fibrin, or tightly woven ECM requires fibroblast-derived
enzyme proteases like plasminogen activator, collagenases, gelatinase A, and
stromelysin that can cleave the matrix to facilitate their migration91.70 Post migration
into the wound, fibroblasts initiates ECM synthesis.92,
93
It is believed that
transforming growth factor β1 stimulates the replacement of the provisional ECM
with a collagenous scar-type matrix92, 93.70 Once an abundant collagen matrix has
been deposited, fibroblasts stop their excessive collagen production triggering the
replacement of granulation tissue by a relatively acellular scar.70
16
1.6 Role of TGFβ in fibrosis and scarring
The TGFβ superfamily contains over 30 proteins found in both vertebrates and
invertebrates and encompassing a wide range of functions throughout the lifetime of
the animal.94 The three mammalian TGFβ gene isoforms are TGFβ1, 2 and 3 which
share 60-80% sequence homology and are encoded by three different genes.95, 96
The three mammalian isoforms of TGFβ are expressed in a cell specific and
developmentally regulated manner.97,
98
For example, TGFβ1 and TGFβ3 are
expressed during the morphogenesis stage of early development followed by TGFβ2
which is expressed later in mature and differentiating epithelium.98 All three
isoforms are highly conserved in mammals, but differ in their binding affinity
towards TGFβ receptors.98,
99
Altered and different phenotypes were resulted as a
result of the deletion of individual isoforms in mice.98-100 TGFβ1-null mice were
mostly embryonically lethal due to abnormal development of the yolk sac, and those
that did survive developed multi-organ autoimmunity and multi-focal inflammation
disease and died within 3 weeks of birth101.102 TGFβ2-null mice showed perinatal
mortality due to cyanotic heart disease and various other developmental
defects103.102 TGFβ3-null mice die due to cranial bone defects especially cleft
palate104.102
Figure 6: Schematic representation of latent TGF-β. The putative transglutaminase-mediated covalent bonding
between the N-terminus of LTBP and the ECM is indicated by a question mark (?). Disulfide bonds between
LAP monomers, between TGF-β monomers and between LAP and LTBP are indicted by thin lines. Carbohydrate
residues on LTBP are not shown. Abbreviations are in the text. Figure reproduced, and caption taken from
Munger 1997.105
17
Newly synthesised TGFβ is secreted in an inactive latent form. The non-covalent
association of TGFβ with a latency associated peptide (LAP) is the reason behind
the latency (Figure 6). LAP is a homodimer formed from the propeptide region of
TGFβ.105 This association of LAP and TGFβ is termed as latent TGFβ (LTGFβ). To
exert its function, TGFβ needs to be cleaved from its latent complex with LAP.105
Two of the N-linked carbohydrate residues contain mannose-6-phosphate (M6P)
groups. There are 3 cysteines in LAP-1 which form interchain disulfide bonds to
dimerise LAP monomers.105 This cysteine interaction is critical to maintain latency
because their replacement with serines at positions 223 and 225 of LAP-1results in
the secretion of active TGFβ1.106 LAP is also known to form a disulfide linkage to
another protein called latent TGFβ binding protein (LTBP).105 LTBP link latent
TGFβ to the ECM.
TGFβ has three high-affinity membrane receptors where type I and II are
transmembrane serine/threonine kinases that coordinate to facilitate each other’s
signalling,98,
107-109
while type III receptor, a transmembrane proteoglycan, has no
specific signalling function because of its highly conserved cytoplasmic domain.98,
110, 111
TGFβ interact with these different receptors to mediate specific functions. For
example, interaction with receptor type I promotes secretion and deposition of ECM,
while interaction with type II receptor mediates cell growth and proliferation. In
order to activate the type II receptor, TGFβ binds either to a type III receptor which
then facilitates its binding to type II receptor or it can bind directly to a type II
receptor.112 Upon activation by TGFβ, type II receptors recruit, bind and
transphosphorylate type I receptors, thereby promoting their protein kinase activity
and initiating its downstream signalling cascade.112 TGFβ signalling within the cells
is mediated by the Smad family of transcriptional activators and has been
extensively reviewed elsewhere.113-118
18
1.7 Mechanisms of activation of extracellular TGFβ
Downstream signalling of TGFβ is regulated by local activation of the LTGFβ
complex in vivo.102 There are various mechanisms of LTGFβ activation which have
been widely studied.2, 102, 119 M6P/IGFII receptor mediated activation of LTGFβ has
been illustrated as one of the most critical mechanism of activation.
One mechanism of activation involves integrins. Integrins are transmembrane
receptors that conjugate the cell cytoskeleton to the ECM and are critical for cell
adhesion, proliferation, migration and differentiation120-122.102 Integrins are made up
of α and β subunits.102 αv integrins recognise and bind to both TGFβ1 and LAP via
their RGD sequence. This RGD mediated interaction between αvβ6 integrin and
LAP-1 cause the activation of LTGFβ1 by altering its conformation.102 However, β6
integrin is required to interact with the actin cytoskeleton of the cell to ascertain
LTGFβ1 activation.102,
123
This alteration in the LTGFβ1 structural conformation
produces mature TGFβ which can interact with the TGFβ II receptor and therefore
allow it to mediate its function124.102
Thrombospondin-1 (TSP-1), a 300 kDa protein found in the α-granules in platelets
and ECM has also been shown to activate both small LTGFβ (LAP-TGF-β complex)
and large LTGFβ (LTBP-LAP-TGF-β) complexes in a non-proteolytic mechanism
both in vitro and in vivo125-127.102 It has been postulated that the TSP-1 interaction
with LAP induces a conformational change in relation to mature TGFβ thereby
unveiling it to the TGFβ receptor recognition site and facilitating TGFβ binding to
its receptor128.102
The proteases such as plasmin, matrix metalloproteinase (MMP) 2 and 9 have also
been reported to induce LTGFβ activation in vitro129,
130 102
.
Plasmin conditioned
media has also been shown to generate active TGFβ in fibroblasts and Chinese
hamster ovary cultures131.102 Alternatively, the activation of LTGFβ can be inhibited
in vitro by plasmin inhibitors or prolonged by neutralising antibody to plasminogen
activator inhibitor-1 as demonstrated in co-culture studies of endothelial cells and
pericytes or smooth muscle cells132.102
19
Another mode of in vivo activation is via the binding of the M6P residue present on
the latency associated peptide (LAP) to the M6P/IGFII receptor. This invokes a
conformational change in TGFβ bound to M6P/IGFII receptor, thus allowing the
proteolytic cleavage of active TGFβ from its latent complex. In vitro activation of
TGFβ has been investigated in detail in activated macrophage cultures and in cocultures of endothelial and smooth muscle cells where latent TGFβ is activated by a
complex process involving recognition and binding of the latent form to the
M6P/IGFII receptor and the concerted action of both transglutaminase and serine
protease plasminogen/plasmin.133, 134
Once activated TGFβ suppresses the inflammatory response and promotes the
formation of granulation tissue. TGFβ promotes keratinocytes migration by
upregulating fibronectin expression and various integrins necessary for keratinocyte
adhesion, whilst simultaneously inhibiting keratinocyte proliferation thereby
regulating reepithelialisation stage of wound healing.135 It has been reported that
overexpression of TGFβ1 in the epidermis results in delayed re-epithelialisation
response in transgenic mice model.136-138 Despite having an inhibitory effect on the
proliferation of endothelial cells, TGFβ has been reported to promote angiogenesis
in vivo.139, 140
In the dermis, TGFβ activates fibroblasts which then produce higher levels of matrix
molecules including collagen, fibronectin and glycosaminoglycans (GAG), matrixdegrading proteases and protease inhibitors thereby resulting in an increase in the
production of matrix proteins and decreasing their proteolysis.141 Whilst these
changes are critical for rapid wound repair, excessive and/or extended matrix
deposition and excess TGFβ stimulation can lead to poor healing and scar outcomes.
1.8 Exogenous mannose-6-phosphate act as an antagonist of
TGFβ1 activation
Purchio et al., used radiolabeled [32P] glycopeptide and demonstrated the presence
of endogenous M6P as a carbohydrate unit on LAP (TGFβ precursor).142 Binding
20
studies demonstrated that TGFβ precursor has high binding affinity towards purified
M6P/IGFII receptor which can be hindered by exogenous M6P.142Miyazono and
Heldin studied the importance of carbohydrate in TGFβ1 activation and
demonstrated that interaction between endogenous M6P present on LTGFβ1
precursor and M6P/IGFII receptor can be result in TGFβ1 activation. In a model to
study potential anti-fibrotic agents Gosiewska et al. used macrophage-based system
for TGFβ1 activation and demonstrated that exogenous M6P inhibits the activation
of LTGFβ.143 Bates et al., used a rabbit flexor tendon in vitro and in vivo models to
investigate the role of exogenous decorin and M6P in TGFβ activation. It was
concluded that both decorin and M6P inhibits the stimulatory effects of TGFβ on
collagen production while even a single low intraoperative dose of M6P
significantly improved the range of motion of the operated tendon.144 Xia et al.,
studied inhibitory potential of M6P in three different cell types; sheath fibroblasts,
epitenon tenocytes and endotenon tenocytes from rabbit flexor tendon post TGFβ
stimulation. They reported that TGFβ induced collagen production was significantly
downregulated by the addition of exogenous M6P in a dose dependent manner in all
three cell types.145 It is clear that exogenous M6P is a potent inhibitor of TGFβ
signalling in multiple systems.
21
Figure 7: Chemical structure of a) mannose-6-phosphate showing the carbon numbers, b) an example of
isosteric phosphonate analogue and c) an example of non-isosteric analogue. Image taken from Vidal 2000.146
1.9 Structural requirements for mannose-6-phosphate recognition
to M6P/IGFII receptor
Tong et al., first elucidated the structural requirements of M6P recognition by both
the cation dependent mannose-6-phosphate receptor (CD/MPR) and cation
independent mannose-6-phosphate receptor (CI/MPR, also known as M6P/IGFII
receptor).147,
148
They investigated various ligands like M6P, pentamannose
phosphate, β-galactosides and a high mannose oligosaccharide with two
phosphomonoesters, and inferred that certain structural features are determinant in
binding to M6P/IGFII receptor such as:
a)
The hydroxyl group at the C2-position (C2-OH) of the pyranose ring must be
axial (Figure 7a) to allow hydrogen bonding interaction between hydroxyl group and
Gln-348 and Arg-391 amino acids located on β-strands 3 and 7 of the receptor.149
22
Similar strong interaction was observed in case of fructose-1-phosphate (F1P) while
glucose-6-phosphate, with C2-OH in equatorial position is weakly recognised (2epimer of M6P).150
b)
In addition to the orientation, C2-OH is also inherently required as its absence
results in diminished interaction with the receptor like in the case of 2-deoxy
glucose-6-phosphate (2dG6P).147
c)
The substitution at anomeric centre does not influence the interaction of the
analogue to the receptor as F1P is recognised in similar propensity as M6P.
Furthermore, replacement of the hydroxyl group at anomeric position with a paranitrophenoxy group, slightly improves its recognition towards the M6P/IGFII
receptor. It is due to the lipophilic interactions between the aromatic moiety of the
ligand and the binding pocket of the receptor.150
d)
Substitution at 5-position of the pyranoside ring does not influence the
interaction of analogue to the receptor.147
e)
The distance between the negative charge (phosphate group) and C5 of the
pyranose ring should be four atoms.146
f)
Only a single negative charge is necessary for the binding to M6P/IGFII
receptor since isosteric M6-phosphonate analogues (example in Figure 7b) are very
well recognised compared to non-isosteric M6-phosphonate analogues (example in
Figure 7c) which are very weakly recognised.151
1.10 Mannose-6-phosphate analogues
Due to the importance of M6P in inhibiting TGFβ and its potential to therefore
ameliorate scarring M6P has been studied in a double blind, placebo controlled,
randomised, phase 2 efficacy clinical trial (Renovo UK, Juvidex®).152, 153 The study
was carried out in almost 200 male and female subjects administering different doses
of the drug to split thickness skin graft donor sites. Intradermal delivery of the drug
was shown to significantly accelerate the rate of wound healing. However, no
significant improvement in the extent of scar formation was observed.154 This is in
part due to the subjective measurement and difficulties in measuring scar
improvement. However, the metabolic vulnerability of M6P against phosphatases
23
has also been postulated to be a major limitation in its clinical translation. Therefore
attempts have been made to develop M6P analogues with greater stability against
enzymatic degradation.
Vidil et al., reported the synthesis of M6P analogues with the aim of developing
analogues with enhanced affinity towards the M6P/IGFII receptor.155
It was
postulated that replacing the P-O bond at C6 position with P-C bond, as is the case
in phosphonate analogues, would enhance the stability of the analogues towards
hydrolases. They showed that the isosteric analogue of M6-phosphonate binds to the
M6P/IGFII receptor with higher affinity compared to the non-isosteric derivative
(with shorter chain length (1 atom) between phosphate group and pyranose ring) and
with similar affinity to M6P.155
Berkowitz et al., reported the synthesis of three mono and bivalent ligands bearing
M6P surrogates (malonyl ether, malonate, and phosphonate) and studied their
binding affinity towards the M6P/IGFII receptor. These surrogates were hydrolase
resistant phosphates (with a methylene bridge at C6 position bridging the pyranose
ring and respective functional groups) and mimic the M6P locked in the αconfiguration.156 It was concluded that phosphonate analogues have greater binding
affinity than malonyl ether and malonate analogues.156 In another example, Jeanjean
et al., developed two sulfonate (M6S) and one unsaturated phosphonate analogue of
M6P in an attempt to study their stability in human serum and their binding affinity
towards the M6P/IGFII receptor.157 The two isosteric sulfonated analogues were
synthesised with the view that higher chemical stability of conjugated (unsaturated)
analogues compared to M6S and unconjugated (saturated) analogue would have
elevated binding efficiency to the receptor. Conjugation studies were carried out to
decipher the binding of these analogues to the M6P/IGFII receptor. Binding studies
confirmed the vulnerability of M6P in human serum which dropped 5 fold while the
binding affinity of the three new analogues remained intact.157
24
Figure 8: Chemical structure of carboxylate analogues of M6P. Image taken from Jeanjean 2006. 158
In a similar study, four new carboxylate analogues of M6P were investigated for
their receptor binding affinity and serum stability (Figure 8).158 Importance of a
negative charge on the M6P analogue was emphasised by the receptor binding assay
based on the ligand dependent elution of the receptor from PMP affinity columns.159
It was demonstrated that unsaturated isosteric carboxylate analogue had similar
binding affinity as M6P while non-isosteric carboxylate analogues had slightly
weaker affinity to the receptor.158 This difference in binding efficiency was
explained by structural proximity of the double bond, which in the case of
unsaturated isosteric analogue, could stabilise the analogue and its geometry would
promote its interaction with M6P binding sites on the receptor, in comparison to the
non-isosteric (saturated) analogue. The three analogues with greater binding affinity
were also shown to have prolonged stability against hydrolases. Analogues were
shown to retain their recognition potential even after 2 days incubation in human
serum, while 6 hour incubation was shown to reduce M6P binding affinity by 32%.
Christensen et al., synthesised a series of glycopeptide derivatives of M6P,
containing two 6′-O-phosphorylated mannose disaccharides linked either α(l → 2) or
α(l → 6) and 3-5 amino acids, in order to study the influence of structural variation
in mannose disaccharides on their binding affinity towards M6P/IGFII receptor.160
The analogue comprised of two 6’-O-phosphorylated α(1→2)-linked mannose
disaccharide showed higher binding affinity compared to α(1→6)- linked
analogue,160, 161 It was concluded that analogue receptor affinity was dependent on
terminal phosphorylated mannose disaccharides and was found independent of the
25
variation in invariant saccharide core from tri-, tetra- to penta-peptides. However,
removal of aminobenzoyl group bound to the lysine tail in one of the analogue was
shown to significantly diminish their receptor binding affinity.159 This demonstrates
that structural rigidity is required to maintain strong affinity of the peptide towards
the receptor160.159
It can be concluded that there are a number of key structural requirements to be
considered in M6P analogue design. These structural constraints are important for
binding affinity and specificity and also for stability in vivo. However, despite the
design of a number of analogues to date with promising data suggesting increased
stability and binding affinity, no clinical treatment targeted at the M6P inhibition of
TGFβ signalling currently exists. Therefore it is likely that other delivery
considerations may also be important.
26
Figure 9: An illustrated representation of raised dermal scar types, as commonly observed after a midsternal incision post‐cardiac surgery. (a) Fine line scar, (b) Hypertrophic scar, (c) Intermediate raised dermal
scar, (d) Keloid scar. Image and caption taken from Sidgwick 2012.162
1.11 Scarring
There are a broad range of scar outcomes found in humans. Clinically scars are
quantified using scar assessment scales.163-165 Normal scars are usually fine line
scars which can be extended with mechanical stretch applied on the scar (Figure
9a).2,
162
Hypertrophic scars are raised dermal scars which are the result of
aggressive proliferation leading to excessive healing and matrix deposition162
(Figure 9b). Contracted scars as the name suggests result in contracture of the
granulation scar tissue around the joints following burn injury or surgery162 (Figure
9c). Finally, keloid scars are similar to hypertrophic scars but crucially extend
27
beyond the boundaries of the initial wound and are commonly considered to be more
similar to a benign tumour than a normal scar (Figure 9d).166-169
1.12 Current treatments for skin wound healing
The etiology of skin damage is diverse, from genetic disorders such as epidermolysis
bullosa through to acute trauma, tumours, chronic wounds (ulcers) and even surgical
intervention.170 One of the main factors behind a major loss of skin is burn injury.171
It is estimated that each year over 300,000 people die from burn-related injuries
worldwide.172 Many more people suffer from burn-related disabilities and
disfigurements.172
With the advent and development of new aggressive surgical interventions the rate
of patient mortality due to burn related injuries has receded considerably in recent
years. This has had a profound effect on patients, with the increase in survival being
mirrored by an increase in poor extensive scar outcomes.
Treatment of skin wounds, including burn injury, is dependent on the extent of the
injury. While superficial or incisional wounds heal with little or no intervention,
deeper wounds often require clinical intervention. The current gold standard for
replacement skin is via a split-skin graft - an autograft where skin is harvested from
an uninjured donor site on the patient, meshed to cover a larger area and patched
onto the injured site of the same patient. This approach is utilised preferentially in
full thickness and deep partial thickness wounds where both epidermal and some of
the dermal layer of the skin is harvested from an uninjured site and grafted on the
damaged site.173 The major limitation to this approach is that as the surface area of
the injury increases the availability of donor tissue decreases. Additionally, tissue
from different body sites is not equivalent and this can lead to poor aesthetic
outcomes (for example pigmentation). Finally, the use of donor sites creates an
additional wound which needs to be healed and can lead to significant donor site
morbidity.173,
174
In order to address these problems it is essential to develop new
28
interventions. Synthetic skin substitutes and engineered replacements are employed
to restrict the burden of disease by reducing the need for graft donor sites.175
1.13 Cell based therapies
Suspended keratinocytes cell based therapies are emerging as a leading therapeutic
strategy, aimed at utilising cells as replacement therapy to repair severely damaged
tissues176.173 In 1975, the first subcultures of human keratinocytes was developed
and manipulated to obtain epithelial sheets for subsequent grafting177.173 Their first
major clinical application was reported in 1981 where cultured autologous epithelial
sheets were used for the treatment of extensive third degree burns178.173 The time to
culture these sheets, approximately 3-5 weeks, was a major limitation leaving
patients susceptible for prolonged periods after the injury179-181.173 Some of the other
limitations included high production costs,182, fragility, variable engraftment rates,
difficult handling, storage and preservation of viable sheets.173,
183-186
It was the
development of an automated membrane bioreactor which brought down the delay
time to 2 weeks187.173 An alternative innovative approach was later developed, to
further reduce fabrication time, where cells extracted from a skin biopsy were not
cultured but rather directly sprayed onto the lesions (ReCell®; Avita Medical,
Australia and Spray®XP; Grace, USA)183, 188, 189.3 The biggest advantage in using
cultured keratinocytes, or non-cultured sprayed cells, was the coverage of larger
surface areas from relatively small biopsies in the range of 2-5 cm2 from an
uninjured site of the patient. It is unsuitable to be used as a sole treatment for
patients with deep dermal or full thickness injuries, as it predominantly caters to the
epidermal cell population while deeper injuries inevitably also demand dermal
support for efficient healing response.183 Nevertheless, autologous cells are in
widespread use today for the treatment of partial thickness burns. More recent
research has involved further developments of cell-based treatments, with the
potential to enhance these treatments through novel delivery mechanisms or the
addition of growth factors or other biologicals to promote better healing.3,
29
190-195
However, these advanced cell therapies remain in the research stage and are not
currently in widespread use for burn treatment.
1.14 Matrix based therapies
1.14.1
Amniotic membrane
Since 1910, allogenic amnion has been used as a biological wound dressing.173 It
draws parallel to normal human skin allografts and was considered the most
effective dressing ever used, especially in preserving the healthy excised wound bed
and providing protection from pathogenic contamination196-199.173 Amniotic
membrane is derived from the innermost layer of the fetal membrane, it is a semitransparent tissue consisting of avascular stroma, thick continuous basement
membrane with a full complement of collagen IV and V, laminin and also contains
several protease inhibitors200.173 Advantages of human amniotic membranes include
promotion of epithelial regeneration by reduction in loss of proteins, electrolytes and
fluids, drug delivery and reducing the need and frequency of dressing changes and
pain associated with the process.173 The key disadvantages are its fragility,
technically it is difficult to handle, the risk of contamination and transmission of
disease and it is inadequate in deep dermal injuries where it disintegrates before
healing occurs201, 202.173
1.14.2
Human cadaver derived allografts and xenografts
Human cadaver allograft skin (HCAS) can be used as a temporary dressing in cases
where the availability of patient donor sites are limited203-205.173 Similar to amniotic
membranes, HCAS may also be used as wound dressing to cover widely meshed
autografts in massive burns.173, 206 There are serious problems associated with HCAS
use including limited supply, variable quality, possible contamination and immune
rejection.204, 206, 207 On the other hand, xenografts are derived from various animal
species including rabbit, dog and pig and have been used as a temporary replacement
30
skin for a long time.173 The major limitation in their use include immunologic
disparities and predetermined rejection over time, despite the obvious benefits of
biologically active dermal matrix they provide.203 These alternatives continue to be
considered strictly as only temporary wound cover with therapeutic strategies to
promote healing focused on living autograft cells and tissue.
Figure 10: (a), Cells are isolated from the patient and may be cultivated (b) in vitro on two-dimensional surfaces
for efficient expansion. (c), Next, the cells are seeded in porous scaffolds together with growth factors, small
molecules, and micro- and/or nanoparticles. The scaffolds serve as a mechanical support and a shapedetermining material, and their porous nature provides high mass transfer and waste removal. (d), The cell
constructs are further cultivated in bioreactors to provide optimal conditions for organization into a functioning
tissue. (e), Once a functioning tissue has been successfully engineered, the construct is transplanted on the defect
to restore function. Image and caption taken from Dvir 2011. 5
31
1.14.3
Tissue engineered skin
In this approach skin cells are expanded in the laboratory before their application in
wound healing (Figure 10).69 In addition, for deeper injuries a number of dermal
scaffolds have also been developed that can be used in conjunction with the cell
based approaches to restore both dermal and epidermal layers.
All tissue engineered materials needs to have some essential characteristics
including low disease risk, promotion of healing, possess physical properties similar
to normal skin, it must attach well to the wound bed, be compatible to the
developing new vasculature, non-immunogenic and be convenient to use.69,
171
To
replace the dermal layer, predominantly collagen based matrices have been used,
with Integra the first demonstration of a bilayered and biocompatible dermal
scaffold used in the treatment of extensive burn injury to successfully induce the
synthesis of neodermis.141, 208
Epidermal tissue grafts consist of in vitro differentiated keratinocytes forming a
stratified epidermal layer.209 Keratinocytes and fibroblasts are cultured on various
biocompatible substrates as implantable scaffolds including both natural and
synthetic materials.209 Epicel® (Genzyme Corporation Offices, Cambridge, MA) is
the first commercialized epidermal autograft which is composed of cultured
keratinocytes extracted from 2 x 6 cm biopsy taken from the patient, which are
expanded in the laboratory and then cocultured in the presence of proliferation
arrested murine fibroblasts to form grafting sheets.209, 210 It is a front line treatment
for full thickness burns in US and Europe.209 The clinical outcome of epidermal
transplants fall short in cases with inadequate dermal tissue, which is required to
provide support for grafted epidermal sheets211. Alternatively, biocompatible and
biodegradable materials, such as benzyl-esterified analogues of hyaluronic acid
(Hyalograft 3D, Fidia Advanced Biopolymers, Padua, Italy) and polyglycolic acid
(Dermagraft, Shire Regenerative Medicine, San Diego, CA) have been used.209
Other examples include Transcyte (formerly Dermagraft-TC, Shire Regenerative
Medicine) which is made up of porcine dermal collagen coated nylon mesh seeded
with neonatal human fibroblast and externally supported by a silicone
membrane212,209 Dermagraft (Shire Regenerative Medicine), utilise a biodegradable
32
scaffold coated with an allogenic human-derived dermal fibroblasts culture, and
capable of producing several growth factors to stimulate angiogenesis, tissue growth
and reepithelialisation from the wound edge even after cryopreservation and
subsequent thawing213,209 and Apligraf (Organogenesis, Canton, MA) is an example
of a clinically approved allogenic dermoepidermal product (combinatorial therapy),
composed of cultured keratinocytes and a fibroblast dermal layer on collagen I
matrix, for the treatment of venous and diabetic ulcers214.209 Clinical trials revealed a
rate of healing of 63% in cases of ulcers as compared to 49% for compression
therapy and a vast improvement in healing times from 181 to 61 days under
Apligraf.215 Guo et al. developed a bilayer dermal equivalent by loading plasmid
DNA (pDNA) encoding vascular endothelial growth factor-165 (VEGF165)/N,N,N-trimethyl chitosan chloride into collagen-chitosan/silicon membrane
scaffold and tested for regenerative properties in a porcine full-thickness wound
model216.3 Full-thickness burn wounds in the pig model when treated with N,N,Ntrimethyl chitosan chloride/pDNA-VEGF dermal substitute showed the best
response towards dermal regeneration with highest density of newly-formed and
mature vessels compared to a blank scaffold and scaffolds loaded with naked
pDNA-VEGF and TMC/pDNA-eGFP.216 In another study, wound healing potential
of five acellular dermal skin substitutes (Integra®, Integra Life Sciences, NJ, USA;
ProDerm®, UDL Laboratories Inc., IL, USA; Renoskin®, Symatese, Ivry-le-Temple,
France; Matriderm® 2 mm, Ideal Medical Solutions, Wallington, UK; and
Hyalomatrix® PA, Addmedica, Paris, France) were investigated in a porcine fullthickness wound model in a two-step process217.3 In the first step, skin substitutes
were implanted followed by the reconstruction of the epidermis using an autologous
split-thickness skin graft or cultured epithelial autograft (after 21 days, current
clinical treatment regime).3 Significant differences between the skin substitutes in
terms of dermis incorporation and early wound contraction was observed. However,
no significant long-term differences were in scar qualities between different dermal
substitutes and the control group.217 Dermoepidermal substitutes mimic both
epidermal and dermal layers of the skin and are considered the most advanced
substitutes available in the clinic.3 They made up of both keratinocytes and
fibroblasts incorporated into the scaffold which can provide temporary cover,
promote simultaneous dermal and epidermal regeneration and also resemble the
normal skin structure.3 The main limitations, however, are the high production costs
33
and their limited efficiency in terms of permanent wound closure due to allogeneic
cell rejection, as well as limited evidence that the tissue substitutes ultimately
improve the final outcome170.3
1.15 Nanofibrous scaffolds for skin tissue engineering
Tissue engineering (TE) and regenerative medicine is the amalgamation of cellular
components of a living tissue and functional biomaterials to develop functional
living tissue of sufficient size for clinical translation.218 Recently, cell patterning,
migration, proliferation and differentiation have been the point of focus for both TE
and regenerative medicine.218 Development of scaffold as ECM substitutes are
aimed to promote cells proliferation and differentiation for skin tissue regeneration
in vivo.218 Cells respond to ECM signalling in a multistep process which is initiated
by cell receptor-ECM ligand interactions, followed by the sequestration of growth
factors by ECM, spatial cues and the mechanical force transduction219.218
Nanofabrication techniques have been employed to develop complex multifunctional
porous, nanometer-sized fibrous scaffolds with surface morphology that can
determine and influence cell fate, regulate the expression of specific proteins and
encourage cell-specific scaffold remodelling.220.218 Techniques including nanoscale
surface pattern fabrication such as lithography, electrospinning and self-assembly
has been explored to fabricate scaffold topography down to nanoscale221.218 They
can incorporate a myriad biological cues such as growth factors, angiogenic factors,
cell surface receptors and spatial cues.218 These approaches can be used to influence
cell proliferation, migration, differentiation and 3D organisation including
ingrowth.3, 218 Electrospinning has gained considerable amount of interest because of
its operational simplicity, versatility, ability to process variety of materials and
propensity in producing nanofibers mimicking the ECM, for the fabrication of
nanoscaffolds for skin regeneration.3, 218
Electrospinning utilise electric field to produce a highly impermeable, non-woven
matrix of sub-micron to nanoscale fibers by pushing a highly charged polymer liquid
jet through a small diameter nozzle of a syringe24, 222.223 Conventionally, a Taylor
34
cone is formed when electrostatic charging of the fluid pulls the polymer solution at
the tip of the nozzle, ejecting a single fluid jet from the apex of the cone. 223 This
polymer jet accelerates and thins in the electric field to form a nanofibrous scaffold.
Electrospun scaffolds with different morphologies from aligned to random have
been fabricated with fiber diameters down to the 100 nm range with various pore
sizes224.209 Electrospun scaffold with high porosity and large surface area to volume
ratio facilitate efficient mass transfer into the 3D structures of the scaffold.209
Nanofibrous scaffolds serve as a combinatorial approach towards regeneration of
various tissues such as skin, bone, cartilage, vascular and neural tissues and
subsequent drug delivery platform.209
Figure 11: Various fibrous structures fabricated via an electrospinning technique. (a) Core-shell structure,
(b) random/align directional fiber and (c) micro-/nano-sized fiber. Image and caption taken from Rim 2013. 225
Nanofibers can be electrospun in various patterns including random, aligned and
core-shell (Figure 11).209 Each type has their own advantages; for example random
35
fibers are used in tissue engineering approaches, while aligned ones have been
explored in solar cells and shown to promote their efficiency225.209 In an another
approach, emulsion electrospinning is used to fabricate polymer fibers with an
integral core-sheath structure, promoting loading efficiency of the cargo, avoiding
the initial burst release and protect and help maintain bioactivity of the loaded
protein by avoiding the exposure of the cargo to the degrading physiological
environment226.209 These fibrous scaffolds can be further functionalised using
surface modification techniques such as plasma treatment, the wet chemical method,
surface graft polymerization, co-electrospinning, and the immobilization of bioactive
ligands227.209 In order to achieve controlled drug release the payload like drugs,
enzymes and cytokines are physically or chemically immobilised within the
fibers228.209
As a dermal substitute, highly porous 3D collagen scaffolds were fabricated using a
electrospinning and were evaluated in a keratinocyte/fibroblast co-culture in vitro
model for skin tissue engineering229.209 It resulted in well dispersed fibroblast
ingrowth, while keratinocytes migrated through the pore structure and differentiated
on the scaffold surface to form a stratum corneum and a 3D dispensed scaffold
similar to normal skin.229 A well interconnected porous network is required in 3D
scaffolds to achieve guided cell adhesion, cell growth favouring tissue development,
cell migration and subsequent transport of solutes.209 This highlights the importance
of mechanical properties of the engineered electrospun scaffolds towards their
functions.209 Parenteau-Bareil et al crosslinked collagen, derived from various
sources, with chitosan to produce ECM mimicking tissue engineering scaffolds230.209
Crosslinking of collagen increased mechanical strength and facilitated a similar
cellular response to normal ECM whilst maintaining the biocompatibility of the
scaffold. Nanofibrous scaffolds provide a high surface area allowing oxygen
permeability and avoid fluid accumulation at the wound site, while preventing
microbial infection by restricting their penetration through its small inter-fiber pores,
thereby making them ideal candidates for wound dressings.209
The aim of using an electrospun scaffold based TE approach is to facilitate and
promote the natural healing response of the body without inducing an immunogenic
response, allowing the body to utilise the scaffold to enhance regeneration of
“neonative” functional tissues.209 Electrospinning provides the flexibility to
36
coelectrospin polymers with drugs or proteins to form nanofibrous scaffolds for
applications in drug delivery.209 Various polymers have been electrospun including
synthetic ones such as poly(ε-caprolactone) (PCL), polylactic acid (PLA),
polyglycolic acid (PGA), PLGA, polystyrene, polyurethane (PU), polyethylene
terephthalate, poly(L-lactic acid)-co-poly(ε-caprolactone) (PLACL) and natural
polymers including collagen, gelatin and chitosan to obtain fibers with diameters
ranging from few nanometers to several microns24, 224, 228, 231.209
1.16 Skin regeneration using electrospun scaffolds
The high surface area to volume ratio, porosity and structural similarity to the ECM
architecture of the dermis makes nanopatterned electrospun fiber meshes ideal
scaffolds for skin regeneration232-234.3 The limitations associated with electrospun
scaffolds include poor mechanical properties, non-uniform thickness distribution and
poor integrity.3 For skin regeneration a myriad of both natural polymers,235-237
synthetic polymers,232,
238-240
and combinations of the two have been successfully
electrospun and tested.3 Poly(ε-caprolactone)/gelatin nanofibrous scaffold was
electrospun on top of a commercial polyurethane wound dressing (TegadermTM; 3M
Healthcare, USA) in an attempt to regenerate dermal wounds241.3 It was shown to
promote cell proliferation, adhesion and growth in human dermal fibroblasts. Yang
et al. employed emulsion electrospinning to fabricate ultrafine core sheath
poly(ethylene glycol)-poly(D,L-lactide) fibers loaded with basic fibroblast growth
factor (bFGF) and demonstrated its gradual release over 4 weeks232.3 An in vitro
study using mouse embryonic fibroblasts showed enhanced cell adhesion,
proliferation and secretion of ECM when cultured on a nanofibrous scaffold.232 In
vivo studies using a dorsal wound model in diabetic rats treated with bFGF/
poly(ethylene glycol)-poly(D,L-lactide) mats showed elevated rates of healing with
complete re-epithelialization and regeneration of skin appendages, while bFGF from
the scaffold promoted collagen deposition and its remodelling to normal
architecture232.3 In another study, potential of bone-marrow derived mesenchymal
stem cells differentiation into the epidermal lineage (keratinocytes) was investigated
37
on poly(L-lactic acid)-co-(ε-caprolactone) nanofibrous scaffold, with and without
collagen242.3 It was concluded that better results were obtained for collagen
containing poly(L-lactic acid)-co-(ε-caprolactone) nanofibrous scaffolds.242
Electrospun nanofiber scaffolds have also been explored as delivery systems for
variety of bioactive molecules including drugs, proteins and even RNA243-247.3
Ignatova et al. reviewed the use of nanofibrous scaffolds loaded with various
antibiotics and antibacterial agents including nanoparticles for wound dressing
applications248.249 Kenawy et al. electrospun poly(ethylene-co-vinyl acetate),
poly(lactic acid) and their blend with tetracycline hydrochloride as a model
drug250.249 It was reported that drug release behaviour was dependent on the polymer
carrier and drug loading. Scaffolds with a 50/50 blend and relatively low drug
loading (5 wt%) showed sustained release for over 5 days. Higher drug loading (25
wt%) resulted in an initial rapid burst release of the surface adsorbed drug that
quickly dissolved in tris buffer.249 Sodium alginate was electrospun with poly(vinyl
alcohol) containing different concentrations of ZnO nanoparticles as an antibacterial
agent in a mouse fibroblasts model251.3 They observed inverse correlation between
cell adhesion and spreading, and ZnO concentration in the scaffolds. The
antibacterial activity was evaluated in both gram positive and gram negative
bacterium cultures through diffusion disc experiments using Staphylococcus aureus
and Escherichia coli. The results showed direct correlation between antibacterial
activity of the scaffold and increasing concentration of ZnO nanoparticles251.3
Various strategies have been employed to control the drug release mechanism from
fibers with the aim to prevent or control the initial burst release of the drug. Dave et
al. developed electrospun enzyme-embedded antibiotic-releasing polymer scaffolds
using antibiotic gentamicin sulfate (GS) and a polymer degrading enzyme (lipase) in
PCL polymer to achieve endogenously triggered controlled drug release252.249The
GS release from the scaffold was shown to be dependent on lipase concentration. In
an another study, PDLLA was electrospun with Mefoxin to develop a nanofibrous
scaffold with the loading efficiency of 90%.249 However, most of the drug was
released in the first 3 h which reached its completion within 48 h, indicating that the
drug was adsorbed on the nanofiber surface.249 In order to block such initial burst
release of Mefoxin from PLGA nanofibers, Kim et al. used an amphiphilic block
copolymer (PEG-b-PLA) and reported prolonged drug release for up to 1 week253.249
38
In an another attempt, Yohe et al. used superhydrophobic polymer made up from
PCL doped with 0-50 wt% poly(glycerol monostreate-co-ε-caprolactone) (PGCC18) as a hydrophobic polymer and trapped air as a barrier to control the rate of
drug release from the scaffold by blocking the fiber pores254,
255 249
.
