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Thesis
Novel micellar systems for the formulation of poorly water soluble
drugs : biocompatibility aspects and pharmaceutical applications
DUMONTET MONDON, Karine
Abstract
Amongst the large number of novel drugs, 95% are lipophilic and poorly water soluble.
Particularly, this renders their aqueous formulation very difficult. In this regard this thesis
focused on polymeric micelles based on novel MPEG-hexPLA copolymers forming a
hydrophilic shell and a very hydrophobic core that favors the incorporation of poorly water
soluble drugs. Although the drug hydrophobicity and water solubility are the main parameters
in respect to their incorporation efficiency, structural parameters showed an important
influence, too. The biocompatibility of the novel MPEG-hexPLA micelles could be proven and
promising pharmaceutical applications can be envisioned. An intravenous application for
ovarian cancer diagnostics was demonstrated in vitro and in vivo in rats with the
administration of a micelle formulation containing the fluorescent compound hypericin. A
topical dermatology application with various antifungal compounds and possible use for oral
delivery were also evidenced.
Reference
DUMONTET MONDON, Karine. Novel micellar systems for the formulation of poorly
water soluble drugs : biocompatibility aspects and pharmaceutical applications.
Thèse de doctorat : Univ. Genève, 2010, no. Sc. 4258
URN : urn:nbn:ch:unige-171669
Available at:
http://archive-ouverte.unige.ch/unige:17166
Disclaimer: layout of this document may differ from the published version.
[ Downloaded 25/10/2016 at 04:54:18 ]
UNIVERSITÉ DE GENÈVE
FACULTÉ DES SCIENCES
Section des sciences pharmaceutiques
Pharmacie galénique
Professeur Robert Gurny
Docteur Michael Möller
Novel Micellar Systems for the Formulation
of Poorly Water Soluble Drugs :
Biocompatibility Aspects and Pharmaceutical Applications
THÈSE
présentée à la Faculté des sciences de l’Université de Genève
pour obtenir le grade de Docteur ès sciences, mention interdisciplinaire
par
Karine Dumontet Mondon
de
Lyon (FRANCE)
Thèse N°:4258
Genève
Atelier de reproduction Repromail
2011
A Maëlou, Léo et Bertrand
A ma famille, belle-famille et amis
Remerciements
Mes premiers remerciements s’adressent au jury qui a accepté d’évaluer mon travail de thèse :
Professeur Jean Marie Devoisselle (Université de Montpellier I, Montpellier, France),
Dr. Barbara Lueckel (F.Hoffman-La Roche AG, Bâle, Suisse) et Dr. Khadija Schwach
(Novozymes, Copenhague, Danemark).
Je tiens ensuite à remercier toutes les personnes qui ont contribué à cette recherche :
Le Professeur Robert Gurny et le Docteur Michael Möller pour avoir proposé un sujet aussi
passionant et touchant à de nombreux aspects pharmaceutiques que j’ai découvert au fur et à
mesure de la thèse. Merci Robert pour votre confiance, votre disponibilité et les conditions de
travail particulièrement favorables dont j’ai pu bénéficier. Merci Michael pour ton soutien
sans faille, tes idées, et aussi tes critiques.
Le Professeur Borkovec et le Docteur Andrea Vaccaro (Université de Genève, Départment de
Chimie Minérale, Analytique et Appliquée) pour la disponibilité et l’utilisation du
goniomètre.
Le Docteur Christoph Bauer (Université de Genève, Département de Biologie moléculaire,
Bioimaging Platform) pour les analyses en Microscopie à Transmission Electronique.
Le Professeur Carrupt, et les Docteur Sophie Martel et Bruno Bard (Université de Genève,
Pharmacochimie) pour la disponibilité du logiciel ACD/Chemsketch ainsi que pour leur aide
scientifique dans la détermination de la solubilité dans l’eau de certaines drogues.
Mes acolytes et ex-acolytes du labo 468, Thomas, Leila, Delia, Benjamin, Sema, Claudia,
Lutz et Sarra. Merci Claudia pour ton soutien, pour les innombrables petits services que tu as
faits pour moi qui m’ont énormément facilité la vie, merci aussi pour ton oreille attentive et
pour ton amitié. Merci Lutz pour ton humour, ta joie de vivre et ta serviabilité. Merci Sarra
pour tes corrections d’anglais, mais surtout merci pour nos petites pauses « papotage » qui
permettaient de souffler un peu pendant l’écriture du manuscrit. Merci à tous. Cela a été un
vrai plaisir de travailler à vos côtés, vous allez tous me manquer.
Mes collègues et ex-collègues, je pense particulièrement à Magali, Marie, Dany et Adriana.
Merci Mag de m’avoir initié au CAM, à la culture cellulaire et aux essais in vivo, qui ont
fortement contribué au « succès » de ma thèse. Merci également pour ton soutien (même à
distance) et aussi pour cette amitié que nous avons liée après de longues heures d’extraction
d’organes et de TP info … Merci Marie pour nos discussions sur comment allier vie familiale
et l’accomplissement d’une thèse, merci aussi d’organiser des petites sorties entre « mamans »
qui font tellement de bien au moral. Merci Dany aussi pour les discussions mais plutôt
«travaux » et « comment gérer nos filles respectives ! ». Merci Adriana pour ton soutien et ton
oreille attentive. J’ai beaucoup apprécié de pouvoir échanger nos sentiments de futures
mamans, et de mamans au cours de ces années de recherche. Je pense aussi au Docteur
Florence Delie pour sa disponibilité à répondre à mes questions d’ordre pharmaceutique mais
aussi pour tous les bons plans qu’elle a pu me conseiller. Merci Flo. Je tiens aussi à remercier
Emilie, pour les expériences de fluorescence avec le Nile Red. Merci pour ta disponibilité et
ton sourire, ainsi que les quelques heures de basket que nous avons pu faire ensemble. Je
n’oublie pas non plus le Docteur Norbert Lange pour avoir son humour sarcastique pendant
les TP info des étudiants. Je pense aussi au groupe PDT, toujours au summum dans
l’ambiance du labo et pour organiser des évènements; le groupe d’Archamps, toujours présent
malgré la distance géographique ; et enfin les nouveaux arrivants avec qui j’ai eu beaucoup de
plaisir à discuter pendant les derniers mois de la thèse (Flo, Julie, Bénédicte, Amandine…).
L’équipe technique et administrative, Myrtha, Florence, Brigitte et Marco. Merci Myrtha pour
ton support administratif et moral, encore plus appréciable en fin de thèse, merci aussi pour
les chocolats ou autres gourmandises que tu nous fais partager et qui sont toujours les
bienvenues à la cafet’. Merci Florence pour toujours avoir répondu à mes questions
administratives. Merci Brigitte pour ta serviabilité et ton aide précieuse dans la commande de
produits, la connaissance des fournisseurs, mais surtout pour les agréables moments à discuter
avec toi dans le labo ou à la cafet’. Enfin un énorme merci à Marco pour son assistance
technique sans limite, merci d’avoir toujours trouvé le temps de réparer la HPLC même si ton
emploi du temps était chargé.
Finalement, je tiens aussi à remercier mes amis, particulièrement Nath et Caro, ma bellefamille, ma sœur et mes parents pour leur présence, leur soutien et leurs encouragements tout
au long de ces 5-6 années.
Enfin, un spécial merci à Bertrand pour son amour qui nous a permis de construire une vie de
famille bien remplie avec la présence de nos deux magnifiques petits bambins.
Table of contents
Table of contents
INTRODUCTION…………………………………………………………………………………
1
CHAPTER I
Colloidal Drug Delivery Systems – Recent Advances with Polymeric Micelles……………....
3
CHAPTER II
Formulation of Poorly Water Soluble Drugs with MPEG-hexPLA Micelles: Investigations on
the
Effects
of
Drug
Physical-Chemical
Parameters
on
the
Incorporation…………………………………………………………………………………… 33
CHAPTER III
Novel Cyclosporin A Formulations Using MPEG-Hexyl Substituted Polylactide Micelles:
A Suitability Study…………………………………………………………………………….. 61
CHAPTER IV
MPEG-hexPLA Micelles as Novel Carriers for Hypericin, a Fluorescent Marker for Use
in Cancer Diagnostics………………………………………………………………………...... 91
CHAPTER V
Novel
Micellar
Formulations
to
Increase
Cutaneous
Bioavailability
of
Azole
Antifungals…………………………………………………………………………………..... 117
CHAPTER VI
MPEG-hexPLA Micelle Formulations for Oral Delivery of Poorly Water Soluble Drugs:
Investigations with Cinnarizine as a Model Drug ………………………………………….…. 139
CONCLUSIONS AND PERSPECTIVES………………………………………………………....... 149
SUMMARY (IN FRENCH)……………………………………………………………………… 153
ABBREVIATIONS……………………………………………………………………………… 155
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Introduction
1
Introduction
Technological advances in drug discovery such as high-throughput screening (HTS) have
considerably increased the number of “lead” compounds for a specific biological activity. The
majority of these compounds are poorly soluble in water, limiting their success in the clinic
because low solubility is often the cause of low bioavailability. Formulating such entities is
therefore a real challenge since stability problems are often encountered. Several strategies
have been developed to overcome this solubility issue, one of which is the incorporation of
the drug into polymeric micelles. The poorly soluble drug is loaded into the carrier core which
has a certain affinity with the drug and around a protective shell which prevents drug
precipitation and exposure to the immune system. To obtain such specific structures,
amphiphilic copolymers with a hydrophobic core forming block and a hydrophilic shell
forming block are used. Amongst a variety of suitable polymers, biocompatible and entirely
biodegradable ones are favoured. Poly(ethylene glycol) (PEG) is often the preferred
hydrophilic block because of its biocompatibility and non-immunogenic properties. For the
hydrophobic block, polylactides (PLA) and their derivatives are often used. However, for
PLA based copolymers, the incorporation of some poorly soluble drugs is limited 1. Alkyl
substituted polylactides were therefore developed to overcome this problem 2. By more
hydrophobic alkyl groups on the PLA backbone, the hydrophobicity of the resulting polymer
increases while preserving the biodegradability properties of PLA. Due to their physicochemical properties, two hexyl- substituted PLA (hexPLA): monohexyl- (monohex) and
dihexyl- (dihex) substituted PLA were selected for the here presented pharmaceutical
investigations. Copolymers of these new PLA derivatives with methoxypolyethylene glycol
(MPEG) were synthesised and able to form polymeric micelles 3;4. In this thesis, such MPEG-
Introduction
2
hexPLA copolymer micelles were investigated towards their biocompatibility aspects and
possible pharmaceutical applications. Recent advances with colloidal drug delivery systems
will be reviewed in chapter I. The incorporation of a poorly water soluble model drug into
MPEG-hexPLA micelles is presented as a proof of concept in this chapter. Chapter II reports
on the incorporation of several poorly soluble drugs into MPEG-hexPLA micelles and
describes the relationship between incorporation results and the physico-chemical properties
of the drugs, such as partition coefficient log P, water solubility, molecular weight, the
number of H bond- donor or acceptor groups, etc. As novel materials for the pharmaceutical
industry, the biocompatibility aspects of MPEG-hexPLA as unimers and as micelles were
investigated in chapter III in terms of in vitro and in vivo toxicity and hemolysis activity. A
possible intravenous application of cyclosporin A is presented as an example. In chapter IV,
intravenously injected MPEG-hexPLA micelles were studied in vivo in rats as a novel ovarian
cancer diagnosis tool. A promising application of these novel micelle formulations for
dermatology applications will be presented in chapter V. At the end the possible oral delivery
of poorly water soluble drugs with MPEG-hexPLA micelles and the initial proof of concept
studies will be presented in chapter VI.
References
(1) Ma, L. L.; Jie, P.; Venkatraman, S. S., Block Copolymer Stealth Nanoparticles for
Chemotherapy: Interactions with Blood Cells In Vitro, Adv.Funct.Mat. 2008, 18, 716-725.
(2) Trimaille, T.; Gurny, R.; Möller, M., Synthesis and ring-opening polymerization of new
monoalkyl-substituted lactides, J.Polym.Sci.Part A: Polym.Chem. 2004, 42, 4379-4391.
(3) Trimaille, T.; Mondon, K.; Gurny, R.; Möller, M., Novel polymeric micelles for hydrophobic
drug delivery based on biodegradable poly(hexyl-substituted lactides), Int.J.Pharm. 2006, 319,
147-154.
(4) Nottelet, B.; Di Tommaso, C.; Mondon, K.; Gurny, R.; Möller, M., Fully biodegradable
polymeric micelles based on hydrophobic- and hydrophilic-functionalized poly(lactide) block
copolymers, J.Polym.Sci.Part A: Polym.Chem. 2010, 48, 3244-3254.
1234567892AB 2CB D2236EB F783B 2656B 35B 7AB 2C8AB B 7B D56B 26E439B 4966B F98B
893B D9C9B 2366B 835853B 7AB B 7B D32499AB 7DD327B 82B D3D73B 7525B
C234567892AB2CB7578B2AA837892ABF98B22B879698EB
26B !"#BD26E439B4966BF98B98B9A37BE32D29B23B73B!D8B
82B 9A23D2378B 93B 7425A8B 2CB E32D29B 35B 87AB 8B 87A73B "#B
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Chapitre I
3
Colloidal Drug Delivery Systems – Recent Advances
with Polymeric Micelles
K. Mondon, R. Gurny, and M. Möller.
School of Pharmaceutical Sciences, University of Geneva, University of Lausanne, 30, Quai
Ernest Ansermet, CH-1211 Geneva 4, Switzerland.
Published in: Chimia, 62 (10), 832-840 (2008)
With the emergence of nanotechnology in drug delivery, colloidal systems and particularly
polymeric micelles have attracted great attention. Polymeric micelles formed by the selfassembly of amphiphilic copolymers adopt a core-shell structure, which can be loaded with
drugs and used as drug delivery systems for various medical applications. The most
interesting aspects involve extended blood circulation times and stability upon dilution, which
enable polymeric micelles to accumulate in tumour or inflammation sites due to the enhanced
permeation and retention effect (EPR). In the first part of this paper polymeric micelles with
different morphologies and different circulating-, active- and passive targeting, and stimuli
responsive properties will be reviewed. Furthermore amphiphilic block-copolymers of
different compositions for pharmaceutical micelle formulations will be discussed. The
hydrophilic block is often composed of the biocompatible polyethylene glycol (PEG),
whereas diverse polymers are used for the hydrophobic block. The biodegradable and
biocompatible polylactide (PLA) is one of the most tested core-forming blocks for micelles,
in part because of approval by the FDA for numerous drug products for use in humans.
However, PLA has limitations with respect to the incorporation of poorly water soluble drugs.
Considering this we will present in the second part of this paper briefly our strategy and
contribution to overcome these limitations and recent results for the envisioned application of
these micelles in the field of cancer treatment. In order to increase the hydrophobicity of PLA
methyl groups are substituted by more hydrophobic hexyl groups. The hexyl-substituted
polylactides in combination with PEG form the amphiphilic block copolymers PEG-hexPLA,
which self-assemble in aqueous solution into stable spherical, homogenous micelles with a
diameter of 20-45nm. The increased hydrophobicity of the hex-PLA micelle core leads to
Recent Advances with Polymeric Micelles
4
higher incorporation rates of hydrophobic drugs, like the poorly water soluble photosensitizer,
meso-tetra(p-hydroxyphenyl)porphine (THPP). THPP water solubility is increased 200-fold
using these micelles. Their application in Photodynamic therapy (PDT), coupled with the
increased accumulation of the photosensitizer in a tumour, followed by irradiation at a special
wavelength, should lead to tumour death. The encouraging drug incorporation results reveal
the potential of hexyl-substituted polylactide micelles as drug carriers for PDT applications.
Keywords: Biodegradable Polymers; Block Copolymers; Drug Delivery; Polymeric Micelles;
Substituted Polylactides
5
Chapitre I
1. Overview on polymeric micelles as drug delivery
systems
In the first part of this article we will give a brief overview on polymeric micelles as drug
delivery systems. Specific properties of polymeric micelles will be outlined with an emphasis
on the different compositions of widely used amphiphilic copolymers. In a second part, recent
developments from our own research on hydrophobic substituted PLA-based copolymers and
their possible use as micellar carriers of poorly water soluble drugs for cancer treatment will
be presented.
One very important aspect of pharmaceutical research is the development of novel drug
carriers, which can deliver drugs to the right pathological site, at the right time, in the right
dose, without affecting surrounding healthy tissues. In order to achieve this, nanotechnology
has gotten into the spotlight with the exploration of various colloidal drug delivery systems
like nanospheres, nanocapsules, micelles, liposomes, nanoemulsions, nanogels, lipid based
nanoparticles, dendrimer nanocomposites, and others 1. Among these systems, micelles have
gained increased interest in the last 15 years. The number of patents and publications is
increasing exponentially. The number of published papers in 1990 dealing with “polymeric
micelles” was about one hundred, whereas in 2007 it was 8 times higher
2
(Figure 1).
Numerous excellent reviews are available in the literature describing micellar carrier systems
for contrast agents in the imaging and diagnostic field 3, as drug carriers 4;5 for cancer 6-8, and
for gene and photodynamic therapy 9.
Number of publications and patents
per year
900
800
700
600
500
400
300
200
100
0
1980
1985
1990
1995
2000
6
2005
Year
Figure 1. Number of yearly publications and patents from 1980 to 2007 using the keyword “polymeric
micelles” (source: Scifinder Scholar, February 2008).
Polymeric micelles for pharmaceutical applications can be formed from amphiphilic
copolymers with different architectures (Figure 2). Composed of a hydrophilic A block and a
hydrophobic B block, these copolymers can either be of the linear AB diblock type, ABA or
BAB triblock- or multiblock type-, or a nonlinear composition having more complex
architectures like star or branched types. Consequently these copolymers form micelles of
different structures.
The self-assembly process is driven by the force to reach the lowest free-energy of the
system10. In solution the amphiphilic copolymers exist first as unimers until their
concentration exceeds the critical micellar concentration (CMC), above which they
spontaneously self-assemble into micelles to form colloidal systems of the nano scale. For a
given polymer concentration micelles form, when the solution temperature reaches the critical
micellar temperature (CMT). The lower critical solution temperature (LCST) is the minimum
temperature beyond which the polymers self-assemble into micelles. Thus both the CMC and
LCST are critical parameters for the existence of stable micelles in the body. The micelle
morphology is controlled mainly by 3 factors: 1) stretching of the core-forming blocks (free
energy of the core), 2) the surface tension between the core-forming block and the solvent
(free energy of the interface), and 3) repulsion interactions of the corona-forming blocks (free
energy of the corona) 10.
7
Chapitre I
Different morphologies are possible depending on the characteristics of the amphiphilic
polymers (molecular weight, physical state and composition) and on the solution parameters
(solvent, polymer concentration, pH, ionic strength, solvent/co-solvent ratio and others).
Amphiphilic diblock copolymers self assemble into basic shapes like spheres (the
predominant form), rods, and wormlike or vesicle structures. When the molecular weight of
the hydrophilic block represents more than 50% of the total polymer molecular weight,
spontaneous spherical micelles are formed, whereas with a hydrophilic block slightly below
50%, the micelles take on a wormlike structure. This latter morphology can also be generated
by sonication from a copolymer forming initially spherical micelles, as it was shown for
poly(ethylene glycol)-b-poly(-caprolactone) (PEG 5000g/mol -PCL 6500g/mol )
11
. In degradation
studies, wormlike micelles transformed into spherical micelles, when the PCL blocks
hydrolysed and the percentage of hydrophobic blocks dropped to less than 50% of the overall
polymer weight. The physical state of the core forming block also influences the micelle
morphology. PEG-poly(-caprolactone-D,L lactide) (PEG-P(CL-DLLA)) with the amorphous
poly(D,L lactide) block self-assembled into spheres, whereas cylindrical structures were
formed with the crystalline poly(L,L-lactide) block (PEG-P(CL-LLA))
12
. A strong
competition between the energy of the crystalline core and the chain stretching corona may
explain the formation of the cylindrical shape. Moreover the influence of the composition and
the block length has been demonstrated in the following examples; PEG 2000g/mol -PCL n formed
a variety of different morphologies with the increase of PCL block length from sphere, rod,
wormlike to lamellae, whereas PEG 5000g/mol -PCL n only self-assembled into spherical micelles.
One polymer, PEG 5000g/mol -PCL 232 , formed a mixture of spheres and lamellae 13. PCL also has
been used to study the effect of the polymer- block-type and -architectures in polymeric
micelles (Figure 2). Classical PEG 5000g/mol -PCL n diblock
13;14
or PCL-PEG-PCL triblock
15
copolymers self-assemble into the common spherical core-shell structure (Figure 2a and 2b,
respectively). BACAB type multiblock copolymers, where C is a less hydrophobic block than
B as in PCL n -(PEG-PPO-PEG) 1900g/mol -PCL n , leads to the formation of a double layer shellcore structure (Figure 2c) 16. Star-shaped copolymers like the 4 arm (PEG 5000g/mol -PCL) 4 17 or
Tetronic®-PCL
18
assemble into spherical micelles, wherein the hydrophobic blocks form the
core and the hydrophilic PEG chains arrange to form the shell (Figure 2d). Graft copolymers
like PAsp-g-PCL can form spherical micelles with a shell, in which hydrophilic chains
overlap, and entangle (Figure 2e) 19. PCL 7000g/mol –PDMA 8000g/mol brush copolymers with their
8
hydrophobic PCL-backbone form a kind a core-shell-“crosslinked” micelles (Figure 2f) 20. All
of the above mentioned micelles are characterised by sizes below 100nm. Yet, not all micelle
types have proved to be suitable drug carriers.
Diblock copolymers
Hydrophilic
block
Hydrophilic
shell
(a)
Hydrophobic
core
Hydrophobic
block
Triblock copolymers
(b)
Multiblock copolymers
(c)
Less
hydrophobic
block
1st Hydrophilic
shell
2nd Shell
More hydrophobic block
Double layer micelles
Star shaped copolymers
(d)
Graft copolymers
(e)
Brush copolymers
(f)
Figure 2. Different amphiphilic copolymer types and architectures leading to different micelle
structures by self-assembly.
Chapitre I
9
The influence of the solvent/co-solvent ratio on micelle shape has been explored by Barghava
et al.
21
. Using 2 different solvent systems, DMF/water and DMF/acetonitrile, PEG-
polystyrene (PEG-PS) copolymers self-assembled first into spheres and with the decrease of
DMF content into cylinders, followed by wormlike, and finally vesicle structures. Mixed
morphologies (spheres and cylinders, cylinders and wormlike structures) were visible at
intermediate DMF concentrations.
In aqueous media, micelles have a core-shell structure with a hydrophobic dense inner core
and a hydrophilic flexible outer shell. It has been shown, that many poorly water soluble
drugs can be incorporated efficiently into the hydrophobic core, and thus within the polymeric
micelles to become water soluble, facilitated by the hydrophilic micelle shell, which is
composed mainly of polyethylene glycol (PEG). Drug loading also influences the micelle
morphology. For example a change from spheres to cylinders occurred, when indomethacin
loading was increased in PEG-poly[2-(diisopropylamino)ethyl methacrylate] (PEG-PDPA)
micelles
22
. Further, wormlike micelles from poly(ethylene glycol)-b-poly(-caprolactone)
(PEG 5000g/mol -PCL 6500g/mol ) possessed the same diameter as the corresponding spherical ones,
but could incorporate twice as much paclitaxel, which is of great interest for an efficient drug
loading of the micellar carrier system 11. Moreover, the stability and cytotoxicity did not differ
between the two morphologies.
The presence of PEG, a hydrophilic, water soluble, nontoxic and non-immunogenic polymer,
at the surface of micelles enables nanocarriers to escape from renal exclusion and from
opsonisation by the mononuclear phagocytic system (MPS). Therefore micelles can protect
the incorporated drug from the biological fluids until it can reach the target site. With a
prolonged circulation time in the blood stream they can accumulate at the inflammation site or
in tumour tissues through the enhanced permeability and retention effect (EPR), which
resembles a passive targeting
23
. Due to their small size (below 100nm), polymeric micelles
can reach the very small vascular vessels and can be internalised into cells more efficiently
unlike bigger carriers such as microparticles
24;25
. The release of the incorporated drug at the
pathological site can occur by simple diffusion, destabilisation of the structure of the micelles
by degradation, change in pH or by other stimuli, which will be discussed in more detail
below. There are several other advantages of polymeric micelles in drug delivery. With their
low CMC in the micromolar range, polymeric micelles generally have good stability upon
10
dilution. Thus the amount of the polymer surfactant can be reduced in comparison to other
classical surfactants in order to formulate a stable drug carrier system. This higher stability of
polymeric surfactant based formulations results in a better shelf-life. The clear appearance and
the good injectability of micellar solutions are also advantages for medical treatments. The
protection of drugs in the micelle core decreases the adverse side effects observed with free
drugs, such as pain and inflammation at the injection site, and systemic side effects, when
administered parenterally 26. Propofol is an example of the improvement of a pharmaceutical
formulation with polymeric micelles. It is a water insoluble anesthetic agent, that was found
to be more stable in poly(N-vinyl-2 pyrrolidone)-block-poly(D,L-lactide) (PVP-PDLLA)
micelle solution than in the currently used water-in-oil emulsion. After reconstitution, the
micelle formulation is stable for 4 days, whereas the standard emulsion can only be used for
6h
27
. As a second example for advantages of polymeric micelles, a formulation with the
potent anticancer drug paclitaxel can be mentioned. Paclitaxel, a chemotherapeutic agent that
is almost insoluble in water (0.3g/mL), is currently marketed as Taxol® and solubilised
therein with Cremophor® EL, a polyoxyethylated castor oil. To replace this surfactant, which
is known for its toxicity, albumin bound paclitaxel nanoparticles (Abraxane®)
28
poly(D,L-lactic acid) polymeric micelles (Genexol® PM) have been developed
and PEG29
. These
formulations lead to a reduction of severe side effects, thus permitting to increase doses of the
chemotherapy. In vivo studies on human ovarian and breast cancer cells in nude mice
demonstrated, that tumour growth was reduced 48h and 14 days, respectively after
administration of Genexol® PM at its maximum tolerated dose (MTD), which is 3 times
higher than in the treatment with Taxol®. Moreover, tumour regrowth which was observed for
Taxol®
30
is not reported for the Genexol® PM treatment. After one month of treatment, this
polymeric micelle formulation showed complete tumour regression.
As mentioned before, drug targeting is a special challenge, since the delivery of a drug to the
right site of action should be both selective and quantitative. Passive targeting by the EPR
effect results from the long–circulation property of micelles. In order to achieve a higher drug
dose at the target site, substantial effort has been put into active targeting research. Active
targeting can be accomplished by modifying the micelle surface with site-specific ligands, or
by creating “immunomicelles” by attaching monoclonal antibodies. One strategy for drug
targeting uses folate functionalization. Receptors recognising the vitamin folic acid are over
11
Chapitre I
expressed on many human cancer cells (breast, ovarian, brain, kidney and lung), consequently
folate conjugated micelles are preferably recognised and bound on the tumour cells, and can
then be internalised by active tumour mechanisms. This mechanism was for example.
observed with adriamycin loaded folated polymeric micelles 31. The presence of these ligands
on the micelle surface increased the in vitro cytotoxicity by increasing the cellular uptake and
the intracellular concentration. The in vivo studies suggested that a defined number of folate
groups on the micelles surface showed an optimized carrier-receptor interaction. Tumour
treatment with these folated polymeric micelles was more efficient than with the non folate
functionalized micelles or even the free drug. Also, other potent drugs like paclitaxel
doxorubicin
35
, and tamoxifen
36
and a multidrug resistance modulator
37
32-34
,
have been
formulated with folate functionalized micelles and yielded improved results.
A recent paper by Beduneau et al. reviews the different active targeting strategies available
for treating brain tumours 38. One major obstacle for efficient drug delivery to the brain is the
blood brain barrier. Different modifications of the surface of nanocarriers (liposomes,
micelles or nanoparticles) by grafting endogenous and chimeric ligands, or by directly
conjugating proteins and peptides through a covalent or non covalent linkage leads to
internalization of the nanocarrier by brain capillary endothelial cells. Other examples of
surface functionalization of micelles for drug delivery can be found in the review of Mahmud
et al. 39.
Stimuli responsive micelles are also an approach of controlled drug release. When
nanocarriers have reached the pathological site, the drug can be released due to destabilisation
by different stimulations. This can be achieved with external sources like light, ultrasound,
hypo- or hyperthermia, or with internal stimuli like pH change or enzymes. pH-sensitive
micelles have been widely tested in drug delivery, because of increased acidity of tumour and
inflammatory tissues compared to healthy tissues. A change in the pH leads to a
demicellization process and the drug is released. It has been reported that doxorubicin loaded
pH-sensitive micelles formed from PEG-poly(-amino esters) disassembled at pH 6.4 leading
to rapid doxorubicin release, whereas at physiological pH they remained stable
40
. A higher
drug concentration at the pathological site was achieved, resulting in higher anti-tumour
efficacy and a higher survival rate compared to treatment with free drug. In other examples,
pH-sensitive polymeric micelles were loaded with adriamycin 41, paclitaxel 36 or tamoxifen 36.
12
Other poorly soluble drugs like triclosan
fenofibrate 44, and progesterone
44
42
, candersartan cilexetil
43
, indomethacin
44
,
could be successfully incorporated in similar micelles and
resulted in pharmacological improvements.
Thermoresponsive micelles can be used as controlled release systems. This is particularly the
case for poly(N-isopropylacrylamide) (PNIPAAm)-based micelles with a LCST of about
40°C
45
. Below this LCST, AB type copolymers with PNIPAAm as the hydrophilic block
form polymeric micelles, which are sterically stabilised by the water soluble flexible chains of
the PNIPAAm block 46. The lipophilic drug is incorporated in the hydrophobic core and can
be stored as a stable formulation. Above the LCST, PNIPAAm becomes insoluble and thus
deforms the micelle structure, which results in a rapid diffusion of the drug from the micelles.
In a cancer treatment study PNIPAAm-based polymeric micelles are accumulated in a solid
tumour site by passive targeting, and upon local heating slightly above the LCST the
encapsulated drug was released 47;48.
Ultrasound-sensitive polymeric micelles based on Pluronic® P-105 were developed by the
groups of Pitt and Rapoport in the late 90’s 49. Over time Rapoport and co-workers optimised
these micelles by mixing them with PEG-distearoylphosphatidylethanolamine (PEG-DSPE).
Mixed micelles demonstrated enhanced stability upon dilution compared to the standard
Pluronic® P-105 micelles. Indeed ruboxil, a paramagnetic labelled anthracyclin, showed
complete release in very diluted concentrations of fetal bovine serum when incorporated in
Pluronic® P105 micelles, whereas 65% of the drug was retained in the mixed polymeric
micelles 50. In addition, the treatment of tumours with doxorubicin as the drug was found to
be more efficient as indicated by a higher uptake and therefore a decrease of the tumour size,
when an ultrasound stimulus was applied to micelles in comparison to non-stimulated
micelles. The efficiency of the tumour treatment is even more pronounced when compared to
the treatment with the free drug
50
. In summary, polymeric micelles with their specific
structure and long circulation in the blood stream are promising carriers for passive and active
targeting, especially for cancer treatment.
13
Chapitre I
For the use of amphiphilic polymers in pharmaceutical applications in general, important
requirements such as non-toxicity and biocompatibility need to be fulfilled. This constrains
the number of available hydrophobic and hydrophilic natural or synthetic polymers.