By using the
model bioactive agent SN-38 (7-ethyl-10-hydroxy campthothecin), PCL electrospun
fibers doped with 10 wt% PGC-C18 was shown to allow to have linear sustained
release over 60 days compared to initial burst release (10 days) observed in case of
PCL fibers alone which plateaued out at 20 days255.249Alternatively, covalent
conjugation of the drug to the fibers can be used as a method to control and regulate
the drug release mechanism from the fibrous scaffolds.249 Jiang et al. explored this
by covalently functionalising ibuprofen onto
poly(ethylene glycol)-g-chitosan
(PEG-g-CHN) polymer and coelectrospinning it with poly(lactide-co-glycolide)
(PLGA) to yield nanofibrous scaffold with prolonged release of the drug for more
than 2 weeks256.249 Zou et al. copolymerised ε-caprolactone with PDLL, to induce
functional ketone groups for later functionalization into a PDLL backbone, and used
them to develop a fibrous scaffold257.249 Subsequently, the scaffold was surface
functionalised with heparin molecules and loaded with bFGF. It was reported that
bFGF release was dependent on the amount and molecular weight of heparin used.
One of the biggest limitations with this approach is the necessity to select carrier
polymers with desired functional groups257.249
In an alternate approach, polymer crosslinking has been explored as a parameter to
control the initial burst release of the drug.249 Some of the methods used include
UV-irradiation,258-260 dehydrothermal treatment,261 and chemical treatment including
glutaraldehyde,262,
prepared
263
formaldehyde264 and carbodiimide265,
Fenbufen-loaded
266 249
poly(D,L-lactide-co-glycolide)
.
Meng et al.
(PLGA)
and
PLGA/gelatin nanofibrous scaffolds267.249 Fenbufen release was found to be
dependent on the concentration of gelatin used, alignment of the fibers in the
scaffold and amount of crosslinking. It was shown that subsequent crosslinking
treatment of the fibers effectively curtailed the burst release of the FBF at the initial
release stage from a PLGA/gelatin (9/1) nanofibrous scaffold267.249 In an attempt to
control the formation of hypertrophic scars (HS), poly(l-lactide) was coelectrospun
with 20(R)-ginsenoside Rg3 (GS-Rg3).268 GS-Rg3 was used for its potential in
inhibiting the formation of HS in vivo as measured using H&E staining and
39
apoptosis of fibroblasts to control the later stage HS hyperplasia by inhibiting
inflammation and down-regulating VEGF expression.268 The in vivo wound rabbit
ear HS model suggested sustained release of the drug for 3 months which was
reported to be dependent on the drug concentration in the electrospun fibers. GSRg3/ poly(l-lactide) scaffold was shown to significantly inhibit HS formation with
decreased expression of collagen fibers and microvessels.268
Bioactive molecules such as proteins, DNA, RNA and growth factors have also been
encapsulated in polymer fibers mostly using modified techniques such as blend
electrospinning and coaxial electrospinning.249 As the name suggests, blend
electrospinning involve premixing of bioactive molecules with polymers prior to
electrospinning increasing the surface localization efficiency of bioactive
molecules.249 In an attempt to study if the activity of an encapsulated protein can be
maintained after electrospinning, human β-nerve growth factor (hNGF) was
electrospun along with BSA as carrier protein, into a partially aligned nanofibrous
scaffold using poly ε-caprolactone (PCL) and poly(ethyl ethylene phosphate)
(PEEP) polymers269.249 While the protein aggregates were heterogeneously
distributed within the fibers, its bioactivity was demonstrated in a neurite outgrowth
assay in PC12 cells. It was concluded that the prolonged release of an active protein
can be achieved over the period of 3 months269.249 The same group also
demonstrated the efficacy of this polymer system in the delivery of small interfering
RNA (siRNA) and transfection reagent (TKO) complexes270.249 The siRNA/TKO
complexes showed better gene knockdown efficiency than siRNA alone both of
which were shown to be able to release from the scaffold for almost a month.
Coaxial electrospinning was developed to fabricate core-shell fibrous scaffolds
where both polymer and biomolecules are coaxially and simultaneously
electrospun.249 The strategy involves use of the shell polymer to protect the cargo
encapsulated in the core from the physiological degradation whilst allowing slow
sustained release of the therapeutic.249 A nanofibrous scaffold encapsulating plasmid
DNA (pDNA) within the PEG core and the non-viral gene carrier poly(ethyleimine)hyaluronic acid (PEI-HA) within the PCL sheath was prepared by the coaxial
electrospinning technique271.249 They also studied the effects of various processing
parameters using fractional factorial design. Extended release of the non-viral gene
delivery vector from the sheath was observed over a period of 60 days quantified
40
using EGFP transfection expression and was shown to be dependent on pDNA
loading into the fibers. Mickova et al. studied if liposome activity can be maintained
post electrospinning process. They studied liposomes blended within nanofibers and
polyvinyl
alcohol-core/poly-ε-caprolactone-shell
nanofibers
with
embedded
liposomes prepared using coelectrospinning technique272.249 They concluded that
liposomes encapsulated in the core/shell fibers retained enzymatic activity of
encapsulated horseradish peroxidase while blended liposome became inactive.
1.17 Electrospun hybrid materials
Despite the evidence of the capabilities of electrospun scaffolds in tissue engineering
and drug delivery this technology has not reached its full potential in terms of
translation. One of the biggest limitations associated with electrospun scaffolds
developed from a single polymer has been the lack of multi-functionality and
universality. They require specific surface modification approaches which are
limited to specific polymers and can be difficult to translate into other polymer
models. In an alternate approach, nanoparticles have shown tremendous potential
and versatility not only in drug delivery applications but also in the fabrication of
various functional materials which has already been translated into commercial
products. Incorporation of these two technologies could open new avenues by
utilising the potential of both technologies. In the following sections the potential of
this combinatorial approach is demonstrated for applications in material science.
This approach can be extrapolated towards tissue engineering using multifunctional
scaffolds.
Among the various fibers prepared using electrospinning, the nanoparticles (NPs)
containing electrospun fibers exhibit a huge variety of potential applications.273 The
composite fibrous mats show flexibility, are free standing,273 and incorporate the
advantages of both starting materials i.e. polymer and the NPs. There are three main
ways specified in the literature to fabricate NPs-electrospun fiber composite
materials.
41
1.
Electrospun fibers with surface functionalised nanoparticles: This is an
efficient method where pre-treated electrospun scaffold is immersed into the
colloidal NP solution to facilitate surface adsorption of NPs on fibrous scaffold.273
Various kinds of NPs such as metal, metal oxide, carbon, polymer, fluorescent NPs
and even cells have been successfully combined with electrospun fibers to form
composite materials15,
274, 275 273
.
For example, a water stable poly(vinyl alcohol)
(PVA)-AuNPs composite fibrous mat was fabricated by pretreating the surface of
the fibers with 3-mercaptopropyltrimethoxysilane, allowing the adsorption of
AuNPs on the surface of the fibers via Au-S bonds276.273 In cases where the
electrospun fibers are stable but do not readily adsorb NPs, an in situ reduction
method is used.273 For example, Au adsorbed TiO2 composite nanofibers were
prepared by organic capping agent mediated photocatalytic reduction of
HAuCl4277.273 It was reported that shape of the nanomaterial is dependent on type
and concentration of the capping reagent.277 It provides the means to fabricate
specific nano-structures on the nanofibers for respective applications in chemical
and biological sensing277.273 Various NPs have been synthesised on electrospun
fibers using this technique including Au, Ag, Pd, Pt, TiO2, WO3 and SnO2278-282.273
2.
NPs encapsulating electrospun fibers: In this method NPs are synthesised
within the polymer fibers.273 Typically, the polymer solution is electrospun with a
metallic or ceramic acetate precursor which is then annealed to reduce the precursor
and form NPs using methods such as gas-solid reaction, calcination or laser
ablation.273 Yang et al. used AgNO3 as a precursor by electrospinning it with
poly(acetonitrile) (PAN) which upon exposure to HCl yielded AgCl NPs both on the
surface and interior of the fibers283.273 Similarly, copper nitrate was used as a
precursor and electrospun with poly(vinyl butyral) (PVB) to give fibrous scaffold
encapsulating copper nitrate. Polymer layer was removed by annealing the scaffold
at 450 °C in an air atmosphere for 2 h. Subsequent heating at 300 °C for 1 h under
hydrogen atmosphere was then applied to obtain copper fibers284.273This technique
can also be used to fabricate more specialised structures.273 For example, Sn NPs
were used as a precursor and were electrospun with porous multichannel carbon
microtubes. The resulting scaffold underwent calcination to yield Sn-Carbon
nanofibers285.273
42
3.
NPs-electrospun fibers: This is a one-pot fabrication technique where NPs
are homogenised in a polymer solution and electrospun together to obtain composite
fibers.273 This technique relies on the homogeneity of the polymer-NP solution to
yield scaffold with uniformly distributed NPs.273
Inorganic NPs can easily be
encapsulated within the electrospun fibers because of their high electron density.273
Heterogeneous solution of NPs results in NP cluster formation within the
electrospun fibers, which can be address by surface treatment of the fibrous
scaffold.273 Both hydrophobic and hydrophilic polymers have been explored and
electrospun. Some of the examples include (poly(ethylene oxide) (PEO), poly(vinyl
alcohol) (PVA), poly(vinylpyrrolidone) (PVP), polyacetonitrile (PAN), PLGA and
poly(L-lactide) (PLLA).273
Magnetite (iron oxide (Fe3O4) nanoparticles), another interesting class of
nanoparticles have been explored for their magnetic resonance properties. These
properties have been shown to be retained in the electrospinning process opening
avenues for various applications.273 It can be used as intelligent fabrics in defence
clothing, ultrahigh-density data storage, sensors and in health care.273 Problem in
their use lies in their aggregation which tends to curb their magnetic response. It is
understood that magnetite nanoparticles form clusters to reduce their energy because
of their high surface area to volume ratio.273 Stabilisers have been employed to
overcome this problem. The other alternatives include the use of electrostatic
surfactant and steric polymers286-288.273 Coaxial electrospinning was used to address
this problem of aggregation of nanoparticles. Core-shell Fe3O4-poly(ethylene
terephthalate) magnetic composite nanofibers were electrospun where Fe3O4NPs
form extended aligned structures within the core of the nanofibers and were further
shown to have retained their super-paramagnetic behaviour289.273 The core-shell
nanofibers demonstrated higher mechanical properties and responded to externally
applied magnetic field due to the dipole-dipole interaction between magnetic NPs in
magneto-rheological fluid.273 Other composites using similar approach include
polyarylene ether nitriles-Fe-phthalocyanine-Fe3O4, Fe3O4-poly(ethylene oxide) and
Fe3O4-poly(ethylene terephthalate)273, 289, 290, 291.
Many metal and inorganic fluorescing NPs have also been electrospun for their
luminescence properties including boron carbonitride oxide, sodium yttrium
43
fluoride, quantum dots and rare earth metals292, 293.273 Some of the examples includes
Sm+3/TiO2, NaYF4:Yb+3, Er+3/SiO2 and YVO4: Eu+3/PEO for rare earth metals,294-296
CdTe and CdS quantum dots. #498;, #499}.273
As demonstrated, electrospun polymer/nanoparticle composite materials showcase
hyphenated properties of the two constituents. This approach has great potential not
only in material science but also in tissue engineering applications. Electrospun
scaffolds can be functionalised with different types of nanoparticles to induce multifunctionality, for example magnetite and upconverting nanoparticles can be
incorporated within the polymer scaffold to achieve sustained drug delivery and
simultaneous imaging using MRI and fluorescence microscopy.
1.18 Summary
With surgery continuing to be the mainstay for treating burn injury and
reconstructing scars and limited evidence for the efficacy of cell therapies alone,
there is a need for the development of alternative treatment modalities. M6P has
been explored as an anti-scarring agent and its exogenous delivery has been shown
to be antagonistic towards the activation of LTGFβ. TGFβ is a key cytokine that
influences scar outcome. Upon activation, TGFβ upregulates the expression and
synthesis of ECM proteins and molecules including collagen I, fibronectin,
hyaluronic acid and α smooth muscle actin. The extent to which these molecules are
expressed closely correlates to the extent of scarring. It has been proposed that
clinical trials for commercial M6P (Juvidex®) could not meet the final end point
because of the delivery mechanism adopted and the metabolic vulnerability of M6P.
The mode of M6P delivery is critical as it gets rapidly metabolised and appropriate
vehicles have to be employed to achieve and maintain requisite tissue levels over the
post-wounding or healing period. This shortcoming of M6P needs to be addressed
for its development as a potent anti-scarring drug and can potentially be solved
through the development of more stable analogues.
44
Phosphonate analogues are metabolically more stable than phosphate analogues and
they also project higher affinity towards the receptor (M6P/IGFII receptor). Two
novel phosphonate analogues of M6P, analogue 1 and analogue 2 were investigated
for their toxicity and efficacy in inhibiting the expression of collagen I under TGFβ1
stimulation in a primary skin dermal fibroblast in vitro model. With analogue 2,
whilst it is able to penetrate cells due to its lipophilicity, this is also a limitation in its
sustained topical delivery and therefore new modalities need to be explored to
achieve local, controlled and consistent administration of the drug.
Conventional skin substitutes are still prevalent in clinic use. However, they suffer
from various disadvantages including low adhesion, creation of a new injury site (in
autografts) and immune rejection (in allografts). To address these problems
alternatives have been developed. Firstly, cell-based epidermal substitutes, efficient
in accelerating the reepithelialisation phase of wound healing but only applicable in
superficial deep wounds without additional therapies have been widely used.
Dermal substitutes, both natural and synthetic have also been developed and used in
deeper injuries. However, these dermal substitutes have problems with inadequate
vascularisation, low mechanical strength, uncontrolled degradation profiles and high
costs. Finally, the most advanced dermo-epidermal substitutes aimed at
simultaneous regeneration of both epidermal and dermal layers by using constructs
incorporating respective cells from both layers of the skin have more recently been
trialled. The major limitations have been high production costs, poor vascularisation
and failure to achieve permanent wound closure due to allogeneic cell rejection.3
Nanotechnology based techniques such as nanoparticles or nanofibrous scaffold
based delivery systems have been shown to facilitate the delivery of hydrophobic
drugs. Nanofibrous scaffolds are preferred because they mimic ECM, providing
congenial environment for cellular ingrowth, and simultaneously deliver drugs to
promote regeneration. In addition, the fractal-like geometries of fibrous scaffolds
have been of interest recently to design and develop personalised diagnostic devices
using fractal models of complex biological processes. For example, fractals have
been used to describe the architecture of tumor microenvironments. Most of the
polymers used thus far have been designed for specific applications and require
specific approaches for surface modifications which are complex and laborious in
order to induce any multifunctionality. PGMA, which has an epoxy group in each
45
monomer unit, would be an ideal polymer to address the concerns with
multifunctionality. A nanofibrous fractal-like scaffold using PGMA was developed
with an ability to incorporate multimodality and multifunctionality.
It has been demonstrated that existing clinical treatments for wound healing have
various shortcomings and there is a need for improved therapeutic approaches. TE
inspired approaches including electrospun scaffolds, although having shown promise
in addressing the shortcomings with conventional approaches, are marred with their
own limitations restricting their clinical translation. These include poor mechanical
strength, inadequate vascularisation and cellular ingrowth, difficult to control drug
diffusion and limited ability to be functionalised. The inadequacy of these scaffolds
for improving scar formation leaves a current unmet medical need. In this thesis, a
multifunctional scaffold has been developed and characterised, with data describing
its use in a range of applications presented. In addition, data showing the potential
efficacy of a stable analogue of mannose-6-phosphate for wound healing has been
demonstrated. This has led to initial work on a combinatorial approach using both
scaffold and analogue to overcome many of the limitations previously discussed and
potentially enhance wound healing and reduce scar formation in patients into the
future.
* PXS25 would be referred to as analogue 1 and PXS64 would be referred as
analogue 2 except in the paper #3
1.19 Hypotheses and Aims
Hypotheses
1.
The epoxy functionalised polymers can be used to develop multifunctional
fractal-like substrates
2.
Stable isosteric analogues of mannose-6-phosphate (M6P) can be used to
reduce the overexpression of collagen I gene
46
Aims
1.
Optimise the parameters for electrospinning PGMA and study its efficacy as a
multifunctional polymeric substrate
2.
Investigate the biocompatibility and potential of two M6P analogues as anti-
scarring agents in human dermal skin fibroblasts
3.
Investigate the potential of a combinatorial scaffold:drug approach to limit
fibrosis
47
Chapter 2
Introduction to the series of papers
2.1 Development of a universal multifunctional scaffold
The first paper presented as part of this thesis presents the preparation of universal
multifunctional fractal-like scaffold. Fractals are structures that maintain their
structural characteristics with successive variation of scale.297 In addition to their
applications in material science, fractals also find applications in cancer biology.
They have been used to understand tumor architecture and morphology with
implications in tumor growth and angiogenesis.298
Polyglycidyl methacrylate (PGMA), which contains a reactive epoxy unit per GMA
monomer, can be functionalized with dyes, other polymers and drugs using S N2
nucleophilic substitution reactions. Amines, alcohols, carboxylic acids, alkyl halides,
thiols and acid anhydrides are examples of highly reactive species that interact with
PGMA.299
Different solvent systems such as acetone, chloroform and methylethyl ketone
(MEK) were tested in the electrospinning of PGMA. Even though PGMA has high
solubility in these solvents, acetone and chloroform were limited in their use because
of their high vapor pressure and low conductivity. Therefore PGMA was electrospun
in methyl ethyl ketone (MEK) to obtain a nanofibrous scaffold matrix (ES-PGMA).
ES-PGMA was heated at 80 °C overnight to allow inter-crosslinking of the polymer.
Fiber morphology was determined using scanning electron microscopy (SEM). The
48
PGMA-MEK system yielded uniform nanofibers with an average diameter of 0.69 ±
0.04 µm over a large area.
PGMA is a hydrophobic polymer which needs to be functionalized to induce
hydrophilicity to the backbone. This is usually carried out by reacting PGMA
polymer backbone with hydrophilic reacting groups such as carboxylic acids and
amines. ES-PGMA film was end-grafted with carboxylic acid-terminated poly (Nisopropyl acrylamide) (PNIPAM-COOH) to introduce ‘switchability’ in terms of
thermally induced hydrophilic-hydrophobic behavior. PNIPAM is a thermoresponsive polymer which undergoes reversible phase transition in water at 32 °C
(lower critical solution temperature (LCST)). This is caused by the transition from a
swollen hydrated state to a shrunken dehydrated state above the LCST. Surface
properties of the scaffolds were examined by contact angle measurements. End
grafting of ES-PGMA with PNIPAM-COOH (ES-PGMA-g-PNIPAM-COOH
nanofibers) resulted in dramatic temperature dependent changes in the contact angle,
confirming the surfacing grafting of the hydrophilic PNIPAM moiety.
Polymer composites comprising both polymers and nanomaterials have been
regarded as a new class of hybrid materials with many functions. Lack of
universality however is the major limitation in the development of a ubiquitous
polymeric functional platform.
PGMA was explored as a potential ubiquitous polymeric platform. PGMA was
coelectrospun with three different types of nanoparticles: (NaGdF4:Yb, Er);
palladium (Pd) and magnetite (Fe3O4) to develop polymer nanoparticle composite
materials. These fibrous composite materials were characterized using SEM,
transmission electron microscopy (TEM) and X-ray microanalysis.
Coelectrospun polymer-nanoparticle composites were tested for their respective
applications in upconversion fluorescent imaging in the case of (NaGdF4:Yb, Er),
hydrogen gas sensing in the Pd composite and magneto-responsive behavior in the
case of
Fe3O4
composite matrix. Upconversion properties
of
the
ES-
PGMA/(NaGdF4:Yb, Er) nanocomposite was evaluated using NIR room temperature
emission spectroscopy. This measures the efficiency of the nanocomposites in
converting near infra-red excitation wavelength into visible emission. ESPGMA/(NaGdF4:Yb,
Er)
nanocomposites
49
showed
prominent
upconversion
properties with three major emission peaks observed in the visible region at 521,
541, and 655 nm.
Pd has been shown to demonstrate high selectivity and significant adsorption of
hydrogen gas. Because of this feature Pd has been used in hydrogen gas sensors. Pd
readily absorbs hydrogen gas which diffuses as atomic hydrogen into the lattice to
form palladium hydride, PdHx, resulting in  to  phase transition and a
corresponding change in the lattice spacing. Such changes in phase and lattice
spacing cause a measurable resistance change in Pd material. Mostly, Pd has been
adsorbed onto functionalized electrospun scaffolds. In this case, however, Pd was
coelectrospun with PGMA to yield the fibrous composite material. ES-PGMA/Pd
composites were evaluated for their hydrogen gas sensing capacity using alternating
concentrations of nitrogen and hydrogen gas. Change in current was measured using
current-voltage (I-V) sweeps. Response time 90 of ~14 seconds was observed for a
hydrogen gas concentration range of 1-10% from 0.16 ng of Pd as ES-PGMA/Pd
fibrous composite mounted on 650 x 900 µm IED platform.
More recently magnetic materials and matrices were developed for potential
applications in monetary currency and defense fabric materials. ES-PGMA/Fe3O4
nanocomposite fibers were evaluated for magnetization properties using SQUID
magnetometry (Superconducting Quantum Interference Device). SQUID analysis
confirmed superparamagnetic behavior of the composite material at room
temperature. The mass specific saturation magnetization (Ms) of the fibers was 4.0
emu g−1.
Results are presented in Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D.,
Gong, P., Ye, W., Singh, R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre,
T., Lorenser, D., Wood, F. W., Fear, M., Sampson, D. D., Zdyrko, B., Smith, N.M.,
Luzinov, I., Iyer, K.S., A Functional Reactive Polymer Nanofiber Matrix, RSC
Advances (Submitted)
50
2.2 Evaluation of mannose-6-phosphate analogues as potential antiscarring agents
M6P has been shown to control and inhibit the secretion of extracellular matrix
proteins including collagen I, one of the key proteins in granulation tissue and
scars.144, 145, 300 M6P inhibits TGFβ1 activation, a primary cytokine regulating wound
healing and subsequent scar formation, by competitively binding to M6P/IGFII
receptor.301 Recombinant M6P (Juvidex®) has been examined in phase II
randomised human clinical trials for its efficacy in wound healing and scar reduction
at split thickness skin graft donor sites.153 Intradermal delivery of Juvidex showed an
accelerated wound healing response. However, no significant reduction was
observed in scar formation post healing. It has been postulated that the metabolic
vulnerability and delivery mechanism for M6P could have been the limiting factors
behind its translation into the clinic.
Two new bioisosteric phosphonate analogues of M6P were investigated (Aim 2,
analogues were kindly provided by Pharmaxis Pvt Ltd and used without any further
modification). Phosphonate drugs were selected because of their inherently higher
stability towards phosphatase enzymes.159 Analogue 1 was designed by replacing the
P-O bond at C6 with the methylene bridge at C6 in M6P to enhance its stability
against phosphatases present in serum. In addition, the anomeric hydroxyl was
substituted with m-xylene group as they have been reported to improve their
recognition by the receptor.150 Analogue 2 was designed as a prodrug by further
derivatization of analogue 1 with bis(pivaloyloxy)methyl (POM) linkers to yield a
neutral product. Neutral analogues are interesting because they get rapidly
internalized into the cell, where the POM linkers are gradually hydrolysed by
microsomal esterases to release an active analogue 1. Hence analogue 2 was
developed to achieve intracellular targeting of the M6P/IGFII receptor.
In vitro behaviour of the analogues was assessed in primary human dermal skin
fibroblasts (HDF) and compared against M6P. Although binding to the M6P/IGFII
receptor is one of the key mechanisms behind LTGFβ activation, there are other
alternate modes of activation of LTGFβ in vivo. Therefore, it is also important to
investigate the efficacy of the new analogues in the presence of recombinant active
51
human TGFβ1. To assess the dose response of the analogues, HDF cells were
incubated with varying concentrations of the analogues both in the presence and
absence of TGFβ1.
The MTS assay was used to assess the cytotoxicity of the analogues. Cells were
incubated with the analogues, both in presence and absence of TGFβ1, over a period
of 72h. At the specified time, MTS reagent was added to the culture and further
incubated to allow metabolically active cells to bioreduce the MTS reagent and form
the coloured formazan product which is measured by spectrometry at 490 nm.
Colour intensity is correlated to the number of metabolically active cells in culture
i.e. proliferating cells. TGFβ1 has been reported to curtail the proliferation of HDF as
was observed here. However, no add on effect of the analogues was observed at any
of the concentrations studied. Both analogue 1 and 2 behaved similarly to M6P in
terms of inherent toxicity. Analogue 2 however, showed toxicity at high
concentrations (> 50 µM) due to the formation of formaldehyde, which is one of the
degradation products produced as a result of esterase hydrolysis of POM linkers in
analogue 2.
A complimentary live/dead assay was carried out to study the effects of the
analogues on cell viability over a period of 72 h. Cells were incubated with varying
concentrations of analogues. At the specified time-point calcein AM/ ethidium
bromide I dyes were added. Calcein stains the intracellular esterase in viable cells
and fluoresces green whereas ethidium bromide stains non-viable cells and
fluoresces red due to the binding of the ethidium homodimer with nucleic acids
penetrated by ethidium through the compromised cell membranes. No loss was
observed in cell viability.
Cell morphology was also assessed. TGFβ1 stimulation significantly inflated the cell
body area compared to the non-treated control cells which was similar to what has
been reported previously. Interestingly, this inflation in cell area came down to
normal levels post analogue treatment. No apparent alteration in cell body area was
observed in cells treated with varying concentrations of analogue alone without
TGFβ1 stimulation.
The potential of the analogues as anti-scarring agents was evaluated in terms of their
ability to inhibit the overexpression of key fibrotic indicators such as collagen I gene
52
expression. Collagen I is a key fibrotic marker which has major implications in
wound healing and formation of granulation tissue. In order to curb the secretion of
collagen matrix it is crucial to control its expression at the transcriptional level.
TGFβ1 is known to up-regulate the expression of collagen I in HDF.
Real time-quantitative polymerase chain reaction (RT-qPCR) analysis was carried
out to quantitate the expression of collagen I post-treatment. TGFβ1 stimulation has
been reported to significantly up-regulate the expression of collagen I gene
expression as was observed here compared to untreated control cells. Efficiency of
the analogues was tested in terms of their potential in inhibiting this increase in
collagen I mRNA. 10 µM analogue 2 significantly reduced the TGFβ1 mediated
increase in collagen I gene expression. This response was similar to that observed
with M6P treatment. The hydrophilic analogue 1 did not show any response.
It can be deduced that analogues behave similar to M6P. Further, intracellular
targeting of the M6P/IGFII receptor, using analogue 2 as a prodrug of analogue 1, is
more efficacious compared to intercellular targeting using analogue 1.
Results are presented in Agarwal, V., Toshniwal, P., Smith, N. E., Smith, N. M., Li,
B., Clemons, T. D., Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B.,
and Iyer, K. S., Enhancing the Efficacy of Cation-Independent Mannose 6Phosphate Receptor Inhibitors by Intracellular Delivery, Angewandte Chemie
International Edition (Submitted)
2.3 Delivery of the lipophilic mannose-6-phosphate analogue PXS64
using an electrospun PGMA scaffold
One of the biggest limitations in the clinical translation of lipophilic drugs has been
their delivery and associated serum stability. Rapid degradation, particularly in a
wound environment, is a common issue. In order to achieve prolonged stability,
drugs are encapsulated in polymer shells to protect them. One of the niche ways to
achieve controlled delivery of sensitive lipophilic drugs is by encapsulating them in
nanofibrous scaffolds developed using electrospinning.302 Drugs get released
53
through a diffusion mechanism in a controlled rate over time sufficient to be taken
up by the cells before degradation.303
Analogue 2 was coelectrospun with PGMA to fabricate a nanofibrous scaffold (Aim
3). Despite proven toxicity at high concentrations, analogue 2 was selected for its
potential in reducing the expression of fibrotic marker, collagen I. In addition, it was
rationaled that the release of the drug would be slow and within the tolerated limit.
The electrospun scaffold was characterised using scanning electron microscopy
(SEM) and drug loading was estimated using HPLC. Release studies were also
attempted using HPLC. It is important to conduct release studies in aqueous
conditions to mimic the in vivo environment. No drug release was detected from the
fibers in the HPLC analysis. It was implicated that either there is no drug in the
fibers or it may have been degraded during the electrospinning process.
Alternatively, it was not being released.
HPLC confirmed the loading of intact drug at ~76 %. Cell-matrix interactions were
assessed by fluorescent microscopy and SEM. In vitro studies were carried out on
human dermal fibroblasts (HDF) using electrospun PGMA scaffold as a control
matrix. It was hypothesised that cells may be more sensitive to the released drug and
that despite the lack of release data obtained there would be a detectable
physiological effect of the drug even at very low concentrations.
Cytotoxicity studies using the MTS assay showed no apparent toxicity of either
control or test scaffolds. Interestingly, no reduction was observed in cell
proliferation incubated on the scaffolds post TGFβ1 stimulation. Cell viability study
using calcein AM/ethidium bromide I dyes confirmed cell viability. Both florescent
and electron microscopy revealed positive interactions between cells and the fibrous
scaffolds. Cells were seen to take their extended spindle shape morphology as
confirmed by SEM images. RT-qPCR studies were carried out to study the effect of
the scaffolds on upregulated expression of collagen I post TGFβ1 stimulation. It was
observed that cells cultured on scaffolds containing drug (ES-PGMA + analogue 2)
showed a tendency for a reduced increase in expression of collagen I after TGFβ1
stimulation. However, this was not statistically significant. Therefore, it is likely that
the drug is restricted within the fibers and is not able to be released. It has previously
54
been reported that DMSO can facilitate drug release. It facilitates drug release by
providing a congenial environment for drug dissolution in an otherwise hydrophilic
physiological environment. When DMSO was incorporated in the culture, a
significant reduction in collagen I gene expression was attained as anticipated. No
significant effect was observed of DMSO alone or DMSO and control scaffold.
Results are presented in Agarwal, V., Wood, F. M., Fear, M. and Iyer, K. S.,
Inhibiting the activation of transforming growth factor-β using a polymeric
nanofiber scaffold, Nanoscale (Submitted)
55
Chapter 3
Series of papers
The results of this thesis are presented in the following series of published papers or
submitted manuscripts, the citations of which are listed below. Supporting
information for these manuscripts can be found in Appendix A, where applicable.
1.
Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D., Gong, P., Ye, W.,
Singh, R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre, T., Wood, F.M.,
Fear, M., Lorenser, D., Sampson, D. D., Zdyrko, B., Smith, N.M., Luzinov, I., Iyer,
K.S., A Functional Reactive Polymer Nanofiber Matrix, RSC Advances (Submitted)
2.
Agarwal, V., Toshniwal, P., Smith, N. E., Smith, N. M., Li, B., Clemons, T.
D., Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B., and Iyer, K. S.,
Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor
Inhibitors by Intracellular Delivery, Angewandte Chemie International Edition
(Submitted)
3.
Agarwal, V., Wood, F. M., Fear, M. and Iyer, K. S., Inhibiting the activation
of transforming growth factor-β using a polymeric nanofiber scaffold, Nanoscale
(Submitted)
56
Journal Name
RSCPublishing
COMMUNICATION
A Functional Reactive Polymer Nanofiber Matrix
Cite this: DOI: 10.1039/x0xx00000x
Received 00th January 2012,
Accepted 00th January 2012
Vipul Agarwal,a Dominic Ho,a Diwei Ho,a Yuriy Galabura,b Faizah M.D. Yasin,a
Peijun Gong,c Weike Ye,d Ruhani Singh,a Alaa Munshi,a Martin Saunders,e Robert C.
Woodward,f Timothy St. Pierre,f Fiona M. Wood,g Mark Fear,g Dirk Lorenser,c David
D. Sampson,c,e Bogdan Zdyrkob, Nicole M. Smith,a Igor Luzinovb,*, and K.
Swaminathan Iyera,*
DOI: 10.1039/x0xx00000x
www.rsc.org/
Synthetic fractal materials have been regarded as a new class
of hybrid materials with many potential applications.
However, the lack of an efficient, reactive large-area fractal
substrate has been one of the major limitations in the
development of these materials as advanced functional
platforms. Herein, we demonstrate the utility of electrospun
polyglycidyl methacrylate (PGMA) fractal-like films as a
highly versatile platform for the development of functional
nanostructured fractal-like materials anchored to a surface.
The utility of this platform as a reactive substrate is
demonstrated by grafting poly (N-isopropyl acrylamide) to
incorporate stimuli-responsive properties. Additionally, we
demonstrate that functional fractal-like nanocomposites can
be fabricated using this platform with properties for sensing,
fluorescence imaging and magneto-responsiveness.
The development of nanostructured polymeric matrices to
obtain organic-inorganic nanocomposites has been actively
researched to produce hybrid materials for applications in
electronics, optics, medical devices, sensors and catalysis. 1-4 Of
the various techniques developed to produce large area
nanoscale polymeric matrices, one of the most researched, cost
effective and facile method is electrospinning. It has been
adapted to cover a wide range of polymers and optimized to
regulate fiber diameter, alignment and shape. 5-7 There have
been numerous reports using this technique to develop matrices
with enhanced mechanical strength, 8 matrices with selective
filtration/permeability, 9 fire retarding material, optoelectronic
devices10 and substrates for catalysis. 11, 12 One of the key steps
involved
in
the
development
of
organic-inorganic
nanocomposites is grafting to achieve excellent integration by
minimizing interfacial tension of the nanoparticles in the
This journal is © The Royal Society of Chemistry 2012
organic nanofiber matrix. Additionally, the ability to modify the
surface of the nanofiber to alter the adhesion, lubrication,
wettability and biocompatibility is pivotal in its customisation
for end-use applications. The achievement of a certain degree
of grafting universality requires the establishment of a
controlled method of introducing the desired functional groups
on a substrate.5 Currently, this is achieved by physisorption. 13,
14
In contrast, chemisorption, which is difficult to achieve on a
polymeric nanofibers, would result in permanent irreversible
surface modification. Polymers containing epoxy groups are
examples of functional polymers that are able to react with a
wide range of substrates through ‘‘grafting to’’ interactions
mediated by the epoxy groups.15, 16 The versatile chemistry of
epoxy groups renders a polymer that is exceptionally suitable as
a universal electrospun nanofiber matrix to provide reactive
groups for further grafting reactions. To this end, poly(glycidyl
methacrylate) (PGMA), which contains an epoxy group in
every repeating unit, has been used extensively as a
macromolecular anchoring layer for grafting of polymers to the
surfaces.17-20 Upon electrospinning, epoxy groups in the
polymer will undergo self-crosslinking upon heating, providing
mechanical integrity to the matrix. 21 Approximately 40% of the
epoxy groups are still available for surface modification
following a 12 hour treatment at 120°C.
The main advantage of using PGMA as a matrix for
electrospinning, as opposed to modifying the surface using
monolayers, is the high mobility of the epoxy groups located in
the “loops” and “tails” of the polymer. The mobility of the free
groups results in the formation of a highly effective
interpenetrating anchoring zone. 22 In this article, we report that
PGMA can be directly electrospun (ES-PGMA) to form large
area nanofibers. We demonstrate that this polymer nanofiber
matrix can be used as an effective platform to graft polymers to
impart switchability, and can be used to produce
J. Name., 2012, 00, 1-3 | 1
COMMUNICATION
nanocomposites with upconverting properties, with hydrogen
sensing capability or with magneto-responsive properties.
The ES-PGMA nanofibers (see Supporting Information for
method of synthesis) were uniform over a large area and had an
average diameter of 0.69 ± 0.04 µm (average ± standard error
mean) (Fig. 1A). The average thickness of the 1 x 1 cm 2 ESPGMA generated was 127 ± 3 µm in 7 hours (Fig. 1B). In order
to test the efficacy of the ES-PGMA nanofiber matrix as an
anchoring platform, carboxylic acid-terminated poly (Nisopropyl acrylamide) (PNIPAM-COOH) was end grafted to
the nanofibers via a ring opening reaction with the epoxy
groups to yield ES-PGMA-g-PNIPAM-COOH.23 PNIPAM is a
thermo-responsive polymer, which has been utilized in various
Fig. 1: (A) SEM secondary electron image of the electrospun PGMA (ESPGMA) fibers, (B) cross-sectional image of ES-PGMA, (C) Water contact
angle θ = 60° at 70°C, (D) Water contact angle θ = 15° at room temperature
respectively measured on ES-PGMA -g-PNIPAM-COOH.
forms, such as thermo-responsive hydrogels, particles, brushes,
spheres and micelles. 24-27 Importantly, PNIPAM exhibits a
2 | J. Name., 2012, 00, 1-3
Journal Name
temperature-sensitive phase transition in water at what is
known as a lower critical solution temperature (LCST), 32 °C. 28
The transition is due to the coil-to-globule transition at the
critical temperature resulting in switching from hydrophilic to
hydrophobic behavior.29 At temperatures below the LCST,
PNIPAM chains arrange into an expanded and hydrated
conformation. Conversely, at temperatures above the LCST,
PNIPAM chains collapse and arrange into a compact,
dehydrated conformation. 29 This thermo-responsive behavior is
retained post-grafting and post-end group functionalization.
The ES-PGMA-g-PNIPAM-COOH nanofibers demonstrated
thermo-responsive behavior as monitored using contact angle
measurements. The contact angle changed from 60 ± 2° at 70°C
(Fig. 1C) to 15 ± 2° at room temperature (Fig. 1D). The contact
angle of the unmodified ES-PGMA remained unchanged at 100
± 2° both at room temperature and at 70°C. The ability of the
ES-PGMA nanofiber matrix to produce nanocomposites was
further evaluated using three distinct classes of nanoparticles:
upconverting fluorescent particles of NaGdF4:Yb, Er (UCNP),
palladium (Pd) and magnetite (Fe3O4). The nanoparticles
synthesized (see Supporting Information for methods) had a
narrow size distribution of 7.4 ± 1.4 nm (average ± standard
error mean) for UCNP, 19.3 ± 0.2 nm for Pd and 6.7 ± 1.4 nm
for Fe3O4 respectively (Fig. 2A, C and E). One of the major
hurdles in developing functional materials by electrospinning
nanocomposites is the lack of control in attaining a
homogeneous distribution of nanoparticles throughout the
polymer matrix. In the present case, electrospinning PGMA
with the aforementioned nanoparticles resulted in relatively
uniform distributions of the nanoparticles throughout the fiber
matrix (Fig. 2B, D and F) which was observed through various
images obtained at similar fields of view. It has been reported
that variations in solution properties such as surface tension and
solution conductivity in the presence of nanoparticles result in
changes in the nanofiber diameter. 30-32 In the present case, the
electrospun fiber diameter increased in the presence of
nanoparticles to 2.56 ± 0.16 µm (average ± standard error
mean) for UCNP, 1.75 ± 0.07 µm for Pd and 4.37 ± 0.44 µm
for Fe3O4 (Insets in Fig. 2G, H and I, respectively). Small fibers
were chosen for TEM analysis because of the contrast problems
related to thicker samples. The ability of the nanocomposites to
be used as functional materials was evaluated by testing the
upconverting properties, hydrogen sensing properties and
magnetic properties of the, UCNP/ES-PGMA, Pd/ES-PGMA
and Fe3O4/ES-PGMA fibers, respectively.
In the case of UCNP/ES-PGMA fibers, the ability to
convert near-infrared excitation into visible emission was
evaluated (see supporting information for methods). UCNP
have
been
successfully
used
as
ultrasensitive
magnetic/upconversion fluorescent dual-modal molecular
probes for MRI and upconversion fluorescence imaging. 33-35 In
the present case the UCNP/ES-PGMA fibers demonstrated
excellent upconversion properties upon 974 nm laser excitation
(Fig. 3A). The three major emissions were located at 521, 541,
and 655 nm. Green emission from 500- 600 nm was attributed
to 2H11/2 → 4I15/2 and 4S3/2 → 4I15/2 transitions, respectively, and
the red emission from 635-670 nm was attributed to the 4F9/2 →
4
I15/2 transition (Fig. 3A). The green to red (G/R) ratio for the
fibers was 1.35:1. The UCNP/ES-PGMA composite fibers
retained the upconversion signal levels to the pure
NaGdF4:Yb,Er nanoparticle samples (G/R ratio = 1.37:1).
This journal is © The Royal Society of Chemistry 2012
Journal Name
Fig. 2: TEM images of the (A) Upconverting nanoparticles (UCNP), (B)
UCNP/ES-PGMA composite fibers, (C) Pd nanoparticles, (D) Pd/ES-PGMA
composite fibers, (E) magnetite (Fe3O4) nanoparticles, (F) Fe3O4/ES-PGMA
composite fibers. X-ray microanalysis spectrum obtained on: (G) UCNP/ESPGMA composite fibers showing the presence of Gd and Yb among other
elements (Inset: SEM micrograph of UCNP/ES-PGMA composite fibers):
(H) Pd/ES-PGMA composite fibers showing the presence of Pd (Inset: SEM
micrograph of Pd/ES-PGMA composite fibers) and (I) Fe3O4/ES-PGMA
composite fibers (Inset SEM micrographs of Fe3O4/ES-PGMA composite
fibers). Scale bars for images A, C and E 10 nm, for images B, D and F 1
µm and for inset images G, H and I 20 µm
The Pd/ES-PGMA nanofibers were evaluated for sensing
hydrogen (see Supporting Information for methods). Palladium
has emerged as an important candidate for hydrogen gas
sensing because of its ability to absorb high quantities of
hydrogen and its highly selective response. 36 Sensing herein is
based on the well-established principle that palladium
spontaneously absorbs H 2 gas as atomic hydrogen which
diffuses into the lattice to form palladium hydride, PdH x,
resulting in an  to  phase transition and a corresponding
change in the lattice spacing. 37 The change in phase and lattice
spacing leads to a measurable resistance change of the
palladium material. However, either replacing precious noble
metals with cheaper materials or alternatively development of
methods that result in the reduction of material used by several
orders of magnitude, especially in applications that require
large amounts of material, would be beneficial. Currently,
hydrogen sensing platforms are based on all-palladium
constructs or hybrids with high Pd loading to stimulate an
effective sensing response. Herein, using the electrospun
polymer/nanoparticle nanocomposite material we demonstrate a
This journal is © The Royal Society of Chemistry 2012
COMMUNICATION
Fig. 3: (A) Upconversion fluorescence spectrum of both UCNP’s (blue) and
UCNP/ES-PGMA fiber composite (black) showing three main emissions
green at 521 and 541 nm and red between 635 and 670 nm upon 974 nm laser
excitation, (B) Current response of the Pd/ES-PGMA matrix sensor to 1-10%
hydrogen gas, with alternating 4 min hydrogen and 20 min. nitrogen
exposure, (C) Zero-field cooled (orange) and field cooled (blue) curves for
Fe3O4/ES-PGMA composite (Inset: hysteresis loop at 5 K (pink) and 300 K
(green) for Fe3O4/ES-PGMA composite) as measured by SQUID
magnetometry.
response is obtainable for as low as 0.6 ng of Pd dispersed
across a 650 m x 900 m area over interdigitated electrodes
(IDE). The ability of Pd/ES-PGMA nanofibers to sense
different hydrogen concentrations (between 1 and 10% in N 2 as
a carrier gas) was tested 38 (Fig. 3B). An increase in resistance
with hydrogen gas-flow and a return to the original state in the
absence of a hydrogen gas flow was observed for hydrogen
concentrations (1-10%) with a response time 90 of ~14 seconds
(Fig. 3B).
Finally, the magnetization properties of the Fe 3O4/ESPGMA fibers were measured by SQUID magnetometry
J. Name., 2012, 00, 1-3 | 3
COMMUNICATION
(Superconducting Quantum Interference Device) (see
Supporting Information). The Fe3O4/ES-PGMA fibers are
superparamagnetic at room temperature with the zero field
cooled/field cooled curves showing a maximum blocking
temperature of 30 K (where the two curves merge) and the
absence of hysteresis at 300K (Fig. 3C). The mass specific
saturation magnetization, M s of the fibers was 4.0 emu g−1.
Particle loading was estimated to be ~7% by weight as
determined from the M s values of the Fe3O4 nanoparticles and
Fe3O4/ES-PGMA fibers.
Conclusions
In summary, we have developed a robust polymeric platform
for the large scale production of electrospun nanofibers based
on poly(glycidyl methacrylate) (PGMA). We have
demonstrated that the epoxy groups of the polymeric matrix can
be effectively used as a grafting platform for surface
modifications and the polymer serves as an excellent platform
to fabricate functional nanocomposites. We believe our findings
presented herein will aid in the design of novel electrospun
materials with tailorable surfaces for applications as scaffolds
in regenerative medicine, optoelectronics, magnetic filtration
and catalysis.
Notes and references
a
School of Chemistry and Biochemistry, The University of Western
Australia, WA 6009, Australia
b
Department of Materials Science and Engineering, Clemson University,
Clemson, SC 29634, USA
c
School of Electrical, Electronic and Computer Engineering, The
University of Western Australia, WA 6009 Australia
d
School of Chemistry and Chemical Engineering, Nanjing University,
China
e
Centre for Microscopy, Characterisation and Analysis, The University
of Western Australia, WA 6009 Australia
f
School of Physics, The University of Western Australia, WA 6009
Australia
g
Burn Injury Research Unit, School of Surgery, The University of
Western Australia, WA 6009, Australia
E-mail: swaminatha.iyer@uwa.edu.au, luzinov@clemson.edu
The authors would like to acknowledge the Australian Microscopy &
Microanalysis Research Facility at the Centre for Microscopy,
Characterization & Analysis, The University of Western Australia,
funded by the University, State and Commonwealth Governments.
Authors would also like to thank Dr C.W.Evans and Dr Peter R. T.
Munro for assistance with valuable experimental and result discussions.
Peijun Gong is supported by The University of Western Australia and the
China Scholarship Council.
† Footnotes should appear here. These might include comments
relevant to but not central to the matter under discussion, limited
experimental and spectral data, and crystallographic data.
Electronic Supplementary Information (ESI) available: [details of
materials and methods]. See DOI: 10.1039/c000000x/
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This journal is © The Royal Society of Chemistry 2012
J. Name., 2012, 00, 1-3 | 5
Intracellular Drug
Delivery
DOI: 10.1002/anie.201((will be filled in by the editorial staff))
Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate
Receptor Inhibitors by Intracellular Delivery**
Vipul Agarwal, Priyanka Toshniwal,# Natalie E Smith,# Nicole M Smith, Binbin Li, Tristan D
Clemons, Lindsay T Byrne, Foteini Hassiotou, Fiona M Wood, Mark Fear, Ben Corry* and K
Swaminathan Iyer*
Abstract: Extracellular targeting of the cation-independent mannose
6-phosphate/insulin-like growth factor II (M6P/IGFII) receptors has
been an attractive approach for the development of antifibrotic
drugs. Several M6P analogues have been developed to regulate the
activity of transforming growth factor-β1 (TGFβ1) by inhibiting its
conversion from the latent to the active form. Herein, we adopt a
combinatorial approach using an in vitro wound healing model and
molecular dynamic simulations, to reveal that the efficacy of
M6P/IGFII inhibitors can be significantly enhanced by adopting an
intracellular approach. We demonstate this using systematic
analysis of a bioisosteric M6P analogue, a lipophilic prodrug which
upon cellular internalization undergoes ester hydrolysis to yield an
[]
V. Agarwal, P. Toshniwal, Dr. N.M. Smith, Dr. T.D. Clemons, Dr. F.
Hassiotou, Prof. Dr. K. S. Iyer
School of Chemistry and Biochemistry, The University of Western
Australia, Crawley, Western Australia, Australia
E-mail: swaminatha.iyer@uwa.edu.au
B. Li
State Key Laboratory of Advanced Technology for Materials
Synthesis and Processing, Wuhan University of Technology,
Wuhan, PR China
Dr. L. T. Byrne
Centre for Microscopy, Characterization & Analysis, The University
of Western Australia, Australia
Dr. N. E. Smith, Dr. B. Corry
Research School of Biology, Australian National University,
Canberra, Australian Capital Territory, Australia
Email: ben.corry@anu.edu.au
Dr. M. Fear, Prof. Dr. F. M. Wood
Burn Injury Research Unit, School of Surgery, The University of
Western Australia, Crawley, Western Australia, Australia
[]
The authors would like to thank the Australian Research Council for
funding the work, Pharmaxis Pvt Ltd for kindly providing analogues
1 and 2. The authors would also like to acknowledge the Australian
Microscopy & Microanalysis Research Facility at the Centre for
Microscopy, Characterization & Analysis, The University of Western
Australia, funded by the University, State and Commonwealth
Governments. Authors would like to thank Dr Bernard Callus, Dr
Cameron Evans and Dr Megan Finch for their assistance with
experimental analysis.
[#]
These authors contributed equally
active M6P analogue to effectively downregulate collagen 1
expression in primary human dermal skin fibroblasts.
In mammalian cells, the cation-independent mannose 6phosphate/insulin-like growth factor II (M6P/IGFII) and cationdependent mannose 6-phosphate (CD-MPR) receptors, have been
identified as pivotal targets that modulate cellular response because
of their role in protein trafficking. Both these receptors are
functionally complimentary and can partially compensate for the
absence of the other.[1] These sorting receptors play an important
role of transporting M6P-bearing glycoproteins from the trans-Golgi
network (TGN) to lysosomes mediated through their M6P binding
sites.[2] Both receptors transport important enzymes to the
intracellular acidic pre-lysosomal compartments where low pH leads
to the release of the enzymes from the complex. The receptor then
gets recycled into the Golgi apparatus.[3] However, only the
M6P/IGFII receptor is anchored to the cell surface membrane and
has been implicated in the internalization of M6P bearing
compounds.[4] Importantly, it modulates the activity of a variety of
extracellular M6P bearing glycoproteins including latent
transforming growth factor-β (LTGFβ) precursor, urokinase-type
plasminogen activator receptor, glycoprotein D of the herpes virus,
granzyme B an essential factor for T cell-mediated apoptosis and
proliferin.[4] This has resulted in an enormous interest in the design
of M6P bearing compounds that target the M6P/IGFII receptor as it
offers an efficient means for internalization of high specificity
therapeutics.[5] This approach has been used to deliver therapeutic
compounds in enzyme replacement therapies in lysosomal diseases
like Fabry disease, aid wound healing, as a treatment for breast
cancer, and to combat viral infections.[4] However, the approach
suffers a major drawback as the phosphomonoester bond of M6P is
prone to hydrolysis by various phosphatase enzymes.[6] This
dramatically reduces its binding efficiency to the receptor thereby
compromising its potency. This problem has been circumvented by
Supporting information for this article is available on the WWW
under http://dx.doi.org/10.1002/anie.201xxxxxx.
Figure 1: a) Chemical structure of mannose-6-phosphate (M6P and the two
analogues, b) Schematic representation of the cLogP of three compounds
1
the design of several isosteric M6P analogues with phosphonate,
carboxylate or malonate groups, which have higher affinity to the
receptor and a stronger stability in human serum than M6P.[7] This
approach is successful in overcoming the issues with hydrolysis of
the phosphomonoester bond, yet falls short as these analogues can
only target the receptors present on the cell surface. In the steady
state, ~90 % of the M6P/IGFII receptors are localized in the
transmembrane compartments while the remainder stays on the cell
surface.[8] The receptor has a relatively long half-life (t1/2 ~ 20
hours) and recycles between the trans-Golgi network, endosomes
and the plasma membrane.[9] In this communication, we report a
novel approach to improve ligand-receptor protein interaction in
cells whilst overcoming stability issues associated with M6P. We
demonstrate this by exploring a prodrug (analogue 2) that undergoes
intracellular chemical modification by esterases to yield an active
M6P analogue (analogue 1) (Figure 1a),[10] resulting in a sustained
and focused therapeutic strategy in an in vitro model of wound
healing.
Table 1. Ligand-Receptor Protein interaction energies obtained for M6P and
each of the two analogues in domain 3 and domain 5 as determined from 100
ns of molecular dynamics simulation. Two ligands were placed into the dimer
binding pocket, because the receptor is secreted as a dimer.
Domain 3 Ligand-Receptor Protein interaction Energy (kcal/mol)
M6P
Analogue 1
Analogue 2
Ligand 1
-368.4
-309.3
-81.6
Ligand 2
-347.4
-304.3
-79.6
Domain 5 Ligand-Receptor Protein interaction Energy (kcal/mol)
M6P
Analogue 1
Analogue 2
Ligand 1
-128.2
-44.1
-43.6
Ligand 2
-118.3
-74.0
-47.8
The design of phosphonate analogue 1 is based on established
principles of bioisosteric M6P analogues by replacing the P-O bond
at C6 by a methylene bridge. Moreover, the replacement of the
hydroxyl group at the anomeric position by an aromatic subtituent
slightly improves recognition by the M6P/IGFII receptor.[11] This
could be due to the hydrophobic interactions between the aromatic
moiety of analogue 1 and the binding pocket of the M6P/IGFII
receptor. Previous studies have demonstrated that neutral ester
prodrugs are relatively benign towards enzymatic degradation,
thereby altering their apparent elimination and half-life.[12] Hence
analogue 2 was designed by masking analogue 1 via esterification of
the phosphate group to yield a non-charged bis(pivaloyloxymethyl)
(POM) derivative. Importantly, derivatization of phosphates
decreases the polarity of the parent drug thereby promoting its
cellular internalization and altering the elimination/distribution
mechanism.[12-13] Notably, the clogP (calculated-logP evaluated
using Chem Draw) for M6P, analogue 1 and 2 are -3.28, 0.10 and
3.29 respectively (Figure 1b). LogP is an estimate of a compound's
overall lipophilicity, a value that influences its physiological
properties such as solubility, permeability through biological
membranes, hepatic clearance, and non-specific toxicity.[14] Polar
compounds with low logP have very low cellular permeability due
to their low affinity for the lipid bilayers. Alternatively, lipophilic
compounds with high logP have high affinity for the phospholipid
phase facilitating their internalization and prohibiting their escape
into the aqueous basolateral side.[14] Herein the lipophilic prodrug,
Figure 2: Cell viability assays showing percentage of live cells in the culture
post incubation with M6P, analogue 1 and 2. First and second column in each
condition is representing 24 h and 72 h respectively. Data presented as average
± SEM (n=4). Significance was set at * p < 0.05 using bonferroni post-hoc test
in one way ANOVA analysis.
analogue 2, will have improved cellular internalization compared to
its charged parent analogue, 1. Once internalized the
bis(pivaloyloxymethyl) linkers of analogue 2 will be gradually
prone to ester hydrolysis by microsomal esterases present within the
intracellular compartments,[15] resulting in the conversion to the
charged parent analogue, analogue 1. Analogue 1 on the contrary
would only target extracellular M6P/IGFII receptors, when
administered directly, due to its low cellular permeability deemed to
its low logP value.
The extracellular region of the M6P/IGFII receptor is comprised of
15 repetitive domains and contains three distinct M6P binding sites
located in domains 3, 5, and 9, with only domain 5 exhibiting
preference for phosphodiesters.[16] In order to assess our strategy to
use the intracellular conversion of the produg analogue 2 to a high
receptor binding phosphonate analogue 1, it is pivotal to examine
the ligand-receptor interactions to validate the hypothesis that
analogue 2 will have minimal interaction with the extracellular
receptors. In the current study, we used six independent molecular
dynamics simulations to study the ligand-receptor protein
interactions of M6P, analogues 1 and 2 with domains 3 and 5 of the
extracellular M6P/IGFII receptor (see Supporting Information for
experimental details, section S8.1). Domain structures were adopted
from previously reported studies and two ligands were placed into
the dimer binding pocket, because the receptor is secreted as a
dimer.[17] Analogue 1 showed similar ligand-receptor protein
interaction energies to M6P in domain 3 (Table 1). Importantly, the
m-xylene ring of analogue 1 was positioned in the middle of the
binding pocket further stabilizing the binding of this compound in
comparison to M6P (see Supporting Information, Figures S1 and
S2). This is in accordance with the previous studies of other
phosphonate analogues of M6P, which are reported to display higher
affinity and stronger stability in human serum than M6P.[6, 7c] The
domain 5 binding pocket is larger than in domain 3, hence all the
Figure 3: Cell body area showing change in cell area post TGFβ1 stimulation
and subsequent analogues treatment. Cell area was measured from the
fluorescent images of live cells taken for viability assay (cells from minimum 40
images per group were measured). Significant increase in cell body area was
observed for cells treated with TGFβ1 (2 ng/mL), however no such increase was
observed in cells treated with analogues +/- TGFβ1 (2 ng/mL). Data presented
as average ± SEM (n > 40). Significance was set at * p < 0.05 using bonferroni
test in one way ANOVA.
2
compounds displayed weaker interactions with the receptor and
occupied more diverse positions in domain 5 due to the increased
space (see Supporting Information, Figure S2). Furthermore, in the
case of analogue 1 in domain 5, the simulations suggested that one
of the two analogue 1 ligands (ligand 1) bound to the protein dimer
has weaker interactions with the protein as it primarily interacts with
the second molecule of analogue 1 (ligand 2). Overall, the
simulations suggested that analogue 1 has high affinity towards
domain 3 similar to M6P whilst the prodrug 2 has weak interactions
with both domains of the receptor (Table 1 and see Supporting
Information, Figure S3). The molecular dynamics simulations
further validated our aforementioned hypothesis that the prodrug
will be internalized with minimal extracellular receptor-ligand
interactions.
We next validated our hypothesis in a well-established in vitro
model for wound healing using primary human dermal skin
fibroblasts (HDF). In mammals, wound healing is not a regenerative
process that restores normal tissue architecture, but a reparative
process that results in scar formation.[18] This process occurs in all
tissues of the body in response to physical, chemical and biological
stressors. Scar tissue is functionally and aesthetically inferior to
normal tissue. It is a result of the excessive production of
extracellular matrix (ECM) that occurs after injury.[19] One of the
most important proteins influencing the ECM architecture during
wound healing is collagen I. Collagen I is synthesized
predominantly by fibroblasts and its synthesis is largely regulated by
cytokine transforming growth factor β1 (TGFβ1).[20] TGFβ1 is
secreted in an inactive form (LTGFβ1), requiring enzymatic
conversion to active TGFβ1 to effect a change in cell function. One
of the methods of TGFβ1 activation involves binding of M6P
residues within the N-linked oligosaccharides on latent TGFβ1 to the
M6P/IGFII receptor.[21] Since the M6P binding sites are involved in
various steps of TGFβ1 activation and inactivation, it is believed that
small molecule inhibitors that block the binding of M6P residues
could present an opportunity to block the activity of TGFβ thereby
reducing overproduction of an important profibrotic extracellular
matrix protein collagen I. Cytotoxicity and cell viability of the
analogues were initially assessed using MTS and live/dead assays
(see experimental details in Supporting Information, sections S2.1,
S3.1 and S4.1). Previous studies characterizing M6P binding affinity
towards the M6P/IGFII receptor reported significant binding affinity
at a concentration of 10 µM.[7b, 22] This concentration was therefore
selected for our in vitro studies. All compounds showed no effect on
cell viability and proliferation both in the presence and absence of
TGFβ1 (Figure 2 and see Supporting Information, Figure S4
respectively). Exposure to TGFβ1 in the absence of analogues 1 and
2 resulted in a reduction in HDF proliferation (see Supporting
Information, Figure S4). This growth suppressive response has been
previously reported in many cell types.[23] The observed change in
cell proliferation upon exposure to TGFβ1 influenced HDF cell
morphology (and see Supporting Information, Figure S5b).
Fibroblasts alter their morphology from dendritic to stellate upon
exposure to various external cues caused by changes in actin
polarisation and focal adhesion.[24] TGFβ1 has been shown to alter
the morphology of many cell types including fibroblasts, potentially
by inducing polymerisation of the actin cytoskeleton from globular
to filamentous.[24a] Different factors such as cell motility and
mechanical strain have also been reported to cause this alteration.[25]
In the present case, we observed a reversal of HDF cell morphology
back to initial cell morphology without TGFβ1 stimulation when
treated with the analogues 1 and 2 (Figure 3 for quantification of
cell body area and see Supporting Information, Figure S5c-e for
Figure 4: Change in Collagen 1 a) mRNA levels and b) protein levels post
TGFβ1 stimulation in the presence and absennce of M6P, analogues 1 and 2
compared to untreated (negative) control. Collagen I protein expression was
normalised against β-actin levels. Data are presented as average ± SEM (n =
3). Significance was set at * p < 0.05 using bonferroni post-hoc test in one way
ANOVA analysis
images). We next assessed if the observed change in morphology is
correlated to collagen I gene expression using qRT-PCR, and if
changes in collagen I gene expression could be altered by inhibition
of TGFβ1 activity by targeting the M6P/IGFII receptor in the
presence of the analogues (refer to Supporting Information for
method, section S5.1). Indeed as previously reported, exposure of
HDF to TGFβ1 (2 ng/mL) resulted in a significant increase in
collagen I mRNA expression at 48 hours post-stimulation.[20, 26] It is
noteworthy that although collagen I gene expression was
upregulated throughout the study period (72 hours), the optimal
response was observed after 48 hours exposure to TGFβ1 (see
Supporting Information, Figure S6). Therefore, the efficacy of the
aformentioned compounds was assessed in the presence of TGFβ1 at
48 hours. TGFβ1 induced collagen I mRNA expression was
downregulated significantly (p < 0.05) with the addition of prodrug
analogue 2 (10 µM) with levels returning to that of normal untreated
cells (Figure 4a). Downregulation was also observed for M6P
however, the change did not reach stastistical significance (Figure
4a). This suggests that the variable responses that have been
reported in the use of hydrolytically unstable M6P may be due to its
realtive instability and that the development of stable analogues may
resolve this issue. Importantly, in the present case we observed no
significant change in collagen I gene expression in HDF cells treated
with analogue 1 (Figure 4a). This was expected given the low
cellular permeability which is believed to affect the ligand-receptor
protein interactions in cells. Next, we investigated if the observed
change at transcription level would have a corresponding influence
on protein translation. Changes in collagen I protein expression were
quantified using immunoblotting (refer to Supporting Information
for method, section S6.1). All protein expression studies were
carried out at 72 h post-stimulation. Significant upregulation in
collagen I protein expression was observed post TGFβ1 stimulation
(Figure 4b; column 2; p < 0.05) which is consistent with previous
reports.[27] Analogue 2 (10 µM) was observed to reduce TGFβ1
3
[8]
mediated upregulation of collagen I protein to non-stimulated levels
(Figure 4b; column 5; p < 0.05). No significant changes were
observed in the case of M6P or analogue 1. This further confirms
that analogue 2 is a potent repressor of TGFβ1 induced collagen I
synthesis and thus can ameliorate the profibrotic effects of TGFβ1 in
human skin dermal fibroblasts.
In summary, we have developed a novel approach using an
intracellular prodrug of M6P, analogue 2, to target M6P receptors.
This approach overcomes the physiological problems associated
with the hydrolysis of M6P whilst successfully targeting the
receptors using an intracellular coversion of the analogue. We
believe that this approach of intracellular drug coversion for
receptor targeting will have far reaching implications in the design
of highly potent drug candidates for enzyme replacement therapies
of lysosomal storage diseases, to aid wound healing and in cancer
therapy.