For the hydrophilic shell-forming block polyethylene glycol (PEG) is the polymer of choice,
because of its superior properties, and its approval by the FDA. Despite its non degradability
PEG is non-toxic and can easily be removed from the body through the excretion pathways,
as long as the molecular weight is less than 15kDa. Thus PEG with molecular weights
between 2 and 15kDa are suitable in polymeric micelles for drug delivery. One of the main
advantages resides in the efficient protection of incorporated drugs in pegylated nanocarriers
51
. Other polymers like poly(N-vinyl-2-pyrrolidone) (PVP) 52;53, polyvinyl alcohol (PVA) and
its derivatives
54
, or recently studied poly(ethyl ethylene phosphate) (PEEP)
55
could be
alternatives to PEG. PVP a non ionic, biocompatible and water soluble synthetic polymer is
often preferred to PEG for freeze drying formulations, because of its better cryoprotectant
properties 56. Polyphosphates were recently investigated as polymers for micellar systems 55,
because of their biocompatibility, degradability and pendant chain functionality, but further
work is needed for a proof of concept.
For the hydrophobic core-forming block, a large number of polymers meet the requirements
for pharmaceutical applications. Here ionic or non-ionic, degradable or non degradable
polymers can be considered. In polyion complex micelles (PICM) a charged core-forming
block can more efficiently incorporate negatively charge drugs like plasmid DNA,
oligodeoxyribonucleotides, some polysaccharides, enzymes or photosensitizers. The typical
positively charged polymers are poly(ethylenimines) (PEI)
lysine)
60
57
58;59
, polyacrylamides
or poly(2-(N,N-dimethylamino)ethyl methacrylate) (PDMAEMA)
, poly(L-
61
. In contrast,
polyanionic polymers, like poly(methacrylic acid) (PMAA) 47 or poly(aspartic acid) (PAsp) 62,
are used for the incorporation of polycationic peptides or lipids 63.
Various non-ionic hydrophobic core-building degradable and non-degradable polymers have
been investigated. Here we will focus on the possible polymers from three different classes
including polyethers, poly(L-amino acids) and polyesters. Certainly there are other important
biodegradable polymers, like polyanhydrides or polyurethanes, but to date they are used more
as implants, microspheres, discs or other matrices for localised drug delivery
64;65
. A few
14
examples of polymeric micelles made from oligoanhydrides and PEG have been reported by
Najafi et al., who demonstrated the ability to incorporate a hydrophilic model drug, calcein 66.
The most interesting polyethers for micellar drug delivery systems are the triblock
copolymers (ABA type) of PEG and poly(propylene oxide) (PPO). They are marketed and
commonly known as Pluronics® (BASF)
67
or as Poloxamers, their non-proprietary name
68
.
Depending on the compositions and block-lengths PEG-PPO multiblock copolymers show
interesting material properties, including the formation of micellar structures or temperature
dependent gels. As for example Pluronic®-based micelles were studied by Gao et al.
Exner et al.
69
50
and
for cancer therapy. In these studies, doxorubicin and carboplatin, respectively
were successfully incorporated into the micelles. For doxorubicin loaded Pluronic® micelles
high drug uptake could be achieved in multidrug resistant cell lines through ultrasonic
irradiation of the tumour 50. An efficient tumour decrease was observed using this method.
Deriving from natural L-amino acids, poly(L-amino acids) are very interesting biocompatible
polymers and were studied as possible hydrophobic core-forming blocks
70
. Several poly(L-
amino acid)-based micelles have been investigated as pH-sensitive polymeric drug carriers.
At acidic pH in the tumour, the protonation of free amine and carboxyl groups in the polymer
chains lead to a destabilization of the micelle structure and resulted in the controlled release of
drug
71
. PEG-poly(L-histidine) micelles loaded with doxorubicin generated a decrease of
human ovarian carcinoma subcutaneously xenografted in mice 72. As another example PEGpoly(L-lysine) and PEG-poly(L-ornithin) micelles were studied for DNA delivery 73. A higher
condensation of DNA and a higher DNA transfection of two mammalian cell lines in vitro
were achieved with these micelle carriers compared to the parent homopolymers, poly(Llysine) or poly(L-ornithine). Disadvantageously the higher haemolytic activity could limit the
intravenous use of these novel vesicles.
As an example for non-pH sensitive poly(L-amino acids) polymeric micelles, benzyl ester of
PEG-poly(aspartic acid) (PEG-P(Asp)) micelles yielded a higher drug loading of the waterinsoluble anticancer agent camptothecin and better stability compared to other esterified
copolymers and to free drug
74
. In these studies mice, transplanted with colon solid tumour
cells, and treated with benzylester-PEG-P(Asp) micelles, had a longer circulation time, an
increased accumulation in the tumour, and thus a more efficient antitumour activity 23. Several
Chapitre I
15
hydrophobic core-forming poly(L-amino acids) could improve micelle stability and showed
promising results for cancer or gene therapy applications.
Poly(hydroxyalkanoic acids) are the most commonly used polymers in PEG-polyester
micelles for drug delivery, because of their outstanding biodegradability and biocompatibility.
Due to the vast number of published articles on PEG-polyester micelles we can only present
some selected examples of recent achievements, and would like to refer the readers to other
more detailed review articles on polymeric micelles 5;6;75;76.
Poly(butyrolactone) (PBL) is preferentially used in its  form, and is derived from the natural
production in microorganisms. -PBL is characterized by a higher hydrophobicity and
crystallinity compared to the synthetic -PBL. Despite a good stability, a low CMC and small
size, the potential of PEG-PBL-PEG triblock micelles (ABA type) was limited due to the
slow biodegradation of the PBL blocks in vivo
77
. Contrary the PBL-PEG-PBL copolymers,
BAB type, showed a faster biodegradation rate, while having similar sizes and CMC values
78
. However, the incorporation of hydrophobic drugs needs to confirm their feasibility for
drug delivery, even if pyrene as a model compound was incorporated with success.
Different studies on the semi-crystalline poly(valerolactone) (PVL) as the hydrophobic coreforming block were reported 79-81. As one example, PEG 2000g/mol -PVL 2000g/mol was used for the
formulation of paclitaxel
79
. With these micelles the water solubility was increased to
9mg/mL. When compared to the actual paclitaxel solubilisation of 6mg/mL in the mixture
Cremophor® EL/dehydrated alcohol in Taxol®, the result suggests that a lower amount of
PEG-PVL surfactant would be needed to achieve the same dose. In vivo studies need to show
that due to the slow degradation of PVL, and the expected longer elimination time from the
body will not be a problem.
Poly(caprolactone) (PCL) is another very suitable polymer for the core-forming block,
because of its proven biocompatible properties, therefore PEG-PCL diblock copolymers have
commonly been investigated for drug delivery. Recently, protoporphyrin IX, a hydrophobic
photosensitizer, was incorporated into PEG 5000g/mol -PCL 4100g/mol micelles, and a higher cellular
uptake and higher photocytotoxicity was obtained compared to the free photosensitizer 82. In
another example reported by Aliabadi et al., cyclosporine A, an immunosuppressive agent
16
was encapsulated in PEG 5000g/mol -PCL 13000g/mol micelles
83
. Here the drug solubility was
increased by 80 fold compared to its water solubility 84. High drug loadings in these micelles
envision the possible replacement of other surfactants, like Cremophor® EL in the current
formulation Sandimmun®, which shows severe side effects. In studies for cancer treatment
PEG-PCL polymeric micelles were loaded with chemotherapeutic agents like cisplatin
doxorubicin
85;86
and curcumin
, paclitaxel
87
and recently tested newer compounds such as -lapachone
20
,
86
88;89
. Other approaches focus on tumour specific targeting PEG-PCL carriers
functionalized with folic acid, heparin or epidermal growth factor (EGF) for the release of a
chemotherapeutics like paclitaxel 32, of both indomethacin and basic fibroblast growth factor
(bFGF)
90
or ellipticine 91. Higher internalisation and higher cytotoxicity were observed, and
thus improved significantly the tumour treatment compared to the non-functionalized
micelles.
Among the various poly(hydroxyalkanoic acids), poly(lactic acid) (PLA) and poly(lactic-coglycolic acid) (PLGA) copolymers are the most outstanding polyesters in drug delivery
applications.
Environmental
friendliness,
synthesis
from
renewable
resources,
biodegradability into non-toxic lactic acid, and excellent biocompatibility have made it the
polymer of first choice for many medical applications. Available stereochemically different
polylactides are characterized by different physical properties. The racemic D,L-PLA is
amorphous, whereas the enantiomeric pure L- or D-PLA (PLLA and PDLA, respectively) are
crystalline materials, resulting in different biodegradation times. Poly(lactide-glycolic acid)
(PLGA), which is a PLA functionalized with glycolic acid units or blocks in the polymer
chain, degrades faster than pure PLA and is often used to tailor the degradation time of PLA
drug carrier systems. Therefore depending on the polymer used these PLA-based micelles
have significantly different properties. PEG n -P(CL-DLLA) 8121g/mol had a lower CMC than
PEG n -P(CL-LLA) 6614g/mol , and both micelles adopted different morphologies 12. Mixing PEGPLLA and PEG-PDLA copolymer led to a stereocomplex formation of the PLA chains in the
micelle core, which had better micelle stability with a lower CMC, smaller size, and higher
incorporation efficacy of rifambin, than PEG-PLLA or PEG-PDLA micelles 93. In Genexol®
PM, PEG-PDLLA micelles are used as the drug carrier for paclitaxel 30. A multicenter Phase
II trial for the treatment of advanced non-small-cell lung cancer in combination with cisplatin
29
has been successfully finished, and the same formulation enters into a Phase II study for
advanced pancreatic cancer treatment 94. Other drugs like 5-fluoroacil 95, doxorubicin
86;96;97
,
17
Chapitre I
-lapachone 86;98, campthothecin 99 and amphotericin B 100 have been incorporated into PEGPLA micelles, and different polymer architectures like linear or star branched, diblock or
triblock polymers, homopolymer or mixed polymer micelles were investigated. Star branched
polymeric micelles showed no difference in comparison to linear polymeric micelles for 5fluoroacil release, whereas paclitaxel was released more rapidly and completely 95. In mixed
micelles with doxorubicin loaded poly(N-isopropylacrylamide-co-methacrylicacid)-graftpoly(D,L-lactide) (P(NIPAAm-co-MAAc)-g-PDLLA) the PEG-PLA did not affect the pHand thermoresponsive properties, but advantageously prevented the adsorption of albumin,
increasing the circulation time in the blood stream
97
. PLA-PEG-PLA triblock copolymers
with acryl end groups self-assembled into micelles, which were transformed into nanogels by
UV irradiation
99
. Loaded with campthothecin these nanogels showed better stability and a
sustained release for at least 20 days in contrast to complete release for non-irradiated
micelles after few hours. Hydrophilic shell-forming blocks other than PEG have been
explored for PLA- or PLGA- based copolymer micelles. Recently PLA-poly(L-glutamic acid)
micelles were developed for magnetic resonance imaging (MRI) diagnostic systems
101
.
Gadolinium (Gd) ions chelated with DTBA were incorporated into these micelles as an MRI
probe. They were stable upon dilution and showed a two times higher relaxivity compared to
the free DTBA-Gd complex. Further in vivo studies should evaluate the feasibility of these
micellar diagnostic systems for their application in patients. Poly(L-lysine) could be another
alternative polymer to PEG as the shell-building block. Grafted on PLGA, the resulting
amphiphilic polymer self-assembled into micelles, which showed lower cytotoxicity and
higher transfection efficiency than poly(L-lysine) itself, and could be a possible carrier for
gene delivery 102. In comparison to PEG, these poly(L-amino acid) based micelles are entirely
biodegradable. Stimuli responsive micelles based on PEG- poly(L-histidine) (PEG-PHis)
copolymer micelles were also investigated. Upon a change of pH in tumour tissue, the
imidazole side groups along the poly(L-histidine) backbone are ionized (pH<pK b ),
transforming the hydrophobic PHis-block into an hydrophilic water soluble one. This is
accompanied by a swelling of the micelles and a concurrent drug release
103
. Loaded with
doxorubicin PLA-PEG-PHis micelles showed an effective treatment of human breast
tumours96. Thermoresponsive micelles made of PLA-based polymers were studied with
(PNIPAAm-co-DMAAm)-PLGA copolymers for the delivery of chemotherapeutic agents like
paclitaxel
104
and doxorubicin 59. These micelles showed a higher drug release upon thermo-
stimulation above the LCST (39.5°C) and a higher cytotoxicity on tumour cells compared to
18
free drug. As an example for tumour targeting micelles, folate conjugated PEG-PLGA
micelles were studied by Yoo et al.
105
and Zhao et al.
35
. A higher uptake, higher
cytotoxicity, and higher apoptosis were obtained with these folate-functionalised micelles
compared to the non-folated micelles and the free drug, respectively 35.
PEG–polyester micelles are potent drug carriers compared to many classical surfactant based
micelles. A higher stability in diluted solutions due to their lower CMC, PEG copolymer
micelles have long circulation properties and can be functionalized as targeting or as stimuli
responsive drug release carrier systems. An increased accumulation, a higher drug uptake by
the tumour or inflammation cells accompanied with a more effective tumour treatment can
often be achieved with polymeric micelles. Regarding this, albeit an incomplete list, PCL,
PLA and PLGA are the most promising polyesters. Because of their approval by the Food and
Drug Administration (FDA) for drug products in human use, new formulations with these
pegylated excipients might lead easier and faster to new applicable therapeutics.
2. Recent developments of hydrophobic substituted PLA
based polymeric micelles
Despite the outstanding applications of PLA in medical applications, limitations for the
incorporation of hydrophobic drugs for pharmaceutical formulations are often encountered.
Since a more hydrophobic core is desirable for a better incorporation into PLA-based
polymeric micelles, we focused in our own research efforts on the controlled functionalization
of PLA with hydrophobic substituents. Therefore methyl groups along the PLA polymer
backbone were substituted by hexyl groups leading to the more hydrophobic hexyl-substituted
poly(lactides) (hexPLA)
106
. PEG-hexPLA copolymers formed micelles in aqueous solutions
and incorporated hydrophobic drugs, e.g. griseofulvin, more efficiently than comparable
PEG-PLA micelles
107
. Recently we investigated the potential of these novel PEG-hexPLA
micelles for possible use in photodynamic therapy (PDT) in cancer treatment. Initial results in
19
Chapitre I
incorporating the hydrophobic photosensitizer meso-tetra(p-hydroxyphenyl)porphine (THPP)
(Figure 3), and stability studies of THPP loaded PEG-hexPLA micelles will be presented.
OH
HO
N
NH
HN
N
HO
OH
Figure 3. Molecular structure of meso-tetra(p-hydroxyphenyl)porphine (THPP).
In order to compare PEG-monohexPLA and PEG-dihexPLA micelles
107
for their increased
hydrophobicity and a possible higher incorporation rate of THPP, a comparable PEG“standard” poly(D,L-lactide) (PEG-PDLLA) was used as a control. All three copolymers were
of comparable molecular weights of around 5200g/mol and a comparable polydispersity of
1.2. The PEG-(hex)PLA polymeric micelles were prepared by the co-solvent evaporation
technique. Both the drug and the copolymer were dissolved in the water-miscible organic
solvent acetone or THF/acetone (1:1). After the drop wise addition of the organic solution into
water under sonication, accompanied by the self-assembly to micelles, the organic solvent
was slowly evaporated. After equilibrium overnight of the remaining aqueous solution the
micelles were characterized by their sizes, morphologies and drug incorporation. Unloaded
PEG-(hex)PLA micelles had sizes between 18-27nm, whereas THPP loaded micelles were
only slightly larger with 21-45nm. Thus the loaded micelle sizes remained below 50nm, a
preferred “sub-100nm” size, facilitating a good internalisation into tumour cells by the EPR
effect. Dynamic light scattering at different detection angles confirmed a unimodal size
distribution for loaded and unloaded PEG-hexPLA and PEG-PLA micelles. The micelles
sizes and morphologies were confirmed by TEM measurements. As shown in Figure 4
unloaded and loaded PLA-based polymeric micelles had comparable sizes with spherical
shapes.
(a)
20
(b)
100 nm
(c)
100 nm
(d)
100 nm
100 nm
Figure 4. TEM images of unloaded (a) and THPP loaded PEG-monohexPLA (b) micelles, and for
comparison unloaded PEG-dihexPLA (c) and PEG-PLA (d) micelles
A study of polymeric micelle solutions stored at room temperature for 10 months proved their
stability and the long shelf-life of this formulation. The amount of THPP in the diverse
hexPLA- and PLA-based micelles was assessed by UV spectroscopy. The results for THPP
incorporation into these micelles are presented in Figure 5.
21
Chapitre I
100% efficiency
Incorporated drug loading
(mgTHPP/gcopolymer)
200
80%efficiency
150
100
50
0
0
25
50
75
100
125
150
175
200
225
250
275
300
Intended drug loading
(mgTHPP/gcopolymer)
Figure 5: Incorporation of THPP in PEG-PLA (○,●), PEG-monohexPLA (□,■) and PEG-dihexPLA
(,) micelles in function of the intended drug loading for the two preparation methods, with acetone
(empty symbol) and THF/acetone (plain symbol), respectively
Shown is the actual achieved loading in dependence of the intended loading, which is based
on the drug concentration of the organic solution used for the micelle preparation.
Furthermore, the two lines indicate a theoretical 80 and 100% incorporation, demonstrating
the efficiency of the THPP incorporation. THPP incorporation of micelles prepared with
acetone and with THF/acetone, respectively as solvent systems are presented. For intended
small loadings up to 50mg THPP/g copolymer no significant differences were observed
between the three different polymers and the organic solvents used. Incorporation rates were
at least 80%. For a desired higher drug loading of 100mg THPP/g copolymer and 300mg
THPP/g copolymer, respectively better incorporation rates were obtained when using a
polymer/drug solution in the THF/acetone (1:1) mixture, which facilitates drug solubility. It is
to point out, that in both cases the incorporation was higher and much more efficient with the
novel PEG-hexPLA micelles than with the standard PEG-PLA micelles. The highest obtained
drug loading was 123mg THPP/g PEG-dihexPLA, although with less efficiency (41%)
compared to 96 and 88 mg THPP for the dihexPLA and the monohexPLA, respectively,
which achieved around 90% efficiency for the intended lower drug loading. Since the
molecular weights and ratios of PEG/hexPLA are not optimized, it is expected, that the
loading capacity for THPP could be higher. Nevertheless the actual results already show a
22
prominent increased water solubility of THPP in PEG-hexPLA micelles. Within the
dihexPLA micelles 480mg THPP/L water could be dissolved compared to only 2.2mg/L in
pure water, which corresponds to an increase of 218 fold! As mentioned previously, this result
still stands for non-optimized conditions, thus further developments in PEG-hexPLA micellar
formulations for THPP delivery could lead to a promising alternative for PDT applications in
cancer treatment. The increased THPP drug loadings within the hexyl-substituted PLA-based
micelles in comparison to standard PLA validate our strategy of improving hydrophobic drug
formulations with “hydrophobized” PLA. As outlined in our previous papers the hexPLA
polymers degrade by hydrolysis to non-toxic lactic acid and 2-hydroxyoctanoic acid, which
has already been approved for topical applications, and which is an important aspect for the
feasibility of this polymer for medical applications. Therefore, on-going studies focus on the
toxicity issues of these novel PEG-hexPLA micelles. Initial cell culture- and human blood
toxicity tests show positive results and will soon be reported.
3. Summary
Polymeric micelles as colloidal drug delivery systems have attracted great interest in the last
15 years due to their favorable properties. Formed by the self-assembly of amphiphilic
copolymers, polymeric micelles have a core-shell structure, which can adopt several
morphologies. The protective hydrophilic shell gives the micelles long circulation times in the
blood stream, and a low CMC a high stability upon dilution. This can enable micelles to
accumulate in tumours by passive targeting and the EPR effect. Active targeting to the site of
action can be achieved by binding specific ligands or antibodies on micelle surfaces. Using
stimulus responsive polymeric micelles can enhance the controlled drug release. For drug
delivery, PEG is the most outstanding and used hydrophilic polymer for the micelle shell.
Considerations regarding toxicity and biocompatibility limit the choice of numerous
hydrophobic polymers for the core-forming block. Next to several polyethers, poly (L-amino
acids), polyesters, particularly PLA has distinguished properties for medical applications.
Nevertheless, the incorporation of hydrophobic drugs into PLA micelles often has limitations.
In order to increase the hydrophobicity along the PLA backbone for better incorporation of
Chapitre I
23
hydrophobic drugs we followed the strategy to substitute the methyl with more hydrophobic
hexyl side groups. Hexyl-substituted lactides copolymerized with PEG lead to amphiphilic
copolymers, which self-assemble in aqueous solution into homogeneous spherical micelles
ranging in sizes from 20 to 45nm. These PEG-hexPLA micelle formulations showed 10
months of stability at room temperature. The incorporation of a poorly water soluble drug,
like the photosensitizer THPP, did not affect the micelle size or the micelle morphology.
However, by incorporation into PEG-hexPLA micelles the water solubility was increased
more than 200 times compared to the water solubility of the free drug. These initial results
suggest that the use of PEG-hexyl-substituted polylactides micelles as drug delivery carriers
for photodynamic therapy is feasible.
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32
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Chapter II
33
Note
Formulation of Poorly Water Soluble Drugs with
MPEG-hexPLA Micelles:
Investigations on the Effects of Drug Physical-Chemical
Parameters on the Incorporation
K. Mondon, R.Gurny, and Michael Möller.
Ernest Ansermet, CH-1211 Geneva 4, Switzerland
Today, over 9 out of 10 pharmaceutical lead products are lipophilic and poorly water soluble,
which render their formulation difficult and limits their clinical use. In this regard,
solubilisation of such drugs into micelles made of methoxypoly(ethylene glycol) hexylsubstituted poly(lactide) (MPEG-hexPLA) was investigated as an alternative pharmaceutical
formulation. The incorporation efficiency of 18 different drugs into MPEG-hexPLA micelles
was studied. Relationships and trends between the known physical-chemical properties of the
18 drugs and their incorporation capacity were examined. All drug loaded micelles had a size
below 200 nm. The incorporation of these drugs was generally higher in MPEG-hexPLA than
in the comparable MPEG-polylactide (MPEG-PLA) micelles. For some drugs the aqueous
solubility could be improved up to 100 times. The physical-chemical parameters, water
solubility, log P, number of H bond acceptor (Ha) and H bond donor (Hd) were found to
strongly influence drug incorporation into the micelles. Concerning the aqueous drug
solubility and the aqueous solubility increase with micelles, the drug water solubility was the
most influencing parameter followed by Hd, Ha, and log P. Hd was the most predominant
parameter for the drug/copolymer ratio in the micelles, followed by log P, Ha, and the drug
water solubility. The drug molecular weight and the surface tension were found as minor
Incorporation of Poorly Water Soluble Drugs into MPEG-hexPLA Micelles
34
influencing parameters. A qualitative trend was observed for drugs with low water solubility
(< 0.055 mM). They showed a high increase of their aqueous solubility when incorporated
into MPEG-dihexPLA micelles. Amongst the variety of investigated drugs, some drugs do not
fit the Lipinski’s “rule of 5” but have suitable physico-chemical parameters for an efficient
incorporation into MPEG-hexPLA micelles. This envisions the use of micelle solutions as an
alternative formulation approach for these poorly water soluble drugs.
Keywords: polymeric micelles; prediction; poorly water soluble drugs; solubilisation,
substituted polylactides.
Chapter II
35
1. Introduction
In the past two decades drug discovery has been accelerated by the use of high throughput
screening (HTS) and combinatorial chemistry methodologies. Amongst the large library of
novel active compounds 95% are lipophilic and poorly water soluble 1, which is often the
critical point to their success in the clinics 2;3. Different strategies to improve solubilisation of
poorly water soluble drugs have been developed. Co-solvency is one approach to enhance
water solubility 4 with the advantage of the protection of the drug from hydrolysis. However,
unstable formulations upon dilution in the blood can lead to local drug precipitation after
administration, inducing serious side effects like pain, inflammation or haemolysis
5;6
.
Alternatively a complexation with carriers like cyclodextrins can be used, but its application
is limited by high toxicity at high concentrations
7
and by the susceptibility of the complex
dissociation upon dilution. Oil-in-water emulsion is another approach for very lipophilic
drugs, however the use of low molecular weight surfactants could lead to some toxicity issues
like it is the case for Cremophor®EL
8;9
. Polymer based drug delivery systems are a further
alternative choice for the formulation of poorly water soluble drugs. Amongst the variety of
applied biocompatible and biodegradable, respectively polymers 10 (poly(ethers), poly(esters),
poly(L-amino acids etc.) with different structure (linear, dendrons, multi block, etc.),
amphiphilic diblock copolymers have the particularity to self-assemble in aqueous media into
a specific micelle core-shell structure. Such nanocarriers can incorporate a drug in their core,
preventing its precipitation in an aqueous environment. The nanosize and pegylation of
studied carrier systems facilitate long-circulation, and ability to cross certain membranes and
passively targeting diseased tissues. The hydrophilic corona most often composed of the
biocompatible poly(ethylene glycol) (PEG) limits protein adsorption at the surface and
therefore warranties long circulating properties
11
. In addition polymeric micelles with sizes
below 200 nm can accumulate at sites of leaky vasculature like inflamed or cancer
tissues 12;13, which make them interesting for a passive targeting
14
. Concerning the
hydrophobic core, usually poly(ethers), poly(L-amino acids) and poly(esters) are used 15. The
micelle core forms a reservoir into which a drug can be incorporated and consequently the
micellar system enhances significantly its aqueous solubility. The amount of drug that can be
incorporated within the core depends principally on the compatibility of the drug with the
36
polymer of the core forming block, as outlined for PEG-poly(caprolactone) (PCL) 16-21, PEGPCL and PEG-poly(lactide) (PLA)
vinyl pyrrolidone-styrene)
16-21
16-21
, poly(ethylene oxide-propylene oxide) and poly(N-
. In our own work, cyclosporin A was incorporated into
MPEG-hexPLA micelles with different degrees of hydrophobicity of the core forming block
21
. The results revealed that the more hydrophobic the core, the higher the drug incorporation.
Better affinity between the core and the hydrophobic drug was created with the presence and
increasing number of hexyl groups in the PLA core forming block, as shown for the di-hexyl
substituted PLA (dihexPLA) compared to the mono-hexyl substituted PLA (monohexPLA)
and PLA. The affinity of a drug to the core forming polymer can be estimated with the
theoretical Flory-Huggins parameters (χsp), via the calculation of the theoretical Hildebrand or
Hansen solubility parameters of both the drug and the polymer
22
. The difference obtained
between the solubilisate (the drug) and the solvent (the polymer) indicates the miscibility of
both components. The closer χsp is to 0, the better is the miscibility reflecting a higher affinity
of the drug and solubility in the given polymer, here resembling its incorporation efficiency
into the micelle core.
To evaluate a possible prediction or estimation of the drug incorporation into MPEG-hexPLA
micelles, 18 poorly soluble drugs were tested. The theory of Flory-Huggins was first applied
to each drug/copolymer system to evaluate the different affinities. Relationships and trends
between the known physical-chemical parameters of the drugs and the experimentally
obtained incorporation results (i.e. drug content in micelle solution, drug aqueous solubility in
micelle solution, aqueous solubility increase) are presented.
2. Materials and methods
2.1. Materials
Chemicals
Tetrahydrofuran (THF) was supplied by SDS (Toulouse, France), and dried by distillation.
Methoxy-poly(ethylene glycol) with a molecular weight of 2000 g/mol (MPEG2000g/mol) was a
gift from BASF (Ludwigshafen, Germany). D,L-lactide, tin(II) 2-ethylhexanoate (Sn(Oct)2),
and acetone p.a. were purchased from Purac Biochem (Gorinchem, The Netherlands), Aldrich
(Buchs, Switzerland), and Fluka (Buchs, Switzerland), respectively and used as received. The
Chapter II
37
monomers, mono-hexyl-substituted lactide (monohexLA) and di-hexyl-substituted lactide
(dihexLA) were synthesized as described previously 23.
Dichloromethane (DCM) (HPLC grade) and toluene were supplied by Fluka (Buchs,
Switzerland). Methanol (MeOH) and acetonitrile (ACN) (HPLC grade) were purchased from
VWR (Nyon, Switzerland). Ammonium dihydrogen phosphate, di-potassium hydrogen
phosphate, sodium chloride, acetic acid, pyrophosphoric acid, phosphoric acid, formic acid
were supplied by Fluka, (Buch, Switzerland). Diethylamine was purchased from Acros
organics (Geel, Belgium)
Drugs
Clotrimazole (CLZ), Danazol (DAN), Diclofenac sodium (DICLO), Econazole nitrate (ECZ),
Etoposide (ETO), Fluconazole (FLZ), Griseofulvin (GF), Ketoconazole (KETO), Piroxicam
(PIRO), Quercetin dehydrate (QUER) and Naproxen sodium (NAPRO) were all supplied by
Sigma (Buchs, Switzerland). Cyclosporin A (CsA), Docetaxel (DTX), Hydrocortisone
(HYDRO), Hypericin (HY), Paclitaxel (PTX), and meso-tetra(p-hydroxyphenyl)porphine
(THPP) were purchased from Fluka, AlfaChem (Kings Point, NY, US), from Aldrich, from
Alexis Corporation (Lausen, Switzerland), Oskar Tropitzsch (Marktredwitz, Germany), and
Frontier Scientific (Carnforth, UK), respectively.
Verteporfin (VER) was extracted from Visudyne (Novartis Pharma, Bern, Switzerland)
purchased from a local pharmacy.
2.2. Synthesis and characterization of MPEG-(hex)PLA copolymers
The MPEG-(hex)PLA copolymers with a molecular weight of around 5000g/mol were
synthesized by ring opening polymerization (ROP) with MPEG 2000 as initiator as described
previously
21;24
, see Scheme 1 for their chemical structures. The copolymers were
characterized by their molecular weight (Mn) and polydispersity index (P.I.) using gel
permeation chromatography (GPC). The GPC setup was composed of a Waters system with
Waters Styragel HR1-3 columns and a Waters 410 differential refractometer (Waters,
Milford, USA). The analysis was carried out using polystyrene (PS) of different molecular
weights as calibration standards (PSS, Mainz, Germany). The degree of polymerization of
MPEG-(hex)PLA was calculated from 1H NMR (CDCl3, Bruker 300 MHz).
38
Please note that in the following text, “MPEG-(hex)PLA” refers to both MPEG-PLA and
MPEG-hexPLA copolymers.
O
O
H3C
45
R1, R2 = CH3
1
R
2
O
O
R
1
O
H
m
O
: MPEG2000-PLA
2
R =C6H13, R =CH3 : MPEG2000-monohexPLA
R1, R2 =C6H13
: MPEG2000-dihexPLA
Scheme 1. Structure of MPEG-hexPLA and MPEG-PLA block copolymers
2.3. Preparation of drug loaded MPEG-(hex)PLA micelles
The preparation and incorporation of drugs into MPEG-(hex)PLA micelles followed the cosolvent evaporation method. Six milligrams of drug and 20 mg of copolymer were dissolved
in 2 mL acetone or THF. The organic solution was added dropwise under stirring into 4 mL of
ultra pure water using a peristaltic pump. The organic solvent was removed under stirring by
evaporation at 200 PSI in a desiccator for 2h under vacuum at 200 mbar. Final micelle
concentrations were adjusted to 5 mgcopolymer/mL by adding ultra pure water. The solutions
were left to equilibrate overnight and then centrifuged at 9500 x g for 15 min to remove nonincorporated drug. Afterwards, drug loaded MPEG-(hex)PLA micelles were analyzed for
their size and their drug content.