[9]
[10]
[11]
[12]
[13]
[14]
[15]
[16]
Received: ((will be filled in by the editorial staff))
Published online on ((will be filled in by the editorial staff))
Keywords: drug development • intracellular drug delivery • Mannose6-Phosphate analogue
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4
Entry for the Table of Contents (Please choose one layout)
Layout 2:
Intracellular Drug Delivery
#
Vipul Agarwal, Priyanka Toshniwal,
#
Natalie E Smith, Nicole M Smith, Binbin
Li, Tristan D Clemons, Lindsay T Byrne,
Foteini Hassiotou, Fiona M Wood, Mark
Fear, Ben Corry* and K Swaminathan
Iyer* __________ Page – Page
Enhancing the Efficacy of CationIndependent Mannose 6-Phosphate
Receptor Inhibitors by Intracellular
Delivery
Intracellular delivery of M6P/IGFII receptor inhibitors exhibits better efficacy than
extracellular inhibitors to regulate TGFβ1 mediated upregulation of profibrotic
marker, collagen I.
5
Journal Name
RSCPublishing
COMMUNICATION
Cite this: DOI: 10.1039/x0xx00000x
Inhibiting the activation of transforming growth
factor-β using a polymeric nanofiber scaffold
Vipul Agarwal,a Fiona M. Wood,b,c Mark Fearb and K. Swaminathan Iyer*a
Received 00th January 2012,
Accepted 00th January 2012
DOI: 10.1039/x0xx00000x
www.rsc.org/
Electrospun poly(glycidyl-methacrylate) (PGMA) nanofibers
were fabricated in the presence of a hydrophobic analogue of
mannose-6-phosphate, (PXS64). The nanofibers were tested
for biocompatibility as a tissue engineering scaffold and for
their efficacy to inhibit transforming growth factor β1
(TGFβ1) activation in human dermal skin fibroblasts.
The principle of tissue engineering involves the development of
advanced functional biomaterials that incorporate biochemical cues
to aid in the regeneration of tissues to restore and maintain normal
organ function.1 This approach has resulted in the fabrication of
several advanced biomaterial platforms that have been used to repair
damaged or diseased tissues and to create therapeutic approaches for
entire tissue replacement.2-4 While tissue regeneration has
successfully been demonstrated in the presence of biocompatible
scaffolds, scarring remains one of the unresolved issues. In
mammals postnatal wound healing results in scar formation,
characterised by excessive collagen deposition and dysfunctional
extracellular matrix formation. This is also a hallmark of fibrotic
disease which occurs in many tissues. Central to wound repair is
transforming growth factor β1 (TGFβ1), a cytokine secreted by
several different cell types involved in wound healing.5 In the case
of skin, scarring during the wound healing process is a result of
TGFβ1 mediated imbalance in the fibroblast activity resulting in an
architecturally disorganised extracellular matrix.6 One of the most
important proteins influencing the ECM architecture during wound
healing is collagen I. Collagen I is synthesized predominantly by
fibroblasts and its synthesis is largely regulated by TGFβ1
signalling.7 Following wound healing in skin, a scar is not only
aesthetically and psychologically detrimental but can also cause
functional disability and pain.8 Importantly, inhibition of TGFβ1
activity has been documented to reduce scar formation9 whereas
subcutaneous delivery of exogenous TGFβ1 in newborn mice was
shown to promote fibrosis and angiogenesis at the site of injection.10
During wound healing TGFβ1 activation from its latent state
involves binding of the mannose 6-phosphate (M6P) residues within
This journal is © The Royal Society of Chemistry 2012
the N-linked oligosaccharides on the latent TGFβ to the M6P/IGFII
receptor.11 Since the M6P binding sites are involved in various steps
of the TGFβ1 activation and inactivation route, it is believed that the
presence of small molecule analogues of M6P that competitively
bind to the receptor could present an opportunity to block the
activation of TGFβ1 thereby reducing overproduction of extracellular
matrix protein collagen and potentially reducing scarring.12 The
major drawback with using M6P as a drug in the reduction of
scarring is that the phosphomonoester bond of M6P is prone to
hydrolysis by various phosphatase enzymes.13 Isosteric M6P
analogues with phosphonate, carboxylate or malonate groups have
been shown to circumvent the aforementioned issues and have been
reported to have greater stability in human serum than M6P.14-17 The
analogue PXS64 reported in the present study is a lipophilic
bioisosteric
phosphonate
analogue
developed
by
[(bis(pivaloyloxymethyl)) (POM)] ester derivatization of M6P.
Importantly, the high lipophilicity of PXS64 limits its solubility in
aqueous solutions thereby reducing its bioavailability.18 In this
communication we report that PXS64 can be incorporated in an
electrospun poly(glycidyl methacrylate) (PGMA) nanofibrous
scaffold. Furthermore, we demonstrate the utility of this scaffold as a
biomaterial platform for wound healing using human dermal skin
fibroblasts (HDF). We demonstrate its effectiveness as a drug
delivery platform to mitigate TGFβ1 mediated upregulation of
collagen I.
Electrospinning is a widely used technique to fabricate large area
nanofibrous scaffolds19 mimicking the architecture of extracellular
matrix (ECM).20 They have been used as tissue engineering scaffolds
to promote cell growth and migration and to achieve controlled
delivery of drugs and growth factors.21 They have been widely used
in skin, bone, cartilage, vascular and neural tissue engineering. 22 For
example, Yang et al. employed emulsion electrospinning to fabricate
ultrafine core sheath poly(ethylene glycol)-poly(D,L-lactide) fibers
loaded with basic fibroblast growth factor (bFGF) and demonstrated
its gradual release over 4 weeks.23 In vitro studies on mouse
embryonic fibroblasts showed enhanced cell adhesion, proliferation
and secretion of ECM when cultured on the nanofibrous scaffold.24
J. Name., 2012, 00, 1-3 | 1
COMMUNICATION
Journal Name
Figure 2: a) Cell viability assay showing percentage of live cells in culture
post incubation on ES-PGMA and ES-PGMA + PXS64, in the presence and
absence of TGFβ1 (2 ng/mL), b) and c) showing HDF cell morphology on
ES-PGMA + PXS64 + TGFβ1 (2 ng/mL) using fluorescent microscopy and
scanning electron microscopy respectively. Red arrows highlighting the cell
attachment and adhere points on the fibers. Scale bars: b) and c) 2 µm.
Figure 1: SEM secondary electron image of electrospun fibers of a) PGMA,
b) PGMA + PXS64, C) showing the chemical structure of PXS64. Scale bars:
a) and b) 10 µm.
‘‘grafting to’’ interactions mediated by the epoxy groups.27, 28 To this
end, poly(glycidyl methacrylate) (PGMA) used in the current study,
has an epoxy group in every repeating unit and has been used
extensively as a macromolecular anchoring layer.29, 30 In the present
study, PGMA was electrospun in the presence of PXS64 and in the
absence of PXS64 as a control (see supporting information for
experimental conditions). PXS64 loading in the electrospun PGMA
fibers was measured using high pressure liquid chromatography
(HPLC) to be ~76 % w/w. Electrospun fibrous scaffolds were
characterized using scanning electron microscopy (SEM). Fiber
diameter was measured from the SEM images to be 0.69 ± 0.31 µm
(average ± standard deviation) for ES-PGMA and 2.24 ± 1.06 µm in
the case of ES-PGMA + PXS64 (Figure 1).
Biocompatibility of the scaffold was investigated using the
colorimetric MTS assay and live/dead cell viability assay (see
Supporting Information for method) and were imaged using
fluorescence and scanning electron microscopy.31 Human dermal
fibroblasts (HDF) cells were cultured on the scaffolds and incubated
In the in vivo studies on dorsal wound model in diabetic rats, bFGF/
poly(ethylene glycol)-poly(D,L-lactide) mats showed elevated
healing with complete re-epithelialization and regeneration of skin
appendages such as hair and sebum, while bFGF promoted collagen
deposition and its remodelling similar to normal architecture.23
In the case of electrospun scaffolds the ability to modify the surface
properties of the nanofibers such as adhesion, wettability and
biocompatibility is pivotal in its integration as a tissue engineering
platform. Currently one of the biggest challenges using this
technique to produce scaffolds is that the polymers used thus far to
develop nanofibers need specific approaches for surface
modification that are complex and laborious.25 The achievement of a
certain degree of grafting universality requires the establishment of a
controlled method of introducing the desired functional groups.
Currently, this is achieved by physisorption.26 In contrast,
chemisorption which is difficult to achieve on polymeric nanofibers
would result in covalent attachment. Polymers containing epoxy
groups are examples of functional polymers that are able to react
with a wide range of biologically relevant molecules through
Figure 3: Collagen I gene expression analysis showing the percentage
change in gene expression as compared to non-treated negative control
(column 1). HDF cells incubated on both scaffolds and plastic tissue culture
plate both in presence and absence of TGFβ1 and release media. Data
presented as average ± SEM (n = 3). Significance was set at * p < 0.05 using
bonferroni post-hoc test in one way ANOVA analysis.
2 | J. Name., 2012, 00, 1-3
This journal is © The Royal Society of Chemistry 2012
Journal Name
both with and without TGFβ1 over a period of 72 hours. No
significant changes were observed on cell proliferation (see
Supporting Information; Figure S1) or cell viability (Figure 2a) on
both scaffolds with or without TGFβ1. To study the cell-matrix
interactions, scaffolds were incubated with HDF cells and imaged
using fluorescence and electron microscopy (see Supporting
Information for method). HDF cells were observed to adopt their
characteristic spindle shape morphology on both ES-PGMA and ESPGMA + PXS64 scaffolds in the presence and absence of TGFβ1
(Figure 2b, c and see Supporting Information; Figure S2 and S3).
Finally, the ability of ES-PGMA + PXS64 scaffold in regulating the
over expression of collagen I gene in HDF cells post TGFβ1
stimulation was evaluated using RT-qPCR (see Supporting
Information). HDF cells cultured on the biocompatible ES-PGMA
control scaffold showed significant increase in collagen I gene
expression in the presence of TGFβ1 (Figure 3). Cells incubated on
ES-PGMA + PXS64 scaffold under similar conditions showed
reduced expression of collagen I gene compared to ES-PGMA
scaffold, which reached significance in the presence of
supplemented release media. The use of DMSO in release media was
adopted from previously reported methodology32 to accelerate the
release of hydrophobic drugs in vitro and was shown to have no
effect on collagen expression alone (Figure 3).
Here we developed a PGMA scaffold and encapsulated an antiscarring drug, PXS64. The biocompatibility of the scaffold was
demonstrated in human dermal skin fibroblasts. Finally, the
efficacy of scaffold and drug in mitigating the increased
expression of collagen I in response to TGFβ1 stimulation was
also demonstrated. We believe that this potential proof of
principle approach can easily be adapted in the design of
scaffolds for tissue regeneration in the presence of other antifibrotic agents.
The authors would like to thank the Australian Research
Council for funding the work. The authors would also like to
acknowledge the Australian Microscopy & Microanalysis
Research Facility at the Centre for Microscopy,
Characterization & Analysis, The University of Western
Australia, funded by the University, State and Commonwealth
Governments. The authors would like to thank Dr Foteini
Hassiotou to kindly provide the RT-qPCR instrument and
Pharmaxis Ltd, Sydney for providing PXS64. MF is supported
by Chevron Australia. This work was partly funded by the
Fiona Wood Foundation.
Notes and references
a
School of Chemistry and Biochemistry, The University of Western
Australia,
Crawley,
Western
Australia,
Australia.
E-mail:
swaminatha.iyer@uwa.edu.au
b
Burn Injury Research Unit, School of Surgery, The University of
Western Australia, Crawley, Western Australia, Australia.
c
Burns Service Western Australia, Department of Health, Perth, Western
Australia, Australia
† Footnotes should appear here. These might include comments relevant
to but not central to the matter under discussion, limited experimental and
spectral data, and crystallographic data.
Electronic Supplementary Information (ESI) available: [details of any
supplementary information available should be included here]. See
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4 | J. Name., 2012, 00, 1-3
This journal is © The Royal Society of Chemistry 2012
Chapter 4
Conclusion and Future Works
In this thesis two complementary approaches were investigated to address the
problem of scarring after injury. The first part of the work focused on a tissue
engineering approach towards promotion of skin regeneration and subsequent drug
delivery application using a nanoscaffold while the latter part focused on the use of
novel anti-scarring therapeutics. The tissue engineering approach involved the
development of a multifunctional scaffold which could have a number of different
applications in addition to a potential drug-delivery scaffold including upconversion
probes, magneto-responsive material and biosensor. Subsequently, two analogues of
the natural anti-scarring molecule mannose-6-phosphate (analogue 1 and 2) were
investigated as potential scar modulators. Finally, the two concepts were blended to
fabricate a nanofibrous scaffold for the delivery of analogue 2 in scar therapy. In this
section, a summary of the work detailed in the publications will be presented.
4.1 Tissue engineering of a nanoscaffold
Tissue engineering has evolved as an innovative approach towards tissue
regeneration which usually relies on a platform to drive the cellular response.
Electrospinning was employed to fabricate a stable nanofibrous universal grafting
substrate using a polyglycidyl methacrylate (ES-PGMA) polymer. PGMA is an
interesting polymer because of the presence of an ester linkage and highly reactive
epoxy ring in each monomer unit. Presence of the epoxy ring provides a thermally
induced self-crosslinking capability to the scaffold which eliminates the need for
external chemical or physical crosslinking methods.
71
Multi-functionality of ES-PGMA was demonstrated first by end grafting it with
carboxy-terminated poly N-isopropyl acrylamide (PNIPAM-COOH) to design a
thermo-responsive scaffold. Contact angle measurements confirmed the surface
grafting of ES-PGMA with PNIPAM-COOH which was demonstrated to undergo
phase transition at the glass transition temperature of PNIPAM. Next, ESPGMA/nanoparticle composite scaffolds using three different types of nanoparticles
((NaGdF4:Yb, Er); Pd and Fe3O4) were generated using the electrospinning protocol
established above.
Nanoparticle loading in the case of hybrid scaffolds was confirmed by x-ray
microanalysis, TEM, and ICP and AAS analysis. SEM analysis however revealed
an increase in the fiber diameters in the hybrid materials. Changes in the solvent
conductivity and surface tension have been claimed to have major influence on fiber
diameter.
One of the major problems encountered during the electrospinning process caused
by the low vapour pressure of solvent used was the coagulation and drying of
Taylor’s cone at the tip of the needle which required periodic cleaning and resulted
in loss of material. Solvents with relatively higher vapour pressure such as THF can
be used to improve the electrospinning process and address the above mentioned
problem.
Applicability of these hybrid materials were examined in gas sensing for Pd,
upconversion imaging in the case of upconverting nanoparticles and magnetoresponsiveness for magnetite nanoparticles. Optimisation of the hydrogen gas
sensing measurement was difficult as the Pd concentration on the surface of the
matrix was unknown. Fiber porosity also limited the sensitivity to high
concentrations of hydrogen. Potentially, surface functionalization of the scaffold
with Pd nanoparticles, in addition to encapsulation, could be attempted to elevate the
sensitivity of the system towards lower concentrations of hydrogen gas. SQUID
analysis on the magnetite/polymer composite showed superparamagnetic behaviour
of the matrix making it ideal for applications in protective fabric material and
currency notes.304
72
4.2 Scar therapy and mannose-6-phosphate analogues
The M6P/IGFII receptor is important in trafficking of cargoes from the cell surface
and trans-golgi network to lysosomes.305-307 It plays an important role in various
diseases including fibrosis and lysosomal storage disorders.307 The trafficking
function of the receptor is exploited mainly for enzyme replacement therapy but has
not been a focus of anti-fibrotic therapeutic approaches 308
M6P is a natural occurring ligand of the M6P/IGFII receptor found in a number of
glycoproteins. The M6P/IGFII receptor consists of 15 repeating domains out of
which domains 3, 5 and 9 are specific towards M6P binding. With the emergence of
the anti-scarring properties of M6P309 and the structural determination of the
M6P/IGFII receptor,305 a great deal of research has focused on the development of
more efficient analogues of M6P.155, 156 This is because the therapeutic potential of
M6P is limited because of its low metabolic stability against phosphatases and
difficulty in retaining high concentrations at the injury site. Research has mainly
focussed on manipulating the structure of M6P in order to enhance its binding
efficiency towards the receptor while improving metabolic stability.157,
310, 311
Bioisosteric phosphonate analogues have been shown to have greater serum stability
than M6P. Taking advantage of this, two bioisosteric phosphonate analogues, 1 and
2, were explored for their potential as anti-scarring agents in this thesis. Analogue 1
was rationalised to have higher metabolic stability than M6P, thereby enhancing its
bioavailability, without lowering its affinity for the receptor. Molecular dynamics
simulation studies confirmed that analogue 1 has a similar interaction towards the
receptor as M6P (carried out in collaboration with Corry group). However,
intercellular delivery of analogue 1 showed no significant response in impeding the
expression of collagen I gene, the model used to study anti-scarring potential of the
analogues.
Neutral lipophilic phosphate compounds have been reported to undergo rapid
internalisation
increasing
their
local
concentration
within
the
cellular
microenvironment.312, 313 Therefore, analogue 2 was designed as a neutral molecule
with high logP to achieve intracellular drug delivery. Analogue 1 was masked by
esterification of the phosphate group to yield a non-charged bis(pivaloyloxymethyl)
73
derivative, analogue 2. Upon internalisation, phosphoester linkers undergo ester
hydrolysis by esterases present within the cellular microenvironment and release
analogue 1, allowing intracellular targeting of the M6P/IGFII receptor. Gene
expression studies showed significant reduction in collagen I expression, similar to
treatment with M6P, upon analogue 2 treatment. Therefore, it was concluded from
this part of the study that intracellular targeting of the receptor is effective using
analogue 2 as a prodrug of analogue 1.
One of the major problems encountered was in the analysis of protein expression.
Analysis of collagen I and fibronectin expression was attempted to supplement the
data on RNA expression levels. Immunoblotting was unreliable and had low
reproducibility, most likely in part due to the large molecular weights of the target
proteins and ineffective protein transfer. Alternative techniques such as HPLC and a
recently developed ‘scar-in-jar’ protocol may be more effective methods for the
quantitation of protein expression, especially for collagen I.314, 315
Despite the positive response in reducing the effects of TGFβ1, analogue 2 has
limitations. It is vulnerable to metabolic degradation mediated by esterases present
in growth serum. In addition, one of the hydrolysis products is formaldehyde, which
is toxic to cells at high concentrations limiting the dose that can be administered.
4.3 Scaffold based delivery of analogue 2
To achieve sustained delivery of the effective analogue 2, the electrospun PGMA
nanofibrous platform was manipulated for a drug delivery application. It was
anticipated that enzymatic vulnerability of analogue 2 could be subdued by
encapsulating it within the polymer scaffold by limiting its direct exposure to serum.
Using the parameters established above, analogue 2 was electrospun with PGMA to
yield a fibrous drug delivery platform with loading efficiency of ~76 % as measured
by HPLC. Despite the high drug loading its subsequent release from the fibers could
not be quantified. It was postulated that either the rate of drug release was below the
limit of detection of the instrument or there is no drug release from the fibers. In an
74
attempt to see if any cellular response could be detected, loaded scaffold was
incubated with HDF cells. Cyto-compatibility of the scaffold was confirmed by
MTS and a fluorescent staining live/dead assay. Interestingly, efficacy of the
released drug from the fibers was evident, in the presence of DMSO (facilitating the
release of drug from the fibers), from the collagen I gene expression studies which
showed a similar significant response as free drug in curtailing the upregulated
expression of collagen I post TGFβ1 stimulation. Therefore, it can be concluded that
slow sustained release of analogue 2 can be accomplished using electrospun PGMA
scaffold.
4.4 Future recommendations
There are still several questions that need to be addressed to translate this
preliminary work into the clinic. For example, the lack of biodegradability of the
PGMA scaffold, the serum stability of the drug and its release kinetics all need
further development.
The general strategy adopted to induce biodegradability is by copolymerising nondegradable polymers with another biodegradable polymer. Polymers have also been
copolymerised with peptides and proteins to not only improve biodegradability but
also to provide biologically active functional cues for their integration into the
matrix. For example, the RGD (arg-gly-asp) sequence peptide motif can be
copolymerised on PGMA using atom transfer living radical polymerisation (ATRP).
The RGD motif is strongly recognised by a subset of integrins which are protein
receptors necessary for cell adhesion and migration.316-318 Wound healing is a
combination of cellular proliferation and migration working in tandem. Copolymers
with RGD motifs are shown to promote cell migration. In addition, PGMA scaffolds
can be covalently functionalised with growth factors such as insulin like growth
factor I (IGF I) and epidermal growth factor (EGF) via an SN2 substitution ring
opening reaction. The ester groups present in the PGMA backbone can undergo
esterase mediated ester hydrolysis to sustainably deliver growth factors. Both IGF I
and EGF are growth inducing proteins which play pivotal roles in wound healing.319
Chemical conjugation of the growth factor to the scaffold allows their passive
75
diffusion which is continued over the life-time of the scaffold.320 Incorporation of
growth factors and the RGD peptide sequence in the electrospun PGMA matrix
could elevate the rate of healing by simultaneously promoting cell growth and
migration. Electrospun scaffolds have found clinical applications because of their
propensity to mimic extracellular matrix thereby providing a suitable environment to
promote regeneration. Matrix based delivery of growth factors are preferred over
their exogenous delivery because it protects the growth factor from degradation by
limiting their exposure to the proteolytic wound environment.321
Despite the efficacy of analogue 2 ([(bis(pivaloyloxymethyl) (POM)] derivative of
analogue 1) in reducing the expression of collagen I gene after TGFβ1 treatment, its
structural integrity in serum limits it’s in vivo translational potential. [Bis(S-acyl-2thioethyl) (SATE)] derivatives of phosphonate analogues have been shown to
possess higher enzymatic stability against serum and gastric juices as compared to
POM derivatives.313,
322
Benzaria et al. synthesised SATE derivatives of 9-[2-
(Phosphonomethoxy)ethyl]adenine (PMEA) and studied their stability in vitro in
comparison to bis[(pivaloyloxymethyl) (POM)]- and bis[dithiodiethyl (DTE)]PMEA
analogues.322 They concluded that although bis(POM)- and bis(tBu-SATE)PMEA
had similar anti-viral activity, bis(tBu-SATE)PMEA had greater stability than
bis(POM)PMEA in human gastric juice and human serum. It would potentially be
beneficial if the replacement of the POM functionality by a SATE derivative in
analogue 2 could improve its serum stability whilst maintaining its functional
efficacy. This could lead to oral or intravenous delivery of the drug. However, SATE
as a linker has its own limitations, with one of the degradation products, ethylene
sulphide, inducing toxicity at high concentrations in cells.
Analogue 2 can be modified at the anomeric centre by replacing the lipophilic group
with a fluorophore which would be beneficial in visualising its internalisation within
the cell microenvironment which can be tracked using confocal microscopy. This
will also allow another platform to study drug release profile from the scaffold.
Further, in a more combinatorial approach, analogue 2 could be combined with a
lysyl oxidase inhibitor. While analogue 2 works predominantly on controlling the
expression of collagen, lysyl oxidase inhibitors reduce collagen crosslinking, making
collagen more prone to degradation. A combined approach could be more effective
76
in reducing fibrosis than the use of a single compound targeting one part of the
fibrotic pathway.
Incorporation of such approaches within the functionalised scaffold would address
the shortfalls with both the scaffold and analogue 2 and could lead to pre-clinical
assessments. The efficacy of the system can be studied on an in vivo porcine wound
healing model. Porcine is a well-established wound healing model which has similar
characteristics to humans i.e. porcine heals with scars unlike rodents which heal
more via contraction instead of granulation tissue formation. The developed scaffold
could potentially be used as a dressing, incorporating growth factors and/or other
anti-scarring agents to be applied directly at the wound site. Scaffold mediated
delivery of the anti-scarring agents would need to be compared against free drug to
confirm the efficacy of this combinatorial approach.
4.5 Final remarks
The three overarching aims of the work compiled in this thesis are set out below:
1.
Optimise the parameters for electrospinning PGMA and study its efficacy as a
biocompatible universal polymeric scaffold
2.
Investigate the biocompatibility and potential of M6P analogues as anti-
scarring agents in an in vitro model using human dermal skin fibroblasts
3.
Investigate the potential of a combinatorial scaffold:drug approach to limit
fibrosis
The data presented within this thesis provides evidence of successful incorporation
of tissue engineering and drug delivery approaches to tackle the problem of skin
scarring. The scaffold developed by electrospinning PGMA can be used in various
other applications, some of which has been demonstrated in this thesis.
Two bioisosteric phosphonate analogues were investigated as potential anti-fibrotic
treatments. Analogue 2, a POM derivative of analogue 1, was shown to function as a
prodrug. Its mode of action was substantiated to be intracellular where its POM
77
linkers get hydrolysed to produce the active analogue 1. Therefore, it was
demonstrated that intracellular targeting of M6P/IGFII receptor using this analogue
is a potentially effective approach to inhibit TGFβ1 signalling and therefore reduce
fibrosis. In a tissue engineering approach, analogue 2 was co-electrospun with
PGMA to fabricate a fibrous scaffold. The biocompatibility of the scaffold was
confirmed in a human dermal skin fibroblast in vitro model. Scaffold mediated
delivery of analogue 2 was shown to regulate collagen I expression. This is the first
example of using this tissue engineering approach in this context and has the
potential with further work to address some of the major limitations with current
treatments aimed at reducing fibrosis.
78
Appendix A
Supporting information for papers
1.
Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D., Gong, P., Ye, W., Singh,
R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre, T., Wood, F.M., Fear, M.,
Lorenser, D., Sampson, D. D., Zdyrko, B., Smith, N.M., Luzinov, I., Iyer, K.S., A
Functional Reactive Polymer Nanofiber Matrix, RSC Advances (Submitted)
2.
Agarwal, V., Toshniwal, P., Smith N. E., Smith, N. M., Li, B., Clemons, T. D.,
Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B., Iyer, K.S., Intracellular
Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor Inhibitors
by Intracellular Delivery, Angewandte Chemie International Edition (submitted)
3.
Agarwal, V., Wood, F. M., Fear, M. and Iyer, K. S., Inhibiting the activation of
transforming growth factor-β using a polymeric nanofiber scaffold, Nanoscale (Submitted)
79
Supporting Information
A Functional Reactive Polymer Nanofiber Matrix
Vipul Agarwal,a Dominic Ho,a Diwei Ho,a Yuriy Galabura,b Faizah M.D. Yasin,a
Peijun Gong,c Weike Ye,d Ruhani Singh,a Alaa Munshi,a Martin Saunders,e Robert C.
Woodward,f Timothy St. Pierre,f Fiona M. Wood,g Mark Fear,g Dirk Lorenser,c David
D. Sampson,c, e Bogdan Zdyrkob, Nicole M. Smith,a Igor Luzinovb, *, and K.
Swaminathan Iyera,*
Materials
Polyglycidyl methacrylate (PGMA) with Mn = 220515 and Mw = 433730 was synthesized
by radical polymerisation as reported previously,1 Carboxy terminated N-isopropyl
acrylamide (PNIPAM-COOH) (Polymer source, Cat. # P5589, Mn = 42000). Methyl ethyl
ketone (MEK, Merck), iron (II) acetylacetonate, tetradecanediol, oleic acid, oleylamine,
dibenzyl ether, 1- octadecene and palladium (II) acetylacetonate purchased from Sigma
Aldrich. Gadolinium chloride, ytterbium chloride and erbium chloride purchased from
GFS chemicals. All other chemicals were purchased from Sigma-Aldrich. All chemicals
used were of analytical grade purity.
Methods
Electrospinning Procedure:
PGMA (15 w%) was dissolved in MEK (10 mL) overnight at room temperature with
constant stirring. In the case of composites, metal nanoparticles resuspended in MEK were
added (magnetite – 35.7 mg, palladium - 40 mg, Upconverting particles - 23.6 mg) to
maintain PGMA concentration at 15 w% the next day and further stirred for 1h to
homogenise the polymer solution.
80
PGMA polymer and polymer composite fibers were obtained via electrospinning (ESPGMA) (Nanofiber Electrospinning Unit, Cat. # NEU-010, Kato Tech, Japan). The
electrospinning parameters after optimisation were a voltage of 9.1 kV, working distance
of 9 cm, syringe pump speed of 0.04 mm/min (1 mL/h). Fibers were annealed at 80 °C for
5 hours post electrospinning. Nanoparticle loading in the ES-PGMA was determined by
ICP-MS and AAS analysis (in the case of the magnetite/ES-PGMA composite).
Magnetite Synthesis:
Magnetite nanoparticles were synthesised using the thermal decomposition method as
described previously.2 Briefly, iron (II) acetylacetonate {Fe(acac)3} (1 mmol),
tetradecanediol (5 mmol), oleic acid (3 mmol) and oleylamine (3 mmol) were dissolved in
dibenzyl ether. The reaction mixture was heated under a N2 atmosphere at 100 °C for 30
min, 200 °C for 2h and further refluxed at 300 °C for another 1h. The black coloured
product was collected via precipitation and washed with 3 x ethanol by centrifugation at
4000 rpm and dispersed in hexane. Nanoparticles were stored under an inert atmosphere.
Palladium Nanoparticles Synthesis:
Palladium nanoparticles were synthesised as per the method described.3 Briefly,
palladium(II) acetylacetonate {Pd(acac)3} (150 mg) was mixed with oleylamine (246 µL)
in toluene (61.5 mL) and stirred vigorously for 10 min at room temperature yielding a
yellow coloured reaction mixture. To this formaldehyde (300 µL) was added and further
reacted for 10 min. The reaction was carried out for a further 8h at 100 °C and the colour
changed to black, indicating the completion of the reaction. The product was brought to
81
room temperature, washed with 3 x ethanol and resuspended in chloroform. Nanoparticles
were stored under an inert atmosphere.
NaGdF4: Yb, Er Upconverting Nanoparticle Synthesis:
NaGdF4:Yb, Er upconverting nanoparticles were synthesised as per the method described.4
Briefly, GdCl3.6H2O (0.80 mmol), YbCl3.6H2O (0.18 mmol) and ErCl3.6H2O (0.02 mmol)
were added to the solution mixture of oleic acid (14 mL) and 1- octadecene (16 mL) and
homogenized under N2 while heated to 150 °C. The solution was then cooled to 50 °C,
methanol (10 mL) containing NaOH (2.5 mmol) and NH4F (4 mmol) was then added
slowly and reacted for another 30 min. Next, methanol was removed by heating the
reaction mixture at 100 °C under vacuum for 10 min. Under atmospheric pressure, the
reaction temperature was raised to 300 °C and the reaction was carried out for 1h under a
N2 atmosphere. The reaction was terminated by cooling to room temperature. The product
was precipitated in ethanol, collected by centrifugation and washed with ethanol. Finally,
the product was resuspended in THF.
PNIPAM- COOH attachment on ES-PGMA:
The PGMA polymer was electrospun onto 12 mm diameter glass coverslips (Cat. # G40112, ProSciTech) and annealed at 80 °C for 5h. Dried polymer was then exposed to carboxy
terminated poly N-isopropyl acrylamide (PNIPAM- COOH) (0.6 w% in water) and
reacted in the oven for 2h at 120 °C. The polymer mat on the coverslip was washed twice
with THF to remove any excess PNIPAM. Finally, the coverslip was dried in the oven
above 40 °C to remove the THF. The contact angle was measured on this coverslip at
room temperature and again at 40 °C.5
82
Scanning Electron Microscopy:
Dried electrospun PGMA fibers both with and without nanoparticles were coated with 4
nm of platinum and imaged using scanning electron microscope (Zeiss 1555, VP-FESEM)
at an accelerating voltage of 4-5 kV. The fiber diameters and width of the matrix were
calculated using the image analysis software ImageJ (NIH). A minimum of 50 random
fibers were measured. The data is reported as an average ± standard error mean.
X-ray Microanalysis:
Elemental analysis was carried out using the Oxford x-ray microanalysis system (Oxford
instrument X-MAX (80mm2)) set up on the SEM (Zeiss 1555, VP-FESEM). Elemental
data was collected at higher accelerating voltage of 15 kV at 10 mm working distance
required for x-ray microanalysis. Data was analysed using the Aztec version 2.1a analysis
instrument software.
Gas Sensing:
The hydrogen sensor setup consists of a test chamber with inlet gas and outlet gas, a
potentiostat and an electronic recorder. See reference for complete detail of the
experimental setup.6 Prior to entering the test chamber, hydrogen and nitrogen gas, which
are tested and carrier gas respectively, were mixed in a cyclonic mixer. Two silver epoxy
electrodes were painted to the Pd/ES-PGMA composite fibers mounted on an IDE and the
whole integration was mounted into the gas chamber subject to current-voltage (I-V)
sweeps. The procedure involved alternating nitrogen gas (20 min) and varying
concentrations of hydrogen gas (4 min). The change in the current was monitored
simultaneously. The total gas flow rate was 1000 mL/min. The voltage applied between
83
the electrodes was 100 mV dc. An IDE consists of 15 fingers, each 15 μm in width and
550 μm in length, with a finger gap of 10 μm. Each individual electrode is connected to a
boding pad (200 μm×250 μm) to provide a sufficient area for subsequent wire bonding.
Transmission Electron Microscopy (TEM):
Nanoparticles were air dried onto carbon-coated copper grids and imaged using a JEOL
3000F TEM operating at 300 kV. The nanoparticle size was determined using Image J
software. A minimum of 200 particles were measured and the data reported as the average
± standard error mean.
Superconducting Quantum Interference Device (SQUID) Magnetometry:
The magnetic properties of the dried magnetite/ES-PGMA composite (88.2 mg) were
measured using a SQUID magnetometer (Quantum Design 7 Tesla MPMS) operating
between 5K and 300K. Magnetite nanoparticles (22.5 mg) were lyophilised prior to their
measurement and their saturation magnetisation was recorded at 70 kOe at 5K. The zero
field cooled and field cooled measurements were conducted in a field of 0.1 kOe.
Contact Angle:
Static contact angles of MilliQ water on the surface of the electrospun polymer matrix
were measured using a home-built goniometer with Rame- Hart scope attachment.7 The
polymer was electrospun on the 12 mm glass coverslips (Cat. # G401-12, ProSciTech)
both with and without nanoparticles (refer PNIPAM attachment on ES-PGMA). 5 µL of
water was pipetted onto the membrane and images were taken after the drop edges came to
rest (~2 min) using a Canon EOS450D and the angle was measured using the calibrated
84
Rame-Hart scope. Measurements were carried out at room temperature and again at 40 °C.
Images were processed using ImageJ software. Measurements were done in duplicates and
reported as average ± standard deviation.
NIR Room Temperature Emission Spectroscopy (λ
exc
= 974 nm):
Upconversion spectra measurements were carried out as previously described.8 Briefly,
upconverting particles were suspended in chloroform and were air dried on glass slides.
The upconversion spectra were obtained with an optical set up incorporating a 980 nm
laser diode. The peak wavelength of the laser diode is 974.5 nm. The optical excitation
intensity for obtaining the spectrum shown in Fig. 3A was 7665 W/mm2
References
1.
V. Tsyalkovsky, R. Burtovyy, V. Klep, R. Lupitskyy, M. Motornov, S. Minko and I.
Luzinov, Langmuir, 2010, 26, 10684-10692.
2.
S. Sun, H. Zeng, D. B. Robinson, S. Raoux, P. M. Rice, S. X. Wang and G. Li, J.
Am. Chem. Soc., 2003, 126, 273-279.
3.
Z. Niu, Q. Peng, M. Gong, H. Rong and Y. Li, Angew. Chem. Int. Ed., 2011, 50,
6315-6319.
4.
C. Liu, Z. Gao, J. Zeng, Y. Hou, F. Fang, Y. Li, R. Qiao, L. Shen, H. Lei, W. Yang
and M. Gao, ACS Nano, 2013, 7, 7227-7240.
5.
G. D. Fu, L. Q. Xu, F. Yao, K. Zhang, X. F. Wang, M. F. Zhu and S. Z. Nie, ACS
Appl. Mater. Inter., 2009, 1, 239-243.
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J. Zou, L. J. Hubble, K. S. Iyer and C. L. Raston, Sensor Actuat.B-Chem., 2010, 150,
291-295.
7.
M. V. Baker and J. D. Watling, Langmuir, 1997, 13, 2027-2032.
8.
M. Challenor, P. Gong, D. Lorenser, M. Fitzgerald, S. Dunlop, D. D. Sampson and
K. Swaminathan Iyer, ACS Appl. Mater. Inter., 2013, 5, 7875-7880.
86
Enhancing the Efficacy of Cation-Independent Mannose
6-Phosphate Receptor Inhibitors by Intracellular
Delivery**
Vipul Agarwal, Priyanka Toshniwal,# Natalie E Smith,# Nicole M Smith, Binbin
Li, Tristan D Clemons, Lindsay T Byrne, Foteini Hassiotou, Fiona M Wood,
Mark Fear, Ben Corry* and K Swaminathan Iyer*
Supporting Information
S1.1 Cell culture
Human primary dermal fibroblast cell cultures from normal skin were cultured in
Dulbecco’s Modified Eagle’s Medium (DMEM/F12 - GlutaMAX; Invitrogen Gibco)
supplemented with 10 % fetal bovine serum (FBS; Invitrogen Gibco) and 1%
penicillin/streptomycin (Invitrogen Gibco). Analogues 1 and 2 were provided by
Pharmaxis Ltd after stringent QC analysis. M6P and analogue 1 were dissolved in
PBS filter sterilized, aliquoted and stored at -20 °C. Each aliquot was used for
maximum 2 freeze thaw cycles. Analogue 2 was dissolved in DMSO and treated the
similar way as other analogues. The cells were incubated at 37 °C in a humidified
atmosphere of 5 % CO2. All experiments were carried out with cells between
passages 3-6.
S2.1 Cell Viability
Cell viability was determined using a LIVE/DEAD viability/cytotoxicity Kit
(Invitrogen, UK) which measures the membrane integrity of cells, as per
manufacturer’s protocol. In brief, 20000 cells were seeded in each well in a 24 well
plate and incubated with analogues at 10 µM concentration in cell culture media
(DMEM F-12 containing 10% FBS and 1% Penicillin/ Streptomycin) for 30 min
(media was added in case of controls). Following which human recombinant TGFβ1
87
(2 ng/mL) (Cat.# 14-8348-62, eBioscience) was added to the wells (fresh media was
added in controls) and plates were incubated for 24 h and 72 h in the humidified
incubator at 37 °C with 5 % CO2. At the stipulated time (24 h and 72 h), cells were
washed with PBS (3 times) and then stained with calcein (100 µL, 1 µM)/ ethidium
bromide (100 µL, 2 µM) in PBS and incubated in the humidified incubator at 37 °C
and 5 % CO2 for 30 min. Images were captured using an Olympus IX71 inverted
microscope with a 20 x objective with fixed exposure time. Both live and dead cells
were counted using Image J software (NIH) with cell counter plug in. Data presented
as mean ± standard error mean (n = 4).
S3.1 Cell Body Area
Cell size was measured using Image J software (NIH).[1] Minimum of 40 cells were
randomly selected from the fluorescence images and their area was measured.
Values reported as mean ± standard error mean.
S4.1 Cell Proliferation
Cell proliferation was measured using the MTS assay (Cell Titer 96 ®Aqueous,
Promega, Madison, USA) as per the manufacturer’s protocol. Briefly, 1500 cells
were seeded in each well of a 96 well plate and treated with analogues at 10 µM
concentration in cell culture media (DMEM F-12 containing 10 % FBS and 1 %
Penicillin/ Streptomycin) for 30 min (media was added in case of controls) prior to
the addition of human recombinant TGFβ1 (2 ng/mL) to each well (fresh media was
added in controls) and the plates (individual for each time point) were incubated for
72 h in the humidified incubator at 37 °C with 5 % CO2. MTS solution (40 µL) was
added in each well the next day, and was considered as 0 h time point, and incubated
for 3 h in the humidified incubator at 37 °C with 5 % CO2. Following which 80 µL
88
from each well was transferred into a new 96 well plates and read under a plate
reader at 490 nm excitation wavelength. Same protocol was followed at every time
point for next 72 h. Data presented as mean ± standard error mean (n = 5).
S5.1 Gene Expression
Gene expression was measured using real time quantitative polymerase chain
reaction. Cells (50000) were seeded in 24 well plates and incubated for 24 h in the
humidified incubator at 37 °C and 5% CO2. Next day culture media (10 % FBS) was
replaced with starve media (0.1 % FBS) and incubated for further 24 h in the
incubator to bring all the cells under same physiological cycle. Next day, cells were
treated with required concentration of M6P analogue 30 min prior to TGFβ1 (2
ng/mL in starve media) stimulation and further incubated for 48 h in the humidified
incubator at 37 °C and 5% CO2. mRNA was extracted using RNeasy mini kit®
according to manufacturers’ protocol (Qiagen GmbH). For reverse transcription 1.5
µg of total mRNA was converted to cDNA using Superscript VILO (Cat.# 11754,
Applied Biosystems) according to manufacturers’ protocol. 150 ng of cDNA was
analysed by ABI 7500 fast analysis real-time PCR system using TaqMan® master
mix and col1a1 probes (Hs01076777_m1, Life Technologies). GAPDH was used as
a reference gene (Cat.# 4326317E, Life Technologies). Analysis was carried out
using the instrument software. Data presented as mean ± standard error mean (n =
3).
S6.1 Protein Expression
89
Protein expression was measured using western blotting. Cells (1 x 105) were seeded
in each well of a 6 well plate and incubated for 24 h in the humidified incubator at
37 °C and 5% CO2. Next day culture media (10 % FBS) was replaced with starve
media (0.1 % FBS) and incubated for further 24 h in the incubator to bring all the
cells under same physiological cycle. Next day, cells were treated with required
concentration of M6P analogue 30 min prior to TGFβ1 (2 ng/mL in starve media)
stimulation and further incubated for 72 h in the humidified incubator at 37 °C and
5% CO2. Whole cell lysates were prepared from treated cells. 35 µg of protein was
denatured and subjected to SDS-PAGE, and transferred to nitrocellulose membrane
(Cat.# 10600007, Amersham, General Healthcare Lifesciences) by standard transfer.
Post transfer, membrane was blocked with 5% skim milk/0.1% TBST for 30 min at
room temperature, then incubated overnight with rabbit anti-human collagen I
antibody (1:2000 in 5% skim milk/0.1% TBST, Cat. # NB600-408, Novus) at 4 °C.
Membranes were washed with 0.1% TBST and incubated with peroxide conjugated
mouse anti-rabbit (1:5000 in 5% skim milk/0.1% TBST, Cat.# NA934VS, GE
Healthcare Lifesciences) for 1 h in 5% skim milk/0.1% TBST at room temperature.
Immunoreactivity was detected using the chemiluminescent HRP substrate (Cat.#
WBKLS0100, Millipore IMMobilon) and the signal was captured with the
Chemidoc (BioRad, Model #731BR02144) and analysed using ImageJ software
(NIH). To confirm equality of protein loading, all membranes were stripped and
reanalysed for β-actin expression using 1° antibody (1:50000 in 5% skim milk/0.1%
TBST, Cat.# A1978, Sigma) and 2° (1:5000 in 5% skim milk/0.1% TBST, Cat.#
NA9310V, GE Healthcare Lifesciences). Data presented as mean ± standard error
mean (n = 3).
90
S7.1 Statistics
The results for cell viability, cell body area, cell proliferation, gene and protein
expression are expressed as mean ± standard error mean (SEM) and analysed for
analysis of variance (ANOVA). Significance was evaluated using Bonferroni and
Turkey’s post-hoc analysis and set at 95% confidence (p < 0.05).
S8.1 Simulation Systems:
Experimental Methods:
6 simulation systems of the M6P/IGFII receptor were investigated; these included
the domain 3 and domain 5 dimers with a ligand in each binding site (2 binding sites
per dimer). These ligands were M6P, analogues 1 and 2. The coordinates for the
domain 3 dimer with M6P bound (pdb accession code 1SYO)[2] and the domain 5
monomer (pdb accession code 2KVB)[3] were obtained from the protein database. In
order to obtain the dimer for domain 5 and position M6P in the binding pocket, the
domain 5 beta sheet regions were aligned with the corresponding beta sheet regions
of domain 3. Analogue 1 and 2 were positioned in each binding pocket by aligning
the mannose ring and phosphate group to the M6P coordinates obtained for domain
3.[2] Each system was then solvated in a TIP3P water box of dimensions 65 x 60 x
114 Å and ionised with 150 mM KCl. All simulations were run with periodic
boundary conditions, constant temperature (310 K) maintained using Langevin
dynamics and constant pressure (1 atm) maintained with a Langevin piston, and the
particle mesh Ewald method was used to compute full system electrostatics.[4] The
CHARMM 36 force field was used for protein, water and M6P parameters.[5] The
ion parameters were obtained from Joung and Cheatham.[6] Missing parameters for
analogues 1 and 2 were obtained using ab-initio techniques
91
[7]
with the program
Gaussian 09.[8] All molecular dynamics simulations were run with the program
NAMD[9] using rigid bonds to hydrogen and 2 fs time steps. Molecular graphics
were generated using VMD.[10]
Water and ions were energy minimized for 5000 steps and equilibrated for 20 ps
with the protein and substrate held fixed. A harmonic restraint with a force constant
of 20 kcal/mol was then applied to the backbone atoms of the protein and on each
ligand and the system was minimized for a further 10,000 steps prior to 500 ps of
equilibration. This step was repeated with gradual reductions in the force constant
with values of 10, 5, 2.5 and 1 kcal/mol. Finally, to replicate the influence of the
surrounding protein domains on the individual domain being simulated, a harmonic
restraint with a force constant of 0.1 kcal/mol was placed on all of the protein Cα
atoms which were located more than 10 Å from the binding pocket in order to
ensure no loss of secondary structure throughout the simulation. The system was
then minimized for a further 10,000 steps prior to 10 ns of equilibration.
Subsequently 100 ns of equilibrium simulation were obtained for each of the 6
systems.
Data Analysis:
Cluster analysis was performed on the final 100 ns of equilibrium simulation for
each ligand in order to determine the most occupied binding positions. Each ligand
was clustered according to the RMSD of its coordinates with a cut-off of 3 Å.
Subsequently the NAMDEnergy plugin of VMD was utilized to determine the
interaction energy of each ligand with the protein for the entire time and for the most
populated clusters.
92
Figure S1: Most populated binding positions (clusters) of each drug in the domain 3 and domain 5 binding domains as determined
from 100 ns of molecular dynamics simulation. In each case the protein is shown in grey, and the most to the least populated
clusters are for ligand 1 (initially bound to dimer 1): cyan, pink, mauve, white and green and for ligand 2 (initially bound to dimer 2):
dark blue, red, orange, yellow and green. a) M6P in domain 3: Ligands 1 and 2 both remain in the vicinity of the binding site
determined in the crystal structure. However as there are multiple hydrogen bond acceptors and donors, both ligands sample
multiple positions as the mannose ring hydroxyl groups form hydrogen bonds with various residues and these change over the 100
ns simulation. b) Analogue 1 in domain 3: Ligand 1 has only one binding position (cyan) and ligand 2 has only two clusters
indicating that in both cases it binds stably to the protein and remains in the binding pocket throughout the simulation. This binding
appears to be stabilized by the interaction between the m-xylene rings of the 2 ligands. c) Analogue 2 Domain 3: Both ligand 1 and
2 are oriented such that the benzyl groups associate in the center of the dimerization domain, hence while the ligands do have more
than one binding position they remain in the vicinity of the predicted binding site. d) Domain 5 M6P: In this case the ligands occupy
positions which are not asymmetrical, while ligand 1 favors a horizontal orientation with one major binding position (cyan) ligand 2
favors a vertical orientation where it has more flexibility to sample new positions. e) Analogue 1 in domain 5: Ligand 1 moves away
from its initial position in the binding site and occupies a position where it is in close contact with ligand 2. f) Analogue 2 in domain 5:
In this case both ligands remain in the vicinity of the original binding position. However, while ligand 1 has one major cluster
implying its motion is restricted, ligand 2 has more flexibility to move occupying multiple clusters.
93
Figure S2: Snapshots from 100 ns molecular dynamics simulations representative of the most heavily occupied binding positions
for each ligand in the domain 3 and domain 5 dimers. The two protein subunits are shown in pink and grey respectively and
residues which form common hydrogen bonds are labelled. a) M6P and b) Analogue 1 in domain 3: Ligand 1 and 2 form the
strongest energetic interactions with Lys350, Lys358, Ser386 and Arg391. c) Analogue 2 in domain 3: Ligand 1 and 2 form the
strongest interactions with Gln348, Arg391 and Glu416. d) M6P in domain 5: Ligand 1 and 2 are oriented differently to each other in
the binding pocket. M6P forms its strongest interactions with Arg687. e) Analogue 1 in domain 5: interacts most favorably with
Asn680 andArg687. f) Analogue 2 in domain 5: Ligand 1 and 2 form the strongest interactions with Gln644, Trp653, Arg687 and
Tyr714.
Figure S3: Average interaction energies obtained for the most occupied positions of each ligand in domains 3 and 5. R391 in
domain 3 is equivalent to R687 in domain 5. M6P (a) and analogue 1 (b) in domain 3 both show similar interactions for each ligand
and the protein subunit it is most closely associated with (for example Protein 1 with Ligand 1 as compared to Ligand 2 with Protein
2). This is because the binding mode is similar for each ligand in its respective binding site (Figure S2). Similarly, in each case the
ligands also interact closely with Lys350 which is actually located on the alternate protein subunit (for example Ligand 1 to Protein
2). When compared to (a) and (b) analogue 2 (c) in domain 3 shows significant decreases in the interaction energies which is
reflected by its lower overall interaction energy. M6P (d) and analogue 1 (e) in domain 5 have two major interactions the first one to
Arg687 and the second to Glu709. Overall the interaction energies for domain 5 are much lower than those observed for domain 3.
Analogue 2 (f) in domain 5 shows a similar pattern with a further reduction in the observed interaction energies.
94
Figure S4: Cell Proliferation assay showing cell growth over the period of 72 h post incubation with analogues both in the presence
and absence of TGFβ1. First and second column in each condition is representing 24 h and 72 h respectively. Significant
proliferation was observed in all groups despite the reduction in proliferation in TGFβ1 treated groups. Data presented as Mean ±
SEM. Significance was set at * p < 0.05 using one way ANOVA with Bonferroni post-hoc analysis.
Figure S5: HDF cell morphology post calcein AM staining imaged using fluorescent microscopy. Cells were treated for 72h, stained
and imaged: a) untreated (control), b) TGFβ1 (2 ng/mL) treatment, c) M6P (10 µM) + TGFβ1 (2 ng/mL), d) Analogue 1 (10 µM) +
TGFβ1 (2 ng/mL) and e) Analogue 2 (10 µM) + TGFβ1 (2 ng/mL). Scale bar 2 µm.
Figure S6: Collagen I gene time response curve post TGFβ1 (2 ng/mL) stimulation. Collagen I gene expression was significantly
upregulated at 24 and 48 h post TGFβ1 stimulation compared to non-treated control. The 48 h stimulation was selected for all
further experiments as it yielded significantly higher response. Data presented as average ± SEM (n = 3). Significance was set at * p
< 0.05 using bonferroni post-hoc test in one way ANOVA analysis.
95
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L. J. Olson, F. C. Peterson, A. Castonguay, R. N. Bohnsack, M. Kudo, R. R.
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U. Essmann, L. Perera, M. L. Berkowitz, T. Darden, H. Lee, L. G. Pedersen, J.
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A. D. MacKerell, D. Bashford, M. Bellott, R. L. Dunbrack, J. D. Evanseck, M.
J. Field, S. Fischer, J. Gao, H. Guo, S. Ha, D. Joseph-McCarthy, L. Kuchnir, K.
Kuczera, F. T. K. Lau, C. Mattos, S. Michnick, T. Ngo, D. T. Nguyen, B. Prodhom,
W. E. Reiher, B. Roux, M. Schlenkrich, J. C. Smith, R. Stote, J. Straub, M.
Watanabe, J. Wiorkiewicz-Kuczera, D. Yin, M. Karplus, J Phys Chem B 1998, 102,
3586-3616.
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I. S. Joung, T. E. Cheatham, 3rd, J. Phys. Chem. B 2008, 112, 9020-9041.
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N. Foloppe, A. D. MacKerell, J Comput Chem 2000, 21, 86-104.
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M. J. Frisch, G. W. Trucks, H. B. Schlegel, G. E. Scuseria, M. A. Robb, J. R.
Cheeseman, G. Scalmani, V. Barone, B. Mennucci, G. A. Petersson, H. Nakatsuji,
M. Caricato, X. Li, H. P. Hratchian, A. F. Izmaylov, J. Bloino, G. Zheng, J. L.
Sonnenberg, M. Hada, M. Ehara, K. Toyota, R. Fukuda, J. Hasegawa, M. Ishida, T.
Nakajima, Y. Honda, O. Kitao, H. Nakai, T. Vreven, J. A. M. Jr., J. E. Peralta, F.
Ogliaro, M. Bearpark, J. J. Heyd, E. Brothers, K. N. Kudin, V. N. Staroverov, R.
Kobayashi, J. Normand, K. Raghavachari, A. Rendell, J. C. Burant, S. S. Iyengar, J.
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Tomasi, M. Cossi, N. Rega, J. M. Millam, M. Klene, J. E. Knox, J. B. Cross, V.
Bakken, C. Adamo, J. Jaramillo, R. Gomperts, R. E. Stratmann, O. Yazyev, A. J.
Austin, R. Cammi, C. Pomelli, J. W. Ochterski, R. L. Martin, K. Morokuma, V. G.
Zakrzewski, G. A. Voth, P. Salvador, J. J. Dannenberg, S. Dapprich, A. D. Daniels,
O. Farkas, J. B. Foresman, J. V. Ortiz, J. Cioslowski, D. J. Fox, 2009, Revision D.01.
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J. C. Phillips, R. Braun, W. Wang, J. Gumbart, E. Tajkhorshid, E. Villa, C.
Chipot, R. D. Skeel, L. Kale, K. Schulten, J Comput Chem 2005, 26, 1781-1802.
[10] W. Humphrey, A. Dalke, K. Schulten, J Mol Graph Model 1996, 14, 33-38.
97
Inhibiting the activation of transforming growth factorβ using a polymeric nanofiber scaffold
Vipul Agarwal,a Fiona M. Wood,b,c Mark Fearb and K. Swaminathan Iyer*a
Supporting Information
S1.
Materials
Polyglycidyl methacrylate (PGMA) with Mn = 220515 and Mw = 433730 was synthesized by radical
polymerisation as reported previously.1 MEK was purchased from Merck. PXS64 was kindly provided by
Pharmaxis Ltd after intensive QC testing.
S2.
Electrospinning Procedure:
PGMA (15 wt%) was dissolved in MEK overnight at room temperature with constant stirring. In case of PGMA
+ PXS64, PXS64 (2.95 mg) was mixed with PGMA and dissolved in MEK overnight at room temperature.
PGMA polymer and PGMA + PXS64 fibrous scaffold were obtained via electrospinning (Nanofiber
Electrospinning Unit, Cat. # NEU-010, Kato Tech, Japan) onto 12 mm and 6 mm glass coverslips (Cat. # G40112, and G401-06 respectively, ProSciTech). The electrospinning parameters after optimisation were a voltage of
9.1 kV, working distance of 9 cm, syringe pump speed of 0.04 mm/min (1 mL/h), and the traverse and collection
drum speeds were set at 0 cm/min and 0 m/min respectively. Fibers were dried and crosslinked by annealing at
80 °C for 5 h and stored at room temperature for maximum of a month. Scaffolds were UV sterilised in the tissue
culture hood for 15 min and further washed with 2 x PBS prior to cell experiments.
S3.
Scanning Electron Microscopy:
Dried electrospun PGMA fibres both with and without PXS64 were coated with 4 nm of platinum and imaged
using scanning electron microscope (Zeiss 1555, VP-FESEM) at an accelerating voltage of 4-5 kV. The fiber
diameter was calculated using the image analysis software ImageJ (NIH). A minimum of 50 random fibers were
measured. The data is reported as an average ± standard deviation. In the cell experiments, cells were incubated
as per cell viability protocol. At the stipulated time point, culture media was removed and coverslips loaded with
scaffold and cells were washed with 2 x PBS and fixed for 10 min using glutaraldehyde (2.5 % in PBS),
98
followed by 2 x PBS washes. Next serial dehydration steps were performed with increasing concentration of
ethanol to replace all the water in the samples with dry ethanol before critical point drying step. Post critical
point drying samples were coated with 4 nm of platinum and imaged using aforementioned parameters.
S4.
High Pressure Liquid Chromatography
Drug loading analysis was carried out using HPLC (solvent A: water with 0.1 % TFA, solvent B: acetonitrile
with 0.1 % TFA, Phenomenex-C18 (2) 100A column (4.6 × 150 mm, 5 microns) at room temperature, 1 mL/ min,
λ = 280 nm, gradient: 100 % solvent B in 17 min) with Water 2695 pumping system and 2489 UV/Vis detector.
Drug loading was calculated by dissolving 35 mg of fibers (PGMA + PXS64) in acetonitrile (0.5 mL) and
calculated from the standard curve developed by dissolving free PXS64 in acetonitrile.
S5.
Cell culture
Human primary dermal fibroblast cell cultures from normal skin were cultured in Dulbecco’s Modified Eagle’s
Medium (DMEM/F12 - GlutaMAX; Invitrogen Gibco) supplemented with 10% fetal bovine serum (FBS;
Invitrogen Gibco) and 1 % penicillin/streptomycin (Invitrogen Gibco). The cells were incubated at 37 °C in a
humidified atmosphere with 5 % CO2. All experiments were carried out with cells between passages 3-6.
S6.
Cell Viability
Cell viability was determined using a LIVE/DEAD Viability/Cytotoxicity Kit (Invitrogen, UK) which measures
the membrane integrity of cells, as per manufacturer’s protocol. In brief, 20000 cells were seeded on the
sterilised electrospun scaffold on glass coverslips in a 24 well plate and supplemented with cell culture media
(DMEM F-12 containing 10 % FBS and 1 % Penicillin/ Streptomycin) for 30 min. Next, human recombinant
TGFβ1 (2 ng/mL, Cat. # 14-8348-62, eBioscience) was added to the wells (fresh media was added in controls)
and plates were incubated for 24 h and 72 h in the humidified incubator at 37 °C with 5 % CO2. At the stipulated
time (24 h and 72 h), cells were washed with PBS (3 times) and then stained with calcein AM (100 µL, 1
µM)/ethidium bromide I (100 µL, 2 µM) in PBS and further incubated for 30 min in the incubator. Images were
captured using an Olympus IX71 inverted microscope with a 20 x objective with fixed exposure time. Both live
and dead cells were counted using Image J software (NIH) with cell counter plug in. Data presented as mean ±
standard error mean (n = 4).
99
S7.
Cell Proliferation
Cell proliferation was measured using the MTS assay (Cell Titer 96 ®Aqueous, Promega, Madison, USA) as per
the manufacturer’s protocol. Briefly, 1500 cells were seeded on the sterilised electrospun scaffold on 6 mm glass
coverslips in each well of a 96 well plate and supplemented with cell culture media (DMEM F-12 containing 10
% FBS and 1 % Penicillin/ Streptomycin) for 30 min followed by the addition of human recombinant TGFβ1 (2
ng/mL) to the wells (fresh media was added in controls) and the plates (individual for each time point) were
incubated for 72 h in the humidified incubator at 37 °C with 5 % CO2. MTS solution (40 µL) was added in each
well the next day, and was considered as 0 h time point, and incubated for 3 h in the humidified incubator at 37
°C with 5 % CO2. Following which 80 µL from each well was transferred into a new 96 well plates and read
under a plate reader at 490 nm excitation wavelength. Same protocol was followed at every time point for next
72 h. Data presented as mean ± standard error mean (n = 5).
S8.
Gene Expression
Gene expression was measured using real time quantitative polymerase chain reaction. Cells (50000, in
DMEM/F12 media containing 10% FBS) were seeded on sterilised scaffolds (electrospun on coverslips) in 24
well plates and incubated for 24 h in the humidified incubator at 37 °C and 5% CO2. Next day culture media
(containing 10 % FBS) was replaced with starve media (containing 0.1 % FBS) and incubated for further 24 h in
the incubator to bring all the cells under same physiological cycle. Next day, TGFβ1 (2 ng/mL in starve media)
was added in the culture and further incubated for 48 h in the incubator. mRNA was extracted using RNeasy
mini kit® according to manufacturers’ protocol (Qiagen GmbH). For reverse transcription 1.5 µg of total mRNA
was converted to cDNA using Superscript VILO (Cat. # 11754, Applied Biosystems) according to
manufacturers’ protocol. 150 ng of cDNA was analysed by ABI 7500 fast analysis real-time PCR system using
TaqMan® master mix and col1a1 probes (Hs01076777_m1, Life Technologies). GAPDH was used as a reference
gene (Cat. # 4326317E, Life Technologies). Analysis was carried out using the instrument software. Release
media was prepared using DMEM/F-12 media (containing 10 % FBS and 1 % Penicillin/ Streptomycin)
supplemented with 0.2 % DMSO. Data presented as mean ± standard error mean (n = 3).
S9.
Statistics
100
The results for cell viability, cell proliferation and gene expression are expressed as mean ± standard error mean
(SEM) and analysed for analysis of variance (ANOVA). Significance was evaluated using Bonferroni and
Turkey’s post-hoc analysis and set at 95% confidence (p < 0.05).
Figure S1: Cell proliferation assay showing cell growth over the period of 72 h post incubation with the two scaffolds both
with and without TGFβ1 (2 ng/mL). Data presented as Mean ± SEM. Significance was set at * p < 0.05 using one way ANOVA
and Bonferroni post-hoc analysis.
Figure S2: A representative florescent images showing cell morphology and the number of live and dead cells in culture. HDF
cells were incubated on a) ES-PGMA scaffold, b) ES-PGMA + TGFβ1 (2 ng/mL) and c) ES-PGMA + PXS64 scaffold. Scale
bar 2 µm. Cells were stained using Calcein AM/Ethidium bromide I staining where live cells fluoresce green while dead cells
fluoresce red.
101
Figure S3: A representative image showing cell adherence on the two scaffolds. HDF cells were incubated a) ES-PGMA
scaffold, b) ES-PGMA + TGFβ1 (2 ng/mL) and c) ES-PGMA + PXS64 scaffold. Scale bar: a) 2 µm, b) and c) 1 µm. Red
arrows highlighting the cells.
References
1.
K. S. Iyer, B. Zdyrko, H. Malz, J. Pionteck and I. Luzinov, Macromolecules, 2003, 36, 6519-6526.
102
Appendix B
Published work not directly included in the thesis
Peer reviewed publications contributed by the candidate are listed below. However, only
published articles are attached in the thesis.
1.
Agarwal, V., Tjandra, E.S., Iyer, K. S., Humfrey, B., Fear, M., Wood, F. W.,
Dunlop, S. and Raston, C. L., Evaluating the effects of nacre on human skin and scar cells
in culture, Toxicology Research, 3, 223-227 (2014)
2.
Eroglu, E., Chen, X., Bradshaw, M., Agarwal, V. , Zou, J., Stewart, S.G., Duan, X.,
Lamb, R.N., Smith, S.M., Raston, C. and Iyer, K. S., Biogenic production of palladium
nanocrystals using microalgae and their immobilization on chitosan nanofibers for
catalytic applications, RSC Advances, 3, 1009-1012 (2013)
3.
Eroglu, E., Agarwal, V., Bradshaw, M., Chen, X., Smith, S. M., Raston, C. and Iyer,
K. S., Nitrate removal from liquid effluents using microalgae immobilized on chitosan
nanofiber mats, Green chemistry, 14 (10), 2682-2685 (2012)
4.
Ho, D., Zou, J., Chen, X., Munshi, A., Smith, N. M., Agarwal, V., Hodgetts, S.I.,
Plant, G. W., Bakker, A., Harvey, A. R., Luzinov, I., Iyer, K. S., Hierarchical patterning of
multifunctional conducting polymer nanoparticles as a bionic platform for topographic
contact guidance, ACS nano, 9 (2), 1767-1774 (2015)
103
Volume 3 Number 4 July 2014 Pages 217–292
Toxicology
Research
www.rsc.org/toxicology
ISSN 2045-452X
Chinese Society Of Toxicology
COMMUNICATION
Colin L. Raston et al.
Evaluating the effects of nacre on human skin and scar cells
in culture
Toxicology Research
Published on 02 May 2014. Downloaded by University of Western Australia on 09/10/2014 10:23:29.
COMMUNICATION
Cite this: Toxicol. Res., 2014, 3, 223
Received 6th January 2014,
Accepted 2nd May 2014
DOI: 10.1039/c4tx00004h
View Article Online
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Evaluating the effects of nacre on human skin and
scar cells in culture†
Vipul Agarwal,a Edwin S. Tjandra,a K. Swaminathan Iyer,a Barry Humfrey,b Mark Fear,c
Fiona M. Wood,c Sarah Dunlopd and Colin L. Raston*e
www.rsc.org/toxicology
Pearl nacre, a biomineralisation product of molluscs, has growing
applications in cosmetics, as well as dental and bone restoration,
yet a systematic evaluation of its biosafety is lacking. Here, we
assessed the biocompatibility of nacre with two human primary
dermal fibroblast cell cultures and an immortalised epidermal cell
line and found no adverse effects.
There are three main types of pearl oysters of the genus
Pinctada: the “Akoya” pearl oyster called Pinctada fucata, the
“Golden lipped” oyster Pinctada maxima and the “Black
lipped” oyster named Pinctada margaritifera. Mollusc shells are
mainly made up of two layers of calcium carbonate, comprising an outer layer of calcite and an inner layer of aragonite.
Nacre (mother of pearl) in all oyster shells is a calcified structure that forms the lustrous inner layer. It is mainly composed
of aragonite (∼95–97%) tablets oriented in multiple layers,
each surrounded by organic matrix.1,2 This organic matrix
makes up ∼5% of the nacre composition and is mainly comprised of polysaccharides and proteins.3 According to a European Commission report published in 2007 the cosmetic and
toiletries industry in the EU, Japan, China and the US had a
total market value of €136.2 billion.4 The cosmetics industry
maintains its edge by constantly developing novel topical skin
treatments. A popular example is the use of all-natural or
organic ingredients, such as fruit and plant extracts to offer
wrinkle relief that mimics the painful and potentially dangerous side effects associated with invasive chemical remedies.5
Clinically, topical treatments containing, for example, aloe
vera, vitamin C, corticosteroids and tacrolimus are used with
a
School of Chemistry and Biochemistry, The University of Western Australia, Crawley,
Australia
b
Pearl Technology Pty Ltd, Geraldton, Australia
c
Burn Injury Research Unit, School of Surgery, The University of Western Australia,
Crawley, Australia
d
School of Animal Biology, The University of Western Australia, Crawley, Australia
e
School of Chemical and Physical Sciences, Flinders University, Bedford Park,
Australia. E-mail: colin.raston@flinders.edu.au; Tel: +61 82017958
† Electronic supplementary information (ESI) available. See DOI:
10.1039/c4tx00004h
This journal is © The Royal Society of Chemistry 2014
the aim of minimizing scarring.6 Recently, there has been
interest in the cosmetics industry in the use of nacre as a key
ingredient.7 Most of the formulations are reported to either
use powdered pearl shell or powdered nacreous layer shell.