2.4. Size determination of drug loaded MPEG-(hex)PLA micelles
The hydrodynamic- (Zav) and number-weight (dn) diameters of drug loaded MPEG-(hex)PLA
micelles were measured by dynamic light scattering (DLS) using a Zetasizer HS 3000
(Malvern, Worcestershire, UK). Analyzes were performed at an angle of 90° at 25°C. For
each sample, the mean diameters were obtained in triplicate.
Chapter II
39
2.5. Drug loading determination of drug loaded MPEG-(hex)PLA micelles
The drug loading of MPEG-(hex)PLA micelles were determined after centrifugation of
micelle formulations. The supernatant was diluted in acetonitrile at a ratio 1:10 to break up
the micelles and to release the drug for quantitative analysis.
The drug loading of MPEG-(hex)PLA was mainly performed by HPLC for the majority of the
micelles (see Supplementary Material for the different HPLC conditions). For VER and
THPP loaded micelles UV absorbance was used to quantify the drug content.
The drug content (DC), the drug solubility in micelles ([Drug]micelles) and the “solubility
increase factor” were calculated using the following equations:
Drug content [mgdrug /g copolymer ] =
[Drug]micelles[mM] =
mass of drug incorporated in micelles [mg]
mass of copolymer used [g]
DC [mgdrug /g copolymer ] × concentration of copolymer [g/L]
Molecular weight of the drug [g/mol]
"Solubility increase factor" =
[Drug]micelles [mM]
Drug water solubility [mM]
(1)
(2)
(3)
2.6. Determination of the Flory-Huggins parameter χsp
The Flory-Huggins parameter reflects the compatibility between the drug and the (hex)PLA
micelle core. It was calculated by the following equation:
χ sp = (δ drug −δ copolymer )2 ×
V
RT
(4)
δdrug and δpolymer are the solubility parameters for the drug, and for the core-forming block of
the copolymer, respectively. R is the gas constant, T the temperature in Kelvin, and V the
molar volume of the drug.
Values for δdrug and δpolymer were calculated by the additive group contribution method as
described by Van Krevelen
25
using the Material Studio Software (Accelrys Inc., USA). The
molar volume V of the drug was calculated by the group contribution method (GCM)
according to Fedors 26 using the same software. The temperature was set at 298 K (25°C).
40
3. Results and Discussion
Due to the large number of potent drugs with low water solubility and high lipophilicity, and
the resulting limitations for pharmaceutical formulations, polymeric micelles have attracted a
lot of attention as possible drug delivery systems
19
. With their specific core-shell structure,
micelles can incorporate a drug in their core, while the hydrophilic corona renders aqueous
solubility. Considering that the affinity of a drug and the core-forming block is affecting the
incorporation of a drug into micelles, novel MPEG-hexPLA based micelles with an increased
hydrophobicity of the micelle core were developed 24. These MPEG-hexPLA micelles showed
higher drug loadings, lower critical micellar concentrations and higher stabilities compared to
the standard MPEG-PLA micelles. As an example, the immunosuppressant drug
cyclosporin A (CsA) was 4 times more incorporated into MPEG-dihexPLA than into MPEGPLA micelles, leading to an aqueous solubility of 1.4 mg/mL
21
. MPEG-dihexPLA micelles
showed also a superior stability at body temperature, both in water and in PBS pH 7.4
27
, in
comparison to the less hydrophobic MPEG-monohexPLA micelles and the standard MPEGPLA micelles. To further generally examine the incorporation efficiency of MPEG-hexPLA
and MPEG-PLA micelles, respectively, 18 poorly water soluble drugs with different physicochemical properties were tested. The experimental results were then correlated with the FloryHuggins parameter and the physico-chemical parameters of the drug regarding the obtained
drug content (drug/copolymer ratio) in micelle solution, the drug solubility in micelle solution
and the aqueous solubility increase (ratio between drug aqueous solubility obtained in
micelles over drug water solubility).
The used copolymers, MPEG-monohexPLA and MPEG-dihexPLA, and the “standard”
MPEG-PLA were synthesized in a controlled manner by ROP with a defined molecular
weight of around 5000 g/mol (± 1 monomer unit) (Table 1), thus allowing a good comparison
of the influence of the hexyl substituents on the drug incorporation.
Chapter II
41
Table 1. Characteristics of MPEG-(hex)PLA copolymers
Copolymer
Mn a (g/mol)
P.I. b
# hexyl groups
MPEG-PLA
5050
1.17
0
MPEG-monohexPLA
5040
1.13
14
MPEG-dihexPLA
5250
1.15
23
a
determined by 1H NMR; b determined by GPC.
The investigated drugs (Figure 1) were selected to cover a wide range of lipophilicity values
(log P), and of different physical-chemical characteristics (Table 2). They are all known to be
poorly water soluble. The chosen drugs are used in various pharmaceutical applications such
as cancer-, transplantation rejection-, dermatology-, or ophthalmology treatments.
42
O
N
N
N
N
HO
O
N
N
N
N
O
N
Cinnarizine
CIN
O
H
Clotrimazole
CLZ
H 3C
O
O
H
N
N
O
N
O
O
Cl
O
H
N
H
N
H 3C
HO
N
N
O
CH
O
Danazol
DAN
Cyclosporine A
CsA
HO
HO
+
NH
O
Na -
O
NH
O
OH
O
O
Cl
O
O
Cl
H
H
O
Cl
N
H
O
O
N
O OH
O
Cl
O
HO
OO
N
OH
O
O
O
O O
H
O
O
HO
O
Cl
HO
Docetaxel
DTX
Diclofenac sodium
DICLO
Econazole nitrate
ECZ
Etoposide
ETO
O
O
Cl
Cl
OH
N
N
Cl
N
MeO
HO
HO
O
N N
F
OMe
N
O
O
HO
O
HO
OMeO
O
N
F
OH
O
OH
N
O
Hypericin
HY
Hydrocortisone
HYDRO
Griseofulvin
GF
Fluconazole
FLZ
H3C
Ketoconazole
KETO
OH
HO
O
O
O OH
OH
O
O
OH
H3C
O
N
H3C HO
OH
O
H
NH
O
H
N
NH
NH
N
O
O
OH
N
O
OO
O
S
HN
N
O
OH
O O
OH
HO
Piroxicam
PIRO
Paclitaxel
PTX
THPP
MeO
OMe
H3C
OH
OH
O
O
HO
O
HO
OH
OH
+
Na
O
Quercetin dihydrate
QUER
O
O
CH3
.2H 2 O
O
O
-
Naproxen sodium
NAPRO
Figure 1. Structure of all tested drugs
O
N
H
N
N
H
N
Verteporfin
VER
O
N
Chapter II
43
Table 2. Characteristics of the investigated drugs
Drug
Mw
[g/mol]
Exp.
logP
CLZ
344.80
6.1 a
Exp.
water
solubility
[mM]
0.0880 b
CsA
1202.60
8.2 c
0.0100 c
23.389
5
17
31.6
DAN
337.5
3.8 a
0.0017 e
20.77
1
2
55.0
DICLO
318.1
4.5 d
0.0560 f
23.21
1
0
57.9
DTX
807.9
2.4 a
0.0068 g
23.66
5
10
66.2
-
1
5
45.4
c
δdrug
Van Krevelen
[MPa 1/2]
Hd bond
donor
Ha bond
acceptor
Surface
tension
[dyne/cm]
20.80
0
2
-
ECZ
444.7
-
1.9316
ETO
588.9
1a
0.1593 h
24.52
3
10
76.4
FLZ
306.3
0.5 b
0.0033 i
24.16
1
8
55.4
GF
352.8
3.18 a
0.0600 c
21.08
0
6
52.5
HYDRO
362.5
1.72 a
0.0003 j
25.11
3
2
58.8
HY
504.4
8.78 a
0.0500 c
33.40
6
2
150.2
KETO
531.5
4.35 d
0.0320 k
21.08
0
8
52.1
PTX
853.9
3a
0.0004 l
23.38
4
11
68.4
PIRO
331.2
3.06 d
0.0695 d
24.11
2
5
79.5
QUER
338.3
1.48 d
0.1774 d
32.73
4
2
114.8
THPP
678.5
-
0.0030 c
25.94
6
2
75.5
VER
718.8
3.74 a
0.0110 c
21.29
3
9
70.1
NAPRO
230.3
2.8
0.0691 c
-
1
1
-
With Mw the molecular weight; Exp. log P, the experimental partition coefficient between octanol and water;
δdrug the solubility parameter calculated with the software Material Studio according to the Van Krevelen group
contribution method; a data taken from http://www.drugbank.ca/drugs; b OSPAR Commission, 2005.
http://www.ospar.org/documents/dbase/publications/p00199_BD%20on%20clotrimazole.pdf; c values obtained
experimentally by Carrupt et al. (University of Geneva, Pharmacochemistry, Geneva, Switzerland, unpublished
data); d data taken from ChemIDplus Advanced http://chem.sis.nlm.nih.gov/chemidplus/. e Erlich et al., Int. J.
Pharm., 1999, 179, 49-53; f Llinas et al., J. Med. Chem., 2007, 50, 979-983; g Ali et al., J. Med. Chem., 1997, 40,
236-241; h Shah et al., Int. J. Pharm., 1995, 113, 103-111; i Loftsson T, Hreinsdóttir D. Determination of
Aqueous Solubility by Heating and Equilibration: A Technical Note. AAPS PharmSciTech. 2006; 7(1): E1-E4. j
Ould-ouali et al., Pharm. Res., 2004, 21, 1581-1590; k BioAqueous Solubilization Services, Dowpharma.
http://www.dow.com/webapps/lit/litorder.asp?filepath=pharma/pdfs/noreg/715-00011.pdf&pdf=true; l Soga et
al., J.Control. Release, 2005, 103, 341-353; m Soga et al., J.Control. Release, 2005, 103, 341-353
44
All drugs were formulated and incorporated into MPEG-(hex)PLA micelles by a co-solvent
evaporation standard procedure. It has to be pointed out that neither the preparation procedure
nor the polymer molecular weight and polymer block-ratio have been optimized for the
specific drugs. It is expected that a systematic optimization could significantly improve the
incorporation results. In general the micelle size obtained for all formulations were similar,
with hydrodynamic diameters (Zav) varying from 70 to 200 nm and number weight diameters
(dn) from 17 to 150 nm (Table 3). This small carrier size could be of importance for a possible
use for passive drug targeting to inflamed or diseased tissues
12;14
. The majority of the drugs
was much more efficient and higher incorporated into MPEG-dihexPLA than into MPEGmonohexPLA micelles. In contrast, the drug incorporation results i.e. the drug content
(mgdrug/gcopolymer), the drug solubility in micelle solution (mM), as well as the “solubility
increase factor” (between drug solubility in micelles and water solubility) were significant
different for the tested formulations.
Various observations are remarkable when comparing drugs with similar log P values or
similar water solubility values. For example, GF (log P =3.18) incorporated at 16.6 mg/g
shows a more than 4 times higher drug incorporation in MPEG-dihexPLA micelles than PIRO
(log P=3.06). For drugs with a significantly different log P, FLZ (log P=0.5) and CLZ (log
P=6.1), a higher drug incorporation in MPEG-dihexPLA micelles was surprisingly obtained
for FLZ with an achieved drug content of 268 mg/g, corresponding to a 4 times higher drug
content than for CLZ (59 mg/g). As another example, PTX, with an experimental water
solubility of 0.0004 mM compared to HYDRO with a similar water solubility of 0.0003mM,
could be 5 times less incorporated in MPEG-dihexPLA micelles.
In consequence, it is not applicable to draw on only one physical-chemical parameter like log
P to predict the incorporation of a selected drug into polymeric micelles. Therefore, the
influence of the various physical-chemical parameters needs to be evaluated in more detail for
a possible prediction or estimation of the drug incorporation ability.
Chapter II
45
Table 3. Incorporation results, size characteristics and Flory-Huggins parameter of all the investigated
drug loaded MPEG-(hex)PLA micelle formulations.
(a) MPEG-PLA micelles
Micelle size
Drug
CLZ
χsp
Drug content
Van Krevelen
[mgdrug/gcopolymer]
1/2
[MPa ]
0.49
129.5
[Drug]micelles
[mM]
1.877
Solubility
increase Zav P.I. dn [%]dn
factor
21.3
56 0.42 25 99.9
CsA
1.14
66.2
0.275
27.5
89 0.39 28
99.9
DAN
15.02
3.5
0.051
30.0
118 0.43 29
99.9
DICLO
1.18
76.8
1.183
21.1
179 0.16 156 100.0
DTX
2.68
4.2
0.026
3.8
194 0.69 28 100.0
ECZ
-
177.1
1.991
1.0
70 0.29 29 100.0
ETO
8.46
229.0
1.937
12.2
154 0.53 26
95.7
FLZ
3.82
250.2
4.051
1240.7
92 0.44 20
99.2
GF
1.34
18.7
0.265
4.4
HYDRO
6.82
47.4
0.649
2163.3
HY
32.42
6.7
0.066
1.3
126 0.52 17 100.0
KETO
1.94
99.1
0.923
28.8
129
PTX
9.09
1.8
0.011
27.5
208 0.69 28 100.0
PIRO
3.87
1.4
0.021
0.3
184 0.33 76
QUER
19.07
9.4
0.139
0.8
100 0.52 23 100.0
THPP
13.68
4.6
0.033
11.0
107 0.50 35 100.0
VER
3.01
11.9
0.151
13.7
111 0.46 26 100.0
-
14.6
0.064
0.9
161 0.24 83
NAPRO
114 0.43 25 100.0
66 0.24 23
99.9
25 100.0
99.9
99.8
With [Drug]micelles, the concentration of drug in micelles; Zav the hydrodynamic diameter, dn the number weight
diameter and [%]dn the percentage of micelles with a given dn.
46
(b) MPEG-monohexPLA micelles
Micelle size
Drug
χsp
Solubility
Drug content
[Drug]micelles
Van Krevelen
increase
[mgdrug/gcopolymer]
[mM]
[MPa 1/2]
factor
Zav
P.I.
dn
[%]dn
CLZ
1.378
34.7
0.503
5.7
33
0.02 19 100.0
CsA
16.711
213.7
0.907
90.7
82
0.54 24 100.0
DAN
1.424
8.1
0.120
70.6
188
0.25 140 96.5
DICLO
2.996
149.9
2.269
40.5
147
0.34 19
99.8
DTX
10.638
11.6
0.072
10.6
178
0.41 45
97.9
ECZ
-
252.7
2.841
1.5
145
0.41 130 97.7
ETO
9.252
275.8
2.375
14.9
183
0.44 24 100.0
FLZ
4.201
271.7
4.452
1363.5
112
0.51 18 100.0
GF
1.598
13.7
0.195
3.3
145
0.51 20 100.0
HYDRO
7.408
83.2
1.143
3810.0
159
0.52 22 100.0
HY
33.732
84.0
0.809
16.2
211
0.25 217 100.0
KETO
2.313
46.6
0.443
13.8
183
0.36 22
PTX
10.121
5.1
0.030
75.0
205
0.48 103 74.7
PIRO
4.252
1.6
0.025
0.4
149
0.56 22 100.0
QUER
19.880
28.3
0.418
2.4
172
0.40 187 100.0
THPP
14.737
11.7
0.086
28.7
1179 0.06 406 66.5
VER
3.548
45.7
0.318
28.9
1243 0.23 112 97.6
-
25.9
0.112
1.6
149
NAPRO
97.1
0.41 25 100.0
Chapter II
47
(c) MPEG-dihexPLA micelles
Micelle size
Drug
χsp
Solubility
Drug content
[Drug]micelles
Van Krevelen
increase Zav
[mgdrug/gcopolymer]
[mM]
[MPa 1/2]
factor
P.I.
dn
[%]dn
CLZ
1.49
59.16
0.858
9.7
75
0.26
30
99.8
CsA
17.45
320.7
1.100
110.0
86
0.27
68
89.9
DAN
1.54
17.5
0.255
150.0
85
0.45
26
100.0
DICLO
3.13
119.5
1.839
32.8
99
0.59
25
100.0
DTX
11.09
32.7
0.201
29.6
135 0.34
35
99.5
ECZ
-
295.1
3.318
1.7
95
0.30
29
100.0
ETO
9.60
67.8
0.609
3.8
183 0.44
23
99.9
FLZ
4.37
268.3
4.381
1341.8
71
0.40
27
98.2
GF
1.71
16.6
0.196
4.4
163 0.37
31
99.9
HYDRO
7.66
78.6
1.067
3556.7
205 0.47
65
94.0
HY
34.29
112
1.122
22.4
101 0.16
17
99.9
KETO
2.48
55.4
0.494
15.4
109 0.26
34
99.7
PTX
10.57
14.1
0.082
205.0
50
0.52
23
98.9
PIRO
4.42
3.4
0.051
0.7
113 0.26
29
99.8
QUER
20.22
12.5
0.182
1.0
205 0.26 110 100.0
THPP
15.19
12.2
0.088
29.3
177 0.17 126 100.0
VER
3.79
85.0
0.591
53.8
193 0.45
29
100.0
NAPRO
-
26.9
0.117
1.7
104 0.27
29
97.2
48
3.1. Examination of Flory-Huggins parameter and incorporation data
It would be desirable to have a fast access to a prediction or estimation of the incorporation of
any drug into the polymeric micelle carrier system by simply calculating this from known
physical-chemical parameters of the drug. As demonstrated above a solely characterization of
the drug lipophilicity and affinity with the hydrophobic micelle core via the log P is not
possible. A generally applied method for the determination of the solubility of a compound in
a solute is given by the Flory-Huggins theory. Here, the Flory-Huggins parameter χsp of a drug
would reflect its solubility and affinity, respectively in the hydrophobic micelle core. It is
assumed that first the drug is only solubilized by the core forming block copolymer and
second that the core of the micelles can be represented as a bulk polymer. The first condition
is a hypothesis since one cannot exclude drug dispersion in the micelle core. The second
condition is met for the hexPLA block as e.g. CsA was found to be dissolved in the MPEGdihexPLA micelle core
28
, and e.g. haloperidol could be completely dissolved in bulk
hexPLA 29. Based on this the theory of Flory-Huggins ought to be applicable here. χsp was
calculated for all drug-copolymer systems using equation 4. The results are presented in Table
3. The closer χsp is to zero, the better is the miscibility between the drug and the copolymer.
The 18 investigated drugs showed significant different χsp-affinities with the 3 different coreforming polymers. With a slightly lower δpolymer (δdihexPLA= 17.19 MPa
δmonohexPLA=17.33 and δPLA=17.64 MPa
1/2
1/2
) compared to
, respectively, dihexPLA micelles should have a
slightly lower affinity/compatibility with those drugs which have a corresponding higher δ
values. However, this is not the case. The majority of the tested drugs e.g. CsA, THPP, VER,
HY, PIRO, NAPRO, PTX, DTX, VER, DAN are better incorporated in MPEG-dihexPLA
micelles rather than in MPEG-PLA micelles. Only KETO and CLZ were found to be better
incorporated in MPEG-PLA micelles. Hy and CsA with a high χsp of 34.29 and 17.45 MPa1/2
with MPEG-dihexPLA, respectively were expected to be less incorporated than drugs with
lower χsp. However, both drugs were among the 5 best incorporated drugs for MPEGdihexPLA micelles (Table 3c), whereas THPP with a χsp of 15.19 yielded one of the lowest
drug contents with only 12.2 mg/g. Hence, no clear relationship between χsp and the drug
content in MPEG-(hex)PLA micelles was found. Thus the Flory-Huggins theory does not
seem to be applicable to predict drug incorporation into MPEG-hexPLA based polymeric
micelles.
As mentioned above so far only non-optimized polymers towards the drugs were used for this
study. Letchford et al. reported for MPEG-PCL copolymers with varying MPEG and PCL
Chapter II
49
block lengths after incorporation of 5 drugs with different χsp
11;18
that (1) drugs with low χsp
were solubilised in higher extent than those with high χsp, indicating a better affinity and thus a
better incorporation, and (2) that copolymers with higher molecular weights of the PCL block
(higher block length) yielded higher drug incorporations. These findings envision the
possibility of improving drug incorporation into MPEG-hexPLA micelles by adjusting the
molecular weight of the core forming block.
3.2. Differentiated views on the drug water solubility and log P and
experimental incorporation results
As an alternative approach the here tested 18 different drugs were classified related to the
Biopharmaceutics Classification System (BCS) modified by the addition of high/low log P
and high/low water solubility as parameters (Figure 2). The set limit for high and low log P
was inspired here by the Lipinski’s “rule of 5”
2;14
. Concerning the solubility in water, the
limit was set for an intermediate value of 0.055 mM. The results of the drug incorporation
into MPEG-dihexPLA micelles (drug content and “solubility increase factor”) are
summarized in Figure 2.
Water
solubility
>0.055 mM
<0.055 mM
Low log P (log P<5)
Drug
QUER
PIRO
NAPRO
GF
ETO
DICLO
KETO
DTX
VER
DAN
PTX
FLZ
HYDRO
High log P (log P>5)
Drug Solubility
Drug Solubility
content increase Drug content increase
[mg/g] [× times]
[mg/g] [× times]
12.5
(×1)
3.4
(×1)
26.9
(×1.7)
CLZ
59.2
(×10)
18.7
(×3)
68.0
(×4)
119.5
(×33)
55.4
(×15)
32.7
(×30)
85.0
(×54)
(×22)
HY
112.0
17.5
(×150)
CsA 321.0
(×110)
14.1
(×205)
268.0
(×1342)
78.6
(×3559)
Figure 2. Classification of the tested drugs by high/low log P and high/low water solubility. (Log P of
ECZ and THPP are not available)
50
As the drug content differed from 4 to 295 mg/g for drugs with a considered high water
solubility and from 12 to 321 mg/g for those with a considered low water solubility, no
tendency for the drug incorporation could be found. However, a trend for the “solubility
increase factor” was observed. For drugs with a considered low water solubility, drug aqueous
solubility (solubilised by micelles) could be increased from 15 to 3559 times, when
incorporated into MPEG-dihexPLA micelles. For drugs with a considered high water
solubility, the increase solubility factor varied only from 1 to 10 (Figure 3). These results
confirm the findings of Letchford et al., who also reported for the drug with the low aqueous
solubility the highest solubility increase 18.
0.055 mM
Solubility increase factor
(F)
4000
3000
2000
1000
500
300
100
60
50
40
30
20
10
0
0.00
F>10
F<10
0.01
0.1
0.2 1.5
2.5
Figure 3. Influence of drug water solubility of the 18 tested drugs on the “solubility increase factor”.
For drugs with a similar log P (2.4 <log P< 3.8) in the group of “low solubility and low log
P”, the drug content increased with the water solubility of 4 selected drugs (VER, DTX,
DAN, and PTX) for the 3 investigated MPEG-(hex)PLA micelle carrier systems (Figure 4).
Chapter II
51
VER
90
Drug content [mg/g]
80
70
60
50
DTX
40
30
DAN
20 PTX
10
0
0
0.002
0.004
0.006
0.008
0.01
0.012
S water [mM]
Figure 4. Influence of the drug water solubility (Swater) in drug content obtained in MPEG-PLA (),
MPEH-monohexPLA () and MPEG-dihexPLA () micelles for drugs with log P between 2.4 and
3.8: PTX (paclitaxel), DAN (danazol), DTX (docetaxel), and VER (verteporfin).
The more hydrophobic MPEG-dihexPLA micelles incorporate considerably higher quantities
of those poorly water soluble drugs in comparison to MPEG-monohexPLA and MPEG-PLA
micelles. This superior incorporation of poorly water soluble drugs into MPEG-hexPLA
micelles compared to MPEG-PLA micelles was also observed for the majority of the tested
drugs (Figure 5).
350
CsA
ECZ
Drug content (mg/g)
300
QUER
FLZ
250
200
DICLO
150
CLZ
VER
100 HYDRO
50
KETO
HY
DTX
GF
DAN THPP
PTX
0
0.000
0.005
0.010
0.03
NAPRO
ETO
PIRO
0.05
0.07
0.09 0.1
1.1
Swater (mM)
2.1
Figure 5. Drug contents of all tested drugs in MPEG-PLA (○) MPEG-monohexPLA () and MPEGdihexPLA () micelles in function of the water solubility (Swater) of the investigated drugs.
52
Looking closer at two drugs with similar χsp and chemical structures, as it is the case for PTX
and DTX, significant different incorporation results were obtained. A drug content of 14.12
mgPTX/g and 32.7 mgDTX/g was found for MPEG-dihexPLA micelles. However, this
corresponds to a “solubility increase factor” of 205 for PTX and only 30 for DTX. The
structure of both drugs differs only in two chemical groups, at the same C positions. PTX has
a benzamid goup at the secondary amine, and a methoxy group, whereas DTX bears a tertbutoxy group at the secondary amine, and has a free hydroxyl group (Figure 1). PTX has a
slightly higher log P (log P=3) compared to DTX (log P=2.4). Interestingly here PTX with its
higher log P yielded a lower drug content, but the higher “solubility increase factor”. This
high increase of aqueous solubility of PTX in micelles might be due to its very low water
solubility value of 0.0004 mM, which is 17 times lower than that of DTX (0.0068 mM). As
another example one can compare the photosensitizer drugs THPP and VER which also have
similar chemical structures. In this case THPP with the lower water solubility of 0.003 mM
has a “solubility increase factor” of 29 whereas VER with its higher water solubility of 0.011
mM leads to a “solubility increase factor” of 54. These results show clearly the importance of
the drug water solubility as such on the “solubility increase factor”.
The, to some extent, contradictory results highlight that the incorporation efficiency under
same preparation conditions can not be predicted by a simplistic look at only one physicalchemical parameter of the drug, especially log P, which is often used in the literature as a
major value to describe the hydrophobic character of a drug and does not reflect consistently
the obtained experimental data.
3.3. Examination of other physical-chemical parameters
Since the incorporation data obtained for the 18 drugs with MPEG-hexPLA could not be
explained via the log P or the water solubility parameter only, further parameters like the
molecular weight (Mw) of the drug, the number of hydrogen bond donors (Hd) and hydrogen
bond acceptors (Ha) in the drug structure were taken into consideration, as well as the surface
tension of the drug, which was calculated with the software ACD/ChemSketch software
(Advanced Chemistry Development, Inc., Canada). Because simultaneous interpretation with
6 chemical-physical parameters (Mw, Hd, Ha, drug water solubility, log P, and surface tension)
becomes difficult to handle manually, the incorporation results (the drug content, the
solubility in micelles, and the “solubility increase factor”) were analyzed by a linear multiple
Chapter II
53
regression in order to find the most influencing parameters. The corresponding regression
coefficients of each chemical-physical parameter were obtained by using the Data Analysis
tool of Microsoft Office Excel 2003. The higher is the absolute value of the coefficient, the
higher is the influence of the parameter on the specific response. The classification of the
influence of the different tested parameters regarding the obtained regression coefficients is
given in brackets in Table 4. Please note that these analyses were done only with those drugs
for which all the defined parameters were available.
Table 4. Classification of the influence of the different chemical physical parameters of the tested
drugs on the incorporation responses in MPEG-dihexPLA micelles after a linear multiple regression.
In brackets the ranking of importance is given, whereby 1 is high and 6 is low importance.
Responses
“Solubility
increase factor”
-6190.6 (1)
Solubility in
micelles [mM]
-6.06 (1)
Hd
0.56 (2)
668.2 (2)
65.3 (1)
Ha
0.34 (3)
39.8 (4)
30.2 (3)
Log P
0.19 (4)
-49.0 (3)
34.4 (2)
Surface tension [dyne/cm]
0.02 (5)
-31.4 (5)
3.4 (5)
Mw [g/mol]
0.01 (6)
-5.4 (6)
0.8 (6)
Chemical – physical parameters
Drug content
[mg/g]
29.5 (4)
The results reveal that the drug content is mainly influenced by Hd, log P, Ha and the drug
water solubility, Mw and the surface tension have less influence. Similarly these parameters
affect the drug solubility in micelles. Here the drug’s water solubility appeared to be the most
influencing parameter, followed by Hd and Ha, and then log P. Like for drug content, Mw and
the surface tension of the drug had only a slight influence. The parameter ranking of “the
solubility increase factor” response is similar to the one of the solubility in micelles,
confirming the observation that for a drug with lower water solubility (< 0.055 mM) a much
higher increase of its aqueous solubility can be achieved when incorporated in MPEGhexPLA micelles. The influence of Ha and Hd on the incorporation results in MPEG-hexPLA
micelles may explain the difference found and discussed above for the similar drugs PTX and
DTX, wherein the latter DTX has 1 more Hd group than PTX. The obtained results show in
54
general that Hd is the more influencing parameter for the drug content, and indeed for DTX a
higher drug content in MPEG-hexPLA micelles was obtained. Thus the presence of chemical
groups with Ha and Hd functions also influences the drug incorporation in MPEG-hexPLA
micelles. Indeed, the drug/copolymer affinity was found to be induced by the formation of Hbonds between the multiple H-bond site of the PCL copolymer and single H-bond site of two
cucurbitacin drugs
30
. Drugs with multiple H bond donors and acceptors have favourable χsp
values and are better solubilised
31
. In consequence, the prediction or estimation of drug
incorporation in MPEG-hexPLA micelles needs also to take structural parameters into
account. It becomes clear that a quantitatively relationship between a drug via its physicalchemical parameters and the experimentally obtained incorporation results is difficult to draw.
From a practical point of view towards suitable pharmaceutical formulations, we applied the
criteria of Lipinski’s “rule of 5” for of the here tested drugs (Figure 6).
“Rule of 5”
PIRO
NAPRO
Mw<500 g/mol
QUER
KETO
Ha<10
Hd <5
Log P<5
GF
PTX
DICLO
VER
ETO
ECZ
DTX
CsA
DAN
THPP
HY
HYDRO
FLU
CLZ
Inverse “rule of 5”
Mw>500 g/mol
Ha>10
Hd >5
Log P>5
Figure 6. The Lipinski’s “rule of 5” applied on the investigated drugs
As shown half of the drugs do not fit the rules (Mw<500 g/mol; Ha<10; Hd<5 and log P<5),
and have in majority a considered low solubility, which would limit their success in clinics.
Amongst those drugs, CsA is a drug that in theory would have only little chances of
bioavailability by all the Lipinski’s criteria. However, as demonstrated here CsA would be a
preferable drug for incorporation in MPEG-dihexPLA micelles with its high number of H
bond donor and acceptor groups, its high log P (8.2) and its very low water solubility (0.01
mM). Indeed, the successful CsA incorporation into MPEG-hexPLA micelles was presented
previously 21.
Chapter II
55
4. Conclusions
Eighteen different poorly water soluble drugs, chosen to cover a wide range of lipophilicity
and of water solubility, were incorporated into MPEG-hexPLA micelles. The calculation of
the Flory-Huggins parameter χsp does not confirm the general better affinity of the
hydrophobic drugs with the most hydrophobic micelle-core forming block (dihexPLA). The
value of log P did not consistently correlate the experimental incorporation data either.
Therefore the influence of other drug physical-chemical parameters on the incorporation
results in micelles was studied. The molecular weight and the surface tension of the drug
show only minor contribution on the drug incorporation data, whereas the log P, the drug
water solubility and the number of hydrogen bond donor Hd– groups and hydrogen acceptor
Ha- groups were found to be major influencing parameters. Drugs with a low water solubility
(< 0.055 mM) showed a high increase of their aqueous solubility when incorporated into
MPEG-dihexPLA micelles. Drugs with a high number of Hd yielded higher drug contents in
MPEG-hexPLA micelles. Therefore the prediction or estimation of drug incorporation into
MPEG-hexPLA micelles depends also on the structural characteristics of the drugs. In order
to take the maximum of factors into consideration, physical-chemical and structural,
molecular dynamics (MD) simulation of the presented results of MPEG-hexPLA
incorporation are currently under investigation. It has to be pointed out that drugs which do
not fit the Lipinski’s “rule of 5” indicating a low oral bioavailability, can efficiently be
incorporated and formulated in MPEG-hexPLA micelles. Thus MPEG-hexPLA micelles can
be envisioned as alternative pharmaceutical formulations for poorly water soluble drugs.
Acknowledgments
The authors thank the Swiss National Science Foundation (SNF) for financial support (SNF
200020-103752).
The authors thank also Aliya Kasimova (University of Geneva, Geneva, Switzerland) for the
determination of the solubility parameters with the Material Studio software.
56
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C.,
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(7) Uchenna Agu, R.; Jorissen, M.; Willems, T.; Van den Mooter, G.; Kinget, R.; Verbeke,
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(9) Tibell, A.; Norrlind, B., Cyclosporin A in fat emulsion carriers: studies on the
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Transplant Int. 1994, 7, 438-441.
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carriers, Eur.J.Pharm.Biopharm. 1999, 48, 101-111.
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Müller, R. H., Stealth corona-core nanoparticles surface modified by polyethylene
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Colloid.Surface B 2000, 18, 301-313.
(12) Park, J. H.; Lee, S.; Kim, J. H.; Park, K.; Kim, K.; Kwon, I. C., Polymeric nanomedicine
for cancer therapy, Prog.Polym.Sci. 2008, 33, 113-137.
(13) Kohane, D. S., Microparticles and nanoparticles for drug delivery, Biotechnol.Bioeng.
2007, 96, 203-209.
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(14) Singh, R.; Lillard, J., Nanoparticle-based targeted drug delivery, Exp.Molec.Pathol.
2009, 86, 215-223.
2006, 12, 4669-4684.
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Block copolymer micelles: preparation, characterization and application in drug
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formulations using MPEG-hexyl substituted polylactide micelles: a suitability study,
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58
(29) Asmus, L. R.; Gurny, R.; Möller, M., Solutions for Lipophilic Drugs - A biodegradable
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(30) Patel, S. K.; Lavasanifar, A.; Choi, P., Roles of Nonpolar and Polar Intermolecular
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1780-1786.
Supplementary material
HPLC conditions developed for the determination of the drug content of each tested drug
Drug
Type of
Column
T [°]
Mobile Phase
Flow rate
(mL/min)
UV
detection
Calibration
Standards
(µg/mL)
CLZ
C18 a
35°C
MeOH:0.05 M K2HPO4
(75:25)
1.5
214 nm
5.5 – 220
CsA
C18 b
65°C
ACN:Water (75:25)
1.2
210 nm
0.15 - 200
DAN
C18 b
25°C
ACN:Water (70:30)
1.0
283 nm
2.1 – 213
DICLO
C18 b
30°C
ACN: 20mM K2HPO4
(40:60)
1.0
275 nm
2.0 – 200
(pH = 3.1 adjusted with PA)
DTX
C18 b
25°C
ACN:Water (50:50)
1.0
227 nm
1.6 – 307
ECZ
C18 a
35°C
ACN:0.05 M K2HPO4
(60:40)
1.5
214 nm
2.0 – 200
ETO
C18 b
25°C
ACN:Water:AA (34:65:1)
1.0
230 nm
10.0 – 200
FLZ
C18 a
350C
MeOH:0.1 M NH4H2PO4
(50:50)
1.5
214 nm
5.0 – 200
Chapter II
GF
C18
a
25°C
ACN:20 mM KH2PO4
(45:55)
59
1.0
293 nm
2.0 – 200
(pH = 3.0 adjusted with PPA)
HY
C-8 c
25°C
MeOH:Water (90:10) +
1% FA
1.0
590 nm
2.0 – 200
HYDRO
C18 a
35°C
ACN:0.03% PA
1.0
243 nm
2.0 – 200
KETO
C18 a
25°C
MeOH:Water:DEA
(74:26:0.1)
1.0
240 nm
2.1 – 207
PIRO
C18
a
30°C
ACN: 20mM K2HPO4
(40:60)
1.0
355 nm
1.9 – 190
PTX
C18 a
25°C
ACN:Water(50:50)
1.0
227 nm
1.0– 200
QUER
C18 a
25°C
ACN:1% AA (45:55)
1.0
230 nm
2.0– 200
NAPRO
C18 a
30°C
ACN: 20mM K2HPO4
(40:60)
1.0
331 nm
2.0– 200
AA, acetic acid; DEA, diethylamine; FA, formic acid; K2HPO4, potassium biphosphate buffer; NH4H2PO4,
ammonium phosphate monobasic; KH2PO4, potassium dihydrogen phosphate; PA, phosphoric acid; PPA,
pyrophosphoric acid.
a
Lichrospher® RP-18 column (124 mm × 4 mm)
b
Nucleosil 100-5 C18 column (250 mm × 4 mm)
c
Zorbax- Eclips XDB-C8 column (150 mm× 4.6 mm)
60
1234562678693A738B63456CADE23FB4865234DD5386532DE6395E23A8B33BAB37E653D33
D793 FF648D53 793 9D7D53 5DE423 F73 63 C653 6AFA6783 D5E8693 A8B3 B6 13
AF6623 793 A7F562A73 2E2878A3 8B6A53 8653 2DEAA8!3 "73 8B63 DDA73 3 D963
1
D5E8AD73D3FFD24D5A7313238628693D53A823676534B5F6E8AF32EA8AA8345D4658A62!3
1
1
1
1
123456789998
8
1
Chapter III
61
Novel Cyclosporin A Formulations Using MPEG-Hexyl
Substituted Polylactide Micelles: A Suitability Study
K. Mondon, M. Zeisser-Labouèbe, R. Gurny, and M. Möller
Ernest Ansermet, CH-1211 Geneva 4, Switzerland
Submitted to: European Journal of Pharmaceutics and Biopharmaceutics
The immunosuppressive agent Cyclosporin A (CsA) has very poor solubility in water and, in
consequence, non-aqueous formulations have been developed for its intravenous
administration to treat patients with transplant rejection. In this article, aqueous micelle
solutions of novel amphiphilic copolymers based on methoxy-poly(ethylene glycol) (MPEG)
and hexyl-substituted poly(lactides) (hexPLA) were studied for possible incorporation and
formulation of CsA, and for their biocompatibility towards novel pharmaceutical applications.
Above the critical micellar concentration (CMC), MPEG-hexPLA block-copolymers selfassemble into unimodal micelles with diameters of around 30 nm, either unloaded or drugloaded. The best shelf-life stability of these formulations was observed when stored at 4°C
with a drug loss inferior to 7% after 1 year. The polymer and micelle toxicities were evaluated
in vitro for three different cell lines and in vivo using the chick embryo chorioallantoic
membrane (CAM) model. The hemolytic property was assessed using human blood samples.
As the studies revealed, MPEG-hexPLAs are non-toxic and do not show hemolysis; the same
was found for the comparable MPEG-PLAs, both as unimers below their CMC and as
polymeric micelles up to copolymer concentrations of 20mg/mL. At this concentration, CsA
was efficiently incorporated into MPEG-hexPLA micelles up to 6 mg/mL, which corresponds
to a 500-fold increase of its water solubility. The current recommended clinical concentration
administered per infusion (0.5-2.5 mg/mL) can be easily achieved and requires four times less
copolymer than with the often-used Cremophor ®EL surfactant. In this regard, MPEGhexPLA micelle formulations can be an applicable formulation in transplant rejection
treatments as an injectable CsA carrier system.
CsA/ MPEG-hexPLA Micelle Formulations : A Suitability Study
62
Keywords: biocompatibility; CAM model; Cyclosporin A; drug delivery; polymeric
micelles; substituted polylactides.
Chapter III
63
1. Introduction
To deliver drugs to the right pathological site, at the right time, in the right dose, and without
altering surrounding tissues is one of the important aims of drug delivery systems. For this
purpose, nanocarriers, such as nanoparticles
1;2
, liposomes
3
and micelles
4-6
, have been
studied in pharmaceutical research. Due to their small size of less than 1 μm, nanosized
vehicles can improve the therapeutic and pharmacological properties of drugs by escaping the
typical response of the immune system and therefore reaching and accumulating into
pathological sites, such as tumors, due to the enhanced permeability and retention effect
(EPR) 7. For other pathological sites, the nanosize of micelles allows for better internalization
in cells and increases the therapeutic effect of the drug compared to larger carriers 8-10.
Polymeric micelles of amphiphilic copolymers, which are composed of hydrophobic and
hydrophilic blocks, are characterized by self-assembled core-shell structures in aqueous
media. For pharmaceutical use, polymeric micelles with such a structure are of interest due to
their ability to incorporate and protect hydrophobic drugs within their hydrophobic core,
while the hydrophilic corona stabilizes and solubilizes the micelles in the aqueous
environment. Usually the hydrophilic block is formed by the water-soluble, biocompatible
and non-immunogenic poly(ethylene glycol) (PEG), which is known for its ability to reduce
opsonization and recognition by the mononuclear phagocytic system (MPS). This
consequently provides the long circulation properties of, for example, pegylated
nanoparticles 11. The hydrophobic block can be composed of various polymers, which must be
non-toxic and biocompatible for pharmaceutical applications, i.e., polyethers
acids
13
, or polyesters
14
12
, polyamino
. Amongst the hydrolysable polyesters, polycaprolactone (PCL) and
polylactides (PLA) are the polymers of highest interest due to their known biodegradability,
biocompatibility and their FDA status of “Generally Regarded as Safe” (GRAS). Due to its
faster degradation time, PLA is often preferred to PCL in medical applications. However,
PLA has limited applicability and performance as a carrier of hydrophobic drugs; the ability
of micelles to solubilize drugs within their core depends on the compatibility of the core
forming block and the drug itself
15-17
. Therefore, novel hydrophobic and biocompatible
polymers based on the biodegradable PLA backbone were developed by T. Trimaille et al.,
who synthesized different alkyl substituted polylactides by ring opening polymerization
(ROP) 18. Due to their interesting physico-chemical properties, viscous and injectable hexyl-
substituted polylactides (hexPLA) were investigated for pharmaceutical applications
64
19
.
HexPLA polymers of the two different structures, that is to say the mono- and di- hexylsubstituted polylactides (replacement of every second methyl group and all methyl groups by
hexyl groups, respectively along the polymer backbone), are named here as monohexPLA and
dihexPLA, respectively (Scheme 1). The presence of hexyl groups results in the highest
hydrophobicity for dihexPLA, followed by monohexPLA, and the smallest for PLA. When
hexPLAs are copolymerized with methoxy-poly(ethylene glycol) (MPEG), amphiphilic
copolymers that self-assemble in aqueous media into polymeric micelles are formed and they
incorporate hydrophobic drugs like griseofulvin
hydroxyphenyl)porphine (THPP)
21
20
or the photosensitizer meso-tetra(p-
much more efficiently than comparable “standard”
MPEG-PLA micelles. In this work, the very hydrophobic drug Cyclosporin A (CsA), which
has a log P of 8.2, was formulated for the first time with MPEG-hexPLA micelles. We also
investigated their biocompatibilities to determine the feasibility of these novel drug
nanocarriers as conceivable pharmaceutical formulations. The neutral, cyclic, undecapeptide
CsA is a potent immunosuppressive agent used to reduce graft rejection after organ
transplantation
22;23
. With its rigid structure and its poor water solubility (0.012mg/mL,
determined experimentally at 25°C), CsA must be solubilized in non-aqueous media with
surfactants like CremophorEL, as is done for the Sandimmune formulation (Novartis
Pharma). Here, MPEG-hexPLA micelles can facilitate an aqueous formulation for the
solubilization and delivery of CsA for intravenous applications. Due to this newly-developed
drug carrier system, the biocompatibility issues of the MPEG-hexPLA micelles needed to be
addressed. Therefore, we first characterized both unloaded and CsA-loaded micelle
formulations and studied their drug incorporations and shelf-life stabilities. Secondly, the
biocompatibilities of the unloaded micelles and their unimers (possible compounds from
disassembled micelles below their CMC that can occur under the dilution conditions in
intravenous applications) were studied. These studies included (a) the in vitro toxicity on
three different cell-lines, which differed in their mammalian origins (murine or human) and
their biological origins (intestinal or cancerous), (b) the in vivo toxicity using the chick
embryo chorioallantoic membrane (CAM) model, and (c) the hemolytic property using human
blood. The obtained favorable and encouraging results are presented here.
Chapter III
65
2.1 Materials
Triton® X-100 was purchased from AppliChem (Gatersleben, Germany). Tetrahydrofuran
(THF) was supplied by SDS (Toulouse, France), and distilled over sodium to make it
anhydrous when needed. Methoxy-poly(ethylene glycol) with a molecular weight of 2000
g/mol (MPEG2000g/mol) was a gift from BASF (Ludwigshafen, Germany). D,L-lactide, tin(II)
2-ethylhexanoate (Sn(Oct)2), and acetone p.a. were purchased from Purac Biochem
(Gorinchem, The Netherlands), Aldrich (Buchs, Switzerland), and Fluka (Buchs,
Switzerland), respectively and used as received. The monomers, mono-hexyl-substituted
lactide (monohexLA) and di-hexyl-substituted lactide (dihexLA) were synthesized as
described previously 18. CremophorEL was supplied by BASF and Sandimmune (Novartis
Pharma, Bern, CH) was purchased from a local pharmacy. Cyclosporine A was supplied by
Fluka (Buchs, Switzerland).
2.2 Synthesis and characterization of MPEG-(hex)PLA copolymers
MPEG-hexPLA polymers.
MPEG-hexPLA and MPEG-PLA copolymers of 5000 g/mol were synthesized by ring
opening polymerization (ROP). Briefly, 3.0 g of the corresponding lactide monomer were
polymerized in bulk at 100°C with the required amount of the initiator MPEG2000g/mol to reach
the targeted molecular weight. The catalyst Sn(Oct)2 was added in an equimolar ratio to the
initiator. The copolymerization reactions were stopped after 1.5 h by adding 5 mL of THF
(1% water) into the reaction mixture. After the removal of THF by evaporation the resulting
copolymers were precipitated dropwise into 120 mL of cold methanol, then filtrated and dried
under vacuum. Finally they were analyzed by 1H NMR (Brüker, 300 MHz) to detect any
residual organic solvent traces.
The copolymers were characterized to determine their molecular weight (Mn) and
polydispersity index (P.I.) using gel permeation chromatography (GPC). The GPC setup was
composed of a Waters system with Waters Styragel HR1-3 columns and a Waters 410
differential refractometer (Waters, Milford, USA). The analysis was carried out using
66
polystyrene (PS) of different molecular weights as calibration standards (PSS, Mainz,
Germany).
2.3 Preparation and size characterization of unloaded MPEG-(hex)PLA
micelles
A MPEG-(hex)PLA copolymer solution was dissolved at 100 mg/mL in acetone (2 mL) and
was added dropwise under sonication into 4 mL of isotonic saline solution. The organic
solvent was slowly removed by evaporation at 15 mbar. Final micelle concentrations were
adjusted to 50 mgcopolymer/mL by adding isotonic saline solution. Afterwards, MPEG(hex)PLA micelles were analyzed for their size. The hydrodynamic- (Zav) and numberweighted (dn) diameters were measured using dynamic light scattering (DLS) with a Zetasizer
HS 3000 (Malvern, Worcestershire, UK). The analyzes were performed at an angle of 90° at
25°C. For each sample, the mean diameters were obtained after three runs of ten
measurements.
2.4 Incorporation of CsA into MPEG-hexPLA micelles
The preparation and incorporation of CsA into MPEG-(hex)PLA micelles followed the
similar co-solvent evaporation method as described for unloaded micelles. Here, CsA and 20
mg MPEG-(hex)PLA were dissolved in acetone (2 mL), and the final micelle concentrations
were adjusted to 5 mgcopolymer/mL by adding ultra-pure water or isotonic saline solution. When
needed, the isotonicity of micelles solution prepared in water was obtained by addition of the
corresponding amounts of sodium chloride. The solutions were left to equilibrate overnight
and then centrifuged at 9500 x g for 15 min to remove non-incorporated, non-soluble CsA.
For drug-loading determination, the supernatant was diluted in acetonitrile at a ratio 1:10 to
break up the micelles and to release the CsA for quantitative analysis.
For more highly-concentrated CsA loadings in MPEG-dihexPLA micelles, the copolymer
concentration was increased up to 20 mg/mL with a targeted drug loading set at 400
mgCsA/gcopolymer. To ensure complete micelle destruction before CsA quantification, a 1:100
dilution in acetonitrile was made.
The CsA concentration was quantified by HPLC using a C-18 column (250 mm × 4.6 mm)
heated at 65°C. The flow rate was 1.2 mL/min and the mobile phase was a mixture of
acetonitrile/water (75:25) with a pH of 3.1 after the addition of phosphoric acid. Samples of
Chapter III
67
20 µL were injected with a running time fixed at 6 min. The CsA peak was detected by UV at
λ = 210 nm and appeared at 4.1 min. The CsA calibration standards from 0.15 to 200 µg/mL
were prepared and the resulting calibration curves were obtained with a regression coefficient
superior to 0.99. All samples were measured in triplicates.
2.5 Preparation of CremophorEL micelles
CremophorEL micelles were prepared by adding 400 mg of the surfactant into 4 mL isotonic
saline solution with stirring to reach a final concentration of 100 mg/mL. They have been
characterized by size as previously mentioned for unloaded MPEG-(hex)PLA micelles.
2.6 Cell culture
A Caco-2 cell line, a human epithelial colorectal adenocarcinoma cell line was maintained in
Dulbecco's Modified Eagle's medium (DMEM) (Gibco Life Technologies, Carlsbad, USA)
supplemented with 10% fetal bovine serum (FBS) (Brunschwig, Amsterdam, The
Netherlands), 1% of non-essential amino acids and 1% of penicillin and streptomycin 24. The
cells were grown for 2 weeks to allow for monolayer formation and cell differentiation 25.
A NuTu-19 cell line, a poorly differentiated Fischer 344 rat-derivative epithelial ovarian
cancer cell line 26, was kindly provided by Dr. A Major (Geneva University Hospital, Geneva,
Switzerland). The cells were cultured in DMEM culture medium supplemented with 10%
FBS and 100 U/mL penicillin–streptomycin (Gibco Life Technologies).
A SKOV-3 (HTB-77) cell line, a human ovarian carcinoma cell line, was purchased from
American Tissue Culture Collection (ATCC, Manassas, USA). The cells were grown in
Roswell Park Memorial Institute medium (RPMI 1640) (Gibco Life Technologies) and
supplemented with 10% FBS, 1% l-glutamine, and 1% penicillin/streptomycin.
All cell lines were cultured at 37°C in a humidified atmosphere containing 5% CO2.
2.7 In vitro toxicity
The in vitro toxicity of unloaded micelles was determined by standard MTT tests on Caco-2,
NuTu-19, and SKOV-3 cells. Briefly, the cells were seeded in a 96-well plate at a density of
1.5×104 cells per well in 100 μL of culture medium and incubated in a humidified atmosphere
with 5% CO2 at 37°C for 24 h. After removal of the culture media and washing with PBS, the
68
cells were incubated with 100 μL MPEG-(hex)PLA and CremophorEL micelle solutions or
isotonic saline solution. After 24 h incubation, 50 μL of a 3-(4,5-dimethylthiazol-2-yl)-2,5diphenyltetrazolium bromide (MTT) solution of 1 mg/mL in PBS were added to each well
and the plate was incubated for 3 h to allow the soluble yellow MTT to be reduced into the
dark-blue, insoluble formazan crystals by the metabolically active cells. Afterwards, the
formazan crystals were dissolved by addition of 200 μL dimethyl sulfoxide (DMSO) in the
incubator at 37°C for 1 h. The UV absorbance of individual wells was measured at 595 nm
with a microplate reader (Model 550, Bio-RAD, Hercules, USA).
Formulation concentrations below the CMC (0.01mgcopolymer/mL) were studied to determine
the toxicity of MPEG-(hex)PLA as unimers, whereas concentrations above the CMC yielded
the toxicity of MPEG-(hex)PLA as micellar structures. Different copolymer concentrations
were obtained by dilution in culture media from the most concentrated MPEG-(hex)PLA
formulation (50 mgcopolymer/mL). The same procedure was carried out for CremophorEL
solutions, which have a CMC of 0.9 mg/mL. The culture media and a 0.5% TritonX-100
solution in 1 N NaOH were used as positive (100% survival) and negative (0% survival)
controls, respectively.
The cell viability was determined by the following formula:
Cell viability (%) =
A sample − A 0%
A100% − A 0%
× 100
where Asample, A0% and A100% are the respective absorbances of the sample, the negative and
positive control.
2.8 In vivo toxicity by the CAM model
The in vivo toxicity of the MPEG-(hex)PLA micelles was assessed by using the chick embryo
chorioallantoic membrane (CAM) model adapted from Lange et al. 27. The egg incubation
procedure followed the one described by Vargas et al.
28
. Briefly, fertilized hen eggs, kindly
provided by the Animalerie Universitaire of the University of Geneva (Geneva, Switzerland)
were placed with the narrow apex down into an incubator Savimat MG 200 at 37°C with a
relative humidity of 65%. The eggs were rotated twice a day until the embryo development
day 3 (EDD3). On EDD4, a 3-mm hole was drilled into the eggshell and the narrow apex was
covered by an adhesive tape. The eggs were incubated in a static mode until EDD12. After the
removal of the tape, a bigger hole of 2-3 cm was drilled, thus allowing the visualization of the
CAM vasculature. Portions of 20 µL of unloaded MPEG-(hex)PLA formulations were
Chapter III
69
injected via the main vessel. The eggs were returned to the incubator and the survival rate was
evaluated at 24 h post-injection.
The unloaded MPEG-(hex)PLA micelles at 0.0001, 5, 10, and 50 mgcopolymer/mL from five
chick egg embryos were investigated. The highest concentrated micelle solution of 50
mgcopolymer/mL was used as prepared and the other three were obtained by dilution with
isotonic saline solution. The isotonic saline solution itself was used as the 100% survival
control.
2.9 Hemolysis test
Blood samples were freshly collected from healthy human volunteers in acid citrate dextrosecoated (ACD) tubes at the Blood Transfusion Center (Geneva University Hospital, Geneva,
Switzerland). The formulation samples were incubated with blood at a 3:1 sample/blood ratio
at 37°C for 24 h. After centrifugation at 770 x g for 10 min, the supernatant was collected in a
96-well plate and the release of hemoglobin was measured by UV absorbance at 575 nm with
four points of measurement per well using a microplate reader (Safire, Tecan, Männedorf,
Switzerland).
Various concentrations (below and above the CMC) of unloaded MPEG-(hex)PLA, CsAloaded MPEG-dihexPLA micelles and Sandimmune were tested. Samples of different
concentrations were prepared by dilution of the highest copolymer or surfactant concentration
with isotonic saline solution. Both the concentrated CsA-loaded MPEG-dihexPLA
formulations and the concentrated Sandimmune were diluted in order to obtain final CsA
concentrations of 0.5 mg/mL and of 2.5 mg/mL. An isotonic saline solution and 1%
TritonX-100 solution were tested as the 0% lysis control and 100% lysis control,
respectively.
The percentage of hemolysis was calculated as follows:
Hemolysis (%) =
A sample − A 0%
A100% − A 0%
× 100
where Asample, A0%, A100% represent the absorbances of the sample, the negative and positive
control, respectively.
70
2.10 Shelf-life stability
The size stability of unloaded and CsA-loaded micelles in ultra-pure water was investigated at
25°C over one year. The shelf-life of CsA-loaded micelles in isotonic saline solution in terms
of drug content and size was studied at 3 different temperatures (4°C, 25°C and 37°C) over 3
months and one year.
The micelle size was determined by DLS at multiple detection angles with a goniometer
ALV/CGS-5 (ALV-GmbH, Langen, Germany) and a power of 0.2 W. Briefly, 100 µL of
unloaded or CsA-loaded MPEG-(hex)PLA micelle formulations were diluted in 2 mL of
isotonic saline solution or ultra-pure water in a clean, clear and capped tube resulting in a
copolymer concentration of 0.24 mg/mL. After initial size measurements, the samples were
stored at their respective storage conditions. At the desired time points, they were allowed to
equilibrate to room temperature for 2 h before being analyzed directly in the sample tubes.
A similar procedure was applied for the formulation stability tests. After the desired storage
time, the CsA-loaded MPEG-(hex)PLA formulation samples were allowed to equilibrate to
room temperature and then centrifuged at 9500 g for 15 min to remove any non-entrapped
CsA. The CsA drug content in micelles was quantified by the HPLC method described above.
3. Results
3.1 Synthesis of MPEG-(hex)PLA copolymers
The MPEG-monohexPLA and MPEG-dihexPLA copolymers were prepared by ring opening
polymerization (ROP) in bulk using Sn(Oct)2 as the catalyst and MPEG2000g/mol as the
initiator. These molecular structures are presented in Scheme 1. The MPEG-PLA was
synthesized by the same method and was used as the “reference”. Copolymers with molecular
weights of 5000 g/mol (± 1 monomer unit) were obtained with low polydispersity indices
(P.I.) smaller than to 1.1 (Table 1). The MPEG-hexPLAs of the same molecular weight have
half and double number of hydrophobic hexyl side groups, respectively, thus permitting us to
investigate the influence of hydrophobicity on the micellization, stability, drug loading and
toxicity of their resulting micelles.
Chapter III
71
O
O
H3C
45
R
O
O
R
R1, R2 = CH3
1
2
1
O
H
m
O
: MPEG2000g/mol-PLA3000g/mol
for m=21
2
R =C6H13, R =CH3 : MPEG2000g/mol-monohexPLA3000g/mol for m=14
R1, R2 =C6H13
: MPEG2000g/mol-dihexPLA3000g/mol
for m=11
Scheme 1. Structure of MPEG-hexPLA and MPEG-PLA block copolymers.
3.2 Unloaded MPEG-hexPLA micelles
Unloaded MPEG-hexPLA micelles were prepared in isotonic saline solution by the co-solvent
evaporation method and had the number weighted (dn) sizes between 18-27 nm with a
population of 99-100% (Table 1). The values obtained for the hydrodynamic diameter (Zav)
showed sizes of 70-90 nm with a polydispersity between 0.22-0.35. This higher polydispersity
is related to a minor population of some larger micelles, which could be filtered off.
Table 1. Characteristics of synthesised MPEG-(hex)PLA copolymers and their corresponding
unloaded micelles prepared in isotonic saline solution.
a
Micelle size
[%]dn
Zav [nm]
P.I.
18
100.0
75
0.35
1.03
26
99.1
70
0.22
1.05
27
99.6
87
0.31
Mwa [g/mol]
P.I.a
dn [nm]
MPEG-PLA
5100
1.07
MPEG-monohexPLA
5000
MPEG-dihexPLA
5300
Copolymer
determined by GPC using PS standards
72
3.3 CsA incorporation in MPEG-hexPLA micelles
Formulations with a copolymer concentration of 5 mg/mL and different targeted drug
loadings from 10 to 1000 mgCsA/gcopolymer were prepared. The amount of incorporated CsA
was assessed by HPLC after micelle disruption by the addition of acetonitrile. The results are
summarized in Fig. 1, where the incorporation efficiency of 100% and 80% are drawn in plain
and dashed lines, respectively.
100% efficiency
Actual drug loading
[mgCsA/gcopolymer]
500
80% efficiency
400
300
200
100
0
0
100
200
300
400
Targeted drug loading
[mgCsA /gcopolymer]
500
600
Figure 1. Incorporation of CsA in MPEG-PLA (○), MPEG-monohexPLA (□), and MPEG-dihexPLA
() micelles in function of the targeted drug loading.
As seen in the figure, the CsA loading increased with the targeted loadings until a plateau at
80 mgCsA/gcopolymer for MPEG-PLA, 230 mgCsA/gcopolymer for MPEG-monohexPLA and 300
mgCsA/gcopolymer for MPEG-dihexPLA copolymer was reached. At the latter angular points,
loadings with efficiencies of 80% and higher could be achieved. For MPEG-monohexPLA
micelles, the highest and most efficient drug loading of 230 mgCsA/gcopolymer was obtained
from the targeted loading of 300 mg/g, with an incorporation efficiency of 77%. For the same
target loading, MPEG-dihexPLA incorporated 265 mgCsA/gcopolymer with an efficiency of 88%.
For the higher targeted loading of 500 mg/g, the highest drug loading of 320 mg/g for a
MPEG-hexPLA formulation could be achieved, still with an efficiency of 65%. Any nonincorporated drug could be removed by centrifugation of the formulation.
Chapter III
73
Micelle sizes, Zav and dn of CsA-loaded MPEG-(hex)PLA micelles were determined by DLS
at a detection angle of 90°. The results for different drug loadings are presented in Fig. 2,
where the sizes are given on the primary Y-axis and the [%]dn on the secondary axis. The
MPEG-(hex)PLA micelles had a number weighted diameter between 16 and 30 nm for all
tested formulations. The sizes were identical to the unloaded micelle solutions (Table 1). It is
notable that the average hydrodynamic diameter Zav increased with the incorporation amount
of CsA in MPEG-hexPLA micelles, which was due to the presence of a small number of
larger micelles in contrast to unloaded micelle formulations. Still, 98-99% of all the micelles
100
90
80
70
60
50
40
30
20
10
0
[%]dn
90
80
70
60
50
40
30
20
10
0
9
19
35
39
67
77
86
10
27
37
68
100
229
234
9
26
33
68
100
265
327
Size [nm]
had sizes dn smaller than 30 nm (Fig. 2).
MPEG-PLA
micelles
MPEGmonohexPLA
micelles
MPEG-dihexPLA
micelles
Drug loading [mg/g]
Figure 2. Size characteristics of CsA loaded PEG-(hex)PLA micelles in dependence of the drug
loading with the hydrodynamic diameter Zav (
), the number weighted diameter dn ( ) and the
percentage of micelles with a given dn ([%]dn (●)).
For the formulations envisioned for intravenous application, isotonic saline solutions were
prepared by first generating MPEG-hexPLA polymeric micelles in water and then adjusting
the solution isotonicity with the addition of NaCl. The average hydrodynamic diameter, Zav,
increased with the isotonicity change in the medium (Table 2), whereas the diameter dn
remained at 25 nm. The CsA solubility in polymeric micelles was found to be a little bit lower
with 1.4 mg/mL, as compared to 1.5 mg/mL in water. The more hydrophobic MPEGdihexPLA micelles incorporated higher amounts of CsA than MPEG-monohexPLA and
MPEG-PLA-based formulations. More highly-concentrated micelle solutions of 6 mg/mL
could be prepared by increasing the MPEG-dihexPLA copolymer concentration in the
medium while keeping the same drug to copolymer ratio.