Powdered shell and powdered nacre comprises of both organic
and inorganic components. It is reported that nacre stores in
its mineral-based organic structure a variety of bioactive molecules. Efficacy of this water soluble matrix (WSM) has been
tested in a porcine burn injury model.8 WSM was obtained by
suspending powdered nacre in ultra-pure water and collecting
the supernatant via precipitation of insoluble components by
centrifugation. It was concluded that the active mineral based
organic component has beneficial effects on the skin with
enhanced wound healing.8,9
Nacre has also attracted attention for its potential in supporting bone grafting and bone regeneration. In culture under
physiological conditions, nacre can transform to hydroxyapatite, the phosphorous rich main constituent of the mammalian bone framework.10,11 Nacre and its WSM can also aid in
osteogenic regeneration.9,12–17 High phosphorous rich domains
have been described at the interface between bone and
implants made from Margaritifera shells which are biocompatible, biodegradable and osteoconductive and thus are thought
to promote bone formation.18 Furthermore, nacre powder has
been used as an implantable material for reconstruction and
regeneration of maxillary alveolar ridge bone in humans.19 In
this example, the implanted nacre dissolves gradually and is
eventually replaced by the mature lamellar bone suggesting
that the nacre acts as a biocompatible substrate for bone replacement.19 The water soluble components of the crushed nacre
have also been investigated for their potential in bone regeneration in a similar vein.20,21 Lee et al., demonstrated the wound
healing potential of WSM component in a deep burn porcine
skin model and showed enhanced collagen secretion and
deposition at the injury site resulting in enhanced healing.8 In
another in vivo study using a rat skin incisional injury model,
powdered nacre was implanted between the epidermis and
dermis at the incisional site, with an aim of studying the effect
of nacre on the synthesis of certain constituents of the dermal
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extracellular matrix. It was concluded that implanted nacre
increased collagen synthesis by dermal fibroblasts.22 While
extensive investigations have been carried out in bone, the
evaluation of the biocompatibility of nacre with human skin
cells is lacking. Thus the growing number of cosmetic formulations in the market with nacre as a key ingredient7,23,24
clearly warrants a thorough assessment with human skin cells.
Since scars are also common, and contain cells with a phenotype distinct from normal skin,25 it is also important to test
potential cosmetic ingredients with both cell types.
In the present study, we use nacre from the inner calcified
layer of the shell of Pinctada margaritifera and report the
in vitro toxicity assessment of the material on three cell types
representing both epidermal and dermal layers of human
skin. These were HaCaT cells, a human derived immortalised
keratinocytes cell line, primary human dermal skin fibroblasts
(HDF) and primary human scar fibroblasts (HSF).
Nacre used in the study was gently scraped26,27 from the
inner layer of the shell to avoid the post processing required in
the case of powdered shell. SEM images (Fig. 1) confirmed
that the nacre was composed of pseudo-hexagonal shaped aragonite tablets which have basal plane dimensions of 2–6 µm,
and a thickness of 300–400 nm.28 This structure is characteristic of previously reported nacre, which is a composite
material consisting of alternating layers of mineral tablets separated by thin layers of biomacromolecular “glue”.29,30
Fig. 1 Top view of the scrapped nacre, imaged using scanning electron
microscopy (SEM). Scale bar (a) 1 µM and (b) 2 µM respectively. Sample
was coated with 4 nm platinum prior to imaging.
224 | Toxicol. Res., 2014, 3, 223–227
Toxicology Research
To test the cytotoxicity of nacre, a live/dead assay was
carried out (see ESI S1.4†). Cells were incubated with nacre in
culture media at physiological conditions for 24 h to 72 h and
were then stained for viability using calcein AM/ethidium
bromide I solutions. Viable cells fluoresce green through the
reaction of calcein AM with intracellular esterase, whereas
non-viable cells fluoresce red due to the diffusion of ethidium
homodimer across damaged cell membranes and binding with
nucleic acids.
Fig. 2 shows live cells as the percentage of the total cells in
human primary dermal skin fibroblast (HDF), human primary
scar fibroblast (HSF) and human derived immortalised HaCaT
cell cultures when exposed to various concentrations of nacre
for 24 h and 72 h. Cytotoxicity of nacre was not observed for
any of the concentrations examined in HDF cells (Fig. 2a).
However, interestingly at a concentration of 2.5 mg ml−1 of
nacre (highest concentration tested) there was a significant
reduction in viability at both 24 and 72 hours in the HSF cells
(Fig. 2b). This underlines the importance of testing both scar
and normal skin cell types for cosmetic application. Toxicity
was also observed at a concentration of 0.5 mg ml−1 in HaCaT
cells (Fig. 2c), although this was only observed at 24 hours and
Fig. 2 Cell viability assays showing percentage of live cells in the
culture post incubation with nacre. (a) Human dermal skin fibroblasts
cells, (b) human scar fibroblasts cells and (c) human derived immortalised HaCaT cells were incubated with various concentrations of
scrapped nacre and treated with calcein AM/ethidium bromide I to stain
for live and dead cells. Both live and dead cells were counted using fluorescence microscopy. ‘None’ is the untreated control. Data presented as
average ± SEM (n = 4). Significance was set at *p < 0.05 using bonferroni
test in one way ANNOVA.
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was no longer present at the 72 hour time point. These results
are in line with the results obtained previously where constituents of nacre were shown to promote wound healing in a rat
model22 and deep burn porcine skin.8 In both these studies,
nacre has been shown to promote the recruitment of fibroblasts for restoration and coverage of the injury site while
showing no apparent signs of cytotoxicity. It has also been
shown to promote bone formation when implanted in the
femur of sheep with midshaft hemidiaphysis resection of their
femur in vivo reiterating the non-cytotoxic advantage of
nacre.31
Fibroblasts have been reported to undergo morphological
changes from dendritic to stellate shapes upon exposure to
external cues caused by changes in actin polarisation and
adhesion.32,33 Cell morphology in fibroblasts is known to be
influenced by cytokines such as transforming growth factor β
which can potentially induce polymerisation of globular to
filamentous actin.34 Fibroblast morphology can also be modulated by extracellular matrix architecture during wound
healing via cell–matrix interaction.32 Such morphological
change has been observed in cells undergoing oxidative
stress.35,36 In our study, we found similar changes in fibroblast
morphology for both HDF and HSF cells at the highest concentration of nacre of 2.5 mg mL−1 (Fig. 3). Similar altered mor-
Fig. 3 Cell morphology post calcein AM staining and imaged using
fluorescent microscopy. Cells were treated with various concentrations
of nacre for 24 h, stained and imaged. (i) untreated (control) primary
human dermal skin fibroblasts (HDF), (ii) HDF treated with 0.05 mg mL−1
nacre, (iii) HDF treated with 2.5 mg mL−1 nacre, and (iv), untreated
(control) primary human dermal scar fibroblasts (HSF), (v) HSF treated
with 0.05 mg mL−1 nacre and (vi) HSF treated with 2.5 mg mL−1 nacre.
Scale bar 1 µm.
This journal is © The Royal Society of Chemistry 2014
Communication
phology was also observed for HaCaT cells (see ESI Fig. S1†). It
could be postulated that the high concentration of nacre
induces cellular stress, resulting in changes in the actin cytoskeleton and a more stellate morphology (Fig. 3i, iii, ii and iv).
Cell area was calculated from the fluorescent images shown in
Fig. 3 using Image J software.37 It was found that both HDF
and HSF had significantly larger cell areas ( p < 0.05) when
treated with 2.5 mg mL−1 of nacre (HDF: 2.79 ± 0.13 µm and
HSF: 3.0 ± 0.19 µm respectively) as compared to the nontreated controls (HDF: 1.56 ± 0.08 µm and HSF: 1.54 ± 0.10 µm
respectively) (see ESI Fig. S2†).
Altered fibroblast morphology has been thought to occur in
response to various factors including aging,38 strength of the
extracellular matrix39 or other etiologies that induce mechanical stress on the cell. Changes in morphology also commonly
indicate oxidative as well as mechanical stress.39 Therefore, we
explored whether the morphological changes and increase in
Fig. 4 Reactive oxygen species (ROS) assay showing ROS levels in cells
stressed with various concentrations of nacre for 24 h. No significant
stress was observed as a result of calcium (from nacre) induced oxidative
stress at the concentrations studied. (a) Human dermal skin fibroblasts
cells, (b) human scar fibroblasts cells and (c) human derived immortalised HaCaT cells were incubated with various concentrations of
scrapped nacre for the specified period of time. Cells were then incubated with 2’,7’-dichlorodihydrofluorescein diacetate (DCFH-DA) solution which fluoresce in the presence of reactive oxygen species. ‘None’
is the untreated control. Data presented as average ± SEM (n = 3). Significance was set at *p < 0.05 using bonferroni test in one way ANNOVA.
Toxicol. Res., 2014, 3, 223–227 | 225
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cell area at the highest concentration of nacre in culture was a
result of, or induced oxidative stress in, the cells. Oxidative
stress was tested using the cell permeable fluorogenic probe
2′,7′-dichlorodihydrofluorescein
diacetate
(DCFH-DA).
DCFH-DA is taken up by cells and is deacetylated by cellular
esterases to non-fluorescent 2′,7′-dichlorodihydrofluorescein
(DCFH) which is rapidly oxidised by reactive oxygen species
(ROS) to highly fluorescent 2′,7′-dichlorodihydrofluorescein
(DCF). The fluorescent intensity is proportional to the ROS
levels within the cytosol (see ESI S1.5†). Cell responsiveness to
the assay was carried out by stressing the cells with the H2O2
solution provided in the kit which was also used to generate
the calibration curve (see ESI Fig. S3†). No changes in levels of
reactive oxygen species were observed in any cell type at any
concentration of nacre (Fig. 4). This is important as oxidative
stress is known to be a significant contributor to skin damage
and excessive scarring in previous studies.40 It has been
known that cells alter their morphology depending on their
environment.41,42 It is therefore hypothesized that the altered
fibroblast morphology in the present case is mainly due to the
regulation of cell motility through geometrical constraint in
the presence of nacre. Indeed, it has been previously reported
that when cells probe their physical surroundings, they acquire
mechanical information or signals that help to determine the
direction of migration, with a consequential change in cell
morphology.43
Conclusions
We have established the biocompatibility of nacre using three
human cell types representing the two primary layers of
human skin, using immortalised keratinocytes from the epidermal layer and two primary human dermal cell cultures. The
nacre used in the present study showed limited cytotoxicity at
high concentrations in scar derived cells, with the morphology
of the cells significantly changed by exposure at such concentrations of nacre. No apparent oxidative stress was evident in
any of the cell types. Overall, the data support the use of low
concentrations of nacre in aesthetic formulations, with the
potential for high concentrations to cause changes in skin
and/or scar cells which may have impact on efficacy.
The authors would like to thank the Australian Research
Council and Pearl Technologies for funding the work under
the grant number LP100100812. The authors would also like
to acknowledge the Australian Microscopy & Microanalysis
Research Facility at the Centre for Microscopy, Characterization & Analysis, The University of Western Australia, funded
by the University, State and Commonwealth Governments.
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Electronic Supplementary Material (ESI) for Toxicology Research.
This journal is © The Royal Society of Chemistry 2014
Evaluating the effects of nacre on human skin and scar cells in
culture
Vipul Agarwal,a Edwin S. Tjandra,a K. Swaminathan Iyer,a Barry Humfrey,b Mark Fear,c
Fiona M. Wood,c Sarah Dunlopd and Colin L. Rastone, *
Supporting Information
S1. Experimental
S1.1
Materials
Pinctada margaritifera shells were provided by Pearl Technologies Pty Ltd, which were grown in the waters of the Abrolhos
Islands, Western Australia. The decalcified organic conchiolin layer was removed by wet sand blasting of the shell followed
by gentle brushing to remove any dust particles that might otherwise contaminate the samples. Inner nacreous layer was then
scraped using a surgical scalpel and stored at room temperature for a maximum of 2 weeks.
S1.2
Scanning Electron Microscopy
Scraped nacre above from the inner layer of the shell was mounted on SEM stubs (ProSciTech, Cat.# G040). Samples were
coated with 4 nm of platinum. Images were taken using scanning electron microscope (Zeiss 1555, VP-FESEM) at 4-5 kV at
30µm aperture. Images were analyzed with the image analysis software ImageJ (NIH).1
S1.3
Cell culture
A human derived immortalized keratinocyte cell line, HaCaT 2 and two human primary dermal (fibroblast) cell cultures from
normal skin and normal scar were used. All three cell types were cultured in Dulbecco’s Modified Eagle’s Medium
(DMEM/F12 - GlutaMAX; Invitrogen Gibco) supplemented with 10% fetal bovine serum (FBS; Invitrogen Gibco) and 1%
penicillin/streptomycin (Invitrogen Gibco). The cells were incubated at 37°C in an atmosphere of 5% CO2. Primary cells
used were between the passages 7-10. Nacre was sterilized using UV sterilization technique in the tissue culture hood for 15
min prior its dissolution in media. Fresh nacre solution was prepared before every experiment.
S1.4
Cell Viability
Cell viability was determined using a LIVE/ DEAD Viability/Cytotoxicity Kit (Invitrogen, UK) which measures the
membrane integrity of cells,3, 4 as per manufacturer’s protocol. In brief, 20000 cells were seeded in each well in a 24 well
plate and treated with scraped nacre at various concentrations in cell culture media (DMEM F-12 containing 10% FBS and
1% Penicillin/ Streptomycin) and incubated for 24 h or 72 h in the humidified incubator at 37°C with 5% CO2. At the
stipulated time (24h and 72h), cells were washed with PBS (3 times) and then stained with calcein (100 µL, 1µM)/ ethidium
bromide (100 µL, 2 µM) in PBS and incubated in the humidified incubator at 37°C and 5% CO2 for 30 min. Images were
captured using an Olympus IX71 inverted microscope with a 20 x objective with fixed exposure time. Both live and dead
cells were counted using Image J with cell counter plug in. Experiments were performed in triplicate. Minimum of fifty
images were captured per condition.
S1.5
Reactive Oxygen Species (ROS)
ROS was measured using the ROS assay kit (Oxiselect ROS assay kit, Cat.# STA 342, Cell Biolabs) following
manufacturer’s protocol. In brief, 6000 cells were seeded in a 96 well plate and incubated in the humidified incubator at
37°C with 5% CO2 for 24h. Next day, wells were washed with PBS (3 times) and incubated with 2’, 7’dichlorodihydrofluorescin diacetate (DCFH-DA) solution (0.1 x, 100 µL/ per well) for 1 h in the humidified incubator at
37°C with 5% CO2. DCFH-DA solution was removed and the wells washed with 3 x PBS. Cells were then treated with
scraped nacre solution in culture media at a specified concentration for 24 h, wells were washed with 3 x PBS and cells were
lysed using the lysis buffer provided (1 x, 100µL/ per well, incubated for 20 min. at room temperature) before reading the
plate at 480 nm excitation/ 530 nm emissions using the plate reader. Experiments were performed in triplicate.
S1.6
Cell Body Area
Cell size was measured using Image J software (NIH).1 A minimum of 25 cells were randomly selected from the
fluorescence images and their area was measured. Values reported as mean ± standard error mean.
S1.7
Statistics
The results for cell viability, ROS experiments and cell area are expressed as mean ± standard error mean (SEM) and
analysed by analysis of variance (ANOVA). Significance was evaluated using Bonferroni and Turkey’s post-hoc analysis
and set at 95% confidence (p < 0.05).
Supporting Figures
Figure S1: Cell morphology post calcein AM/ ethidium bromide staining and imaged using fluorescent microscopy. HaCaT cells were
treated with various concentrations of nacre for 24h, stained and imaged. a) Untreated (control), b) HaCaT cells treated with 2.5 mg/mL
nacre. Scale bar 1µm.
Figure S2: Cell area size showing increase in the cell area post incubation with nacre. Cell area was measured from the fluorescent images
of live cells taken for viability assay. ‘None’ is the untreated control. Data presented as average ± SEM (n>25). Significance was set at * p <
0.05 using bonferroni post hoc test in one way ANNOVA
Figure S3: Reactive oxygen species (ROS) assay showing increase in ROS levels in cells stressed with various concentrations of H2O2 in
dose dependent manner. Human dermal skin fibroblasts cells were incubated with various concentrations of H2O2 for the specified period of
time to generate the standard curve. Cells were then incubated with 2’, 7’- dichlorodihydrofluorescin diacetate (DCFH-DA) solution which
fluoresce in the presence of reactive oxygen species. ‘None’ is the untreated control. Data presented as average ± SEM (n=3).
References
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RSC Advances
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COMMUNICATION
Cite this: RSC Advances, 2013, 3, 1009
Received 4th October 2012,
Accepted 12th November 2012
DOI: 10.1039/c2ra22402j
www.rsc.org/advances
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Biogenic production of palladium nanocrystals using
microalgae and their immobilization on chitosan
nanofibers for catalytic applications3
Ela Eroglu,ab Xianjue Chen,a Michael Bradshaw,a Vipul Agarwal,a Jianli Zou,a Scott
G. Stewart,c Xiaofei Duan,d Robert N. Lamb,d Steven M. Smith,*b Colin L. Raston*a
and K. Swaminathan Iyera
Spherical palladium nanocrystals were generated from aqueous
Na2[PdCl4] via photosynthetic reactions within green microalgae
(Chlorella vulgaris). Electrospun chitosan mats were effective for
immobilizing these biogenic nanocrystals, as a material for
recycling as a catalyst for the Mizoroki–Heck cross-coupling
reaction. This photosynthetically-driven metal transformation
system can serve as a good candidate for an environmentallyfriendly method for the synthesis of metal nanocatalysts.
Biogenic systems can control the phase, structure, and topography
of inorganic nanocrystals with a level of precision similar to
synthetic approaches.1 Various organisms such as bacteria,2
yeast,3 fungi4 and algae5 are capable of processing a wide range
of metals. Such bioprocessing has been effectively used in the
reduction of environmental pollution and also for the recovery of
metals from waste.6 The formation of metal nanoparticles in the
presence of microorganisms primarily involves the reduction of
metal ions in solution by enzymes generated by microbial cell
activities, which can be intracellular or extracellular. This depends
on the nanoparticle crystallization site, as being either inside the
cell or on the cell surface.7,8 When considering palladium, the
biosynthesis of nanoparticles of the metal have been reported in
bacteria (Desulfovibrio desulfuricans, Shewanella oneidensis, Bacillus
sphaericus),9 cyanobacteria (Plectonema boryanum UTEX 485),10,11
plants,12–14 and viruses (e.g., tobacco mosaic virus).15,16 Herein we
report a biogenic synthesis of palladium nanocrystals in the
presence of Chlorella vulgaris with the photoautotrophic microa
Centre for Strategic Nano-Fabrication, School of Chemistry and Biochemistry, The
University of Western Australia, M313, 35 Stirling Highway, Crawley, WA 6009,
Australia. E-mail: colin.raston@uwa.edu.au; Fax: +61 8 6488 8683;
Tel: +61 8 6488 3045
b
ARC Centre of Excellence in Plant Energy Biology, The University of Western
Australia, M313, 35 Stirling Highway, Crawley, WA 6009, Australia.
E-mail: steven.smith@uwa.edu.au; Fax: +61 8 6488 4401; Tel: +61 8 6488 4403
c
School of Chemistry and Biochemistry, The University of Western Australia, 35
Stirling Highway, Crawley, WA 6009, Australia
d
Surface Science & Technology Group, School of Chemistry, The University of
Melbourne, VIC 3010, Australia
3 Electronic supplementary information (ESI) available: Experimental details,
characterization, and an additional TEM image. See DOI: 10.1039/c2ra22402j
This journal is ß The Royal Society of Chemistry 2013
algal metabolism most likely providing the necessary reducing
agents. Studies on the biogenic reduction of palladium by various
organisms are usually limited to the generation of Pd(0). We have
used a non-toxic and environmentally-available microorganism to
achieve this, and then collected the particles from the liquid
culture via immobilizing on electrospun chitosan nanofibers. In
addition, we have established the utility of this composite material
as a recyclable catalyst which is functional even at very low loading
rates. The overall integrated approach is without precedent, as are
the individual components.
Algae are a large group of photosynthetic organisms that are
ubiquitous in various aquatic habitats including sea, freshwater,
wastewater and also in moist solids.17 Microalgal remediation has
been reported for several metal ions,18,19 and involves both
intracellular (polyphosphates) and extracellular (polysaccharides
on algal cell wall) metal binding groups.20–22 Based on these
observations, we initially evaluated the ability of a common green
microalga, Chlorella vulgaris, to reduce palladium(II) to Pd(0)
starting with a solution of Na2[PdCl4]. Various concentrations of
Pd salt were investigated (100, 50, 25, 12.5, 0 mg L21) in order to
detect their effects on cell viability. The cells were grown in
solutions of Na2[PdCl4] added to algal freshwater media (MLA
media)23 at 25 uC and a rotation speed of 120 rpm, under diurnal
illumination of 16 h light/8 h dark cycles, with the total chlorophyll
content used to validate the viability of the cell cultures.24
Accumulation of total chlorophyll pigment (Chl a + Chl b) as a
function of growth time was monitored (Fig. 1a), and the
threshold concentration for toxicity on the cells established
amongst 25–50 mg L21. Within the first five days, the culture
flasks containing less than this concentration clearly showed an
increase in cellular growth with solution retaining the green color
which is associated with the chlorophyll content of the cells. In
contrast, higher concentrations of Na2[PdCl4] resulted in the color
fading and loss of cell viability (Fig. 1b). The solution containing
25 mg L21 of Na2[PdCl4] had the highest amount of palladium
precursor while showing little reduction in growth rates compared
to the control. Thus, it was chosen as the prototype for further
experiments.
RSC Adv., 2013, 3, 1009–1012 | 1009
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Communication
Fig. 1 (a) Total amount of accumulated chlorophyll (Chl a + Chl b) measured as a
function of growth time for Chlorella vulgaris cultures in the presence of various
concentrations of Na2[PdCl4] (100, 50, 25, 12.5 mg L21), and control cultures in the
absence of Na2[PdCl4] (0 mg L21, hollow circles), with the total volume of solution in
all flasks at 40 mL during the growth experiments. (b) Chlorella vulgaris culture flasks
containing various concentrations of Na2[PdCl4].
Transmission electron microscopy (TEM) of the liquid culture
media revealed the presence of crystalline spherical palladium
nanoparticles with an average particle size of around 7 nm, with a
range of 2 to 15 nm in diameter (Fig. 2a). The corresponding
electron diffraction pattern further confirmed the presence of
elemental Pd nanocrystals with a characteristic face-centered cubic
(fcc) structure and average d-spacing values of 0.22, 0.19, 0.14, and
0.12 nm for the (111), (200), (220) and (311) planes, respectively
(Fig. 2b). High resolution TEM analysis further confirmed the
presence of single palladium nanocrystals (Fig. 2c and d).
Reducing agents produced by photosynthesis are the key
components for the reduction of Pd(II) into Pd(0) nanoparticles
(Scheme 1). Photosynthetic processes in green microalgae can take
place either under oxygenic or anoxic environments.17,25 Green
algae have chlorophyll-a and chlorophyll-b pigments, and can
Fig. 2 (a) TEM image of palladium nanoparticles precipitated in the four week old
microalgae solution (scale bar: 20 nm). (b) Selective area electron diffraction pattern
of the palladium nanoparticles in 2(a), giving d-spacings of 0.22, 0.19, 0.14, and 0.12
nm which correspond to 111, 200, 220, and 311 reflections, respectively, for Pd(0).
(c) High resolution TEM image of palladium nanoparticles. (d) Pd(0) nanocrystals
with (111) lattice spacings of around 0.22 nm. Inset shows the Fast Fourier
Transform (FFT) pattern corresponding to the area shown within the yellow
rectangle in 2(c), indicating the crystal structure of palladium nanoparticles.
1010 | RSC Adv., 2013, 3, 1009–1012
RSC Advances
Scheme 1 Palladium nanoparticle synthesis by photosynthetic green microalgae,
and their uptake on an electrospun chitosan mat for use as a catalyst in Mizoroki–
Heck reactions. The left stage shows a combination of mechanisms taking place
within the photosynthetic organisms, resulting in the production of reducing
agents.17,25,26 (ADP: adenosine diphosphate, ATP: adenosine triphosphate, Fd:
ferredoxin, NADP+: oxidized form of nicotinamide adenine dinucleotide phosphate,
NADPH: reduced form of nicotinamide adenine dinucleotide phosphate, PGA:
phosphoglycolic acid, RuBisCO: rubilose biphosphate carboxylase). NADPH is likely
one of the main reducing agents for the reduction of Na2[PdCl4], which is partially
oxidised as a result of aerobic culture conditions.
accomplish oxygenic (oxygen-evolving) photoautotrophic reactions
while using H2O as an electron donor.17 Oxygenic photoautotrophic processes include two sets of photochemical reactions: (i)
light reactions conserving chemical energy in the form of
adenosine triphosphate (ATP), and the reduced form of nicotinamide adenine dinucleotide phosphate (NADPH), and (ii) ‘dark’
reactions in which CO2 is reduced to organic compounds using
ATP and NADPH. Light reactions have two separate sets of
photosystems (PS), namely PS I and PS II, in which the electron
transfer follows a Z-scheme between these two photosystems. PS II
is mainly responsible for the splitting of water (H2O A KO2 + 2e2
+ 2H+) as a first stage of cyclic electron flow. While electron flow
follows the Z-scheme between PS II and PS I, a proton motive force
is created and used for the generation of ATP. CO2 is fixed by the
enzyme ribulose bisphosphate carboxylase (RuBisCO) and reduced
in the Calvin cycle using NADPH.17 Reducing equivalents can be
exported from the chloroplast in the form of carbon metabolites,
particularly triose phosphates (triose-P).26 Their oxidation by
dehydrogenases in the cytosol can also generate NADH and
NADPH for other reduction reactions. We postulate that such
reducing agents promote the reduction of Pd(II) into crystalline
Pd(0) nanoparticles, which is further partially oxidized due to the
aerobic culture conditions. An hypothesis for the reduction of
palladium(II) by photoheterotrophic bacteria (Rhodobacter sphaeroides) cultures involves the reduced electron carriers (such as
ferredoxin, NADH, and FADH) as the electron donors.27
Consistent with this hypothesis, we found that addition of
NADPH to the MLA growth medium23 resulted in reduction of
Na2[PdCl4] to produce Pd nanocrystals in solution which were
effectively recovered on exposure to cross-linked electrospun
chitosan nanofiber mats (see ESI3, Fig. S1).
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RSC Advances
Chitosan is derived from chitin present in the exoskeletons
of crustaceans,28 and is an attractive, renewable feedstock
which is effective in binding a number of metal ions,29 and
effective as a support for Pd, Ni, Cu, Cr and Zn based catalysts.30
Cross-linked chitosan mats (3 6 2 cm rectangular shape)
generated using electrospinning31,32 were placed into a four
week old microalgal culture containing 25 mg L21 Na2[PdCl4]. A
control experiment involved exposing chitosan mats to 25 mg
L21 Na2[PdCl4] solution in the absence of microalgae. Chitosan
mats were kept in the solution for about two weeks as to achieve
sufficient particle adsorption from the liquid media. Scanning
electron microscopy (SEM) revealed the nanofiber structure of
the chitosan mats with an average diameter of 100 nm, which
was maintained after exposure to the growth media (Fig. 3a). It
is noteworthy that there was a detectable difference in color of
the recovered chitosan mats from the microalgae and the
control solutions. The initial light yellow chitosan mat turned
dark brown after being incubated with algae (Fig. 3b), whereas
no color change was observed for the control experiment
without algae (Fig. 3c). The darker color is considered to be
consistent with the immobilization of the reduced Pd nanocrystals from the growth media, and corresponds to approximately 1.03 wt% Pd loading on each rectangular piece of
electrospun chitosan mat (3 6 2 cm), which was established
using inductively coupled plasma-optical emission spectroscopy (ICP-OES). X-ray photoelectron spectrometric (XPS)
analysis of the Pd loaded chitosan mats showed a dominating
peak in the PdO 3d5/2 region (binding energy of 337.9 eV) and
also the presence of the Pd(0) peak in the Pd 3d5/2 region
Fig. 3 (a) SEM image of the as-prepared electrospun chitosan nanofibers before
being exposed to palladium solution (scale bar: 1 mm). Mats of this material placed
into 4 week old 25 mg L21 Na2[PdCl4] solutions with and without microalgae which
were collected after two weeks are shown in (b) and (c) respectively. Curve-fitted Pd
3d XPS spectra for (d) as collected and (e) etched internal surface of a chitosan mat
under argon ions for 60 s, both showing Pd(0) (blue curve) and PdO (red curve).
This journal is ß The Royal Society of Chemistry 2013
Communication
(binding energy of 336.0 eV) (Fig. 3d). Analysis of the sample
following argon ion etching primarily showed a dominant Pd(0)
peak in the Pd 3d5/2 region (binding energy of 335.9 eV) with an
additional lower peak of PdO in the 3d5/2 region (binding energy
of 337.7 eV) (Fig. 3e). Thus the XPS data establish that the
surface of the palladium nanoparticles is partially oxidized
under the aerobic culture conditions. Slight oxidation was also
reported when palladium containing samples were exposed to
an oxygen containing atmosphere before XPS analysis.33
Chlorine was not detected during these XPS analyses, indicating
the absence of palladium(II) chloride species. The binding
energies of both Pd 3d5/2 and PdO 3d5/2 were found to be slightly
higher than the values for pure Pd metal,34 and PdO samples.35
This slight shift is reported to be common when the Pd/PdO
nanoparticles are embedded in an insulating substrate,36 which
is the chitosan mat in the present case. In their study,
Schildenberger and his colleagues also reported similar peaks
for Pd 3d5/2 (336.0 eV) and PdO 3d5/2 (337.9 eV) due to the
isolated arrangement of metal nanoparticles on the layers of
oxidized silicon wafers.36
The utility of these palladium loaded chitosan mats were tested
as catalyst supports for the standard Mizoroki–Heck reaction.37,38
Six dried mats (a total of 0.23% mol Pd per mol of iodobenzene)
were introduced into a solution containing iodobenzene, butyl
acrylate, triethylamine and dimethylformamide (DMF). The
reaction temperature was kept constant at 80 uC for 16 h, and
after each reaction, the supported catalyst was recovered for
recycling studies after washing with DMF under nitrogen gas to
avoid any oxygen induced regrowth of Pd(0) nanoparticles.39 The
catalyst can be recycled at least four times with quantitative
conversion for each cycle, with the conversion yields of: 68% (1st
cycle), 62% (2nd cycle), 45% (3rd cycle) and 36% (4th cycle) by
weight. For the control comparison, a mat exposed to a Na2[PdCl4]
solution without algae resulted in a conversion yield of only 5%.
We have previously reported that a chitosan mat containing
palladium nanoparticles generated by reduction of pre-absorbed
Na2[PdCl4] is also an effective catalytic support with no appreciable
leaching of the metal.31 Conventional palladium catalysts usually
require a 1–5 mol% loading rate for effective Mizoroki–Heck crosscoupling reactions.40 Our biogenic palladium nanocatalysts have
high catalytic activity (68%) with respect to commercial material
(5%), even at low palladium loadings (0.23 mol%), which is
significant for applications in the fine chemical industries.
In conclusion we have established a biogenic synthesis of
palladium nanocrystals in the presence of Chlorella vulgaris and
their subsequent immobilization on an electrospun chitosan mat
as a novel support for application in catalysis. In addition, we have
demonstrated that NADPH involved in the photoautotrophic
microalgal metabolism is likely to play a role in the biogenic
synthesis. Utilization of easy-to-grow, nontoxic and environmentally-available microalgae for the synthesis of palladium has
potential for the development of green chemistry processes for
other metals.
RSC Adv., 2013, 3, 1009–1012 | 1011
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Communication
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Acknowledgements
We kindly acknowledge support of this work by the Australian
Research Council (ARC) and internal grants of The University
of Western Australia. The microscopy analysis was carried out
using the facilities in the Centre for Microscopy,
Characterization and Analysis (The University of Western
Australia). The authors also would like to thank Dr L. Byrne for
his kind help during the NMR analysis.
Notes and references
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7 K. Simkiss and K. Wilbur, Biomineralization. Cell Biology and
Mineral Deposition, Academic Press, Inc., San Diego, 1989.
8 S. Mann, Biomimetic Materials Chemistry, VCH, Weinheim, New
York, 1996.
9 P. Yong, N. A. Rowson, J. P. G. Farr, I. R. Harris and L.
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Y. Su, J. B. Opiyo, L. Hong, Y. Wang, N. He and L. Jia, J.
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L. Raston, Chem. Commun., 2011, 47, 12292–12294.
32 K. Ohkawa, D. I. Cha, H. Kim, A. Nishida and H. Yamamoto,
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This journal is ß The Royal Society of Chemistry 2013
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Supplementary Information
Biogenic production of palladium nanocrystals using microalgae and
their immobilization on chitosan nanofibers for catalytic applications
Ela Eroglua,b, Xianjue Chena, Michael Bradshawa, Vipul Agarwala , Jianli Zoua, Scott G. Stewartc,
Xiaofei Duand, Robert N. Lambd, Steven M. Smithb*, Colin L. Rastona*, and K. Swaminathan Iyera
a
Centre for Strategic Nano-Fabrication, School of Chemistry and Biochemistry, The University of Western
Australia, Crawley, WA 6009, Australia, E-mail: colin.raston@uwa.edu.au
b
ARC Centre of Excellence in Plant Energy Biology, The University of Western Australia M313, 35 Stirling
Highway, Crawley, WA 6009, Australia, E-mail: steven.smith@uwa.edu.au
c
School of Chemistry and Biochemistry, The University of Western Australia, 35 Stirling Highway, Crawley,
WA 6009, Australia
d
Surface Science & Technology Group, School of Chemistry, The University of Melbourne, VIC 3010, Australia
S1. Biosynthesis of Palladium Nanoparticles
Wild type Chlorella vulgaris cultures (from the Australian National Algae Culture Collection
at CSIRO, Tasmania) were used as the green microalgae source for the biosynthesis of
palladium. Sterile algal freshwater media (MLA media)[1] containing standard micronutrients,
nitrate, phosphate, carbonate buffer and vitamins was used as the substrate source. The
microalgae cultures were mixed with Na 2 [PdCl 4 ] solution at various concentrations (100, 50,
25, 12.5, 0 mg/L), towards reaching an initial total-chlorophyll content of around 1.8 mg/L
(Figure 1a).