74
Table 2. Characteristics of CsA loaded MPEG-dihexPLA micelles with different copolymer concentrations in different aqueous media.
MPEG-dihexPLA
concentration [mg/mL]
Aqueous Medium
CsA loading
[mgCsA/gcopolymer]
CsA concentration
[mgCsA/mLmicelles]
Size dn
[nm]
[%]dn
Size Zav
[nm]
P.I.
5
Water
307 ± 1
1.51 ± 0.01
26
100.0
40
0.24
5
Isoton. Saline Sol.
286 ± 11
1.43 ± 0.06
25
99.7
83
0.50
10
Water
318 ± 4
3.23 ± 0.05
26
100.0
40
0.24
20
Water
295 ± 5
5.97 ± 0.11
27
99.7
41
0.28
Chapter III
75
3.4 In vitro toxicity of MPEG-hexPLA solutions
To evaluate the in vitro toxicity of the unloaded MPEG-(hex)PLA micelles, three different
cell lines were tested: human intestinal epithelial Caco-2 cells, murine NuTu-19 and human
SKOV-3 ovarian cancer cells. Various copolymer concentrations were investigated, both
below the CMC for studying the toxicity of MPEG-(hex)PLA as unimers, and above the
CMC for studying the micelle solutions. The isotonic saline solution and CremophorEL
solutions were tested for comparison (Fig. 3). Triton X-100 and culture media were used as
0% survival and 100% survival controls, respectively.
After 24 h incubation at 37°C on Caco-2 cells (Fig. 3a), MPEG-(hex)PLA polymeric micelles
were non-toxic with a cell survival above 80% for copolymer concentrations ranging from
0.0001 to 25 mgcopolymer/mL, whereas CremophorEL micelles showed no toxicity only for
solutions with low concentrations (<1 mgcopolymer/mL). Negative cell survivals were observed
for higher concentrations, which can be explained by a washing-out of the cells from the test
plates. On NuTu-19 cells, MPEG-(hex)PLA micelles were non-toxic for concentrations up to
a 10 mgcopolymer/mL for MPEG-dihexPLA and MPEG-PLA solutions, and up to 1
mgcopolymer/mL for the comparable MPEG-monohexPLA solutions (Fig. 3b). CremophorEL
micelles showed no toxicity for low concentrations (<0.5 mg/mL). On the human SKOV-3
ovarian cancer cells, the three MPEG-(hex)PLA micelle solutions were non-toxic up to
concentrations of 10 mgcopolymer/mL, whereas CremophorEL micelles started to show a toxic
effect for a concentration of 1 mg/mL (Fig. 3c). At higher concentrations MPEG-hexPLA
micelles showed a slightly better cell viability than compared to MPEG-PLA micelles.
76
Figure 3. In vitro toxicity on (a) Caco-2, (b) Nutu-19 and (c) SKOV-3 cells of MPEG-PLA ( ),
MPEG-monohexPLA (
different concentrations.
), MPEG-dihexPLA ( ), CremophorEL ( ) and NaCl 0.9% () for
Chapter III
77
3.5 In vivo toxicity of MPEG-hexPLA micelle solutions
In vivo toxicity of MPEG-(hex)PLA micelles was evaluated using the CAM model. The chick
embryo survival was determined 24 h after the intravascular injection of 20 µL micelle
solution for four different copolymer concentrations of 0.0001, 5, 10 and 50 mg/mL. All
tested concentrations of MPEG-(hex)PLA formulations were non-toxic, since 100% of the
tested chick embryos were still alive at this crucial time point.
3.6 Hemolysis test of MPEG-hexPLA micelle solutions
The hemolytic property of MPEG-(hex)PLA solutions was assessed on fresh human blood
after incubation at 37°C for 24 h. The method was adapted from a hemolysis test protocol
described earlier by Mottu et al. 29. The percentage of hemolysis was determined for various
copolymer concentrations (Table 3). As compared to the control of an isotonic saline solution,
the unloaded MPEG-(hex)PLA micelle solutions showed less than 1% of red blood cell lysis
up to a copolymer concentration of 20 mgcopolymer/mL. Likewise, under much diluted
conditions that were below the CMC, the non hemolytic activity of MPEG-(hex)PLA unimers
was proven (Table 3a). For the highest concentration of 50 mgcopolymer/mL, which is much
higher than the formulation concentration that is considered to be practical, a hemolysis of
2.0% for MPEG-dihexPLA, 9.2% for MPEG-monohexPLA and 9.7% for MPEG-PLA were
observed. The incorporation of CsA into MPEG-dihexPLA micelles did not induce any red
blood cell lysis (Table 3b). A formulation of CsA in MPEG-dihexPLA (20 mg/mL) with a
CsA concentration of 6 mg/mL affected less than 1% lysis (Entry 1). Also, the formulations
with lower concentrations that are similar to those used in current clinical CsA doses (Entry 2
and 3) do not show hemolysis.
78
Table 3. Haemocompatibility of (a) unloaded and (b) CsA loaded MPEG-hexPLA micelles of
different copolymer concentrations after 24h incubation with human blood.
(a)
% lysis
Copolymer
concentration
[mg/mL]
50
0.01 to 20
0.0001 to 0.005 a
a
MPEG-dihexPLA
MPEG-monohexPLA
MPEG-PLA
2.0% ± 0.9%
<1%
<1%
9.2% ± 3.6%
<1%
<1%
9.7% ± 2.3%
<1%
<1%
Concentrations below CMC
(b)
Entry
1
2
3
4
a
CremophorEL
CsA
Surfactant
concentration
concentration
% lysis
[mg/mL]
[mg/mL]
6
78
<1%
2.5
32.5
<1%
0.5
6.5
<1%
<0.09 a
<1%
MPEG-dihexPLA
Copolymer
concentration
% lysis
[mg/mL]
20
<1%
8.3
<1%
1.7
<1%
<0.008 a
<1%
Concentrations below CMC
3.7 Shelf-life stability
The sizes and possible size changes of MPEG-hexPLA micelles were monitored using DLS at
different detection angles over a period of one year. At small detection angles of 30-40°, large
particles scatter more light than small particles, whereas at larger detection angles of 50-140°,
mainly the small particles are detectable. Therefore, it is important to measure micelle sizes at
the full range of detection angles in order to observe a possible micelle enlargement or
aggregation with time. The results obtained for MPEG-hexPLA micelles prepared in ultrapure water and stored at 25°C for one year are given in Fig. 4a. Directly after preparation, the
unloaded MPEG-hexPLA micelles had an average hydrodynamic diameter (observation
between 50°-140°) of around 30 nm for MPEG-monohexPLA and of around 40 nm for
MPEG-dihexPLA micelles, respectively. Only a few micelles larger than 50 nm were visible
at the small angles of 30°-40°. After one year of storage, the size curves of the initial and
stored solutions were similar, except at the low detection angle of 30°, where some larger
micelles in the stored solutions were observable. For all other detection angles, size changes
Chapter III
79
of only 10 nm for MPEG-dihexPLA and of 17 nm for MPEG-monohexPLA were observed.
The formation of larger micelles at small angles was also found for CsA-loaded micelles (Fig.
4b). This was more pronounced for MPEG-monohexPLA micelles, for which a substantial
increase of the diameter occurred at detection angles between 30° to 60°, showing the
existence of 1-2% of larger micelles in the stored solutions. For all other detection angles the
difference in diameter values after one year at 25°C varied in the same range as for the
unloaded micelles. The results from intermediate time points (data not shown) revealed that
micelle size enlargement or aggregation started only after five months for CsA-loaded
MPEG-monohexPLA and MPEG-PLA micelles.
Hydrodynamic diameter [nm]
(a)
250
200
150
100
50
0
30
50
70
90
110
130
110
130
Detection angle [°]
(b)
250
200
150
100
50
0
30
50
70
90
Figure 4. Hydrodynamic diameters Zav at different detection angles of (a) unloaded- and (b) CsA
loaded- MPEG-monohexPLA (,) and MPEG-dihexPLA (,) micelles in ultra pure water directly
after preparation (filled symbols) and after 1 year (unfilled symbols) at 25°C.
80
MPEG-(hex)PLA micelle formulations in isotonic saline solution were assessed for their
shelf-life stability, including a size and drug content evaluation.
(a)
1000
750
500
425
375
325
300
200
100
0
30
50
70
90
110
130
110
130
110
130
(b)
1000
750
500
425
375
325
300
200
100
0
30
50
70
90
(c)
1000
750
500
425
375
325
300
200
100
0
30
50
70
90
Figure 5. Hydrodynamic diameters Zav at different detection angles of CsA loaded MPEG-PLA
(●,○), MPEG-monohexPLA (,) and MPEG-dihexPLA (,) polymeric micelles in isotonic saline
solution directly after preparation (filled symbols) and after 3 months (unfilled symbols) at (a) 4°C, (b)
25°C and (c) 37°C.
Chapter III
CsA concentration
[µg/mLmicelles]
(a)
81
1200
1000
800
600
400
200
0
CsA concentration
[µg/mLmicelles]
(b)
0
2
4
0
2
4
0
2
4
6
8
Time [months]
10
12
14
10
12
14
10
12
14
1200
1000
800
600
400
200
(c)
1200
CsA concentration
[µg/mLmicelles]
0
1000
6
8
Time [months]
800
600
400
200
0
6
8
Time [months]
Figure 6. Formulation stability of CsA incorporated in MPEG-PLA (), MPEG-monohexPLA ()
and MPEG-dihexPLA () micelles in isotonic saline solution (a) at 4°C, (b) 25°C and (c) 37°C (n=3,
standard deviation for most time points smaller than symbols).
82
Fig. 5 shows the micelle hydrodynamic diameters of the three micelle solutions at (a) 4°C, (b)
25°C and (c) 37°C. Here the formation of larger micelles was visible at detection angles
between 30° to 60° after three months of storage time. The size change in these formulations
was influenced by the storage temperature. Larger micelles were observed at 37°C (200-450
nm) than at 25°C (100-750 nm) and at 4°C (150-220 nm). The drug content of CsA-loaded
MPEG-(hex)PLA formulations in isotonic saline solution at the same storage temperatures
was followed up over time (Fig. 6). The CsA content in micelles remained most stable at 4°C,
with a drug loss of less than 7% for all tested formulations after one year. At 25°C, the drug
loss increased to 26%, 82% and 100% for MPEG-dihexPLA, -monohexPLA, and MPEGPLA micelles, respectively. At 37°C, the entire incorporated amount of CsA was released
after three months in MPEG-PLA micelles, after seven months in MPEG-monohexPLA and
13 months in MPEG-dihexPLA micelles.
4. Discussion
Polymeric micelles based on MPEG-hexPLA are characterized by an increased
hydrophobicity of the micelle core in comparison to the standard MPEG-PLA. It has been
demonstrated that this enables higher drug loadings of poorly water-soluble drugs
21
. In this
paper, we investigated the biocompatibility of these novel micelles towards formulations of
the poorly water-soluble drug Cyclosporin (CsA) for intravenous applications.
The novel copolymers, MPEG-monohexPLA and MPEG-dihexPLA, and the “reference”
standard MPEG-PLA were synthesized in a controlled manner by ROP with a defined
molecular weight of 5000 g/mol (± 1 monomer unit), thus allowing a good comparison of the
influence of the introduced hexyl substituents along the PLA backbone. The increase of
hydrophobic interactions between copolymer chains favors micellization at lower critical
micellar concentrations (CMC). Indeed, the CMCs of MPEG-hexPLA micelles decrease with
the increase of hexyl groups in the core-forming block, leading to a value of 1.6×10-6 M and
1.7×10-6 M for MPEG-dihexPLA and -monohexPLA compared to a value of 2.0×10-6 M for
MPEG-PLA micelles
20
. These values in the micro molar range are in agreement with data
reported in the literature 5;30. Compared to MPEG-PLA and to other typical surfactants, which
Chapter III
83
have CMCs with 10 to 1000 times higher concentrations 4, the low CMC of MPEG-hexPLA
micelles facilitates better stability upon dilution, which can improve these formulations for
envisioned intravenous applications.
Moreover, MPEG-hexPLA micelles are very small and truly nanocarriers. Around 99% of
MPEG-hexPLA micelles had number weighted diameters (dn) between 18-30 nm, confirming
a unimodal size distribution in the unloaded (Table 1) or CsA-loaded state (Fig. 2). The
incorporation of CsA into the micelles did not influence the micelle size, dn. As a nanosized
drug delivery system, MPEG-hexPLA micelles should have the same advantages as described
for other polymeric micelles, in particular the ability to escape the mononuclear phagocyte
system (MPS) uptake and renal clearance, allowing a long circulation in the body, which
should result in a higher probability of reaching the target 30.
The potential of MPEG-hexPLA micelles to solubilize the hydrophobic drug CsA was herein
demonstrated in comparison to MPEG-PLA micelles and the CremophorEL surfactant.
Comparing the obtained CsA loadings for MPEG-hexPLA with the corresponding MPEGPLA formulations, it becomes obvious that the increased hydrophobicity of the new hexPLAbased micelles improve the loading and the solubility capacity. In aqueous MPEG-hexPLA
formulations, CsA was solubilized very efficiently up to the concentration of 1.5 mg/mL with
5.0 mg/mL MPEG-dihexPLA excipient (Table 2). Increasing the concentration of the MPEGhexPLA copolymer while keeping the drug/copolymer ratio constant leads to higher
concentrated formulations. At comparable polymer concentrations of 10 mg/mL, MPEGhexPLA could incorporate twice the amount of CsA as MPEG-polycaprolactone
(MPEG5000g/mol–PCL13000g/mol) micelles (~1.3 mg/mL) as reported by Aliabadi et al.
31
.
Formulations with 20 mg/mL MPEG-hexPLA facilitated a CsA concentration of 6 mg/mL,
corresponding to an increase in water solubility of CsA by a factor of 500. For comparable
CsA concentrations, MPEG-dihexPLA formulations would require four times less excipient
than a currently marketed formulation with CremophorEL as the surfactant.
Toxicity studies carried out on three different cell lines showed that MPEG-hexPLA micelles
had a lower toxicity than formulations with CremophorEL micelles on human ovarian cancer
SKOV-3 cells. On the two other cell lines, CremophorEL formulations at higher
concentrations than 0.5 mg/mL washed out cells from the test plates, and thus cell viabilities
could not be determined (Fig. 3). This was not the case for MPEG-hexPLA micelle
84
formulations, for which non-toxicity could be found for concentrations at least up to 10
mg/mL.
For possible intravenous pharmaceutical applications, isotonic micelle solutions of MPEGhexPLA micelles were prepared first in water and then adjusted by the addition of NaCl
crystals before use. By this procedure, CsA formulations with a concentration of 1.4 mg/mL
could be more efficiently prepared and the loss of drug observed in other procedures could be
significantly reduced. For example, by direct preparation in isotonic saline solution, MPEGhexPLA micelles could solubilize around 1 mg/mL CsA, which corresponds to an
incorporation of 48% (results not shown) as opposed to 73% when prepared in water and
adjusted afterwards to isotonicity. This was also observed for other copolymers like
MPEG5000g/mol-PCL13000g/mol micelles, which incorporated CsA with 28% or 37% efficiency
when prepared in isotonic solution, whereas in water a 37% or 64% efficiency could be
achieved, depending on the initial applied drug loading
32
. The choice of the medium is not
only of importance for the drug loading efficiency, but it is also an important factor regarding
the stability of the micelle formulations. In isotonic saline solutions, the formation of larger
micelles was observable after three months, whereas in water, no size increase could be
detected for MPEG-dihexPLA formulations after one year of storage (Fig. 4). A micelle size
enlargement after addition of salt has also been observed by Jain et al. and has been explained
by the dehydratation effects on the PEG units of the micelle shell
33
. For the investigated
formulations, the increase in micelle diameter was dependent on the storage temperature and
the copolymer. A storage temperature of 4°C showed the best formulation shelf-life stability,
whereas in the isotonic saline solution only a very few larger MPEG-hexPLA micelles were
observed after three months (Fig. 5), still with diameters below 200 nm, which is the
maximum size for particles with long-circulating properties
34
. To avoid their potential
consequences, larger micelles or aggregates could be simply filtered off with common sterile
filters after extended storage times. Regarding the CsA concentrations in the MPEG-hexPLA
formulations, the CsA amounts remained very constant over one year when stored at 4°C,
with a final drug loss of only 7% (Fig. 6). In contrast, at 25°C and 37°C, the initial drug
content in the micelles decreased more rapidly. At 37°C, a complete CsA release was
observed at earlier time points; the MPEG-hexPLA micelles showed a higher stability than
the MPEG-PLA micelles. A recent paper of Nottelet et al. have also demonstrated the
superior stability of MPEG-hexPLA micelles in PBS modeling intravenous conditions (pH
Chapter III
85
7.4 at 37°C) compared to MPEG-PLA micelles 35. These results could be associated with the
degradation of the polymers under these conditions. Previous studies on hexPLA polymers
have shown that, despite a similar degradation profile, the homopolymer PLA had a faster
molecular weight decrease and weight loss than hexPLA in PBS at pH 7.4 after 50 days
19
.
Thus, the slower polymer degradation and the lower CMC increase the excipient stability in
the formulations and stabilize the solubilized CsA drug content. Nevertheless, the data
obtained here for the higher temperatures indicate that the carrier system has the properties of
a certain shelf-life stability next to its ultimate degradation and drug release, when applied in
the human body.
The biocompatibility aspects of both MPEG-hexPLA unimers and micelles were investigated
by studying their toxicity in vitro and in vivo and by assessing their hemolytic activity on red
blood cells. Practically, the biocompatibility of unimers was carried out by modeling the
dilution effect after intravenous application and thus the disassembly of micelles. Unloaded
MPEG-hexPLA micelles were tested on three cell lines with different mammalian origins and
from different sources: human epithelial intestinal cells (Caco-2), murine (NuTu-19) and
human ovarian cancer cells (SKOV-3). Non-toxicity was found for MPEG-dihexPLA unimers
and for their polymeric micelles at least up to 10 mgcopolymer/mL on all tested cell lines and
was comparable to the results obtained for MPEG-PLA copolymers as controls (Fig. 3). In
addition to the MTT tests, the in vivo toxicity of MPEG-hexPLA micelles was studied on the
CAM model. This model is an alternative to mammalian models and has been proven helpful
for testing intravenous formulations due to its well-developed vasculature network
36
.
Injection volumes of 20 µL into the vasculature of the chick embryos were found to be welltolerated
28
and were therefore chosen for the toxicity studies of the MPEG-hexPLA micelle
solutions. The survival rate of chick embryos was assessed 24 h after injection, since longer
times usually give identical results. The CAM model results showed neither a toxicity for
MPEG-hexPLA unimers nor for the MPEG-hexPLA micelle formulations of the three tested
concentrations. All chick embryos survived the injection of the formulations with the novel
excipient. It has to be pointed out that MPEG-hexPLA micelles were non-toxic up to a
concentration of 50 mgcopolymer/mL, the same as was found for the control MPEG-PLA micelle
solution, whose biocompatibility has been reported in the literature 37. To add to these toxicity
studies, the compatibility with human blood was assessed. The hemolytic property of MPEGhexPLA unimers and unloaded micelles was investigated at different concentrations and
86
showed no hemolytic activity (<1% lysis) up to a copolymer concentration of 20 mg/mL. A
slight hemolytic activity was observed for the three different MPEG-(hex)PLA micelles for
the “extreme” copolymer concentration of 50 mgcopolymer/mL, which is far above the
concentrations possibly needed for an envisioned intravenous application. Also, under these
conditions, the novel hexPLA based micelles did not differ in their hemolytic activity from
the standard PLA based micelles.
Considering the practical application of CsA MPEG-hexPLA formulations, the corresponding
drug concentrations of 0.5 mg/mL up to 2.5 mg/mL, as used in the current CremophorEL
based products, were found to not induce hemolysis. The high drug loading capacity of
MPEG-hexPLA micelles allows a significant reduction of polymeric excipient for formulating
the same amounts of CsA. This increases the potential intravenous use of such micelle
formulations, which may improve the maximum tolerated dose (MTD) in the treatment. In a
previous study, MPEG-PLA micelles could increase four times the MTD of paclitaxel in nude
mice compared to the current treatment with Taxol, which is also formulated with
CremophorEL
38
. In addition, regarding the influence of the physical state of the
hydrophobic core forming block on the micelle stability, MPEG-hexPLA micelles are
expected to retain the drug more efficiently than CremophorEL micelles 31;37.
5. Summary and Conclusions
In this paper, we demonstrated the biocompatibility and non-toxicity of polymeric micelles
based on MPEG-hexPLA. The amphiphilic MPEG-hexPLA copolymers self-assembled in
aqueous media into polymeric micelles of 30 nm. The resulting micelle formulations show no
toxicity and no hemolytic activity in MTT tests, the CAM model and on human blood tests.
All results are comparably good or better than those obtained for the controls of standard
MPEG-PLA. By increasing the MPEG-hexPLA copolymer concentration while keeping the
drug:copolymer ratio constant, the immunosuppressive and hydrophobic drug Cyclosporin A
could be solubilized and formulated with concentrations up to 6 mg/mL, which is equivalent
to a 500-fold increase of the drug’s water solubility. The current clinically-used CsA
Chapter III
87
concentration administrated intravenously with Sandimmune can be prepared in a MPEGhexPLA micelle formulation by a simple procedure and would need four times less copolymer
than is used with the surfactant CremophorEL. MPEG-hexPLA micelles have the potential
to be very interesting non-toxic injectable nanosized drug carriers for improved formulations
of poorly soluble drugs.
Acknowledgments
The authors thank the Swiss National Science Foundation (SNF) for financial support (SNF
200020-103752).
The authors are grateful to Professor Michal Borkovec and Dr. Andrea Vaccaro (University of
Geneva, Geneva, CH) for the availability and use of the goniometer ALV/CGS-5.
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1234563746489AB93CDE69EF6CEFC93CDE6379367416C4AA469546EDE3DC69EF6FD6ED36
7D6 74DAC6 93C8C36 9E6 CE35984EDB6 9AC93CDE6 D26 7416 C4AA46 C376 3746
1
2ABD544E36 745CCE6 95456 96 3BFC4F6 96 96 E46 FC9ED3C6 2D5BA93CDE6 2D56 D895C9E6
9E4569EF643939C6F4343CDE 66
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1
Chapter IV
91
MPEG-hexPLA Micelles as Novel Carriers for
Hypericin, a Fluorescent Marker for Use in Cancer
Diagnostics
K. Mondon, M. Zeisser-Labouèbe, R. Gurny, and M. Möller
School of Pharmaceutical Sciences, University of Geneva, University of Lausanne, 30 quai
Ernest-Ansermet, CH-1211 Geneva 4, Switzerland.
Submitted to: Photochemistry and Photobiology
Ovarian cancer is the most common gynecological cancer diagnosed in Western countries.
Detection of micrometastases at an early stage of the disease could lead to a cure rate of 90%
by limiting the spread of the disease outside the ovaries. In this article, hypericin (Hy), a
hydrophobic photosensitizer used for the photodynamic diagnosis of ovarian cancer, was
efficiently incorporated into a core of micelles made from methoxy-poly(ethylene glycol)
(MPEG) and hexyl-substituted poly(lactides) (hexPLA) copolymers. The fate of these
micelles following intravenous injection was studied in vivo in two ovarian tumor-bearing
animal models. In the chick embryo chorioallantoic membrane (CAM) model, 17 times more
Hy accumulated in tumor nodules when Hy was delivered with micelles than when Hy was
delivered as an ethanol solution. Studies of the biodistribution of Hy in Fisher rats revealed
escape of these nanosized micelles (<32 nm) from the mononuclear phagocyte system. Hyloaded micelles showed maximal accumulation in tumors and demonstrated the best
tumor/muscle contrast visible 3h after injection in the rat model. The rapid and highly
selective accumulation of Hy in tumors that we demonstrated in this study suggests that these
micelle formulations could be used for the photodynamic diagnosis of ovarian cancer in the
future.
Keywords: diagnosis; drug delivery; hypericin; imaging; ovarian cancer; polymeric micelles
Hy/ MPEG-hexPLA Micelle Formulations for Ovarian Cancer Diagnostics
92
Chapter IV
93
1. Introduction
In the current era, cancer is overtaking cardiovascular disease and becoming the leading cause of
death worldwide. According to the World Health Organization, cancer-related deaths will reach
12 million in 2030 as compared to 7.4 million in 2004. Early detection of many types of cancers
can reduce fatality rates by as much as 30%. Additionally, a precise localization of tumor tissues
is essential before surgery, radiotherapy or chemotherapy is considered 1. Despite the progress
that has been made in imaging technology (ultrasound, endoscopy and radiography), ovarian
cancer is still the leading cause of death from gynecological malignancies in Western countries 2.
Due to the absence of symptoms during the early stages of the disease, the majority of affected
patients have already developed metastases when their cancer is diagnosed. As a consequence, the
5-year survival rate very much depends on the stage of the disease at the time of diagnosis. For
stage I (localized) disease, the 5-year survival rate is 93%, whereas for stage IV (distant) disease,
it decreases drastically to 31% 2. Early surgery combined with taxane- or platinum-based
chemotherapy has improved the overall survival rate of ovarian cancer patients, but survival rates
are still unacceptably low for patients with advanced disease. Half of these patients will relapse
within five years after surgery 3, mainly due to non-detected residual metastases in the peritoneal
cavity 4. A newly developed selective photodynamic diagnosis (PD) imaging technology has
recently been shown to efficiently detect bladder cancer
5
. This technique consists of
administrating a photosensitizer (PS) or a precursor for the in situ formation of a PS, which
preferentially accumulates in diseased tissues 6. Due to its fluorescence properties, the PS is easily
located when excited at the appropriate wavelength, enabling easy detection of malignant tissues.
Of the PSs that have been studied, 5-aminolevulinic acid (5-ALA)-mediated protoporphyrin IX
(PP IX), Photofrin®, Temoporfin (Foscan®) and Metvix® have been successfully used for the
detection of various cancers 7, such as bladder 5, esophageal 8, skin 9 and head and neck cancer 10.
In ovarian cancer, 5-ALA-mediated PP IX has shown feasibility and promise in detecting
metastases in both animal models 11 and humans 4. Another PS that has been investigated for use
in cancer detection is hypericin (Hy). Hy is an interesting PD compound due to its low
photobleaching 12 and its natural origin as well as its photoactivity in several cancer cell lines 13-15
and rodent models
13;14;16
. Recently, Hy has been found to be effective at detecting oral cancer 17
as well as ovarian tumors in rats
22
challenge due to its high degree of hydrophobicity (log P=8.78
23
and bladder tumors in humans
18-21
. Formulating Hy is a
). Polylactide (PLA)
94
nanoparticles (Nps) loaded with Hy have recently been evaluated and have been shown to
specifically reach tumor nodules in rats following intravenous (i.v.) administration. Because these
polymeric Nps had a longer half-life as compared to the drug in solution as well as a small size
(~200 nm), they were able to extravasate from the neovascularature and accumulate in diseased
tissues through the enhanced permeation retention (EPR) effect
22;24
. Their small size, long
circulation time and specific accumulation and transport of the drug into tumor tissues are the
reasons that there has been increasing interest in nanocarrier systems in cancer research 25-27.
Amongst the different nanocarrier systems that have been studied, polymeric micelles have
gained special interest because of their nanoscale small size, in vivo stability, high drug-loading
capacity and good biocompatibility
28
. Polymeric micelles have a specific core-shell structure
formed by the self-assembly of amphiphilic copolymers. The outer shell, which is composed of a
hydrophilic polymer, most often poly(ethylene glycol) (PEG), reduces interactions with blood
proteins, facilitating a long circulation time. The inner core, which is formed by the hydrophobic
block, can successfully be loaded with lipophilic drugs, such as PSs
29-33
. Biocompatible and
biodegradable polymers like PLA are preferred for use in the hydrophobic core polymer, but
some hydrophobic drugs can only be incorporated into PLA micelles in a limited fashion 34. Thus,
it would be ideal to identify and create polymers with an increased degree of hydrophobicity that
would allow more efficient drug loading, which, as a result, could lead to better clinical results.
Our group recently demonstrated the efficiency with which hydrophobic drugs, such as the PS
meso-tetra(p-hydroxyphenyl)porphine (THPP) could be incorporated into (hexPLA)-based
micelles. These micelles were shown to load this hydrophobic drug more efficiently than standard
PLA. Successful loading of 123 mgTHPP/gcopolymer was achieved with MPEG-hexPLA, as
compared to 62 mgTHPP/gcopolymer that was achieved with MPEG-PLA micelles
35
. To the best of
our knowledge, there are no previous reports in the literature examining the potential
pharmaceutical applications of polymeric micelle-incorporated Hy. In this paper, we investigated
the potential utility of Hy formulated in MPEG-hexPLA micelle solutions in ovarian cancer
diagnostics. Hy-loaded MPEG-hexPLA micelle formulations with a high drug-loading capacity
were prepared and investigated in two different ovarian tumor-bearing animal models, the chick
embryo chorioallantoic-membrane (CAM) and the female Fisher rat F-344.
Chapter IV
95
2.1. Materials
Hypericin (Hy) was purchased from Alexis Corporation (Lausen, Switzerland). Tetrahydrofuran
(THF) was supplied by SDS (Toulouse, France). Methoxy-poly(ethylene glycol) (MPEG) with a
molecular weight of 2000 g/mol was a gift from Union Carbide Corporation (USA). D,L-lactide,
tin(II) 2-ethylhexanoate (Sn(Oct)2) and acetone p.a. were purchased from Purac Biochem
(Gorinchem, The Netherlands), Aldrich (Buchs, Switzerland) and Fluka (Buchs, Switzerland),
respectively, and used as received. The monomers mono-hexyl-substituted lactide (monohexLA)
and di-hexyl-substituted lactide (dihexLA) were synthesized as described in a previous
publication 36.
2.2. Synthesis and characterization of MPEG-hexPLA copolymers
The synthesis of MPEG-hexPLA copolymers with a molecular weight of approximately
5000g/mol has been described previously (34-35). Briefly, MPEG-hexPLA copolymers were
synthesized by ring opening polymerization (ROP) with a MPEG of 2000g/mol as the initiator
and tin octanoate as the catalyst. The obtained copolymers were characterized by their molecular
weight (Mn) and polydispersity index (P.I.) by gel permeation chromatography (GPC). The
structures of MPEG-hexPLA are presented in Scheme 1.
O
H3C
O
45
R
O
O
R
1
2
O
H
m
O
R1=C6H13, R2=CH3 : MPEG2000g/mol-monohexPLA3000g/mol for m=14
R1, R2 =C6H13
Scheme 1. Structure of MPEG-hexPLA block copolymers
for m=11
96
2.3. Preparation of Hy-loaded MPEG-hexPLA micelles
Hy-loaded MPEG-hexPLA micelles were prepared by the co-solvent evaporation method
described by Mondon et al.
37
. Briefly, 100 mg/mL solutions of MPEG-hexPLA copolymers in
acetone and a 10 mg/mL solution of Hy in acetone were prepared. Next, 500 µL or 200 µL of Hy
solution were gently mixed with 400 µL of MPEG-dihexPLA or 800 µL of MPEG-monohexPLA
solution, respectively. Acetone was added to the organic phase in order to obtain a final volume
of 1 mL. The organic mixture was poured into 2 mL isotonic saline solution under sonication and
then slowly removed by evaporation at 15 mbar. The final micelle concentrations were adjusted
by adding isotonic saline solution to reach a copolymer concentration of 20 mgcopolymer/mL for
MPEG-dihexPLA and 40 mgcopolymer/mL for MPEG-monohexPLA (see structural difference in
Scheme 1). The solutions were left to equilibrate overnight and then centrifuged at 9500 x g for
15 min to remove non-incorporated Hy. Both MPEG-hexPLA micelles were then analyzed for
size and drug loading.