Na 2 [PdCl 4 ] solution was prepared by dissolving Na 2 [PdCl 4 ] powder (Sigma-Aldrich) in
distilled and sterilized water overnight, and subsequently filtered through a sterile Pall®
Acrodisc® 32 mm syringe filter (0.2 µm membrane) for further sterilization. Concentrations
of Na 2 [PdCl 4 ] in microalgae solutions are reported here as the initial concentrations
measured by the gravimetric analysis before the filtration process. The experiments were
conducted under batch conditions and cyclic diurnal conditions (16 h light/8 h dark) at a
constant temperature (25 °C). Algae cultures (total liquid volume of 40 mL) were grown in
250 mL Erlenmeyer flasks, under continuous cool-white fluorescent illumination at an
incident intensity of around 200 µmol photons m-2s-1(PAR) upon orbital shaking (Thermoline
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Scientific) at 120 rpm.
Algal growth was investigated by measuring the total chlorophyll content (Chl a + Chl b)
with respect to time, following the spectrophotometric method involving methanol
extraction.[2] Chlorophyll content was also used to validate the viability of the cell cultures.
All experiments were conducted in triplicates, with the standard deviation of each value given
in the form of error bars within the related figure.
S2. Characterization
A JEOL 2100 TEM instrument operating at 80 kV was used for determining the size and
morphology of chitosan nanofibers and palladium nanoparticles. Samples were prepared by
inserting the solutions on top of carbon-coated 200 mesh copper grids and allowing them to
dry. High resolution images were obtained using a JEOL 3000F instrument operating at 300
kV. A Zeiss 1555 VP-FESEM with a 3 kV accelerating voltage was used to image samples
coated with platinum (~3nm). XPS data was acquired using a VG ESCALAB220i-XL X-ray
Photoelectron Spectrometer equipped with a hemispherical analyzer. The incident radiation
was monochromatic Al Kα X-rays (1486.6 eV) at 220 W (22 mA and 10kV). Survey (wide)
and high resolution (narrow) scans were taken at analyzer pass energies of 100 eV and 50 eV,
respectively. Survey scans were carried out over 1200-0 eV binding energy range with 1.0 eV
step size and 100 ms dwell time. Narrow high resolution scans were run over a 20 eV binding
energy range with 0.05 ev step size and 250 ms dwell time. Base pressure in the analysis
chamber was 4.0x10-9 mbar and during sample depth profile analysis 1.5x10-7 mbar. A low
energy flood gun (~6 eV) was used to compensate the surface charging effect. Argon ions at
3 keV beam energy were used to sputter off approximately 18 nm surface layers at a rate of
~3 Angstrom/second. The ion source gave a crater of approximately 3x3 mm. The energy
calibration was referenced to the C 1s peak at 284.7 eV.
S3. Cross Coupling Reactions
A rectangular piece of electrospun chitosan mat (3 x 2 cm per each piece) was placed into
four week old 25 mg/L Na 2 [PdCl 4 ] containing microalgae solutions (total volume: 40 mL per
flask) for two weeks. The amount of palladium was established using inductively coupled
plasma - optical emission spectroscopy (ICP-OES), at ~1.03 % w/w per each rectangular mat
(3x2 cm). ICP-OES analysis was acquired with ARL-3520B sequential scanning ICP-OES,
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with a 20mm torch, 1200W incident power, and a MDSN nebuliser. For the Mizoroki-Heck
reactions iodobenzene (55 µL; 0.49 mmol), butyl acrylate (85 µL; 0.59 mmol), triethylamine
(171 µL; 1.23 mmol) in dimethylformamide (DMF, 0.5 mL) were added to 6 hybrid
palladium nanoparticle-chitosan mats with a total of 0.23% mol Pd per mol iodobenzene.
The reaction mixture was heated to 80°C for 16 hours, whereupon the catalyst mat was
filtered and recovered for its recycling after washing with DMF five times under nitrogen
gas. Quantitative conversion yields were assessed gravimetrically upon the confirmation of
the final product (butyl cinnamate) using 1H and
13
C NMR spectroscopy (Varian® 400
NMR).
S4. Electrospinning of Chitosan
Pd was collected by electrospun chitosan mat following the electrospinning procedure
optimized by Bradshaw et al. (2011)[3], which was slightly modified from the original
protocol given by Ohkawa et al. (2004)[4]. The variables of electrospinning processes were as
follows: (i) syringe pump speed: 0.1 mm/min, (ii) voltage: 18 kV, (iii) distance between the
target and the tip of the syringe: 11 cm, (iv) target speed: 1 m/ min, (v) traverse speed: 0.5
cm/min.[3] Chitosan (2-amino-2-deoxy-(1-4)-β-D-glucopyranose), 75-85% deacetylated
(Sigma-Aldrich), powder (6% wt) was mixed with TFA (trifluoroacetic acid): DCM
(dichloromethane) solution (70:30 v/v), and stirred overnight for its complete dissolution.
These ratios were taken to be optimum for attaining ultrafine chitosan nanofibers (Bradshaw
et al. 2011).[3] TFA (trifluoroacetic acid) and DCM (dichloromethane) were obtained from
Chem Supply. Before the electrospinning process, the mixture was sonicated for 15 min and
5.4% v/v glutaraldehyde (25% in H 2 O, Sigma-Aldrich) was immediately added to the
chitosan/TFA/DCM solution for an effective crosslinking. Electrospinning of this crosslinked solution created insoluble nanofibers. Two rectangular pieces of electrospun chitosan
mat (3 x 2 cm per each piece) were placed into one flask of (four weeks old) 25 mg/L
Na 2 [PdCl 4 ] containing microalgae solutions (total volume of 40 mL per flask) and left inside
for two weeks. The amount of palladium was established by using ICP-OES, yielding ∼1.03
% w/w per each rectangular mat (3x2 cm).
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S5. Reduction of Palladium with NADPH
Figure S1. High resolution TEM images of Pd(0) nanoparticles after the inoculation of
Na 2 [PdCl 4 ] precursor with an excess amount of NADPH, dissolved inside MLA algal
growth-media,[1] indicating high crystallinity with (111) lattice constant of around 0.22 nm
(inset: FFT pattern corresponding to the area shown with red rectangle).
References
1. R.A. Andersen, in Algal Culturing Techniques, Elsevier Academic Press, 2005.
2. H.K. Lichtenthaler and C. Buschmann, in Current protocols in food analytical chemistry, eds:
R.E. Wrolstad RE, John Wiley & Sons Inc., New York, 2001, pp. F4.3.1- 8.
3. M. Bradshaw, J. Zou, L. Byrne, K.S. Iyer, S.G. Stewart, C.L. Raston, Chem. Commun. 2011, 47,
12292-12294.
4. K. Ohkawa, D. I. Cha, H. Kim, A. Nishida, H. Yamamoto, Macromol. Rapid Commun. 2004, 25,
1600–1605.
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Nitrate removal from liquid effluents using microalgae immobilized on
chitosan nanofiber mats
Ela Eroglu,a,b Vipul Agarwal,a Michael Bradshaw,a Xianjue Chen,a Steven M. Smith,*b Colin L. Raston*a
and K. Swaminathan Iyera
Received 25th June 2012, Accepted 8th August 2012
DOI: 10.1039/c2gc35970g
Mats of electrospun chitosan nanofibers were found effective
in immobilizing microalgal cells. These immobilized microalgal cells were also used as a durable model system for
wastewater treatment which has been demonstrated by the
removal of around 87% nitrate from liquid effluents ([NO3−N]initial 30 mg L−1). Virtuous nitrate removal rates were
derived by superimposing physicochemical adsorption and
biological nutrient consumption phenomena for chitosan and
microalgae, respectively.
Wastewater treatment focuses on eliminating unwanted chemicals and/or biological impurities from contaminated water. In
general the treatment methods are mainly based on the separation
of pollutants from the wastewater with a requirement for a
further processing stage to eliminate these pollutants.1 Integrated
wastewater treatment processes are important in eliminating
undesired species, ideally converting them into valuable products. As relatively recent bioprocesses, algal cultivation in
wastewaters has a combination of several advantages such as
wastewater treatment and simultaneous algal biomass production, which can be further exploited for biofuel production,
food additives, fertilizers, cosmetics, pharmaceuticals, and other
valuable chemicals.2 However, there are inherent difficulties
associated with algae-based bioprocessing in the harvesting,
dewatering and processing of the algal biomass.
Immobilization of the cells on solid surfaces confer advantages over free cells in suspension, namely the immobilized cellular matter occupy less space, require smaller volume of growth
medium, are easier to handle, and can be used repeatedly for
product generation.3,4 Moreover, immobilization can also
increase the resistance of cell cultures to harsh environmental
conditions such as salinity, metal toxicity and variations in
pH.3,4 Entrapment is one of the most common immobilization
methods which involves capturing the cells in a three dimensional gel matrix, made of polymeric materials or inorganic
spheres.4 Both synthetic polymers (e.g., acrylamide,
a
Centre for Strategic Nano-Fabrication, School of Chemistry and
Biochemistry, The University of Western Australia, M313, 35 Stirling
Highway, Crawley, WA 6009, Australia. E-mail: colin.raston@uwa.edu.au;
Tel: +61 8 6488 3045; Fax: +61 8 6488 8683
b
ARC Centre of Excellence in Plant Energy Biology, The University of
Western Australia M313, 35 Stirling Highway, Crawley, WA 6009,
Australia. E-mail: steven.smith@uwa.edu.au; Tel: +61 8 6488 4403;
Fax: +61 8 6488 4401
2682 | Green Chem., 2012, 14, 2682–2685
polyurethane, polyvinyl) and natural polymers (e.g., collagen,
agar, cellulose, alginate, carrageenan) have been used for this
purpose.5 Several studies have been reported on wastewater treatment involving the entrapment of microalgae cultures inside
alginate beads, porous glass, and several synthetic polymers.6–8
However, most attempts to immobilize viable algae cells inside
such insoluble materials have limitations, with the encapsulating
materials having volume/surface ratios usually orders of magnitude larger than thin films. As a consequence, algal viability is
mostly reported to decrease which relates to the need for the
nutrients or reactants to diffuse far into the material to reach the
algal cells.5 We have developed a new technique to overcome
these problems using electrospun nanofiber mats as the matrix
for immobilizing the algal cells, with an overall strategy to
combine wastewater treatment processing with algal harvesting
in a single process.
Electrospun nanofibers of chitosan were employed as a
polymer/matrix support for green microalgae in the current
study. Chitosan is composed of D-glucosamine and N-acetyl-Dglucosamine, and is formed by the deacetylation of chitin (β-Nacetyl-D-glucosamine polymer).9 Chitin is usually extracted from
the exoskeletons of crustaceans (e.g., crab, lobster and shrimp)
and even from the cell walls of fungi.9,10 Chitosan is non-toxic
and biodegradable, and can be used as an animal feed.9 Furthermore, it can be used as a coagulant for wastewater treatment and
for the recovery of waste sludge.11
The removal of nitrate ions is regulated by law which relates
to its hazardous effects on human health and the environment.
Several methods have been reported for the removal of nitrate
from water bodies, including biological denitrification,12 chemical reduction,13 electrodialysis,14 and a combined bioelectrochemical/adsorption process.15 In this study, we aimed to
superimpose the treatment efficiencies of microalgae in reducing
nitrate and electrostatic binding of the nitrate ion by chitosan, i.e.
combining biological and chemical processing.
Chlorella vulgaris cultures were used as the green microalgae
which were originally obtained from the Australian National
Algae Culture Collection at CSIRO, Tasmania. Cell growth was
carried out under artificial diurnal-illumination (16 h light/8 h
dark cycle) at around 22 °C. Electrospun chitosan mats were
fabricated following the procedure optimized in previous
studies.16,17 This involved dissolving chitosan powder (6 wt%)
in a mixture of trifluoroacetic acid (TFA) and dichloromethane
(DCM) (70 : 30 v/v), with 5.4% (v/v) addition of glutaraldehyde
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solution (25% in H2O) immediately prior to electrospinning.17
This was deemed necessary to cross-link the chitosan to avoid
polymer breakup and dissolution. The experimental settings of
electrospinning processes were as follows: (i) syringe pump
speed: 0.1 mm min−1, (ii) voltage: 18 kV, (iii) distance between
the target and the tip of the syringe: 11 cm, (iv) target speed:
1 m min−1, (v) traverse speed: 0.5 cm min−1. Fiber mats were
annealed overnight to remove any remaining solvent and stored
until required. After electrospinning, 2 mL of algae solution in
its exponential phase of growth with a total chlorophyll content
(Chl a and b) of ∼2 mg L−1, was placed onto a cut out rectangular chitosan-mat (3 × 2 cm) and left at room temperature for
approximately 48 hours to allow sufficient attachment of algae
cells to the surface of the mat. Microalgae cell walls contain
various polysaccharides, which are compatible with the surface
of the chitosan nanofibers.18,19 The presence of negative surface
charge on the surface of Chlorella cells, arising from dissociation
of uronic acid groups, and/or the presence of sulfate groups for
example,18 provide electrostatic attraction to the positively
charged primary amine groups of chitosan not involved in the
above cross linking. Moreover, the negatively charged surface of
the microalgae can also result in binding metal ions, thereby providing an opportunity for biosorption applications, along with
the removal of nitrate ions as functional algal cells.18–20
This bionano-composite material was then placed into nitrate
containing artificial growth medium which contained mainly
phosphates, nitrates, carbonate buffer, micronutrients and vitamins,21 with an initial nitrate-nitrogen concentration of around
30 mg L−1. Nitrate-nitrogen (NO3−-N) term refers to the amount
of nitrogen (N) in liquid solutions coming from nitrate ions
(NO3−). This nitrate-nitrogen concentration is within the range
of other algal nitrate removal studies.2,22 Furthermore, it simulates the range of nitrate content present in ground-waters
(∼0.1–50 mg L−1 NO3−-N)23 and sewage treatment plant
effluents. The regulatory limit for the maximum contaminant
levels of [NO3−-N] in public drinking water is 10 mg L−1, as
established by the United States Environmental Protection
Agency (EPA).24,25
During the algal growth, the pH of the medium was kept
around 6.5–7.0 by the addition of dilute hydrochloric acid (HCl)
when necessary. The colorimetric “cadmium reduction method”
was employed for the nitrate-nitrogen analysis, using chemicalkits in the form of powder pillows (HACH®, NitraVer Nitrate
Reagent) and a colorimeter (HACH® DR/870).26 For comparative purposes, a control experiment with a chitosan nanofiber
mat devoid of algae culture was also treated under the same conditions. Scanning electron microscopy (SEM) analyses were
acquired using a Zeiss 1555 VP-FESEM, while the accelerating
voltage was changed between 3 to 5 kV. The air-dried samples
were coated with approximately 3 nm layer of platinum before
imaging. A NanoMan AFM system (Veeco Instruments Inc.)
was used for the atomic force microscopy (AFM) analysis, operating under the tapping mode. Chlorophyll content of the cells
was analyzed using spectrophotometric measurements of methanol extracts obtained from the algal culture pellets.27
A challenge in the present work was to fabricate an insoluble,
fibrous structure with sufficient porosity, which can facilitate the
diffusion of nutrients and cellular products between the environment and the algae. Fig. 1a and b show scanning electron
This journal is © The Royal Society of Chemistry 2012
Fig. 1 SEM images of as prepared electrospun chitosan nanofibers at
low and high magnifications with scale bars of (a) 1 μm, and (b) 10 μm,
respectively.
Fig. 2 (a) SEM images of porous and swollen chitosan nanofiber mats,
after exposure to nitrate containing media for two days, (b) SEM images
of chitosan nanofibers surrounding individual, and (c) multiple algae
cells. Scale bars are given as 1 μm.
microscopic (SEM) images of the nanofiber structure of the
chitosan mats after the electrospinning process. The diameter of
the fibers was between 50 to 180 nm, with an average diameter
of around 91 nm. After placing the nanofiber mat into aqueous
solution, the fibers gradually swelled with a significant increase
in porosity of the material, Fig. 2a, and became an effective
support matrix for the C. vulgaris cells, which have the expected
diameter from SEM images around 3–4 μm, Fig. 2b and c. This
porous structure has an advantage for facilitating the diffusion of
materials such as nutrients and waste products between the
environment and the algae, while the replication of algal cells is
accomplished on the surface of the nanofiber mat, Fig. 2c. The
height profile measurements obtained by AFM analyses established the thickness of the chitosan mat at close to 400 nm,
Fig. 3a, with the overall thickness for the algae attached chitosan
mat at 4.3 μm, Fig. 3b. The difference in height is consistent
with the attachment of a single layer of individual C. vulgaris
cells on the nanofiber mats.
Fig. 4 shows the optical images of algae cells attached to the
surface of chitosan mats (a) initially, (b) after 3 days, and (c)
after 10 days from the start of the growth experiments in 40 mL
liquid media. Note the increasing green color on the surface is
due to the increased concentration of algal cells on the chitosan
mat with respect to time. Detailed imaging is given as SEM
images in Fig. 5. The amount of algae cells on a chitosan mat
with same dimensions yielded around 4 cells per 100 μm2 for a
3 day old sample, Fig. 5c, whereas this increased to around
Green Chem., 2012, 14, 2682–2685 | 2683
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Fig. 3 AFM topographic mapping of chitosan nanofiber mats
(4 × 4 μm) without, (a) and with, (b) algal cells.
Fig. 6 Nitrate-nitrogen (NO3−-N) concentration (mg L−1) of algal
medium versus time. Chitosan nanofibers devoid of algal cells are represented with triangles, whereas those with immobilized algal cells are
shown as rectangles.
Fig. 4 Progress of the algal growth on the surface of chitosan mats: (a)
initially, (b) after 3 days, (c) after 10 days of the growth experiment.
Fig. 5 SEM images of immobilized C. vulgaris cells on the surface of
chitosan nanofiber mats after different time intervals. (a and c) are for
mats after 3 days, and (b and d) are for mats 10 days old.
20 cells per 100 μm2 by the 10th day of the treatment process,
Fig. 5d.
Fig. 6 shows the nitrate-nitrogen concentration versus time in
the absence of algae (triangles) or algae attached (rectangles)
nanofiber mats. After the insertion of this bionano-composite
into the liquid media (Vtotal: 40 mL), around 30% of the initial
nitrate value was decreased within the first 2 days. This reduction
in nitrate is mainly caused through the uptake by the chitosan
nanofibers rather than the algal cultures, as a physicochemical
adsorption process. A similar pattern was also observed for the
mats devoid of algal cells where after the second day there was
no further nitrate removal. In contrast the algae containing mats
continued their nitrate uptake, being used in their cellular metabolism for replication, building more biomass and energy
products.
2684 | Green Chem., 2012, 14, 2682–2685
Amino groups in chitosan are protonated at acidic to neutral
pH conditions,28 which enhance the adhesive properties of chitosan by increasing its tendency to attach negatively charged entities which in this case are algal cell walls and nitrates. Chlorella
vulgaris cell walls are known to be highly negatively charged
with a zeta potential of around −30 mV in neutral water.29 On
the other hand, the zeta potential of the positively charged chitosan nanofibers is +20 mV at neutral pH. Matsumoto et al.30
reported the zeta potential values of chitosan nanofibers to be
highly dependent on the pH of the media, with it increasing
to +30 mV in pH around 5–6, whereas it drops to zero for pH
values above 8.30 Algal growth tends to alkalify its medium, as
the cellular uptake of anions (such as nitrates, phosphates, carbonates, etc.) is stabilized with equivalent amounts of hydroxyl
(OH−) anion efflux.31 For this reason, we maintained the pH of
the culture around 6.5–7 by the regular addition of dilute HCl
during the current study. Due to the nature of HCl, several other
acidifying agents (such as CO2) can be considered for any future
developments and advanced scale-up processes for municipal
and/or industrial wastewater samples.
Clearly the presence of the nanofiber mat in the liquid
environment is responsible for the initial removal of nitrate while
the continued growth of algae subsequently consumes the
remaining nitrate in further stages with a slower rate. Overall
nitrate removal rates were calculated as 32 ± 3%, and 87 ± 4%,
for the “microalgae-absent” and “microalgae-attached” chitosan
mats, respectively. Several studies have already been reported on
wastewater treatment with immobilized microorganisms. Fierro
et al.22 investigated the effect of nitrate removal by Scenedesmus
spp. cyanobacterial cells immobilized within spherical chitosan
beads. They achieved 70% nitrate removal for the immobilized
cultures, while 20% of the initial nitrate content was removed by
the blank chitosan beads. In another study, Mallick and Rai32
also achieved relatively higher nitrate removal rates (73%) by
Anabaena doliolum and Chlorella vulgaris cells immobilized in
chitosan beads compared to the cells immobilized in other types
of gels made of alginate, carrageenan or agar. De-Bashan et al.33
achieved only 15% nitrate removal for co-immobilized
This journal is © The Royal Society of Chemistry 2012
Published on 08 August 2012. Downloaded by University of Western Australia on 09/10/2014 10:19:02.
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microorganisms (Chlorella vulgaris with a growth-promoting
bacterium Azospirillum brasilense) within alginate beads. At the
other end of the scale, Tam and Yong34 reported complete nitrate
removal using immobilized C. vulgaris cells within calcium alginate beads. The treatment efficiency of our current method is
comparable with that of these aforementioned methods, although
large variations among the experimental parameters; including
wastewater composition, microbial species, duration of the
process, type of bioreactor, chemical composition and shape of
the immobilization matrix, make direct comparison difficult.
In summary, we have established the use of cross-linked chitosan nanofiber mat as a water-insoluble and non-toxic support for
algal growth and nitrate removal from waters. Algal growth on a
support material can lead to combine algal harvesting, dewatering, and processing steps in a single stage. This bionano-composite material is potentially an attractive, simple and highly
durable polymer, with the mats still retaining their integrity after
six months in contact with an aqueous solution, and has promise
for industrial and/or municipal wastewater treatment processes.
Acknowledgements
This work has been supported by The University of Western
Australia and the Australian Research Council. We would like
to acknowledge the facilities of the Australian Microscopy &
Microanalysis Research Facility at the Centre for Microscopy,
Characterization & Analysis, The University of Western Australia, which was funded by the University, State and Commonwealth Governments. AFM images were performed at Curtin
University.
Notes and references
1 Metcalf and Eddy, Inc, Wastewater Engineering: Treatment and Reuse,
McGraw-Hill, New York, 4th edn, 2003.
2 N. Mallick, BioMetals, 2002, 15, 377.
3 L. Hall-Stoodley, J. W. Costerton and P. Stoodley, Nat. Rev. Microbiol.,
2004, 2(2), 95.
4 Y. Liu, M. H. Rafailovich, R. Malal, D. Cohn and D. Chidambaram,
Proc. Natl. Acad. Sci. U. S. A., 2009, 106(34), 14201.
5 M. S. A. Hameed and O. H. Ebrahim, Int. J. Agri. Biol., 2007, 9, 183.
This journal is © The Royal Society of Chemistry 2012
6 K. Abe, E. Takahashi and M. Hirano, J. Appl. Phycol., 2008, 20, 283.
7 G. Schumacher and I. Sekoulov, Water Sci. Technol., 2002, 46, 83.
8 E. Delahaye, R. Boussahel, T. Petitgand, J. P. Duguet and A. Montiel,
Desalination, 2005, 177, 273.
9 A. Lavoie and J. de La Noue, J. World Maricul. Soc., 1983, 14, 685.
10 A. Zamani, L. Edebo, B. Sjöström and M. J. Taherzadeh, Biomacromolecules, 2007, 8, 3786.
11 W. A. Bough, Process Biochem., 1976, 11, 13.
12 E. Wasik, J. Bohdziewcz and M. Blaszczyk, Process Biochem., 2001, 37,
57.
13 H.-Y. Hu, N. Goto and K. Fujie, Water Res., 2001, 35, 2789.
14 A. Elmidaoui, F. Elhannouni, S. M. A. Menkouchi Sahli, L. Chay,
E. Elabbassi, M. Hafsi and D. Largeteau, Desalination, 2001, 136, 325.
15 Z. Feleke and Y. Sakakibara, Water Sci. Technol., 2001, 43, 25.
16 K. Ohkawa, D. I. Cha, H. Kim, A. Nishida and H. Yamamoto, Macromol.
Rapid Commun., 2004, 25, 1600.
17 M. Bradshaw, J. Zou, L. Byrne, K. S. Iyer, S. G. Stewart and
C. L. Raston, Chem. Commun., 2011, 47, 12292.
18 D. Kaplan, D. Christiaen and S. M. Arad, Appl. Environ. Microbiol.,
1987, 53, 2953.
19 R. H. Crist, K. Oberholser, N. Shank and M. Nguyen, Environ. Sci.
Technol., 1981, 15, 1212.
20 B. Volesky and Z. R. Holan, Biotechnol. Prog., 1995, 11, 235.
21 C. J. S. Bolch and S. I. Blackburn, J. Appl. Phycol., 1996, 8, 5.
22 S. Fierro, M. del P. Sanchez-Saavedra and C. Copalcua, Bioresour.
Technol., 2008, 99, 1274.
23 U. N. Dwivedi, S. Mishra, P. Singh and R. D. Tripathi, in Environmental
Bioremediation Technologies, ed. S. N. Singh and R. D. Tripathi,
Springer, New York, 2007, ch. 16, pp. 353–389.
24 EPA, National Pesticide Survey: Project Summary, U.S. Environmental
Protection Agency, Washington DC, 1990.
25 S. Ghafari, M. Hasan and M. K. Aroua, Bioresour. Technol., 2008, 99,
3965.
26 APHA, Standard Methods for the Examination of Water and
Wastewater, American Public Health Association, Washington, DC,
18th edn, 1992.
27 H. K. Lichtenthaler and C. Buschmann, in Current Protocols in Food
Analytical Chemistry, ed. R. E. Wrolstad, John Wiley & Sons Inc.,
New York, 2001, pp. F4.3.1–F4.3.8.
28 L.-Q. Wu, P. Gadre Anand, H. Yi, M. J. Kastantin, W. Rubloff Gary,
E. Bentley William, F. Gregory Payne and R. Ghodssi, Langmuir, 2002,
18, 8620.
29 B.-M. Hsu, Parasitol. Res., 2006, 99, 357.
30 H. Matsumoto, H. Yako, M. Minagawa and A. Tanioka, J. Colloid Interface Sci., 2007, 310, 678.
31 J. Naus and A. Melis, Plant Cell Physiol., 1991, 32, 569.
32 N. Mallick and L. C. Rai, World J. Microbiol. Biotechnol., 1994, 10,
439.
33 L. E. de-Bashan, Y. Bashan, M. Moreno, V. K. Lebsky and J. J. Bustillos,
Can. J. Microbiol., 2002, 48, 514.
34 N. F. Y. Tam and Y. S. Wong, Environ. Pollut., 2000, 107, 145.
Green Chem., 2012, 14, 2682–2685 | 2685
)
ARTICLE
Hierarchical Patterning of
Multifunctional Conducting Polymer
Nanoparticles as a Bionic Platform for
Topographic Contact Guidance
Dominic Ho,†,‡ Jianli Zou,§ Xianjue Chen,†,0 Alaa Munshi,† Nicole M. Smith,†, Vipul Agarwal,†
Stuart I. Hodgetts,‡ Giles W. Plant,^ Anthony J. Bakker,‡ Alan R. Harvey,‡ Igor Luzinov,# and K. Swaminathan Iyer*,†
School of Chemistry and Biochemistry, The University of Western Australia, Crawley, Western Australia 6009, Australia, ‡School of Anatomy, Physiology and Human
Biology, The University of Western Australia, Crawley, Western Australia 6009, Australia, §Institute for Integrated Cell-Material Sciences (iCeMS), iCeMS Complex 2,
Kyoto University, Yoshida-Honmachi, Sakyo-ku, Kyoto, 606-8501, Japan, Experimental and Regenerative Neurosciences, School of Animal Biology, The University of
Western Australia, Crawley, Western Australia 6009, Australia, ^Stanford Partnership for Spinal Cord Injury and Repair, Department of Neurosurgery,
Stanford University School of Medicine, Stanford, California 94305, United States, and #School of Materials Science and Engineering, Clemson University, Clemson,
South Carolina 29634, United States. 0Present address: Centre for NanoScale Science and Technology, School of Chemical and Physical Sciences, Flinders University,
Bedford Park, Adelaide, SA 5042, Australia.
)
†
ABSTRACT The use of programmed electrical signals to influence
biological events has been a widely accepted clinical methodology for
neurostimulation. An optimal biocompatible platform for neural
activation efficiently transfers electrical signals across the electrode
cell interface and also incorporates large-area neural guidance
conduits. Inherently conducting polymers (ICPs) have emerged as
frontrunners as soft biocompatible alternatives to traditionally used
metal electrodes, which are highly invasive and elicit tissue damage
over long-term implantation. However, fabrication techniques for the
ICPs suffer a major bottleneck, which limits their usability and medical translation. Herein, we report that these limitations can be overcome using colloidal
chemistry to fabricate multimodal conducting polymer nanoparticles. Furthermore, we demonstrate that these polymer nanoparticles can be precisely
assembled into large-area linear conduits using surface chemistry. Finally, we validate that this platform can act as guidance conduits for neurostimulation,
whereby the presence of electrical current induces remarkable dendritic axonal sprouting of cells.
KEYWORDS: multimodal nanoparticles . conducting polymers . capillary force lithography . neurostimulation
E
xogenous electrical stimulation has
been effectively used both in clinical
practice and in laboratory research
to regulate cell-type-dependent adhesion,
differentiation, and growth.1 This phenomenon of introducing programmed electrical
signals locally to influence biological events
has resulted in the pursuit of sophisticated
medical bionic devices.2 An important property that dictates the performance of most
bionic electrodes is the electrode/cellular
interface and its ability to transmit charge
across the biointerface.3 Traditionally metallic electrodes made of platinum, gold, iridium
oxide, tungsten, their alloys, and more
recently carbon fibers have been effectively
HO ET AL.
employed in bionic devices.4 They have
been employed for deep brain stimulation,
as cochlear implants, for vagus nerve stimulation to treat epilepsy, and for stimulating
regeneration in the central nervous system.2 However, stiff metal electrodes suffer
a major drawback of eliciting tissue damage
over long-term implantation.2 Importantly,
it is now recognized that nanoscale patterns
provide topographic guidance cues for
cells. This has been widely exploited to
engineer sophisticated regenerative platforms for nerves, muscles, skin, and bones.5
The need to incorporate large-area nanoscale patterns for bionic applications coupled
with the demand toward miniaturization of
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* Address correspondence to
swaminatha.iyer@uwa.edu.au.
Received for review November 20, 2014
and accepted January 26, 2015.
Published online January 26, 2015
10.1021/nn506607x
C 2015 American Chemical Society
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HO ET AL.
sustained release from the nanoparticles once patterned and multimodal imaging of the nanoparticle
constructs once implanted.
RESULTS AND DISCUSSION
Patterned Multifunctional PEDOT:PSS Nanoparticle Arrays.
In this study poly(glycidal methacrylate) (PGMA) is
used as a reactive macromolecular anchoring platform
both on the substrate as a nanoscale layer and as a
colloidal nanoparticle to enable multilayer assembly
(Figure 1). A polymer with epoxy functionality was
chosen, since the reactions of epoxy groups are universal and easily transferable to various substrates,
affording ease of attachment of functional molecules.
Furthermore, the epoxy groups of the polymer can
cross-link to provide structural integrity to the pattern
and nanoparticle constructs.11 The mobility of the
reactive loops of PGMA ensures greater access to
anchoring, resulting in a 23-fold greater grafting
density when compared to a monolayer of epoxy
groups on a nanoparticle surface of similar dimension,
enabling high loading using a layer by layer approach
that is adopted in the current study.11 Polymer nanospheres were initially prepared using an oil in water
emulsion methodology from PGMA modified with a
rhodamine-B (RhB) dye, encapsulated with magnetite
(Fe3O4) nanoparticles to form the core platform
(Figure 1a,b). Not only does the incorporation of
magnetite and RhB render these constructs multimodal for both MRI and fluorescence imaging, but
importantly in the present case magnetite provides a
means to separate, wash, and purify the nanoparticles
using a magnetic fractionation column during each
step of layered assembly. Polyethylenimine (PEI) was
then covalently bound to the RhB-PGMA core to facilitate a cationic layer for electrostatic conjugation of an
anionic conducting polymer, PEDOT:PSS (Figure 1c,d).
Capillary force lithography (CFL) was then used
to generate large-area nanoscale conduits in which
PEDOT:PSS nanoparticles are electrostatically directed
to self-assemble as linear channels from solution
(Figure e,f). Capillarity allows the polymer melt to fill
up the void space between the polymer and the
applied mold when the temperature is above the
glass-transition temperature (Tg), thereby generating
a large-area pattern that depends on the size of stamp.
Importantly, the technique needs no specialized
instrumentation for generation of large-area patterns.