2.4. Characterization of Hy-loaded MPEG-hexPLA micelles
The hydrodynamic (Zav) and number-weighted (dn) diameters of micelles were measured by
dynamic light scattering (DLS) after centrifugation. Analyzes were performed at 25 °C with a
Zetasizer HS 3000 system (Malvern, Worcestershire, UK) at an angle of 90°. For each sample,
mean diameters were obtained after 3 runs of 10 measurements. For drug-loading determination,
the micelles were broken up by a 10-fold dilution with acetonitrile. The released Hy was
quantified in triplicate by reversed-phase HPLC as previously described by Zeisser-Labouèbe
et al. 38. The HPLC was calibrated with standard solutions of 2.0–200 µg/mL of Hy dissolved in
acetone. The resulting calibration curves were obtained with regression coefficients greater than
0.99.
The incorporation efficiency and drug loading were calculated using equation (1) and (2),
respectively:
Incorporation efficiency (%) =
Drug loading (% w/w) =
mass of drug incorporated in micelles (g)
× 100
mass of drug introduced (g)
× 100
mass of copolymer used (g)
(1)
(2)
Chapter IV
97
2.5. Animals
Female Fisher rats F-344 (150-200g) were purchased from Charles River Laboratories
(L’Arbresle, France) and housed in a temperature-controlled room with a 12 h light/dark cycle.
They were given a commercial basal diet and water ad libitum. All animal experiments and
animal husbandry were carried out in compliance with national regulations and approved by the
cantonal veterinary office of Geneva, Switzerland (Certificate number 31.1.1020/3065/2).
2.6. Cell culture and cell preparation for in vivo inoculation
The NuTu-19 cell line, a poorly differentiated Fischer 344 rat-derivative epithelial ovarian cancer
cell line
39
, was kindly provided by Dr. A. Major (Geneva University Hospital, Geneva,
Switzerland). The cells were cultured in Dulbecco's Modified Eagle's medium (DMEM) (Gibco
Life Technologies, Carlsbad, USA) supplemented with 10% fetal bovine serum (FBS)
(Brunschwig, Amsterdam, The Netherlands) and 100 U/mL penicillin–streptomycin (Gibco Life
Technologies) at 37 °C (atm. 5% CO2). Before tumor implantation onto the CAM or into rats,
NuTu-19 cells were washed twice with phosphate buffered saline (Gibco Life Technologies,
Carlsbad, USA), harvested using 0.5% Trypsin-EDTA (Gibco Life Technologies, Carlsbad,
USA) and counted. After centrifugation, an equal mixture of complete culture medium and
Matrigel matrix (BD Biosciences, Bedford, USA) was prepared. A suspension of 108 cells/mL
was prepared for the CAM experiments and a suspension of 5x106 cells/mL was prepared for the
rat experiments.
2.7. Chick embryo chorioallantoic membrane (CAM) study
The chick embryo chorioallantoic membrane (CAM) model was adapted from the model
described by Lange et al. 40. Egg incubation was performed according to the procedure described
by Vargas et al.
24;41
. Briefly, fertilized hen eggs, which were kindly provided by the Geneva
University Animal House (Geneva, Switzerland) were placed with the narrow apex down in an
incubator (Savimat MG 200, Chauffry, France) at 37 °C with a relative humidity of 65%. The
eggs were rotated twice a day until embryo development day 3 (EDD3). On EDD4, a 3-mm hole
was drilled into the eggshell and the narrow apex was covered with an adhesive tape. The eggs
were incubated in a static mode until EDD8, when the hole in the eggshell was enlarged by 2 or 3
mm to allow the placement of a silicon O-ring (Apple Rubber products inc., Lancaster, USA) on
98
the CAM. Next, 20 µL of the ovarian cancer NuTu-19 cell suspension (108 cells/mL) were
inoculated topically to allow tumor growth inside the O-ring area. The open hole was then sealed
with a plastic film (Parafilm, Pechiney Plastic Packaging, Chicago, USA) to avoid contamination
and desiccation of the CAM. Eggs were returned to the static incubator until EDD12, the day of
the test formulation administration. A total of 3 formulations were tested on 5 eggs each: (a) an
Hy-loaded MPEG-monohexPLA formulation with a concentration of 0.69 mg/mL Hy; (b) an Hyloaded MPEG-dihexPLA formulation with a concentration of 1.0 mg/mL Hy; and (c) a Hy
solution (1.0 mg/mL Hy in a mixture of ethanol, PEG 400 and water). These formulations were
injected via the main vessel of the CAM at a dose of 2 mg/kg (i.e., 20 ng/egg). The eggs were
returned to the incubator and the fluorescence of the tumor nodules was evaluated at 1 min, 1 h, 3
h and 6 h post-injection by fluorescence microscopy with an attenuation coefficient of 4 [see
details of the set-up in reference 24.
The relative tumor fluorescence (Frel) was calculated using equation (3):
Frel =
(Ft - F0 )
,
F0
(3)
where Ft is the tumor fluorescence intensity at the different time points and F0 is the tumor
autofluorescence prior to drug injection.
2.8. Biodistribution of Hy-loaded MPEG-dihexPLA micelles in rats
Ovarian tumor-bearing female Fisher rats (F-344) were used for Hy biodistribution studies.
The rats were inoculated intraperitoneally with 1 mL of ovarian cancer NuTu-19 cell
suspension (5x106 cells) to induce tumor growth. After five weeks, ascites were palpable and
tumors had grown sufficiently to allow us to perform the study. The rats were injected
intravenously in the tail vein with Hy formulations at a dose of 2 mg/kg. Three formulations
were investigated: (a) Hy loaded MPEG-dihexPLA micelles; (b) Hy solution (2 mg/mL in a
mixture of ethanol, PEG 400 and water); and (c) isotonic saline solution (which was
administered to two rats as a control). After intravenous injection of the Hy formulations, 4
rats per formulation were sacrificed by CO2 asphyxia at 1, 3, 6 and 24 h after injection for the
micelle formulations and 1 h after injection for Hy in ethanol solution. In order to visualize
the tumors by fluorescence imaging, the abdominal cavity was opened and images were
compared under white and blue light, respectively using the set-up described previously by
Chapter IV
99
Zeisser-Labouèbe et al. 22. Blood samples were obtained by cardiac puncture with heparinized
tubes. The liver, spleen, lung, tumor and muscles surrounding the tumors were excised. All
samples were weighed and stored at -20 °C until analysis. Tissue samples were extracted with
tetrahydrofuran (THF) complemented with a tissue homogenizer (Eurostar digital, IKA®Werke, Staufen, DE). Blood samples were prepared for analysis by THF extraction and
sonication (5 times for 5 s with a sonifier S-450D®, Branson Ultrasonic S.A, Geneva, CH).
After sample centrifugation, the supernatants were evaporated under nitrogen. The residues
were dissolved in 0.3 mL DMSO and their Hy fluorescence was determined with a microplate
reader (Safire®, Tecan, Salzburg, AT) at the excitation and emission wavelengths of 530 and
645 nm. The obtained fluorescence intensity was corrected by subtraction of the fluorescence
of control samples. The respective Hy concentration was calculated from the calibration curve
(1.95−25 ng/mL), which had a regression coefficient higher than 0.999.
2.9. Stability of Hy-loaded MPEG-dihexPLA micelles in blood plasma.
Human blood plasma from an anonymous AB donor (Geneva University Hospital,
Switzerland) was used for the study of the stability of Hy loaded MPEG-dihexPLA micelles
in the presence of plasma. A volume of 187 μL Hy loaded MPEG-dihexPLA micelles was
added and mixed with 12 mL plasma for 1, 3, 6 and 24h in the dark at 37°C under orbital
shaking to mimic the experimental conditions of the in vivo studies. A control experiment
with Hy solution (2 mg/mL in a mixture of ethanol, PEG 400 and water) was also carried out
under the same conditions, except that 94 μL of the Hy solution were mixed with 6 mL
plasma. After shaking, 4 mL of the plasma/micelles mixture were centrifuged for 5 min at
10 000 rpm (Beckman, AvantiTM, Fullerton, USA) in order to remove non-entrapped drug.
One milliliter of the supernatant was dissolved in 2 mL THF under sonication. The resulting
organic mixture was afterwards centrifuged for 5 min at 10 000 rpm to remove blood
components that have precipitated with the addition of THF. The entire remaining supernatant
was collected, dried, and dissolved in 400 μL acetone. The acetone solution was finally
centrifuged to remove all other precipitates (5 min, 10 000 rpm), and directly analyzed (with
no dilution) by HPLC for the Hy drug content.
100
Please note that the used volume of blood plasma was chosen to simulate the in vivo
concentrations in the animal model, and was calculated from the following equation (4) for
the blood volume and the body weight of the rats 42:
Blood volume (mL) = 0.06 × body weight (g) + 0.77
(4)
Here, the body weight corresponded to 187 g, the average body weight calculated from the 22
rats studied in the in vivo experiments.
2.10. Statistical analysis
Results are expressed as means ± SDs (standard deviations). The significance of betweengroup differences was determined using Student’s t-test. All p-values <0.05 were considered
to be statistically significant.
3. Results
3.1. Hypericin/MPEG-hexPLA micelle formulations
Hypericin (Hy)-loaded micelle formulations were prepared for in vivo evaluation in the CAM
model. Table 1 shows the size characteristics and Hy incorporation of these micelles. MPEGmonohexPLA and MPEG-dihexPLA micelles had number weighted diameters (dn) of 32 nm
and of 19 nm and mean hydrodynamic diameters (Zav) of 133 nm and 85 nm, respectively.
The high Hy incorporation efficiency that we observed (>70%) yielded a Hy solubility of 0.69
mg/mL for MPEG-monohexPLA micelles and 2.02 mg/mL for MPEG-dihexPLA micelles.
These values correspond to a drug-loading of 1.7% and 9.8 % (%w/w), respectively. It is
notable that with the MPEG-dihexPLA micelles, half of the amount of excipient was needed
to create a formulation with double the Hy concentration in comparison to the less
hydrophobic MPEG-monohexPLA.
Chapter IV
101
Table 1. Hy loaded MPEG-monohexPLA and MPEG-dihexPLA micelle formulations and characteristics.
Polymeric micelles
Mn
(g/mol)
Mn/Mw
Copolymer
Hy
Encapsulation
concentration
concentration
Efficiency
Drug
loading
(%w/w)
Zav
(nm)
P.I.
dn
[%]dn
MPEG-monohexPLA
5600
1.15
40 mg/mL
0.69 mg/mL
69.6%
1.7
133
0.1
32
100
MPEG-dihexPLA
5351
1.12
20 mg/mL
2.02 mg/mL
80.3%
9.8
85
0.3
19
100
P.I.: Polydispersity index
102
3.2. Chick embryo chorioallantoic membrane (CAM) study
Four days after topical inoculation of NuTu-19 cells onto the CAM (which occurred on
EDD8), visible tumor nodules with extensive neovascularization were observed at the
membrane surface. Hy-loaded micelles and the Hy solution were injected at the same dose
(2 mg/kg). The fluorescence intensity of the accumulated Hy in the nodules was observed and
photographed 1 min, 1 h, 3 h and 6 h following the administration. The corresponding
fluorescence images showed the increase in fluorescence intensity over time within the
ovarian tumor nodules as compared to the non-tumoral surrounding tissues (Fig. 1).
Compared to Hy in solution (upper row), both Hy-loaded MPEG-hexPLA micelle
formulations (lower row) led to a much higher fluorescent signal in the nodules.
Figure 1. Fluorescence images of nodules after injection of Hy in solution and Hy loaded MPEGdihexPLA micelle formulation at a dose of 2 mg/kg in chick embryo (CAM model).
Chapter IV
103
The higher accumulation of Hy within the tumors was confirmed by the plot of the relative
tumor fluorescence intensity versus time (Fig. 2). No statistically significant differences were
found between the two micelle formulations. However, the micelle formulations led to a
significantly higher fluorescence as compared to the drug in solution at all time points
(p<0.05). Six hours after injection, the fluorescence intensity in the tumor nodules that was
achieved by the MPEG-hexPLA micelle formulations was 13 to 17 times higher than that
achieved by the Hy in ethanol solution.
15
Corrected
Fluorescence intensity (AU)
*
*
*
10
*
*
*
5
*
*
0
0
1
2
3
Time (h)
4
5
6
Figure 2. Relative tumor fluorescence intensity over time after i.v. administration of Hy loaded
MPEG-monohexPLA micelles (▲), MPEG-dihexPLA micelles (■) and Hy solution () at a dose of 2
mg/kg in chick embryos (mean ±S.D., n=5). * Significantly different from chick embryos injected
with Hy solution (Student’s t test, p<0.05).
104
3.3. Biodistribution of Hy-loaded MPEG-dihexPLA micelles
The biodistribution of Hy-loaded micelles was investigated following injection into Fisher rats at
a dose of 2 mg/kg body weight. In this experiment, the Hy-MPEG-dihexPLA formulation with
the higher drug-loading efficiency (at a concentration of 1 mg/mL) was chosen. Hy in a clear
ethanol-PEG 400-water solution was used for comparison. No sign of precipitation was visually
observed at the site of injection for both tested solutions. The plasma Hy concentration profiles
observed for the Hy solution and Hy-loaded micelles are shown in Fig. 3. The results for the Hy
solution concerning the 3, 6 and 24h time points were taken from the published work of ZeisserLabouèbe et al. (who performed their studies on Fisher 344 rats) 22. At all tested time points, the
Hy plasma levels that were obtained were significantly higher when Hy was incorporated into
micelles as compared to Hy in ethanol solution. At 1 h after injection, Hy loaded in micelles
showed a plasma concentration that was 6 times higher concentration than that was achieved
when the rats were injected with ethanol solution. After 3 h, the plasma concentration achieved by
the Hy-loaded micelles was 4 times higher than was achieved with Hy in ethanol solution and
after 6 h, it was 2 times higher. After 24 h, the plasma concentration achieved by the Hy-loaded
micelles was 3 times higher than for the drug in solution, indicating that the circulation time of
Hy was significantly prolonged when it was incorporated in the micelle carriers.
8.00
Hy concentration
(µg/ml)
*
6.00
4.00
*
2.00
*
*
0.00
1
3
Time (h)6
24
Figure 3. Plasma concentration profile of Hy after i.v. injection in rats of Hy solution (white bars) and
of Hy loaded MPEG-dihexPLA micelles (dashed bars) at a dose of 2 mg/kg. (Mean ±SD, n=3-4). *
Significantly different from rats injected with Hy solution (Student’s t test, p<0.05). Results of Hy in
solution after 3, 6 and 24h were taken from reference 22.
Chapter IV
105
The distribution of both micelle formulations that were tested in rat livers, spleens and lungs
is illustrated in Fig. 4. The uptake of Hy following administration of Hy in solution was
higher in the spleen than in the lungs and liver. Hy delivered in micelles was found in higher
concentrations in the spleen and lungs than in the liver. A total of 24 h after administration,
Hy that had been loaded in micelles was eliminated by these organs at the same rate as the
drug solution. Both formulation profiles were similar for the liver and led to a final Hy
concentration of 0.60 and 0.43 µg/g tissue when Hy was delivered in solution and in micelles,
respectively. Hy-loaded micelles exhibited a slower elimination from the spleen than Hy in
ethanol solution at 1 h, 3 h and 6 h after injection, but showed the same concentration after 24
h. In the lungs, Hy loaded in micelles showed a similar profile as in the spleen, except that
after 24 h, Hy that had been administrated in micelle formulations was more rapidly
eliminated than the drug solution.
Hy concentration
(µg/g)
25.00
20.00
15.00
10.00
5.00
0.00
1h
3h
6h 24h 1h
Liver
3h
6h 24h 1h
Spleen
3h
6h 24h
Lung
Figure 4. Biodistribution of Hy in liver, spleen and lung over time after injection of Hy solution
(white bars) and of Hy loaded MPEG-dihexPLA micelles (dashed bars) at a dose of 2 mg/kg in rats
(Mean ±SD, n=3-4). Results of Hy in solution after 3, 6 and 24h were taken from reference 22.
106
In addition to examining the tissue distribution of Hy, we also determined the accumulation of
Hy-loaded MPEG-dihexPLA micelle in the tumors and surrounding muscle tissue. For both
tested formulations, the Hy concentration in the surrounding muscles was very low as
compared to the concentration that was observed in the other investigated organs. However, a
slightly higher Hy accumulation in muscles was observed for Hy-loaded micelles as
compared to Hy in solution. In the tumors, a significant accumulation of Hy-loaded micelles
was found for the first 6 h after injection, with a maximum level detected at 3 h (Fig. 5). After
24 h, the Hy accumulation decreased to a concentration that was similar to the concentration
that was obtained with the drug in solution.
1.25
*
*
Hy concentration
(µg/g)
1
0.75
*
*
0.5
*
*
*
0.25
0
0
6
12
Time (h)
18
24
Figure 5. Biodistribution of Hy in tumors (open symbols) and the surrounding muscle (plain symbols)
over time after injection of Hy solution (,) and of Hy loaded MPEG-dihexPLA micelles (,■) at a
dose of 2 mg/kg in rats (Mean ±SD, n=3-4). * Significantly different from rats injected with Hy
solution (Student’s t test, p<0.05). Results of Hy in solution after 3, 6 and 24h were taken from
reference 22.
Chapter IV
107
For the application of Hy-loaded micelles for tumor imaging to be successful, a high contrast
between healthy tissues (muscle) and diseased tissues (tumor) is essential. The visual
distinction between these tissues is possible by using endoscopy with blue light. Under white
light, no difference between muscle tissues and tumor nodules could be observed (Fig. 6).
However, under blue light, the red fluorescence of Hy accumulated in the tumor nodules
became clearly visible, while the absence or near absence of red fluorescence in the healthy
muscle indicated the absence or very low accumulation of Hy.
Figure 6. Pictures of ovarian metastasis in the peritoneal cavity of rats under white light and blue light
after injection of Hy loaded MPEG-dihexPLA micelles. Below the given contrast was obtained
between diseased tissues and surrounding healthy muscle.
Visually, the maximum contrast between healthy and diseased tissues was observed 3 h after
micelle injection. To achieve more quantitative measurements, the tumor-to-muscle ratio was
calculated from the extracted and analyzed Hy concentrations. This analysis also found that
the highest contrast between tumor and muscle tissue was present 3 h after injection of Hyloaded micelles (Fig. 7). It is also worth noting that this time was also the point at which
maximum concentration of the fluorescent marker could be detected in the tumor tissue
(Fig. 5).
Hy Ratio Tumour/Muscle
108
7.0
6.0
5.0
4.0
3.0
2.0
1.0
0.0
1
3
6
24
Time (h)
Figure 7. Tumor to muscle contrast over time after injection of Hy solution (white bars) and of Hy
loaded MPEG-dihexPLA micelles (dashed bars) at a dose of 2 mg/kg in rats. Results of Hy in solution
after 3, 6 and 24h were taken from reference 22.
3.4. Stability of Hy-loaded MPEG-dihexPLA micelles in blood plasma
The stability of Hy loaded micelles has been studied in blood plasma and has been performed
under similar conditions as in the in vivo experiments, i.e. with the same dilution factor and
the same dose. Hy loaded micelles retained more than 80% of the introduced Hy quantity in
plasma at all the tested time points (1, 3, 6 and 24h), whereas for the control of Hy in ethanol
solution only 10-15% of Hy was found in the plasma (Fig. 8).
Figure 8. Percentage of Hy recuperated from human blood plasma when added as an ethanol solution
() and when added within MPEG-dihexPLA micelles (■), respectively.
Chapter IV
109
4. Discussion
Over the last decades, nanosized polymeric micelles have become a topic of increasing interest in
the field of drug delivery. Because of their specific core-shell structure, they can carry potent
hydrophobic drugs within their core, while their hydrophilic shell facilitates water solubility.
Their ability to accumulate at tumor sites via the EPR effect envisions localized cancer
treatments 43.
In the present study, the objective was to investigate novel MPEG-hexPLA micelles for their
potential use in ovarian cancer diagnosis. We examined their ability to achieve efficient drug
loading, a long circulation time and selective accumulation in tumor tissue in a series of in vivo
experiments. As demonstrated, the relatively water insoluble fluorescent marker Hypericin was
successfully incorporated into MPEG-hexPLA micelles, reaching an aqueous solubility of
~2 mg/mL, which represents a high increase in its water solubility as compared to the drug alone.
In addition, incorporating Hy in MPEG-hexPLA micelles increased its circulation time in the
bloodstream, which is a major advantage for passive tumor targeting. Even 24 h after injection
into ovarian tumor-bearing rats, Hy loaded MPEG-dihexPLA micelles were still present in the
plasma. The distribution of Hy has been proven in vitro to be due to a stable incorporation of Hy
into micelles for 24h and not from a gradually release of the drug from the micelles. The loss of
Hy when added within micelles in human blood plasma (<20%) observed after 1h can be
explained either by the multiple steps needed for the extraction of the drug, by a possible
equilibrium needed to be reached during the first hour between the micelles and the plasma
components, or by both phenomena. The observed low Hy plasma concentration even after 1h
can be attributed to a probably very fast decrease of the Hy loaded micelle concentration in the
blood between the injection and the first studying time point. Indeed, a study with comparable
MPEG-PLA micelles (molecular weight around 5000g/mol with MPEG 2000g/mol) and with
paclitaxel as the incorporated drug demonstrated a biphasic nature of the plasma profile with an
initial fast distribution phase completed within the first hour 44. The presence of inert hydrophilic
PEG on the nanocarrier surface provided good steric hindrance, limiting blood serum protein
binding and as expected increasing the carrier’s circulation time
45;46
. In addition, the increased
hydrophobicity of the hexPLA core, in comparison to a PLA or PCL micelle core, allowed the
novel MPEG-dihexPLA micelles to have a lower critical micelle concentration
47
, thus a higher
110
stability upon dilution, and a higher stability in PBS modeling intravenous conditions (pH 7.4 at
37°C) 48, consequently increasing the life time of the Hy loaded in micelles in the blood stream.
For cancer diagnosis, increased accumulation of the fluorescent marker at the tumor sites
enables better tumor detection. In both in vivo studies, Hy loaded micelles primarily
accumulated in ovarian tumors and thereby meeting this requirement. The accumulation of
free Hy is often associated with aggregation, and a possible loss of fluorescence intensity.
Here, this quenching phenomenon is not likely to occur as Hy released from the micelles at
the site of interest would not aggregate due to the presence of proteins like LDL, HDL or
albumin49;50 to which Hy rapidly binds and does not influence the fluorescence properties. In
the CAM studies, accumulated Hy-loaded micelles showed a much higher intensity of
fluorescence as reported for Hy-loaded nanoparticles (Nps)
24
. A 5-fold increased
accumulation of Hy in tumors was obtained with Hy-loaded MPEG-hexPLA micelles as
compared to Hy-loaded Nps 3 h after injection. In ovarian tumor-bearing rats, the
biodistribution study confirmed that Hy micelles rapidly accumulated in tumor tissues at high
concentrations. Hy loaded in MPEG-dihexPLA micelles reached a maximum concentration
and a maximum tumor-to-muscle contrast 3 h after injection. In contrast Hy-Nps showed its
highest accumulation 24 h after injection 22. The differences that were observed in the tumor
accumulation profiles between the two animal models (rat versus CAM) can be attributed to
the absence of an elimination process in the CAM model
24
. In the rat experiments, the
differences between the two Hy nanocarriers (i.e., micelles and Nps) can be explained by
differences in the size of the two carrier systems. It has been reported that the accumulation
effect is primarily influenced by the size of the carrier
51
. Because of their small size,
nanocarriers can easily pass through the gaps in leaky endothelial walls and are efficiently
taken up by tumor cells. Nagayasu et al. reported that for liposomes, a diameter of <100 nm
seemed to be most suitable in order to achieve a long half-life in the blood circulation, tumorspecific drug accumulation and in vivo drug release 52. Moreover, another biological process
of micelle uptake in tumor tissues has recently been described by Kawaguchi et al. 53. They
discovered by histological tracking that labeled PEG-poly(aspartate) copolymer micelles
accumulated more rapidly (after 1 h) via the blood vessels at the tumor periphery than via
tumor blood vessels due to the EPR effect. The Hy-loaded MPEG-hexPLA micelles (Zav= 90
nm and dn< 30 nm) accumulated faster and more efficiently in tumors 3 h after injection than
Hy-PLA Nps (Zav= 200-300 nm), which needed at least 24 h to reach their highest
Chapter IV
111
accumulation 22. For the MPEG-hexPLA formulations, it was found that the disappearance of
Hy from the bloodstream coincided with Hy uptake in tissues and, more specifically, tumor
uptake. The decrease in plasma Hy concentration 6 h after injection demonstrates the
elimination of Hy and Hy-loaded MPEG-hexPLA micelles from the bloodstream. Similarly,
Burt et al. showed that MPEG-PLA copolymers were rapidly eliminated from the body
through the urine 44. The degradation of the copolymer in acidic conditions, which are present
in tumor tissues, is one plausible explanation
54;55
. The emerging lower molecular weight
components are also more rapidly filtered and eliminated by the kidneys 29;56. This suggests a
low risk of chronic accumulation of a micelle formulation in the human body if it were to be
administered in a clinical setting.
With regard to the need for an early detection of ovarian cancer and any metastases that are
present that was discussed above, it can be concluded that Hy-loaded MPEG-hexPLA
formulations lead to a rapid (3 h) accumulation of the fluorescent marker in rat tumors with an
excellent contrast to surrounding healthy tissues. When one considers the procedures that are
currently practiced in the clinical setting, it becomes apparent that laparoscopic examination using
blue light following intravenous administration of Hy micelle formulations could increase the
visualization of tumors and micrometastases in the pelvis. This would allow for a efficient, early
diagnosis of ovarian cancer. It would also lead to more accurate staging. Both of these factors
could result in improved patient prognosis due to earlier identification of the appropriate
treatment regimen for each patient. Follow-up laparoscopic examinations should be possible with
the same efficacy, since the very small size MPEG-hexPLA micelles should not induce the
accelerated blood clearance (ABC) phenomenon that is often observed with pegylated
nanoassemblies >50 nm following a second injection
57;58
. In addition to cancer diagnosis,
laparoscopy under blue light could be performed during surgery, facilitating complete removal of
diseased tissues. For both of these possible clinical applications, the fast and efficient micelle
accumulation that can be achieved with these formulations is a great advantage, since it can
reduce the time needed for a routine patient check-up or preparation prior to a necessary surgery.
The proof of concept that we achieved in the CAM and rat model showed promising results. In
view of their possible clinical applications, MPEG/dihexPLA molecular weights and ratios could
be further tailored to modulate and optimize their circulation time, loading capacity and
diagnostic properties in future studies.
112
Acknowledgments
The authors would like to thank the Swiss National Science Foundation (SNF) for its financial
support (SNF 200020-103752)
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Next to intravenous applications of the aqueous MPEG-hexPLA formulation the versatility
of the novel drug delivery systems was investigated for the topical delivery of azole
antifungals.
CHAPTER V
Chapter V
117
Novel Micelle Formulations to Increase Cutaneous
Bioavailability of Azole Antifungals
Y.G. Bachhava, K. Mondona, Y.N. Kalia, R. Gurny, and M. Möller
School of Pharmaceutical Sciences, University of Geneva, University of Lausanne, 30 quai
Ernest-Ansermet, CH-1211 Geneva 4, Switzerland.
a
both authors have equally contributed to this work
Published in: Journal of Controlled Release (2011)
Efficient topical drug administration for the treatment of superficial fungal infections would
deliver the therapeutic agent to the target compartment and reduce the risk of systemic side
effects. However, the physicochemical properties of the commonly used azole antifungals
make their formulation a considerable challenge. The objective of the present investigation
was to develop aqueous micelle solutions of clotrimazole (CLZ), econazole nitrate (ECZ) and
fluconazole (FLZ) using novel amphiphilic methoxy-poly(ethylene glycol)-hexyl substituted
polylactide (MPEG-hexPLA) block copolymers. The CLZ, ECZ and FLZ formulations were
characterized with respect to drug loading and micelle size. The optimal drug formulation was
selected for skin transport studies that were performed using full thickness porcine and human
skin. Penetration pathways and micellar distribution in the skin were visualized using
fluorescein-loaded micelles and confocal laser scanning microscopy. The hydrodynamic
diameters of the azole loaded micelles were between 70 and 165 nm and the corresponding
number weighted diameters (dn) were 30 to 40 nm. Somewhat surprisingly, the lowest
loading efficiency (< 20 %) was observed for CLZ (the most hydrophobic of the three azoles
tested); in contrast, under the same conditions, ECZ was incorporated with an efficiency of
98.3% in MPEG-dihexPLA micelles. Based on the characterization data and preliminary
transport experiments, ECZ loaded MPEG-dihexPLA micelles (concentration 1.3 mg/mL; dn
< 40 nm) were selected for further study. ECZ delivery was compared to that from Pevaryl®
cream (1% w/w ECZ), a marketed liposomal formulation for topical application.
ECZ
deposition in porcine skin following 6 h application using the MPEG-dihexPLA micelles was
ECZ/MPEG-hexPLA Micelle Formulations for Topical Application
118
>13-fold higher than that from Pevaryl® cream (22.8 ± 3.8 and 1.7 ± 0.6 µg/cm2,
respectively). A significant enhancement was also observed with human skin; the amounts of
ECZ deposited were 11.3 ± 1.6 and 1.5 ± 0.4 µg/cm2, respectively (i.e., a 7.5-fold
improvement in delivery). Confocal laser scanning microscopy images supported the
hypothesis that the higher delivery observed in porcine skin was due to a larger contribution
of the follicular penetration pathway. In conclusion, the significant increase in ECZ skin
deposition achieved using the MPEG-dihexPLA micelles demonstrates their ability to
improve cutaneous drug bioavailability; this may translate into improved clinical efficacy in
vivo. Moreover, these micelle systems may also enable targeting of the hair follicle and this
will be investigated in future studies.
Keywords: Antifungal; azole; polymeric micelle; skin deposition; substituted polylactides,
follicular delivery.
Chapter V
119
1. Introduction
The incidence of mycoses especially superficial fungal infections is increasing and according
to a recent report more than 25% of the world’s population is affected 1;2; disease progression
is more rapid and severity increased in patients with compromised immune function 3. Host
immunity can be impaired during infancy, in old age, by pregnancy, by disease, e.g. diabetes
mellitus, or through the administration of antibiotics and glucocorticoids 4. Azole antifungals
such as clotrimazole (CLZ), econazole nitrate (ECZ) and fluconazole (FLZ) are the first line
treatments for various fungal infections 5. Topical therapy is desirable since, in addition to
targeting the site of infection, it reduces the risk of systemic side effects. In general, azole
antifungals tend to be highly lipophilic (although there are exceptions (e.g., FLZ)) and they
can readily partition into the lipid-rich intracellular space in the stratum corneum; the
challenge is to develop a simple stable formulation that facilitates drug release into the skin 6.
Given the desirable properties of aqueous formulations and the lipophilic character and poor
water solubility of azoles, it was decided to investigate polymeric micelles as a drug carrier
system. Due to their stability, size and ability to incorporate significant amounts of
hydrophobic drugs in their core, these systems seem to be well-suited for use with azole
antifungals. In previous studies, micelle formulations using two novel amphiphilic methoxypoly(ethylene glycol)- hexyl-substituted poly(lactides) (MPEG-hexPLA) block copolymers,
mono- and di-hexyl-substituted (MPEG-monohexPLA and –dihexPLA, respectively)
demonstrated their ability to incorporate several poorly water soluble drugs with high loading
efficiencies
antifungals,
7-9
. The present study investigated the micelle formulations of three azole
clotrimazole,
econazole
nitrate
and
fluconazole,
possessing
different
physicochemical properties (Table 1).
The specific objectives were (i) to develop and to characterize micelle formulations for CLZ,
ECZ and FLZ using the novel excipients – MPEG-monohexPLA and -dihexPLA – and to
compare them with formulations using standard MPEG-polylactide (MPEG-PLA), (ii) to
select the best drug candidate (as determined by incorporation efficiency and micelle
properties)
and to optimize the formulation for skin deposition, (iii) to quantify drug
deposition in full thickness porcine and human skin and to compare delivery to that from a
120
commercial formulation and (iv) to visualize micellar transport pathways using fluoresceinloaded micelles and confocal laser scanning microscopy.
Table 1. Physico-chemical characteristics of the three antifungal agents
MWa
(g/mol)
log Po/w b
Aqueous solubility
(g/L)
pKa
Clotrimazole (CLZ)
344.84
5.9 10
0.030 11
5.83 12
Fluconazole (FLZ)
306.27
0.4 c
0.001 c
-
Econazole nitrate (ECZ)
444.70
5.2
d10
a
Molecular weight
b
Experimental partition coefficient between octanol and water
c
Data taken from http://www.drugbank.ca/drugs
d
Data taken from econazole
e
Determined experimentally by the shake-flask method (24h at 25°C)
0.800
e
6.65 d 10
2. Materials and Methods
2.1 Materials
Clotrimazole (CLZ), econazole nitrate (ECZ), fluconazole (FLZ), dipotassium hydrogen
phosphate, monobasic ammonium phosphate, acetone, and fluorescein acid were purchased
from Sigma Aldrich (Buchs, Switzerland). Methanol and acetonitrile (Chromasolv HPLC
grade) and nylon membrane filters (0.22 μm) were purchased from VWR (Nyon,
Switzerland). Pevaryl® cream (1% w/w ECZ ) was purchased from Janssen-Cilag (7HB5P01);
it contains PEG-6 stearate, glycol stearate and PEG-32 stearate (Téfose 63), liquid parafin, a
polyoxyethylated kernel oil (Labrafil M 1944 CS), benzoic acid (E 210), perfumes (essential
oils of rose, jasmine, iris, sandalwood, coriander, ylang-ylang, vétyver, linalol, cinnamic
alcohol, cinnamic aldehyde), butylhydroxyanisole (E 320), purified water.
Chapter V
121
O
H3C
O
45
R
O
O
R
1
2
O
H
m
O
for m=21
: MPEG2000g/mol-PLA3000g/mol
R1, R2 =CH3
2
1
R =C6H13, R =CH3 : MPEG2000g/mol-monohexPLA3000g/mol for m=14
R1, R2 =C6H13
for m=11
Scheme 1. Structure of MPEG-(hex)PLA copolymers
The block copolymers, methoxy-poly(ethylene glycol) −di-hexyl-substituted lactide (MPEGdihexPLA), −mono-hexyl-substituted lactide (MPEG-monohexPLA) and −polylactide
(MPEG-PLA) were synthesized with MPEG2000g/mol as initiator as described previously 7;9
The structures of these three block copolymers are presented in Scheme 1 and their molecular
weights and polydispersity indices are shown in Table 2.
Table 2. Characteristics of MPEG-hexPLA and MPEG-PLA copolymers
Copolymer
MPEG-PLA
MWa
(g/mol)
P.I.b
5050
1.17
MPEG-monohexPLA 5040
1.13
MPEG-dihexPLA (1) 5554
1.13
MPEG-dihexPLA (2) 4881
1.11
a
Determined by 1H NMR (Bruker, 300 MHz)
b
Polydispersity index (P.I.) determined by GPC
(1) Copolymer used only for comparison of CLZ, ECZ and FLZ micelle formulations prepared by Method 1
(2) Copolymer used for the optimization of ECZ MPEG-dihexPLA micelle formulations, skin experiments and
for the fluorescein micelle formulation
122
Please note that in the following text, the term “MPEG-hexPLA” is used to refer collectively
to the MPEG-monohexPLA and MPEG-dihexPLA polymers.
2.2 Preparation of drug loaded MPEG-hexPLA and MPEG-PLA micelles
micelles by stirring (method 1)
The micelles were prepared by a co-solvent evaporation method. Briefly, 6 mg of drug were
dissolved in 1 mL acetone and mixed with 1 mL copolymer solution (20 mg/mL) in acetone.
The organic solution was added dropwise every 5 s, using a peristaltic pump, into 4 mL of
ultra-pure water under continuous stirring. Acetone was then slowly removed by evaporation
(with stirring) in a desiccator under vacuum (2 h, 200 mbar). The final micelle concentration
was adjusted by adding ultra-pure water in order to reach a copolymer concentration of 5
mg/mL. After overnight equilibration, the solution was centrifuged at 9500 x g for 15 min to
remove non-incorporated drug.
2.3 Preparation of ECZ loaded MPEG-dihexPLA micelles by sonication
(method 2)
Econazole nitrate (ECZ) loaded MPEG-dihexPLA micelle formulations with copolymer
concentrations of 5 and of 10 mg/mL were also prepared by the co-solvent evaporation
sonication method 7. ECZ and MPEG-dihexPLA copolymer were dissolved in 2 mL acetone
and added drop wise every 5 s into 4 mL ultra-pure water under sonication. The remaining
acetone was removed by evaporation with a rotavapor at 15 mbar. The concentration of
copolymer was adjusted and the non-incorporated drug was removed by centrifugation.
2.4 Preparation of fluorescein loaded MPEG-dihexPLA micelles
Fluorescein loaded MPEG-dihexPLA micelles were prepared by method 2.
Chapter V
123
2.5 Size characterization of drug loaded MPEG-(hex)PLA micelles
Dynamic light scattering measurements (at 25°C at an angle of 90°) using a Zetasizer HS
3000 (Malvern Instruments Ltd; Malvern, UK) were made to determine the hydrodynamic
diameter (Zav), the number-weighted diameter (dn) and the percentage of micelles having the
number-weighted diameter ([%]dn). All measurements were done in triplicate.
2.6 Morphology determination of drug loaded micelles
The morphology of drug loaded micelles was determined by transmission electron
microscopy (TEM) (EM 410, Philips, 60kV) using the negative staining method. Briefly, 30
µL of the micellar solution were dropped onto an ionised carbon-coated copper grid (0.3 torr,
400 V for 20 s). The grid was then deposited for 1 s onto a 100 µL drop of uranyl acetate
solution (400 µL of a saturated uranyl acetate solution dissolved in 600 µL of distilled water)
and afterwards onto a second 100 µL drop for 30 s. Excess of the staining solution was
removed and the grid was dried at room temperature prior to the measurement.
2.7 Measurement of drug content in micelles
Drug loading was quantified by HPLC analysis. To ensure complete micelle destruction and
release of incorporated drug, a 1:10 dilution in acetonitrile was made for micelles with a
copolymer concentration of 5 mg/mL and a 1:50 dilution for those with a copolymer
concentration of 10 mg/mL. The azole content was quantified using a Lichrospher® RP-18
column (124 mm × 4 mm) thermostated at 35°C. The mobile phases comprised (i) for CLZ, a
75:25 mixture of methanol/0.05M potassium biphosphate buffer , (ii) for ECZ, a 60:40
mixture of acetonitrile/0.05 M potassium biphosphate buffer, and (iii) for FLZ, a 50:50
mixture of methanol/0.1M monobasic ammonium phosphate buffer. All analyses were
performed using a flow rate of 1.5 mL/min and a detection wavelength of 214 nm. All
samples were measured in triplicate. Sharp peaks for ECZ, CLZ and FLZ were obtained at
4.4, 5.0 and 4.6 min, respectively. The HPLC methods were linear over the concentration
range of 0.5-100 µg/mL. The limits of detection for ECZ, CLZ and FLZ were found to be
0.15, 0.43 and 0.25 µg/mL, respectively. The corresponding limits of quantification were
0.45, 1.32 and 0.75 µg/mL for ECZ, CLZ and FLZ, respectively.
124
The incorporation
efficiency and drug content were calculated using equation (1) and (2), respectively:
Incorporation efficiency (%) =
Drug content
× 100
mass of drug introduced (g)
(1)
mass of drug incorporated in micelles (mg)
mass of copolymer used (g)
(2)
(mg drug /g copolymer ) =
2.8 Skin preparation
Full thickness porcine and human skin was used to perform skin transport experiments.
Porcine ears were supplied by a local abattoir (CARRE; Rolle, Switzerland) shortly after
sacrifice. After cleaning under cold running water, the whole skin was removed carefully
from the outer region of the ear and separated from the underlying cartilage with a scalpel.
The skin samples were wrapped in ParafilmTM and maintained at −20°C and used within 2
months of harvesting. Human skin samples were obtained shortly after surgery from the
Department of Plastic, Aesthetic and Reconstructive Surgery, Geneva University Hospital
(Geneva, Switzerland), fatty tissue was removed and the skin was wrapped in Parafilm™
before storage at −20°C for a maximum period of 3 days. The donation was approved by the
Central Committee for Ethics in Research (CER: 08-150 (NAC08-051); Geneva University
Hospital).
2.9 Evaluation of ECZ deposition in the skin
Porcine skin
Full thickness porcine skin samples were equilibrated in 0.9% NaCl solution for 30 min and
then mounted on standard Franz diffusion cells (area = 2.0 ± 0.1 cm2); silicone grease was
applied to ensure a watertight seal. The receptor compartment (volume 10 mL) was filled with
phosphate buffered saline (PBS, pH 7.4). After equilibration, 1 mL of micelle formulation
(containing 1.3 mg ECZ) or 140 mg of Pevaryl® cream (1% w/w ECZ corresponding to 1.4
mg ECZ) was added to the donor compartment. The receptor compartment was stirred at 250
rpm at room temperature; 0.6 mL of receptor phase was sampled at the end of the experiment
(after 1, 3 or 6 h).
Chapter V
125
Upon completion of the permeation experiments, the diffusion cells were dismantled and the
skin surface washed in running water to remove the residual formulation. The amount of ECZ
deposited in the skin was extracted by cutting the skin samples into small pieces and soaking
in 10 mL of acetonitrile for 4 h with continuous stirring at room temperature. The samples
were filtered through 0.22 µm nylon membrane filters (VWR, Nyon, Switzerland) prior to
HPLC analysis.
Human skin
Deposition of ECZ after application of either the micelle formulation or Pevaryl® cream in
human skin was determined as before for porcine skin except that specially designed Franz
diffusion cells (area = 0.66 cm2) with a receptor volume of 4 mL were used and skin
extraction was achieved using 4 mL of acetonitrile as the extraction medium.
2.10 Confocal laser scanning microscopy
Full thickness porcine skin samples were equilibrated in 0.9 % NaCl and mounted in standard
Franz diffusion cells; 1 mL of a fluorescein loaded micelle formulation was placed in the
donor compartment and kept in contact with the epidermal surface for either 6 or 24 h after
which the diffusion cells were dismantled and skin samples were washed with 20% ethanol
solution. The samples were then placed on a glass slide with the stratum corneum side up and
covered with distilled water. A cover slip was placed on the skin sample and fluorescein
fluorescence was visualized by using a Laser Scan Microscope 510 Meta (Carl Zeiss; Jena,
Germany) at a power of 0.95W with a pinhole of 84µm and with a master gain of 541 for 6 h
experiments and of 671 for 24 h experiments. The fluorescence excitation and emission
wavelengths were set with a filter 505-530 nm, respectively using an Ar laser at an excitation
lines at 488 nm. Samples of skin in contact with a solution of fluorescein in water for 6 h and
24 h (2 mg fluorescein was partly dissolved in 6 mL water, then the solution was filtered
using a 0.22 μm filter) were also tested as controls. The confocal images were obtained with
an Achroplan 63× objective and analyzed using Zeiss LSM Image Browser software. Each
image was the average of eight repeated scans.
126
3. Results and Discussion
3.1 Characterization of CLZ, ECZ and FLZ micelle formulations
The simpler and faster Method 1 was used to prepare the first series of formulations in order
to enable a more rapid screening. The incorporation of the three azoles yielded MPEGhexPLA and MPEG-PLA micelles of homogeneous size with a number weighted diameter
(dn) between 30-40 nm for more than 97% of the micelles (Table 3). The “nanosize” of the
micelles was confirmed by TEM – an image of ECZ loaded MPEG-dihexPLA micelles is
shown in Figure 1. The high values of the hydrodynamic diameter (Zav) from 70 to 165 nm
and the relatively high polydispersity index can be attributed to the presence of a few larger
micelles, which can be seen in the TEM image. Since these larger micelles scattered more
light than their smaller counterparts, the mean diameter deduced from the overall intensity
gives a higher mean value (Zav) than dn and consequently a high polydispersity index. Thus,
the number weighted diameter dn and the corresponding percentage [%]dn were considered to
be the best parameters to describe micelle size and confirmed that these micelles are true
nanocarriers with diameters of 30-40 nm and a narrow distribution. Earlier studies have
suggested that carrier size can influence skin deposition. Smaller liposomes with sizes of 110120 nm increased cumulative permeation of cyproterone acetate through split thickness
porcine abdominal skin 13 and skin deposition of the fluorescent marker, caroboxyfluorescein
in human abdominal skin 14. However, no clear size dependence was observed for delivery of
liposomes containing cyclosporin A across porcine skin 15.
Figure. 1. TEM picture of drug loaded micelles with econazole nitrate
Chapter V
127
The MPEG-hexPLA micelles containing FLZ and ECZ showed much higher incorporation
efficiencies than those with CLZ (84-98% cf. 11-19%) (Table 3 and Fig. 2). Moreover, the
less hydrophobic MPEG-PLA standard (control), which usually incorporated less
hydrophobic drug than the MPEG-hexPLA micelles
7-9
incorporated twice as much CLZ,
although the incorporation efficiency remained relatively modest at only 36%. Possessing the
highest log P of the three drugs, CLZ might have been assumed a priori to be the best
candidate for incorporation into the hydrophobic core of the MPEG-hexPLA micelles.
However, previous results with other drugs have shown that, in addition to log P, the aqueous
solubility, the number of H-bond donor (Hd) and H-bond acceptor (Ha) groups can also affect
incorporation efficiency into MPEG-hexPLA micelles 16.
Multiple linear regression analysis of data from an investigation into the incorporation of
eighteen different drugs in MPEG-hexPLA micelles, suggested that the incorporation
efficiency was influenced in decreasing order by Hd, log P, Ha and the aqueous solubility. The
molecular weight and the surface tension of the drug had less effect. The influence of Ha and
Hd groups on the incorporation results in MPEG-hexPLA micelles may explain the
differences in incorporation found for CLZ, FLZ and ECZ. Both FLZ and ECZ contain one
Hd group; in contrast CLZ has none – the presence of Hd was found to be the most important
factor governing drug incorporation into MPEG-hexPLA micelles – thus, it is consistent that
FLZ and ECZ show manifold higher incorporation efficiencies than CLZ. Moreover, CLZ has
fewer Ha groups (two) in comparison to ECZ (five) and FLZ (eight), again consistent with its
poor incorporation efficiency. Drug/copolymer affinity was also found to be favoured by the
formation of H-bonds between the multiple H-bond sites of PCL copolymer and the single Hbond site of two cucurbitacin drugs 17. Drugs with multiple H bond donors and acceptors have
also been found to be better solubilised 18. Similarly, Liu et al. considered the overall system
when selecting the most suitable polymer for an ellipticine micelle formulation 19.
128
Table 3. Characteristics of MPEG-(hex)PLA micelles loaded with clotrimazole (CLZ), fluconazole
(FLZ) and econazole nitrate (ECZ), prepared by method 1.
Micelle size
Drug
Copolymer
Drug content
(mgdrug/gcopolymer)
Incorp.
efficiency
(%)
(nm)
MPEG-PLA
108.3±2.2
36.1
72
0.28
37
99.9
33.3±0.5
11.1
163
0.51
20
100.0
MPEG-dihexPLA
58.0±0.2
19.3
95
0.30
29
98.7
MPEG-PLA
250.2±2.7
83.4
92
0.44
20
99.2
271.7±2.7
90.7
112
0.50
18
100.0
MPEG-dihexPLA
268.3±2.7
89.4
71
0.40
27
98.2
MPEG-PLA
177.1±3.1
59.0
70
0.29
29
100.0
252.7±1.0
84.4
145
0.41
30
97.7
295.1±5.8
98.3
95
0.30
29
100.0
CLZ MPEG-monohexPLA
FLZ MPEG-monohexPLA
ECZ MPEG-monohexPLA
MPEG-dihexPLA
a
Hydrodynamic diameter
b
Polydispersity index
c
Number weighted diameter
d
Percentage of micelles having the diameter dn
Zava
b
P.I.
dnc
(nm)
[%]dnd
FLZ was successfully encapsulated into the different micelles, achieving drug contents
between 250 and 268 mg/g and incorporation efficiencies of 83 to 91%. ECZ showed similar
high drug loadings and an even higher incorporation efficiency (98%) in MPEG-dihexPLA
micelles; this was an advantage since the hydrophobic dihexPLA core increases the storage
stability of the micelle formulations
9;20
. Based on the loading efficiency, the ECZ MPEG-
dihexPLA micelle formulation was selected for further optimization prior to skin transport
studies.
Chapter V
129
350
Drug loading in micelles
(mg/g copolymer)
300
250
200
150
100
50
0
Clotrimazole
Fluconazole
Econazole nitrate
Figure 2. Drug loading in MPEG-PLA (), MPEG-monohexPLA () and MPEG-dihexPLA (
)
micelles for 3 antifungal agents: clotrimazole, fluconazole and econazole nitrate (Mean ±SD, n=3)
3.2 Optimizing the preparation of ECZ MPEG-dihexPLA micelle formulations
Evaluation of the preparation method
As mentioned earlier, micelle size may influence skin deposition. Previous experiments
showed that the evaporation step in the co-solvent evaporation method influenced both the
size and size distribution of the resulting micelles [Mondon et al., unpublished data]. Here,
two micelle preparation methods were tested which differed in the type of mixing during the
addition and evaporation steps. Method 1 involved a simple stirring and a slow evaporation (2
h at 200 mbar), whereas Method 2 involved sonication during the addition step and a faster
evaporation (10 min) at 15 rpm. MPEG-dihexPLA micelles with smaller Zav, dn, and a
narrower polydispersity index were obtained by Method 2 (Table 4). In addition, ECZ
incorporation was also slightly higher with Method 2, 79.0% (cf. 76.5% for Method 1).
Therefore, based on these results, the sonication method was chosen to prepare the ECZ
MPEG-dihexPLA micelle formulations for testing in the skin deposition experiments.
130
Table 4. Characteristics of MPEG-dihexPLA micelle formulations prepared by method 1 (stirring,
slow evaporation) and method 2 (sonication, faster evaporation).
Micelle size
Batch
Micelle prep.
method
1
Method 1
2
Method 2
a
Incorp.
Zav
dn
[ECZ]mica
(g/L)
(nm)
76.5
1.12±0.02
133
0.22
50
99.9
79.0
1.20±0.03
91
0.16
42
99.8
Efficiency
(%)
P.I.
(nm)
[%]dn
ECZ concentration in micelles
Zav, dn and [%]dn defined in Table 3
Improving ECZ drug loading
Micelle preparation using Method 2 was also used to optimize ECZ loading and to maximize
the drug/copolymer ratio (mgdrug/gcopolymer). Different target drug loadings (represented by
different amounts of drug dissolved in the initial acetone phase) were tested using two
copolymer concentrations: 5 and 10 mg/mL. In previous studies, increasing MPEG-dihexPLA
concentration in the micelle formulation yielded formulations with higher drug concentration
without changing micelle size as confirmed by DLS and TEM analysis 7. However, in these
studies higher ECZ drug content was observed at the lower copolymer concentration (~285
and ~182 mgECZ/gcopolymer at 5 and 10 mgcopolymer/mL, respectively) (Table 5). Incorporation
efficiencies were >90% with respect to the target drug loadings of 300 and 200 mg/g for
micelles with 5 and 10 mgcopolymer/mL, respectively. As seen for other drugs, incorporation
efficiency decreased at higher target drug loadings 8. A higher drug / copolymer ratio of 28%
(w/w) with 95% incorporation efficiency and higher shelf-life stability were achieved for
micelles with 5 mgcopolymer /mL; thus, this copolymer concentration was chosen for the next
experiments. Since 74 nm ≤ Zav ≤ 95 nm and 22 nm ≤ dn ≤ 38 nm for the different micelles, it
is clear that the sonication method was capable of preparing truly nanosized drug carrier
systems with high ECZ loading.
Chapter V
131
Table 5. Characteristics of MPEG-dihexPLA micelle formulations of two copolymer concentrations
(5 and 10 mg/mL) for different drug contents.
Incorp.
Micelle size
Target ECZ
content
(mgECZ/gcopolymer)
Actual ECZ
content
(mgECZ/gcopolymer)
Efficiency
(%)
200a
186.5±2.2
92.5
0.93±0.01
86
0.26
38
98.3
300a
285.8±6.4
94.9
1.42±0.03
80
0.27
33
99.8
400a
251.4±5.9
65.6
1.27± 0.03
74
0.27
22
90.5
500a
221.1±6.4
46.1
1.15 ±0.03
75
0.36
27
99.9
200b
181.8±8.1
92.0
1.84±0.08
95
0.31
24
98.8
300b
172.3±12.4
56.6
1.69±0.12
81
0.34
25
100.0
[ECZ]micelles
(g/L)
Zav
(nm)
PI
dn
(nm)
[%]dn
[ECZ]micelles defined in Table 4
Zav, dn and [%]dn defined in Table 3
a
Micelles of 5 mgcopolymer/mL
b
Micelles of 10 mgcopolymer/mL
3.3 ECZ skin deposition studies using MPEG-dihexPLA micelles and Pevaryl®
cream
Porcine skin
An ECZ MPEG-dihexPLA micelle formulation was prepared by sonication with a copolymer
concentration of 5 mg/mL and an ECZ concentration of 1.3 mg/mL with Zav < 90 nm and dn <
40 nm. Skin deposition of ECZ from the MPEG-dihexPLA micelle formulation was
investigated as a function of application time and compared to that from Pevaryl®, a
commercially available liposomal formulation, in which the liposomes are reported to have
diameters (Zav) of 160-200 nm 21. A similar dose of ECZ was applied using each formulation
(1.3 mg MPEG-dihexPLA micelles and 1.4 mg for Pevaryl®, respectively). No ECZ was
detected in the receiver compartment from any of the tested formulations. Skin deposition of
ECZ from the MPEG-dihexPLA micelles was significantly higher than that observed with
Pevaryl® after formulation application for 1, 3 and 6 h (4.7 ± 1.7, 14.7 ± 1.7 and 22.8 ± 3.8
µg/cm2, cf. 1.0 ± 0.3, 1.4 ± 0.2 and 1.7 ± 0.6 µg/cm2, respectively; Fig. 3) (Student’s t-test,
α=0.05). This corresponded to 5-, 10- and 13-fold improvements in ECZ deposition using the
132
MPEG-dihexPLA micelle formulation at the 1, 3 and 6 h time-points, respectively. Indeed,
despite the presence of permeation enhancers such as labrafil M 1944 CS, linalol, cinnamic
alcohol and cinnamic aldehyde in this formulation, Pevaryl® application resulted in lower skin
deposition of ECZ than with the much simpler MPEG-dihexPLA micelles (which contained
only water, copolymer and drug). The increase in ECZ skin deposition may be attributed to
the difference of the carrier size; the MPEG-dihexPLA micelles are approximately half the
size of the liposomes and may have a larger contact area with the skin surface. In addition,
their viscosity may facilitate film formation resulting in a depot at the skin surface and so
increase contact time and drug deposition. Another factor that may affect delivery is the
relative thermodynamic activity of ECZ in the two formulations since this will determine
partitioning from the formulation into the stratum corneum. It has also been reported that the
physical state of the surfactant used in the preparation of niosomes or liposomes had an
influence on the deposition of finasteride into sebaceous gland region
22
. The liquid state
vesicles showed higher permeation than gel state vesicles. This was attributed to a better
penetration of the liquid state surfactant molecules into the stratum corneum compared than
those forming a gel-state. By analogy, the physical state of the polymeric micelles is rather
fluid due to the low Tg of the dihexPLA core 23 and this may also facilitate drug deposition.
Skin deposition of ECZ (µg/cm2)
30
*
25
20
*
15
10
*
5
0
1
3
Time (h)
6
Figure 3. Skin deposition of ECZ from the MPEG-dihexPLA micelle formulation () and the
Pevaryl® (■) after 1, 3 and 6h application on porcine skin (Mean ± SD, n=5-6). * Significantly
different from Pevaryl® (Student’s t-test, α= 0.05).
Chapter V
133
Human skin
ECZ delivery from the MPEG-dihexPLA micelle formulation and Pevaryl® was also
investigated using human skin (Fig. 4); for these experiments skin deposition of ECZ was
determined only at the 6 h time-point. ECZ deposition from the micelle formulation was again
almost an order of magnitude superior to that from Pevaryl® (11.3 ± 1.6 and 1.5 ± 0.4 µg/cm2,
respectively). As with the porcine skin experiments, no ECZ permeation was observed across
the human skin from either formulation.
ECZ deposition following application of Pevaryl® was equivalent for porcine and human skin.
In contrast, a 2-fold difference in ECZ deposition from the MPEG-dihexPLA micelle
formulation was observed between human and porcine skin. The human skin used for these
experiments was obtained following breast reduction surgery and, in contrast to porcine ear
skin, was devoid of hair follicles. Thus, the higher ECZ deposition observed using the MPEGdihexPLA micelle formulation was tentatively attributed to penetration of the smaller micelles
via the hair follicles.
30
Skin deposition of ECZ (µg/cm2)
*
25
20
15
*
10
5
0
Pevaryl®
Micelle formulation
Figure 4. Comparing deposition of ECZ using the MPEG-dihexPLA micelle formulation Pevaryl®
after application for 6h on porcine skin () and human skin (■) (Mean ± SD, n=5-6). * Significantly
different from Pevaryl® (Student’s t-test, α= 0.05).
134
Visualizing micellar penetration pathways using confocal laser scanning microscopy
In order to determine the role of the follicular pathway, fluorescein loaded MPEG-dihexPLA
micelles (dn~28 nm) were prepared using the same method as that for the ECZ MPEGdihexPLA micelle formulation and applied to porcine skin for 6 and 24 h. A solution of
fluorescein in water was applied on the skin for 6 h and 24 h, as control. The skin samples
were subsequently analyzed using confocal laser scanning microscopy in order to visualize
micelle deposition. In contrast to the untreated skin, (Fig. 5a) and to the skin treated with
fluorescein solution (Fig. 5b and 5c), the images in the XY-plane following treatment with
fluorescein loaded micelles for 6 and 24 h show considerable fluorescence (Fig. 5d and 5e).
Figure 5. Confocal laser scanning microscopy images of porcine skin in the XY-plane show that in
contrast to (a) the untreated porcine skin and skins treated for (b) 6 h and (c) 24 h with a fluorescein
solution, considerable fluorescence was observed after contact with fluorescein loaded micelles for (d)
6 h and (e) 24 h.
Chapter V
135
Similarly, no fluorescence is observed in the XZ-plane in skin treated only with a fluorescein
solution (Fig. 6a and 6b). However, the images following treatment with fluorescein loaded
micelles show appreciable fluorescence at depths of up to 40 µm (Fig. 6c and 6d). The image
recorded in the XZ-plane at 24 h (Fig. 6d) shows greater fluorescence intensity and also
shows the presence of localized fluorescence in the follicles penetrating deeper into the skin.
The follicular pathway has been reported as the primary penetration route for nanoparticles 24;
it is thought to be a size dependent process, where smaller nanoparticles show highest
penetration
25;26
. Moreover, nanoparticles of similar sizes to the MPEG-dihexPLA micelles
were shown to cross the skin barrier via hair follicles 27. Thus, the MPEG-dihexPLA micelles
may target delivery to the follicles and their accumulation, depot formation, and subsequent
release of a drug may be used to provide more effective, sustained treatments.
Figure 6. Confocal laser scanning microscopy images of porcine skin from the XZ-plane show that in
contrast to the skin treated with a fluorescein solution for (a) 6 h and (b) 24 h, fluorescence appeared
to accumulate in epidermis via the hair follicles (white squares) when skin was treated with
fluorescein loaded micelles for (c) 6 h and (d) 24 h .
136
4. Conclusion
The results demonstrate that hydrophobic MPEG-dihexPLA copolymers encapsulated
econazole nitrate and fluconazole with high incorporation efficiencies. The preparation
method was shown to impact on micelle properties; faster evaporation and addition under
sonication yielded smaller micelles. Econazole nitrate delivery using MPEG-dihexPLA
micelle formulations resulted in significantly higher drug deposition in both porcine and
human skin as compared to Pevaryl®, a marketed liposomal formulation. Confocal laser
scanning microscopy studies using similarly-sized fluorescein loaded micelles provided
support for the potential role of the follicular pathway and suggested that the MPEGdihexPLA micelles may facilitate targeted follicular delivery
References
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(13) Valenta, C.; Janisch, M., Permeation of cyproterone acetate through pig skin from
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(14) Verma, D. D.; Verma, S.; Blume, G.; Fahr, A., Particle size of liposomes influences
dermal delivery of substances into skin, Int.J.Pharm. 2003, 258, 141-151.
(15) du Plessis, J.; Ramachandran, C.; Weiner, N.; Muller, D. G., The influence of particle
size of liposomes on the deposition of drug into skin, Int.J.Pharm. 1994, 103, 277-282.
(16) Mondon, K., Novel Micellar Systems for the Formulation of Poorly Water Soluble
Drugs: Biocompatibility Aspects and Pharmaceutical Applications, PhD Thesis 2010,
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Increasing PCL Content for Two Hydrophobic Cucurbitacin Drugs, Biomacromolecules
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(18) Patel, S. K.; Lavasanifar, A.; Choi, P., Molecular dynamics study of the encapsulation
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Chapter VI
139
MPEG-hexPLA Micelle Formulations for Oral Delivery
of Poorly Water Soluble Drugs:
Investigations with Cinnarizine as a Model Drug
K. Mondon, R. Gurny, and M. Möller.
Ernest Ansermet, CH-1211 Geneva 4, Switzerland.