Patterns can easily be generated using polydimethylsiloxane (PDMS) stamps, which in turn can be fabricated using the ubiquitous optical storage discs as a
master. An optical data storage disc is typically made of
a polymer (polycarbonate) disc, on which a single spiral
track is drilled. The typical width and depth of each line
in the spiral track are 800 and 130 nm, respectively, and
the periodicity of the track is ∼1.5 μm (Figure S1). In the
present study, an indium tin oxide (ITO) substrate was
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biocompatible implantable devices has resulted in
the emergence of inherently conducting polymers as
frontrunners for fabricating flexible organic electrode
materials. However, advances in the applicability of
patterned surfaces of inherently conducting polymers
in bionic devices have been limited due to the difficulties of transferring printing techniques and their
integration under physiological conditions. In this
article, we report a transferable method to fabricate multifunctional poly(3,4-ethylenedioxythiophene)poly(styrenesulfonate) (PEDOT:PSS) nanoparticles and
direct their self-assembly by electrostatic interactions
into large-area patterns. Using the rat pheochromocytoma cell line (PC12), we demonstrate the suitability
of the assembly as a bionic platform for exogenous
electrical stimulation.
The three primary classes of conducting polymers
that have been studied are polyanilines, polypyrroles,
and polythiophenes.6 The ease of functionalization of
polythiophenes and maintenance of conductivity under physiological conditions has made them primary
candidates for multifunctional organic bionic devices.7
The most widely explored processes for the fabrication
of organic conducting polymer patterns are electropolymerization, extrusion printing, inkjet printing,
microcontact printing, electrospinning, and more
recently high-precision Dip Pen Nanolithography
(DPN).4,6 Electropolymerization has been widely used
for coating metal/carbon substrates, following which
patterning is achieved by top-down lithography on
polymer thin films covering larger area electrodes. This
technique results in controlled, high-resolution nanoscale patterns but is limited by the ability to regulate
polymerization of monomers on nanoscale implantable electrodes.8 Similarly, printing techniques have
achieved significant advances in recent years, reaching
high-throughput patterns, but are limited in resolution
by the liquid dispensing techniques, which operate
within the limit of tens of micrometers.9 Electrospinning techniques have offered simple processable solutions to generate 3D scaffolds at resolutions mimicking
the extracellular matrix architecture but are limited
by the inability to generate patterned conducting
conduits for the development of bionic guidance
channels.4 The aforementioned shortfalls have been
recently overcome by the advances of DPN, which
enables precise deposition, patterning down to nanoscale resolution, and most importantly applicability
over a wide range of substrates.10 However, advances
are limited by their cost, need for specialized equipment, and low throughput. In the present paper we
adopt a bottom-up self-assembly process to precisely
pattern conducting polymer nanoparticles into patterns as conduits for guidance. The approach is easily
adoptable over multiple substrates, needs no specialized equipment, and affords large-area patterns. Importantly, this approach enables drug encapsulation and
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Figure 1. Schematic illustration of the fabrication protocol to pattern multifunctional PEDOT:PSS nanoparticle arrays for
exogenous electrical stimulation. (ad) Multilayer assembly of conducting PEDOT:PSS nanoparticle fabrication via nonspontaneous emulsification. (a) An organic phase is initially formed by dissolving RhB-modified PGMA (yellow) and Fe3O4
(purple) in a 1:3 mixture of CHCl3 and MEK. (b) Colloidal fluorescent PGMA-Fe3O4 nanoparticles are fabricated upon dropwise
addition of the organic phase to an aqueous solution of Pluronic F-108. (c) Cationic second layer via covalent attachment of
PEI (green) to the PGMA-Fe3O4 core. (d) Anionic conducting polymer layer via electrostatic attachment of PEDOT:PSS (blue).
(eg) Patterning of the multilayered PEDOT:PSS nanoparticles for exogenous electrical stimulation of PC12 cells. (e) Linear
nanoparticle conduits patterned on a substrate via capillary force lithography (CFL) using charge complementarity. A detailed
schematic of the CFL procedure can be found in Figure S2. (f) PC12 cells (green) were cultured onto the biocompatible
platform, followed by (g) exogenous electrical modulation.
modified first by spin coating a thin film of PGMA
followed by a second spin-coated layer of polystyrene
(PS) using previously reported conditions.12 The PS
layer acts as a chemical resist to selectively react the
epoxy groups of PGMA following patterning. The
PS/PGMA bilayer was annealed with the PDMS mask
at 130 C (T > Tg of PS) to induce patterning via capillary
flow. The reusable PDMS stamp was peeled off following heat treatment to obtain a patterned surface resulting in alternating PGMA and PS stripes.
HO ET AL.
Ethylenediamine (EA) was then grafted to PGMA to
result in cationic linear patterns. PEDOT:PSS nanoparticles were then electrostatically assembled onto
the patterned surface, followed by washing steps to
remove PS to obtain linear arrays of assembled PEDOT:
PSS nanoparticles. A detailed schematic of the fabrication process is shown in Figure S2. The nanoparticle
and the patterns were characterized at each step of the
assembly (Figure 2). The PEDOT:PSS nanoparticles
were an average size of 200 nm (Z-average) with a
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Figure 2. Characterization of the multilayered PEDOT:PSS conducting nanoparticles and their assembly as linear conduits.
(a) TEM micrograph of the multilayered PEDOT:PSS nanoparticles. Scale bar = 200 nm. (Inset: high-magnification TEM image
of PEDOT:PSS-coated nanoparticles showing encapsulated Fe3O4 nanoparticles. Scale bar = 10 nm.) (b) DLS particle size
distributions of the PEDOT:PSS nanoparticles in solution. (c) Zeta potential distributions of the nanoparticles: PGMA-Fe3O4
core (black) with an average zeta potential of 3.9 ( 1.3 mV, cationic PEI-coated (red) with an average zeta potential of 37 (
1.2 mV, and anionic PEDOT:PSS-coated (blue) with an average zeta potential of 29 ( 6.15 mV. (d) Current vs voltage
response for the nonconducting PEI-coated nanoparticles (red) and conducting PEDOT:PSS-coated (black) nanoparticles.
(eg) Tapping mode AFM topography images of the nanoparticle patterns at each stage of fabrication: PGMA and PS stripes
(e), EA-modified PGMA and PS stripes (f), PEDOT:PSS nanoparticle patterns (g). (hj) Corresponding height profiles of the
nanoparticle patterns at each stage of fabrication.: PGMA and PS stripes (h), EA-grafted PGMA and PS stripes (i), PEDOT:PSS
nanoparticles patterns (j). The AFM line scans corresponding to the height profiles are indicated on the topography images in
(e)(g). (k, l) SEM micrographs of the nanoparticle patterns at a magnification of 25k (k) and 11k (l) indicating the
formation of tightly packed and highly ordered nanoparticle arrays. (m) Confocal fluorescence image of the RhBfunctionalized PEDOT:PSS nanoparticle arrays at 20 magnification.
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Figure 3. Biocompatibility of the PEDOT:PSS nanoparticle arrays with PC12 cells. (a) Cell viability determined using MTS
calorimetric assay obtained at 72 h after an initial exogenous electrical stimulation for 2 h and in the absence of stimulation
showing no significant changes. (b) SEM micrograph demonstrating preferential cell adhesion to the pattern area (yellow
box). Image acquired at 323 magnification 72 h after the addition of NGF without exogenous electrical stimulation. (c) Highmagnification (12k magnification) SEM images demonstrating specific and preferential interactions of neurites (white
arrows) with the PEDOT:PSS linear conduits (red arrows).
polydispersity index (PDI) of 0.07, a zeta potential of
29 ( 6.15 mV, and a conductivity of 2.5 1012 S/cm.
The measured conductivity is in accordance with other
values reported in the literature for polymer blends.13
Importantly, this low conductivity is important under
physiological conditions to induce local cellular stimulation and avoid tissue damage due to toxic overstimulation.14 The final self-assembled linear arrays
of PEDOT:PSS nanoparticles were of large-area highdensity packing, as confirmed at various length scales
using AFM, SEM, and fluorescence imaging.
Biocompatibility Assessment of the PEDOT:PSS Nanoparticle
Arrays. Topographic modulation of tissue response is
one of the most important considerations in developing bionic implants. Topographic contact guidance
using micropatterns has been widely exploited to
influence cell migration, adhesion, and proliferation.15,16 One of the pivotal first steps in the present study was to establish biocompatibility of the
patterned structures. PC12 cells were chosen in the
present case, as they have been demonstrated to show
HO ET AL.
enhanced neurite outgrowth and spreading upon
exogenous stimulation on a conducting polymer
substrate.17 MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium, inner salt) assays, cell viability assays, and SEM
imaging were performed after exogenous electrical
stimulation and in the absence of electrical stimulation
to determine effects on cell viability and cell adhesion
(Figure 3a,b and Figure S3). The stimulation conditions
used in the present study involved a monophasic
pulsed current at a frequency of 250 Hz with a 2 ms
pulse width and an amplitude of 1 mA for 2 h, similar to
protocols previously reported for similar cell lines.18,19
Importantly, we observed no changes in cell viability
upon exogenous stimulation and observed preferential adhesion of the PC12 cells to the patterned surface over a nonpatterned surface in both cases
(( stimulation). High-magnification SEM imaging (no
stimulation) further revealed preferential interaction of
the PC12 cells to the PEDOT:PSS nanoparticle arrays,
confirming not only biocompatibility with the large
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Figure 4. Exogenous electrical stimulation induced dendritic sprouting of the PC12 cells guided by the PEDOT:PSS linear
conduits. (a) Significant increase in the average cell area is observed 72 h after exogenous electrical stimulation on the
nanoparticle platform in comparison to unstimulated and nonpatterned controls. (b) Corresponding decrease in PC12 cell
proliferation observed 72 h after exogenous electrical stimulation on the nanoparticle platform in comparison to
unstimulated and nonpatterned controls. (cf) Representative confocal images (40 magnification) of β-III tubulin
immunohistochemically stained cells 72 h after the following treatments: {(þ) pattern, () stimulation} (c); {(þ) pattern,
(þ) stimulation} (d); {() pattern, () stimulation} (e); {() pattern, (þ) stimulation} (f), demonstrating modulation of cell
morphology. (g) High-magnification SEM image (magnification 8k) indicating the formation of extensive dendritic networks
(white arrows) guided by the PEDOT:PSS arrays. Inset: The corresponding low-magnification image of the area (yellow box)
analyzed (magnification 3k).
area of the pattern but also potential applicability
of the nanoscale linear arrays as guidance conduits
(Figure 3c).
Exogenous Electrical Stimulation Induced Dendritic Sprouting
of the PC12 Cells. Electrical stimulation has been effectively used to modulate growth and differentiation of
HO ET AL.
anchorage-dependent cells such as neurons, fibroblasts, and epithelium cells.17,20,21 In the central nervous system, brief stimulation to the proximal end of
transected peripheral nerves has been shown to augment preferential motor reinnervation,22 improve the
specificity of sensory reinnervation,23 and accelerate
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METHODS SUMMARY
Nanoparticle Synthesis. The conducting nanoparticles were
prepared via a nonspontaneous emulsification route. Briefly,
rhodamine B was attached to PGMA in MEK at 80 C under N2 for
5 h. The modified PGMA was then precipitated in diethyl ether
and dried under N2. This was dispersed in a 1:3 mixture of CHCl3
and MEK along with 25 mg of Fe3O4 to form the organic phase.
This organic phase was added dropwise into a rapidly stirring
aqueous solution of Pluronic F-108. The emulsion was homogenized with a probe-type ultrasonic wand for 1 min. The
organic solvents were then evaporated off under N2. Large
aggregates of Fe3O4 and excess polymer were separated via
centrifugation. The nanoparticles in the supernatant were then
mixed with PEI and heated to 80 C for 16 h to facilitate
attachment. The PEI-coated nanoparticles were isolated and
washed on a magnetic separation column. Next, a diluted
solution of PEDOT:PSS was added dropwise under rapid stirring
to nanoparticles at a concentration of 0.5 mg/mL to facilitate
electrostatic attachment. This was followed by sonication for
10 min and stirring for 18 h. The nanoparticles were then washed
multiple times in water before being stored at 4 C for further use.
Platform Fabrication. To direct the self-assembly of the nanoparticles, a template was fabricated by CFL. A 0.2% w/v PGMA in
CHCl3 solution was spin coated on ITO coverslips and annealed
at 120 C for 20 min. Next, 1.3% w/v PS in toluene was spin
coated onto the PGMA surface. A PDMS stamp was then placed
onto the PS layer, followed by heat treatment in an oven at
130 C for 1 h. Once cooled, the stamp was peeled off. This was
followed by exposure to EA at room temperature for 5 h.
The pattern was next washed multiple times with water to
remove unreacted EA. A 50 μL amount of 4 mg/mL nanoparticle
suspension was drop casted onto the patterned area of the
coverslip. The setup was then placed in a sealed vial, facilitating
HO ET AL.
extension during NGF-induced PC12 cell differentiation.29,30 Using immunohistochemical staining
for β-III tubulin it was determined that stimulation on
the patterned surface resulted in a significant increase
in the cell area and lower number of cells per unit area,
indicating exogenous electrical stimulation induced
differentiation of PC12 cells (Figure 4af, Figure S4).
High-magnification SEM (Figure 4g) also revealed
that stimulation resulted in an extensive dendritic
network guided by the linear conduits of PEDOT:PSS
nanoparticles.
ARTICLE
the reinnervation of distal target tissues.24 These have
been reported to depend on depolarization of the
neuronal soma and its axon, involvement of axon
guidance factors such as polysilylated neural cell
adhesion molecule,25 the L2/HNK-1 carbohydrate,26
and brain-derived neurotrophic factor.27 Finally electrical stimulation induced neurite outgrowth was
recently reported to be dependent on calcium influx
through L- and N-type voltage-dependent calcium
channels and calcium mobilization from IP3R and
RYR-sensitive calcium stores.28 In the present case,
we analyzed the morphological modulation of PC12
cells following electrical stimulation having determined no change in cell viability using the MTS assay.
Nerve growth factor (NGF) induces PC12 cells to
change their phenotype and acquire a number of
properties that are similar to sympathetic neurons.
Importantly, although they can acquire properties
similar to sympathetic neurons upon NGF treatment,
they do not develop definitive dendritic axons or form
true synapses with each other in the absence of
exogenous stimulation.29 This change in phenotype
upon NGF treatment is associated with a retardation in
proliferation, the extension of neurites making them
electrically excitable. Monitoring the cell numbers and
cell area can assess this change from the proliferation
state to a differentiation state. Furthermore, microtubule levels correlate precisely with the neurite
CONCLUSION
In summary, we have demonstrated a practical and
transferable protocol to fabricate self-assembled largearea patterns of conducting polymers from solution.
This overcomes some of the shortfalls in the current
fabrication techniques in developing patterned organic
bionic devices. The patterns generated have demonstrated excellent biocompatibility. At the same time,
they have been shown to induce exogenous electrical
stimulation under physiological conditions to elicit a
measurable and consistent cellular response. Importantly the methodology permits the design of bionic
devices capable of inducing local electrical stimulation
for in vivo applications while integrating multimodal
imaging and simultaneous drug delivery capabilities of
nanoparticles.
controlled evaporation, which allowed for electrostatic nanoparticle attachment onto the EA surface. The PS mask was then
removed by washing with toluene. The resulting patterned
PEDOT:PSS nanoparticle array was then used for further
experimentation.
Electrical Stimulation Protocol. For electrical stimulation experiments, two silver epoxy electrodes were painted onto the ends
of the patterned nanoparticle arrays. Prior to cell culture, the
whole platform was UV and ethanol sterilized. Wells were
coated with poly(L-lysine) and 15 μg/mL of laminin followed
by cell seeding at a density of 50 000 cells/well. Cells were left to
adhere for 18 h. Immediately prior to stimulation, the proliferation media was replaced with low-serum nerve growth factor
containing differentiation media. For stimulation, the cells were
subjected to a monophasic pulsed current at a frequency of
250 Hz with a 2 ms pulse width and an amplitude of 1 mA for 2 h,
after which they were left for an additional 72 h before analysis.
Cell Viability Assessment. Cell viability was measured using the
MTS assay as per the manufacturer protocols (Invitrogen, UK).
For measurements, 80 μL from each well was transferred into a
new 96-well plate and read under a plate reader at 490 nm
excitation wavelength. To analyze cell morphology, cells were
immunohistochemically stained for β-III tubulin.
Material Characterization. AFM was performed on a Dimension
3100 AFM system with a Nanoscope IV controller used to obtain
the AFM images in tapping mode, using Pt/Ir-coated contact
mode probes with a spring constant of 0.2 N/m (type SCM-PIC,
Bruker). TEM was performed on a JEOL 2100 transmission
electron microscope at an accelerating voltage of 80 kV. SEM
was performed on a Zeiss 1555 VP-FESEM, and all samples were
coated with 5 nm of Pt. Biological samples were initially fixed in
2.5% glutaraldehyde and dehydrated in increasing concentrations of ethanol followed by critical point drying prior to
Pt coating. Immunohistochemically stained samples were
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Conflict of Interest: The authors declare no competing
financial interest.
Supporting Information Available: Detailed materials and
methods: synthesis, characterization (TEM, SEM, AFM), cell
culture, and electrical stimulation experiments. This material is
available free of charge via the Internet at http://pubs.acs.org.
Acknowledgment. D.H., I.L., and K.S.I. designed the experiments, developed the concept, and analyzed the data. D.H., J.Z.,
and N.M.S. optimized the capillary force lithography experiments. D.H., X.C., V.A., and A.M. performed image acquisition
using confocal microscopy, transmission electron microscopy,
scanning electron microscopy, and atomic force microscopy.
D.H., A.R.H., G.W.P., S.I.H., and A.B. optimized and designed the
electrical stimulation experiments. This work was funded by
the Australian Research Council (ARC), the National Health &
Medical Research Council (NHMRC) of Australia, and the National
Science Foundation (CBET-0756457). The authors acknowledge
the Australian Microscopy & Microanalysis Research Facility at the
Centre for Microscopy, Characterization & Analysis, and The
University of Western Australia, funded by the University, State
and Commonwealth Governments. The authors also wish to
thank Margaret Pollett and Chrisna LeVaillant for their invaluable
contribution in assisting with the PC12 cell cultures and immunohistochemistry, and Ella Marushchenko (www.scientificillustrations.com) for assistance with Figure 1.
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analyzed using a Leica TCS SP2 AOBS multiphoton confocal
microscope.
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Hierarchical Patterning of Multifunctional Conducting Polymer Nanoparticles as a Bionic
Platform for Topographic Contact Guidance
Dominic Ho,1,2 Jianli Zou,3 Xianjue Chen,1,‡, Alaa Munshi,1 Nicole M. Smith,1,4 Vipul Agarwal,1
Stuart I. Hodgetts,2 Giles W. Plant,5 Anthony J. Bakker,2 Alan R. Harvey,2 Igor Luzinov6 & K.
Swaminathan Iyer1*
1
School of Chemistry and Biochemistry, The University of Western Australia, Crawley, WA
6009, Australia;
2
School of Anatomy, Physiology and Human Biology, The University of Western Australia,
Crawley, WA 6009, Australia;
3
Institute for Integrated Cell-Material Sciences (iCeMS), iCeMS Complex 2, Kyoto University,
Yoshida-Honmachi, Sakyo-ku, Kyoto, 606-8501, Japan;
4
Experimental and Regenerative Neurosciences, School of Animal Biology, The University of
Western Australia, Crawley, WA 6009, Australia;
5
Stanford Partnership for Spinal Cord Injury and Repair, Department of Neurosurgery, Stanford
University School of Medicine, Stanford, CA 94305, USA;
6
School of Materials Science and Engineering, Clemson University, Clemson, South Carolina,
29634-0971, USA.
‡
Present Address: Centre for NanoScale Science and Technology, School of Chemical and
Physical Sciences, Flinders University, Bedford Park, Adelaide, SA 5042, Australia.
* Correspondance: swaminatha.iyer@uwa.edu.au
S1
Supplementary Information
Materials. All chemicals were purchased from Sigma-Aldrich unless otherwise stated: iron(III)
acetylacetonate (97 %), benzyl ether (98 %), oleic acid (90 %), oleyl amine (70 %), 1,2tetradecanediol (90 %), rhodamine B (Fluka), methyl ethyl ketone (99 %, Fisher), chloroform
(99 %, merck), toluene (99 %, Fisher), diethyl ether (90 %, Asia Pacific Speciality Chemicals),
polyethylenimine (50 % solution, Mn 1200, Mw 1300), Poly(3,4-ethylenedioxythiophene)
Polystyrene sulfonate (1.3% solution, Mw 10355), ethylenediamine (99.5 %, Fluka) and Pluronic
F-108. All tissue culture reagents were purchased from Gibco unless otherwise stated.
Dulbecco's Modified Eagle's medium (DMEM), PBS, foetal calf serum (Sigma), horse serum
(Sigma),
penicillin/streptomycin,
L-glutamine,
non-essential
amino
acids
(NEAA),
trypsin/EDTA (Sigma), laminin (#L2020, Sigma) and nerve growth factor (β-NGF, PeproTech).
Magnetite Synthesis. Fe3O4 was synthesized by the organic decomposition of Fe(acac)3 in
benzyl ether at 300 oC, in the presence of oleic acid, oleyl amine, and 1,2- tetradecanediol, as
previously described by Sun et al.1 The method to synthesise 6 nm Fe3O4 nanoparticles was
followed.
Synthesis of RhB-Modified PGMA: PGMA was synthesized by radical polymerization
according to a published procedure.2 Briefly, glycidyl methacrylate was polymerized in methyl
ethyl ketone (MEK) to give PGMA (Mn = 220515, Mw = 433730), using azobisisobutyronitrile
as initiator. The polymer was purified by multiple precipitations from MEK solution using
diethyl ether. To attach the dye to the polymer, a solution of rhodamine B (RhB, 20 mg) and
PGMA (100 mg) in MEK (20 mL) was heated to reflux under N2 for 18 h. The solution was
reduced in vacuo before the modified polymer was precipitated with diethyl ether (20 mL). The
S2
polymer was redissolved in MEK and precipitated with ether twice to remove ungrafted RhB.
PEDOT:PSS Multilayer Nanoparticle (NP) Synthesis. To prepare the organic phase of the
emulsion, the dried RhB-PGMA polymer was initially dissolved in 2 mL of CHCl3 and dried
under N2, leaving a sticky residue. This was redissolved in a 1:3 mixture of CHCl3 (2 mL) and
MEK (6 mL) along with 25 mg of Fe3O4. This organic phase was added drop wise to a rapidly
stirring aqueous solution of Pluronic F-108 (12.5 mg/mL, 30 mL). The emulsion was
homogenised with a probe-type ultrasonic wand at the lowest setting for 1 min. The organic
solvents were evaporated off overnight under a slow flow of N2. The suspension was purified via
centrifugation at 3000 g for 45 mins. The supernatant was transferred to a 50 mL flask
containing PEI (50 wt % solution, 100 mg) and heated to 80 oC for 16 h. The magnetic polymer
nanoparticles were collected on a magnetic separation column (LS, Miltenyi Biotec) in 3 mL
batches, washed with water (5 mL) and then flushed with water until the filtrate ran clear. This
purified product produced 10 mL of nanoparticle suspension at a concentration of 1 mg/mL.
Next, PEDOT:PSS was electrostatically attached to the nanoparticles. 60 µL of PEDOT:PSS (1.3
wt % dispersion in H2O) was diluted in 2 mL of water and added drop wise under rapid stirring
to NPs at a concentration of 0.5 mg/mL. The PEDOT:PSS was further dispersed under sonication
for 10 mins to ensure complete dispersion and then left to stir for 18 h. After 18 h, the mixture
was again sonicated for 2 mins. Excess PEDOT:PSS was then removed via centrifugation 16800
g for 20 mins). NPs were then washed twice in water before being stored at 4 oC at a
concentration of 4 mg/mL for further use.
Nanoparticle Conductivity Measurement. The nanoparticle conductivity was determined using
4-point probe measurements. 80 µL of nanoparticle solution with a concentration of 1 mg/mL
S3
was selectively dried on a square area 0.5 cm x 0.5 cm. Electrodes were placed at the 4 corners
of the square and subject to current-voltage sweeps.
Fabrication of PDMS Stamp. The metal layer of a blank compact disc (CD) was peeled off and
the CD washed with ethanol. The remaining polycarbonate structure was used as a master for the
PDMS stamp. The polymer base and curing agent from a Sylgard® 184 (Dow Corning) silicone
elastomer kit were mixed at a 10:1 ratio by weight in a glass vial. The glass vial was placed in a
vacuum desiccator to remove trapped bubbles from the mixture. Following vacuum treatment,
the elastomer was restored to atmospheric pressure slowly several times until it was free of
bubbles. The PDMS mixture was then cast onto the surface of the grooved side of CD and cured
at 80 ºC for 2 hours.
CFL Procedure. Prior to the CFL procedure, the indium tin oxide (ITO) coverslips were first
clean in acetone and isopopanol under sonication. 0.2 % w/v PGMA in CHCl3 was spin coated
onto the conducting surface of the ITO coverslips. Coverslips were then placed in an oven at 120
ºC for 20 mins to anneal the PGMA. Unreacted PGMA on the coverslip surface was removed by
washing in CHCl3. Next, 1.3 % w/v PS in toluene was spin coated onto the PGMA surface. A
PDMS stamp was then placed onto the PS layer, followed by heat treatment in an oven at 130 ºC
for 1 hr. The assembly was then cooled down at room temperature for another hour before the
PDMS stamp was peeled off. Next, the substrate was exposed to ethylenediamine (EA) and left
at room temperature for 5 h. The substrate was then wasted with water to remove unreacted EA.
Next, 50 µL of 4 mg/mL nanoparticle solution was drop casted onto the patterned area of the
coverslip. The setup was then placed in a sealed vial, facilitating controlled evaporation which
allowed for electrostatic nanoparticle attachment onto the EA surface. The PS mask was then
S4
removed by toluene, leaving the patterned nanoparticle array.
Cell Culture. The rat pheochromocytoma cells (PC12 cells) used here were obtained from
Flinders University (Adelaide, Australia) courtesy of Professor Jacqueline Phillips (Macquarie
University, Sydney, Australia). PC12 cells were cultured in P75 flasks in a humidified
atmosphere containing proliferation media: 5 % CO2 at 37 oC and maintained in DMEM medium
containing horse serum (10 % v/v), fetal calf serum (5 % v/v), penicillin/streptomycin (0.5 %
v/v), L-glutamine (1 % v/v) and nonessential amino acids (1 % v/v). For PC12 differentiation,
cells were cultured in differentiation media consisting of DMEM, L-glutamine (1 % v/v), horse
serum (1 % v/v) and nerve growth factor (50 ng/mL).
Electrical Stimulation Experiments. Prior to stimulation experiments, two silver epoxy
electrodes were painted onto the ends of the prepared NP array and platinum wires attached to
allow for connections with the stimulator. Next, a cell culture well was created by first cutting a
1.5 mL microcentrifuge tube in half and then sealing the capped end with silicon vacuum grease.
This was stuck onto the ITO glass with the patterned arrays in the centre of the well. This was
done to ensure that the electrodes did not come into direct contact with the cell culture media.
The array was then placed in a Petri dish to maintain sterility throughout the course of the
experiment (Fig S5a). Prior to culturing cells on the arrays, the coverslips were UV sterilised (20
mins) and then washed with 70 % ethanol three times. Wells were then coated with poly-(Llysine) and 15 µg/mL of laminin. Cells were then seeded at a density of 50 000 cells/well and left
to adhere for 18 h. Prior to stimulation, the proliferation media was replaced with differentiation
media. The cells were then stimulated according to protocols as listed below. Following
stimulation, the cells were left for a further 72 h with fresh differentiating media added every 48
S5
h. Photographs of the electrical stimulation setup are described in Figure S5.
Stimulation Protocol. The electrical signals were supplied by Grass S44 Stimulator (Quincy,
Massachusetts, USA). The stimulation regime is similar to that used by Wallace et al.3-5 Briefly,
the cells were subjected to a monophasic pulsed current at a frequency of 250 Hz with a 2 ms
pulse width and an amplitude of 1 mA for 2 h.
Cell Viability Assays. Cell viability was measured using the (3-(4,5-dimethylthiazol-2-yl)-5-(3carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium, inner salt) (MTS) assay as per the
manufacturer protocols (Invitrogen, UK). Cells were plated as per “electrical stimulation
protocol” stated above. Viability was to be determined at 3 time points: (i) 0 h (immediately
prior to electrical stimulation), (ii) 72 h after the addition of differentiation media and (iii) 72 h
after the addition of differentiation media and electrical stimulation. For measurements, 80 µL
from each well was transferred into a new 96 well plate and read under a plate reader at 490 nm
excitation wavelength. The same protocol was followed for every sample and each measurement
was carried out in triplicate.
Immunohistochemical Staining. The PC12 cells were immunohistochemically stained for ß-III
tubulin. The cells were fixed in 4 % paraformaldehyde for 10 mins. Cells were first incubated
with a primary antibody solution containing PBS, 10 % Normal Goat Serum, 0.1 % Triton X-100
and the anti-β-III tubulin antibody (1:1000, anti-rabbit, Covance) at room temperature for 30
mins. After 3 PBS washes, the antibody binding was visualised with anti-rabbit FITC (1:100,
Sigma) following incubation for 30 mins at room temperature. Coverslips were mounted on glass
slides covered with Dako Fluorescent Mounting Medium (Dako, USA). All experiments were
S6
performed in triplicate.
Confocal and Fluorescence Microscopy Analysis. Immunohistochemically stained samples
were analysed using confocal and fluorescence microscopy. Confocal microscopy was carried
out using a Leica TCS SP2 AOBS Multiphoton Confocal microscope and fluorescence
microscopy with a Diaplan fluorescence microscope.
Image and Statistical Analysis. To determine the effects both stimulation and the NP arrays had
on the PC12 cells, the average area of each cell was determined. 3 randomly selected areas on
each sample was visualised at 40 x magnification. The average area covered by each cell was
assessed using Image J analysis software (version 1.48a, NIH). All immunohistochemical
analyses were conducted by a single investigator, ensuring constant selection criteria, and results
expressed as means ± SD. Data were analysed using Origin data management software to
conduct ANOVA on groups of data. Statistically significant differences between each treatment
were determined using Bonferroni/Dunn post hoc tests (p≤0.05).
Scanning Electron Microscopy (SEM). Prior to SEM imaging, samples without cells were
coated with 5 nm of Pt. Samples with cells were fixed in 2.5 % glutaraldehyde for 2 h at 4 oC and
dehydrated. Samples were washed with deionized water and dehydrated in a microwave in serial
concentrations of ethanol (50 %, 70 % and 90 % once then 3x in absolute ethanol), before critical
point drying with carbon dioxide for 1h and then coating with 5 nm of Pt. Samples were imaged
using a Zeiss 1555 VP-FESEM.
Transmission Electron Microscopy (TEM). Synthesized polymer nanoparticles were dropcasted on carbon coated TEM grids and imaged with an accelerating voltage of 100 kV on a
S7
JEOL 2100 transmission electron microscope.
Atomic Force Microscopy. A Dimension 3100 AFM system (Bruker) with a Nanoscope IV
controller (Bruker) was used to obtain the AFM images in Contact Mode, using Pt/Ir coated
contact mode probes with a spring constant of 0.2 N/m (type SCM-PIC, Bruker). The scan
parameters were adjusted to ensure reliable imaging with the smallest possible contact force
setpoint. Data analysis was performed using the SPM analysis freeware Gwyddion
(http://gwyddion.net).
Dynamic light scattering (DLS) and zeta potential measurements. DLS experiments were
performed using a Malvern Zetasizer Nano series. For measuring the size distribution, 5
measurements were taken and in each measurement there were 10 data acquisitions. Zeta
potential (ζ) measurements were performed using the same instrument. Measurements for each
sample were recorded in triplicate and 100 data acquisitions were recorded in each measurement.
All measurements were recorded at 25 oC in Malvern disposable clear Folded Capillary Cells.
S8
Supplementary Figures
(a)
(b)
Figure S1. (a) The PDMS stamp used in the study. SEM micrograph of the grooved structure of
the PDMS. Image taken at 6k x magnification; (b) Photograph of the polycarbonate disc peeled
from a CD. PDMS was cast on the grooved surface and stamps of the desired size were cut out.
S9
T > Tg (PS)
(130 oC)
Peel off
PDMS Stamp
EA grafting
onto PGMA
Conducting NPs
Electrostatic
attachment
PS removal
Figure S2. Schematic of CFL procedure. Briefly, ITO substrate was modified with a thin layer
of PGMA followed by second layer of PS; a PDMS stamp was placed over the PS film and heat
treated at 130 oC; PDMS stamp was peeled off after cooling; EA was selective reacted to the
exposed PGMA stripes to produce cationic stripes to enable charge complementarity to assemble
the anionic PEDOT:PSS nanoparticles. The PS mask was removed by washing with toluene, to
obtain linear PEDOT:PSS conduits.
S10
Figure S3. SEM micrograph of PC12 cells 18 h after plating and immediately prior to electrical
stimulation. PC12 cells on the patterned surface (yellow box) were evenly spread out, in
comparison to the rounded cells on non-patterned areas of the substrate demonstrating
preferential adhesion. Image taken at 1000 x magnification.
S11
(a)
(b)
Figure S4. Representative low magnification (magnification 400 x) SEM micrographs of PC12
cells 72 h after the following treatments: (a) (+) pattern, (-) stimulation and (b) (+) pattern, (+)
stimulation demonstrating lower coverage due to reduction in proliferation upon stimulation.
.
S12
(b)
(a)
(c)
Figure S5. Photographs of the electrical stimulation setup: (a) A sterile Petri dish containing the
modified cell culture well (red arrow) and the Platinum wires which allow for connections to the
stimulator (green arrows), (b) The stimulator (yellow arrow) was placed next to an incubator and
the wires from the machine leading into the stimulator (blue arrows), (c) The wires from the
stimulator were connected to the platinum wires via alligator clips.
S13
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