1. Introduction
Oral drug delivery remains the most patient-friendly way of medication. The drug discovery
with the present use of high throuput screening methods gives leads of many novel potential
drug candidates, which tend to have a high lipophilicity and low water solubility 1. However,
the Lipinski’s “rule of 5” 2, which suggests that an orally active drug should not have more
than 5 hydrogen bond donors, not more than 10 hydrogen bond acceptors, a molecular weight
below 500 daltons, and an octanol-water partition coefficient log P of less than 5, is often not
matched. Next to the poor water solubility, the poor permeability of the drug is another major
cause of the low bioavailability of drugs in the gastrointestinal (GI) tract 3. Different methods
to improve the biopharmaceutical properties of poorly water soluble drugs have thus been
studied. Amongst other drug delivery systems, polymeric micelles have been investigated for
oral applications
4
. With their specific core-shell structure micelles can protect the
hydrophobic drug by incorporating it within its core, while the hydrophilic shell facilitates
aqueous solubility to the formulation. Regarding the range of pH encountered in the GI tract
(pH of 1.2 in the empty stomach, 5-7 in the small intestine and 6-7.5 in the colon), the
biocompatible micelle carriers need to be stable enough to keep the drug in their core at the
low pH during the passage of the stomach before reaching or releasing the drug if the
intestinal epithelium is targeted.
MPEG-hexPLA Micelles for Oral Delivery
140
In this work the possible oral application of MPEG-hexPLA micelle formulations was
investigated using the antihistaminic class II drug, cinnarizine (CIN). Cinnarizine is a
lipophilic piperazine derivative (log P ~5.3 5) with a low water solubility of 2 mg/L at 37°C 6.
CIN was chosen for these studies because it is reported as an oral drug model of choice7;8.
The stability of MPEG-(hex)PLA micelles was assessed under gastric pH. Therefore the
intactness of the micelle structure was studied via the fluorescence of incorporated Nile Red.
A Nile Red fluorescence intensity can only be detected if this probe remains in the
hydrophobic environment of the micelle core 9. The stability of CIN loaded MPEG-hexPLA
micelles regarding the drug concentration, and the molecular weight of the block copolymer,
was further studied towards possible oral formulations.
2.
Materials and methods
2.1. Materials
Cinnarizine (CIN), acetone, acetonitrile (HPLC grade), hydrochloric acid (HCl) were supplied
by Sigma-Aldrich (Buch, Switzerland). THF and ammonium dihydrogen phosphate were
purchased by SDS (Toulouse, France) and by Fluka (Buch, Switzerland), respectively.
2.2.
Preparation of MPEG-(hex)PLA block copolymers
MPEG-hexPLA block copolymers were prepared by ring opening polymerization of hexyl
substituted lactide as described previously by Mondon et al. 10. Methoxypoly(ethylene glycol)
(MPEG2000) of 2000 g/mol was used as initiator. Copolymers with the molecular weights of
the (hex)PLA block of 4000g/mol (MPEG2000-hexPLA4000) and of 8000g/mol (MPEG2000hexPLA8000) were prepared. The overall molecular weights of the block copolymer were thus
of 6000g/mol and 10000g/mol, respectively.
MPEG-hexPLA block copolymers.
Chapter VI
2.3.
141
Preparation of CIN-loaded MPEG-(hex)PLA polymeric micelles
CIN loaded micelles were prepared by dissolving 1 mL copolymer solution (20 mg/mL in
acetone) with 100 µL of a CIN solution (20 mg/mL) and 900 µL acetone. The resulting
acetone solution was added drop wise every 5s into 4 mL ultra pure water under sonication.
Acetone was removed by evaporation and water was added afterwards to get a final micelle
solution concentration of 5 mgcopolymer/mL. After overnight equilibration, CIN loaded micelles
were centrifuged at 9500 x g for 15 min to remove non-incorporated drug.
2.4.
Preparation of unloaded polymeric micelles
Unloaded micelles were prepared following the method described above, whereby pure
acetone was used instead of the drug/acetone solution.
2.5.
Determination of CIN loading by HPLC
The determination of the CIN loading in micelles was performed after centrifugation of the
micelle solution. One hundred microliter of the supernatant was taken and dissolved with 900
µL of acetonitrile to break the micelles and to release the drug. CIN was then quantified by
HPLC using a Nucleosil 100-5 C18 column (250 mm × 4 mm). The mobile phase was a
mixture of acetonitrile/ 20 mM ammonium dihydrogen phosphate buffer (50/50) eluting at a
flow rate of 1.0 mL/min. The CIN signal was detected by UV at λ= 250 nm and appeared at
around 8 min. Calibration curves with a regression coefficient superior to 0.99 were obtained.
2.6.
Gastric-pH-tests of CIN loaded MPEG-hexPLA micelles
The gastric-pH-tests were performed by adding CIN loaded polymeric micelles under slow
stirring to 10 mL of 0.1N HCl (pH=1.2) at 37°C. After 30 min and 1h stirring, 500 µL of the
acidic solution were removed and centrifuged at 9500 x g for 15 min. The supernatant was
then diluted in acetonitrile in a 1:1 ratio and analysed by HPLC for the drug content.
2.7.
142
Stability studies of Nile Red loaded MPEG-(hex)PLA micelles in 0.1N
HCl
For the stability studies, Nile Red loaded MPEG-(hex)PLA4 micelles were prepared by
incorporating 10µL Nile Red solution (6 10-4 M) into 300 µL unloaded MPEG2000(hex)PLA4000 micelles (ccopolymer= 1.66 mg/mL) under stirring for 3h. The resulting Nile Red
loaded micelles were diluted in 1.5 mL 0.1N HCl at 37°C. After complete dilution the Nile
Red fluorescence intensity was studied for 1h. The Nile Red fluorescence was detected at an
excitation wavelength of 537 nm and an emission wavelength of 600 nm using a Fluoromax
spectrofluorometer (Spex, Stanmore, UK).
The Nile Red fluorescence and its intensity change was visualized by diluting 1 mL of Nile
Red loaded MPEG2000-(hex)PLA4000 micelles into 5 mL 0.1N HCl under slow stirring at 37°C
and taking pictures just after dilution, after 30 min and after 1h under acidic pH. A control
sample with 0.1N HCl containing the same quantity of Nile Red was prepared and analysed in
parallel.
2.8.
Degradation tests of the block copolymers at gastric pH
The copolymer degradation was assessed by determining the copolymer molecular weight and
the polydispersity index by GPC (Waters, Milford, USA) before and after gastric-pH-tests (1h
at 0.1N HCl). A calibration curve with a regression coefficient superior to 0.99 was obtained
using 6 polystyrene standards ranging from 3460 to 277000g/mol (PSS, Mainz, Germany).
The copolymer characteristics after the gastric-pH-tests were determined after evaporation of
the remaining acidic solution containing CIN loaded micelles and its dilution into 1.5 mL
THF.
3.
Results and Discussion
The effective absorption of an orally administered drug depends on the dissolution within the
gastrointestinal (GI) fluids and on the permeation through the intestine
11
. Drugs with low
water solubility and/or low permeability are unfavoured for oral administration. Indeed, poor
drug water solubility is often the cause of an early drug precipitation into the GI tract
concomitant with unreached therapeutic doses. The incorporation of such drugs into MPEG-
Chapter VI
143
hexPLA polymeric micelles was shown to be an efficient strategy to increase their aqueous
solubility. Promising results for the solubilisation in such MPEG-hexPLA micelles for the
intravenous application of cyclosporin A 10 and hypericin (Mondon et al. see Chapter IV), or
for the topical applications of econazole nitrate (Mondon et al. see Chapter V) were
demonstrated. In the current study, the possible use of MPEG-hexPLA micelles in oral
delivery was evaluated incorporing the model drug cinnarizine (CIN).
3.1.
CIN loaded MPEG-(hex)PLA micelle characteristics
Two series of MPEG-(hex)PLA micelles with block copolymers of 2 different molecular
weights (Mw) were investigated. The used block copolymers, MPEG2000-(hex)PLA4000 of
6000g/mol and MPEG2000-(hex)PLA8000 of 10000 g/mol, were synthesized in a controlled
manner by ring opening polymerization as described previously
10;12
. All block copolymers
were obtained with a narrow polydispersity index inferior to 1.2. The resulting CIN loaded
micelles obtained a relative narrow size distribution with a polydispersity index below 0.3
(Table 1).
Table 1. Characteristics of CIN loaded MPEG-(hex)PLA micelle formulations
Micelle size
Copolymer
MPEG2000–PLA4000
MPEG2000-monohexPLA4000
MPEG2000-dihexPLA4000
MPEG2000-PLA8000
MPEG2000-monohexPLA8000
MPEG2000-dihexPLA8000
Drug content [drug]micelles
[mgCIN/gcopolymer]
[mg/L]
87.0
43.09
89.5
44.89
90.6
45.32
57.53
28.69
104.50
52.28
109.00
55.67
Zav
P.I.
30
51
61
69
74
82
0.07
0.22
0.30
0.035
0.16
0.28
dn [%]dn
25
34
32
54
39
44
100.0
100.0
99.5
100.0
100.0
99.9
With [drug]micelles, the concentration of the drug in micelles; Zav, the hydrodynamic diameter; P.I., the
polydispersity index; dn the number weighed diameter, and. [%]dn the percentage of micelles with a given dn
The micelle size described by the hydrodynamic diameter Zav and the number-weight
diameter dn was slightly higher for copolymers with higher Mw. This has also been observed
with MPEG-poly(caprolactone) polymeric micelles
13
. By increasing the Mw of the core
forming block, the hydrophilic/hydrophobic ratio of the block copolymer decreases and thus
favors the structure of micelles towards bigger structures like i.e. wormlike micelles.
144
However, for the here tested formulation of the high Mw micelles a small amount of larger
spherical micelles in the formulation are assumed for the slight increase in size. Only 1% of
micelles have a number weighed diameter superior to 40-55 nm for micelles of MPEG2000(hex)PLA8000, and to 25-35 nm for those with MPEG2000-(hex)PLA4000. In general, all the
tested MPEG-(hex)PLA micelles had a hydrodynamic diameter below or around 80 nm, the
range of the size generally observed for polymeric micelles
14
, and are thus real nanosized
drug carriers.
CIN was successfully incorporated into MPEG-(hex)PLA micelles. For the 3 tested
copolymers the maximum drug loading was achieved with efficiencies superior to 80% for a
targeted drug loading of 100 mgCIN/gcopolymer. A superior incorporation of CIN in the more
hydrophobic hexyl-substituted PLA micelle core compared to the standard PLA was revealed.
For example, with copolymers of 6000g/mol MPEG-dihexPLA micelles a CIN concentration
of 45.3 mg/L compared to 44.9 mg/L for the less hydrophobic MPEG-monohexPLA micelles
and 43.1 mg/L for the even less hydrophobic standard MPEG-PLA micelles was obtained.
When using MPEG-hexPLA copolymers with the higher Mw of 10000g/mol, corresponding to
the increased hexPLA block, the incorporation of CIN was increased by more than 15%. A
CIN aqueous concentration of 55.6 mg/L in MPEG2000-dihexPLA8000 micelles in comparison
to 45.3 mg/L in MPEG2000-dihexPLA4000 micelles (Table 1) was reached. This shows that an
optimization of the polymer molecular weight and hydrophilic/hydrophobic block ratio can
significantly improve the incorporation results.
3.2.
Stability of MPEG-(hex)PLA micelle formulations under gastric pH
As mentioned before, an oral administrated formulation should be stable enough in gastric pH
to pass to the intestine. This is even more true for the oral formulation of CIN which has its
aqueous solubility decreasing with the pH increase of the GI tract 6. In consequence, the
MPEG-hexPLA micelle formulations are required to be stable at least 30 min at gastric pH.
This stability means that the intactness of the micelle core-shell structure and the CIN drug
content should be preserved during this time. MPEG-(hex)PLA micelle formulations were
therefore assessed for their stability under gastric pH (pH=1.2) for 1h. The fluorescent Nile
Red probe was incorporated into the MPEG2000-(hex)PLA4000 micelles to evidence the
Chapter VI
145
intactness of the micelle core-shell structure. A Nile Red fluorescence intensity can only be
detected if this probe remains in the hydrophobic environment of the micelle core
15
. The
obtained constant fluorescence intensity confirms that the core-shell structure was almost
unaffected for 1h at pH 1.2. In addition this result also proves that the hydrophobic dye does
not diffuse out of the micelles under these conditions (Figure 1).
Nile Red Fluorescence Intensity
1.2E+07
1.0E+07
8.0E+06
6.0E+06
4.0E+06
2.0E+06
0.0E+00
0
0.5
Time in HCl [h]
1
Figure 1. Nile Red fluorescence in function of time after dilution of Nile Red loaded MPEG2000dihexPLA4000 (), MPEG2000-monohexPLA4000 () and MPEG2000-PLA4000 (ο) micelle formulations in
HCl 0.1N (pH=1.2).
The stability of the formulations within this time in gastric pH was also visualized by photos,
which show the stable violet colouring of the Nile Red micelle formulations (Figure 2).
1
(a)
2
3
4
1
(b)
2
3
4
1
2
3
4
(c)
Figure 2. Fluorescence observations of Nile Red loaded MPEG2000-dihexPLA4000 (1), MPEG2000monohexPLA4000 (2), and MPEG2000-PLA4000 (3) micelle formulation diluted in 0.1N HCl at t =0 min
(a), 30 min (b) and 1h (c). The same Nile Red quantity was added in 0.1N HCl as control (4).
146
The stability of the CIN loaded MPEG-hexPLA micelle formulations under gastric pH was
investigated by following the CIN loading over time after dilution of CIN loaded MPEG(hex)PLA micelles in 0.1N HCl solution (pH=1.2). After 30 min the drug content was the
same as the initial loading for the two tested series of micelle formulations. However, a drug
loss of maximum 12% was observed after 1h (Figure 3). In contrast no drug loss was
observed for Nile Red. This can be attributed to the difference of the CIN and Nile Red,
respectively water solubilities. CIN with a higher water solubility at gastric pH than the Nile
Red might easier diffuse out of the micelles into the aqueous environment. Moreover the
degradation study of the MPEG2000-(hex)PLA4000 block copolymer revealed that the (hex)PLA
micelle core decreased by 2-3 monomer units after 1h in gastric pH (Figure 4), possibly
slightly decreasing the micelle stability and increasing drug diffusion. Nevertheless, the block
copolymer degradation did not affect the micelle core-shell structure as such, as evidenced by
the observed constant Nile Red fluorescence intensity (Figure 1).
As the result MPEG-hexPLA micelles could be considered for the oral delivery of CIN as
only a minimal drug loss and no change in the core-shell micelle structure were observed at
gastric pH after 1h.
a) MPEG2000-(hex)PLA4000
b) MPEG2000-(hex)PLA8000
100
100
%CIN loading
in micelles
120
% CIN loading
in micelles
120
80
60
40
20
80
60
40
20
0
0
0
0.5
Time in 0.1N HCl [h]
1
0
0.5
Time in 0.1N HCl [h]
1
Figure 3. Percentage of drug loading in (a) MPEG2000-(hex)PLA4000 and (b) MPEG2000-
(hex)PLA8000 micelle formulations in function of the time at gastric pH for MPEG-dihexPLA (),
MPEG-monohexPLA () and MPEG-PLA (○) micelles.
Chapter VI
P.I. : 1.26 ; 1.31
1.17 ; 1.21
147
1.15 ; 1.16
Figure 4. Molecular weight and polydispersity index (P.I.) of MPEG2000-(hex)PLA4000 before (white
bars) and after 1h under gastric pH (0.1N HCl) (black bars)
4.
Conclusion
It could be demonstrated that the novel MPEG-hexPLA micelles successfully incorporate the
poorly water soluble cinnarizine. The stability study of these micelle formulations at gastric
pH (pH=1.2) revealed that MPEG-hexPLA micelles showed no drug loss at 30 min and only
12% loss after 1h, which could be due to the slight degradation by 2-3 (hex)PLA monomer
units of the block copolymer and a micelle destabilization, and the relative higher CIN water
solubility at gastric pH. In conclusion, these initial in vitro investigations prove the stability of
the MPEG-hexPLA micelle formulations for bypassing the stomach into the intestine for a
possible oral application.
148
References
(1) Cheng, W. P.; Thompson,
No.2010/0,029,544.
C.,
Composition,
U.S.Patent
2010, Application
(2) Lipinski, C. A., Lead- and drug-like compounds: the rule-of-five revolution, Drug
Discovery Today: Technologies 2004, 1, 337-341.
(3) Helen Chan, O.; Stewart, B. H., Physicochemical and drug-delivery considerations for
oral drug bioavailability, Drug Discovery Today 1996, 1, 461-473.
(4) Gaucher, G.; Satturwar, P.; Jones, M. C.; Furtos, A.; Leroux, J. C., Polymeric micelles
for oral drug delivery, Eur.J.Pharm.Biopharm. 2010, 76, 147-158
(5) Data taken from http://www.drugbank.ca/drugs. 2010.
(6) Stella, V. J., Why Compounds are Insoluble and What We Can Do About It!, AAPS 45th
Annual Pharmaceutical Technologies Arden Conference: Formulation Strategies for
Poorly Soluble Drugs 2010, West Point, NY, USA.
(7) Ogata, H.; Aoyagi, N.; Kaniwa, N.; Ejima, A.; Kitaura, T.; Ohki, T.; Kitamura, K.,
Evaluation of beagle dogs as an animal model for bioavailability testing of cinnarizine
capsules, Int.J.Pharm. 1986, 29, 121-126.
(8) Boyd, B. J.; Khoo, S. M.; Whittaker, D. V.; Davey, G.; Porter, C. J. H., A lipid-based
liquid crystalline matrix that provides sustained release and enhanced oral
bioavailability for a model poorly water soluble drug in rats, Int.J.Pharm. 2007, 340, 5260.
(9) Davis, M. M.; Helzer, H. B., Titrimetric and Equilibrium Studies Using Indicators
Related to Nile Blue A, Anal. Chem. 1966, 38, 451-461.
(11) Lennernäs, H., Human intestinal permeability, J.Pharm.Sci. 1998, 87, 403-410.
(13) Du, Z. X.; Xu, J. T.; Fan, Z. Q., Micellar Morphologies of Poly(e-caprolactone)-bpoly(ethylene oxide) Block Copolymers in Water with a Crystalline Core,
(14) Kwon, G. S., Polymeric micelles for delivery of poorly water-soluble compounds.,
Crit.Rev.Ther.Drug Carrier Syst. 2003, 20, 357-403.
(15) Davis, M. M.; Helzer, H. B., Titrimetric and Equilibrium Studies Using Indicators
Related to Nile Blue A, Anal. Chem. 1966, 38, 451-461.
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Conclusions and Perspectives
149
Amongst the large number of potent drugs for pharmaceutical applications, 95% are lipophilic
and poorly water soluble. However, these special properties render the aqueous formulation of
those drugs very difficult. Early drug precipitation is often encountered due to poor stability
and the therapeutic dose is rarely reached. Therefore, there is an unmet need to find stable and
sufficient drug loading formulations for such drugs.
One strategy to overcome this formulation issue is the incorporation of poorly water soluble
drugs into colloidal carriers such as liposomes, nanoparticles, and micelles. This thesis work
focused on novel polymeric micelles as promising delivery systems for poorly water soluble
drugs.
The presented novel polymeric micelles are composed of MPEG-hexPLA copolymers. The
high number of hexyl groups in the core forming block polymer compared to standard
MPEG-PLA micelles, was expected to increase the hydrophobicity of the micelle core, thus
the incorporation of hydrophobic drugs.
In a first step, the incorporation of the hydrophobic model drug, THPP, was presented as a
proof of concept. A higher incorporation was achieved with the most hydrophobic MPEGdihexPLA micelles followed by MPEG-monohexPLA and MPEG-PLA micelles. This trend
was generally observed for drugs with a low water solubility (< 0.055 mM). However, the
drug hydrophobicity and the drug water solubility were not the only parameters controlling
the incorporation efficiency into MPEG-hexPLA micelles. Structural parameters such as the
presence of H-bond donor or -acceptor groups in the drug structure showed an influence, too.
Interestingly, poorly water soluble drugs which do not fit the Lipinski’s “rule of 5”, indicating
poor bioavailability and limited applications in clinics, showed to have suitable physicochemical parameters for an efficient incorporation into MPEG-hexPLA micelles. This
envisions the use of these novel micellar carriers as an alternative formulation approach, and
was demonstrated for some selected poorly water soluble drugs, e.g. cyclosporin A, THPP,
hypericin, and econazole nitrate, for very different pharmaceutical applications.
150
As a novel material for a possible clinical use, MPEG-hexPLA micelles were investigated for
their biocompatibility. No toxicity in vitro and in vivo or hemolytic activity on human blood
was found for both, unimers (below CMC) and polymeric micelles, up to concentrations of 20
mg/mL. In consequence, MPEG-hexPLA micelle formulations could be envisioned as
suitable for pharmaceutical applications.
One considerable pharmaceutical application is the possible use in intravenous
administrations. The micelle formulation of cyclosporin A did not show any hemolytic
activity, and efficiently achieved therapeutic doses with 4 times less excipient than the
surfactant in the marketed formulation. Obviously, the in vivo efficiency of such a delivery
system should be evaluated in animal models to confirm its potential for pharmaceutical
applications.
The use of MPEG-hexPLA micelle formulations for ovarian cancer diagnostics was studied.
Intravenous administration to ovarian tumor bearing rats showed a high and fast accumulation
of the incorporated poorly water soluble fluorescent hypericin, in tumour nodules by passive
targeting. For future applications it could be very interesting to combine the cancer diagnosis
with a direct treatment by making use of the photosensitizer properties of hypericin. To
improve the carrier system an active tumor targeting by conjugating MPEG-hexPLA
copolymers with ligands such as folic acid could be envisioned.
A promising topical application of MPEG-hexPLA micelles for antifungal treatment was
evidenced in this thesis. The superior delivery of ECZ into human and porcine skin with
micelles to the commercially available formulation was demonstrated. Intriguingly, the
deposition into the skin may be attributed to a follicular penetration pathway. The presented
findings may help to formulate and deliver various potent hydrophobic drugs into the skin.
Furthermore, the oral delivery of drug loaded MPEG-hexPLA micelles could be a possible
pharmaceutical application. After 1h at gastric pH, MPEG-hexPLA micelles conserved 90%
of the drug content of the model class II drug, cinnarizine, and maintained an intact core-shell
structure with only a slight degradation of the copolymer. This could facilitate a drug
transport through the stomach into the intestine.
151
In summary, these novel biocompatible MPEG-hexPLA micelles are promising for
pharmaceutical formulations for intravenous, topical and oral applications of poorly water
soluble drugs. Still, drug incorporation into MPEG-hexPLA micelles and biologic activity
could be improved by adjusting the copolymer properties, e.g. by changing MPEG/hexPLA
ratios and molecular weights, or attaching specific ligands relevant to the desired application.
152
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1
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Summary (In French)
153
Résumé et Conclusions
Parmi les substances bioactives les plus prometteuses pour des applications pharmaceutiques
à faible poids moléculaire, 95% sont lipophiles et faiblement solubles dans l’eau, rendant leur
formulation difficile et leur application clinique limitée. Après application chez l’homme, une
précipitation précoce est souvent observée due à la faible solubilité dans l’eau, par conséquent
les concentrations thérapeutiques sont rarement atteintes. Une des stratégies utilisée pour
pallier ces problèmes de formulation est l’incorporation de ces substances dans des micelles.
De part leur structure particulière « cœur-couronne », les micelles polymériques peuvent
incorporer des médicaments lipophiles au sein de leur cœur hydrophobe, tout en étant solubles
dans l’eau grâce à leur couronne hydrophile.
Le travail de cette thèse s’est orienté sur l’étude de nouvelles micelles polymériques : les
micelles à base de copolymères de methoxypolyethylène glycol (MPEG) et de polylactides
(PLA) substitués par des groupes hexyls (hex), nommés MPEG-hexPLA. La présence et le
nombre élevé de groupes hexyls dans les micelles de MPEG-hexPLA, comparés aux micelles
de MPEG-PLA standard, augmentent l’hydrophobicité du cœur et donc l’incorporation de
produits hydrophobes.
L’incorporation du composé modèle hydrophobe THPP a été testée et a validé cette stratégie.
Une plus grande incorporation a été mise en évidence avec les micelles les plus hydrophobes,
celles composées de copolymères di-substitués avec des groupes hexyls (MPEG-dihexPLA),
suivent ensuite les micelles composées de copolymères mono substitués de groupes hexyls
(MPEG-monohexPLA) et enfin les micelles de MPEG-PLA. Cette tendance a été
principalement observée pour les substances dont la solubilité dans l’eau dépassait 0.055mM.
Cependant la solubilité dans l’eau des drogues et leur log P ne sont pas les seuls paramètres
influençant leur incorporation dans les micelles. Des paramètres structuraux tels que le
nombre de groupes donneurs et d’accepteurs de liaisons hydrogènes dans la structure
chimique de la drogue ont aussi démontré leur influence. De façon intéressante, les drogues
de faible solubilité dans l’eau qui ne suivent pas la « règle des 5 » de Lipinski, souvent
synonyme de faible biodisponibilité et d’applications limitées en clinique, possèdent des
propriétés physico-chimiques appropriées pour une incorporation efficace dans les micelles de
Summary (In French)
154
MPEG-hexPLA. Ceci suggère l’utilisation de ces systèmes micellaires comme formulation
alternative pour des substances faiblement solubles dans l’eau.
Pour de possibles utilisations pharmaceutiques, les micelles de MPEG-hexPLA ont été tout
d’abord étudiées pour leur biocompatibilité. Aucune toxicité in vitro et in vivo ou activité
hémolytique sur le sang humain n’a été trouvée, à la fois pour les unimères et pour les
micelles polymériques jusqu’à une concentration de 20 mg/mL. Par conséquent, ces nouvelles
formulations micellaires s’avèrent intéressantes pour des applications pharmaceutiques.
L’application intraveineuse a été une des applications envisagées. En effet, les formulations
micellaires de cyclosporine A n’ont montré aucune activité hémolytique et ont pu atteindre
des doses thérapeutiques avec 4 fois moins d’excipient que le surfactant utilisé dans la
formulation commerciale, Sandimmune ®.
De plus, l’injection intraveineuse de formulations micellaires d’hypericine a présenté chez les
rats atteints de tumeurs ovariennes une accumulation rapide et élevée dans les nodules
tumoraux par un ciblage passif. Ainsi les micelles de MPEG-hexPLA peuvent être envisagées
pour le diagnostique du cancer de l’ovaire.
Une application topique pour des traitements antifungiques a également été évaluée dans cette
thèse. En effet, les micelles de MPEG-hexPLA ont montré leur supériorité lors de la
pénétration du nitrate d’éconazole à l’intérieur de la peau porcine et humaine par rapport à la
formulation commerciale. Ce dépôt a été curieusement attribué à une pénétration des micelles
par les follicules. Les micelles de MPEG-hexPLA peuvent être considérées comme vecteurs
colloidaux potentiels pour la libération d’autres produits hydrophobes dans la peau.
Une possible application par voie orale des micelles a aussi été étudiée. Après 1h à pH
gastrique, les micelles de MPEG-hexPLA ont conservé environ 90% de leur contenu en
cinnarizine et ont gardé leur structure de « cœur-couronne » intacte malgré une légère
dégradation du copolymère. Les micelles s’avère donc suffisamment stables pour passer
l’estomac et atteindre l’intestin, pour une application orale.
En résumé, les micelles de MPEG-hexPLA présentent de prometteuses formulations
pharmaceutiques biocompatibles pour des applications intraveineuses, topiques, et orales pour
les substances faiblement solubles dans l’eau.
ABBREVIATIONS
Abbreviations
Abbreviations
ACN
Acetonitrile
BCS
Biopharmaceutics classification system
bFGF
Basic fibroblast growth factor
CAM
Chick chorioallantoic membrane
CDCl3
Chloroform-d
CIN
Cinnarizine
CL
ε-Caprolactone
CLZ
Clotrimazole
CMC
Critical micellar concentration
CMT
Critical micellar temperature
CsA
Cyclosporine A
DAN
Danazol
DC
Drug content
DCM
Dichloromethane
DICLO
Diclofenac sodium
dihexPLA
di-hexyl-substituted polylactides
DLLA
D,L-lactide
D-PLA
Poly(D-lactide)
D,L-PLA
Poly(D,L-lactide)
DLS
Dynamic light scattering
DMAAm
N,N-Dimethylacrylamide
DMF
Dimethylformamide
dn
Number- weighted diameter
[%]dn
Percentage of micelles with a given dn
DSPE
Distearoylphosphatidylethanolamine
DTBA
p-Aminobenzyldiethylenetriaminepenta(acetic acid)
DTX
Docetaxel
ECZ
Econazole nitrate
EDD
Embryo development day
155
Abbreviations
156
EGF
Epidermal growth factor
EPR
Enhanced permeation and retention effect
ETO
Etoposide
FDA
Food and drug administration
FLZ
Fluconazole
Gd
Gadolinium
GF
Griseofulvin
GI
Gastro-intestinal
GRAS
“Generally Regarded As Safe”
Ha
H bond acceptor
HCl
Hydrochloric acid
Hd
H bond donnor
HexPLA
Hexyl-substituted poly(lactide)
(hex)PLA
Hexyl-substituted poly(lactide) and poly(lactide), respectively
HTS
High throughput screening
HY
Hypericin
IE
Incorporation efficiency
KETO
Ketoconazole
LCST
Lower critical solution temperature
LLA
L,L-lactide
LogP
Partition coefficient between octanol and water
L-PLA
Poly(L-lactide)
MAAc
Methacrylic acid
MeOH
Methanol
MD
Molecular dynamics
monohexPLA
mono-hexyl-substituted poly(lactide)
MPEG
Methoxy-poly(ethylene glycol)
MPS
Mononuclear phagocytic system
MRI
Magnetic resonance imaging
MTD
Maximum tolerated dose
MTT
3-(4,5-dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide)
Mw
Molecular weight
NaCl
Sodium chloride
Abbreviations
NAPRO
Naproxen sodium
NR
Nile Red
PAsp
Poly(aspartic acid)
PBL
Poly(butyrolactone)
PBS
Phosphate buffer saline
PCL
Poly(ε-caprolactone)
PDLA
Poly(D-lactide)
PDLLA
Poly(D,L-lactide)
PDMA
Poly[2-(N,N-dimethylamino)ethyl methacrylate]
PDMAEMA
Poly(2-(N,N-dimethylamino)ethyl methacrylate)
PDPA
Poly[2-(diisopropylamino)ethyl methacrylate]
PDT
Photodynamic therapy
PEEP
Poly(ethyl ethylene phosphate)
PEG
Poly(ethylene glycol)
PEI
Poly(ethylene imines)
PG
Poly(L-glutamic acid)
PHis
Poly(L-histidine)
P.I.
Polydispersity index
PICM
Polyion complex micelles
PIRO
Piroxicam
PLA
Poly(lactic acid) or poly(lactide)
PLGA
Poly(lactic-co-glycolic acid)
PLLA
Poly(L-lactide)
PMMA
Poly(methacrylic acid)
PNIPAAm
Poly(N-isopropylacrylamide)
PPO
Poly(propylene oxide)
PS
Poly(styrene)
PTX
Paclitaxel
PVA
Poly(vinyl alcohol)
PVL
Poly(valerolactone)
PVP
Poly(N-vinyl-2-pyrrolidone)
QUER
Quercetine dihydrate
ROP
Ring opening polymerization
157
Abbreviations
SC
Stratum corneum
Sn(Oct)2
Tin octanoate
THF
Tetrahydrofuran
THPP
meso-Tetra(p-hydroxyphenyl)porphine
VER
Verteporfin
Zav
Hydrodynamic diameter
158

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