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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5915_Библиотеки_им_академика_М_И_Перельмана
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(Graveland-Bikker and de Kruif 2006), nanoparticle sunscreens (Wissing and
Muller 2002), and drug delivery systems, the topic of this chapter. Drug delivery
systems range from nanostructures to particles in the sub-visible range (microparticles) to visible implantable devices. The purpose of this chapter is to describe the
materials, methods, and challenges in designing nanoparticulate delivery systems.
Toward the end of this chapter, alternative delivery systems including microparticles
and implants are also discussed.
As early as 1974, Dr. V.F. Smolen discussed methods to design, develop, and
evaluate drug delivery systems to enhance drug efficacy in terms of bioavailability
and drug response (Smolen et al. 1974). More than three decades later, innovative
drug delivery systems are being marketed by several pharmaceutical companies
including Allergan, Inc. (OzurdexTM), Bausch & Lomb, Inc. (Retisert®), and Abraxis
BioScience, Inc. (Abraxane®). However, engineering drug delivery systems is a
continuous process, given the challenges of new diseases and therapeutic agents.
For this reason, several industrial agencies and academic institutions are continuously engaged in designing novel drug delivery systems. Dr. Ellis Meng and group
at the University of Southern California (USC) has conducted pioneer research in
developing novel implantable and refillable microelectromechanical system
(MEMS)-based engineered ocular drug delivery systems (Li et al. 2008; Saati et al.
2009). This system accurately administers a finite amount of drug at certain time
intervals and is also capable of easily being refilled by injection into the device. In
fact, Dr. Meng is working with several collaborators including Replenish, Inc.,
California Retina Consultants and Doheny Eye Institute to develop a novel implantable,
refillable pump for intraocular drug delivery into the vitreous by attaching the device
externally to the vitreous. Novel drug delivery systems will continue to push the
insights and forefront of pharmaceutical biotechnology and is the key to enhanced
ocular therapeutics in terms of safety and efficacy.
This chapter will primarily focus on nanomaterials and systems that are biodegradable or bioresorbable, which will be collectively referred to as nanosystems.
Various materials and methods are available to fabricate nanosystems of varying
size, morphology, and composition (see Fig. 11.1). Size can range from just a few
nanometers to thousands of nanometers and the morphology can range from spheres
to highly ordered structures (e.g., rods, disks, cubes, and diamonds). In addition, the
nanosystems can be solid (e.g., nanoparticles or nanospheres), fluid filled (e.g., nanoliposomes), gel-like (e.g., hydrogels), or soluble (e.g., water soluble drug-polymer
conjugates). Common materials used to fabricate nanoparticles include both metallic and organic compounds. Examples of metallic nanosystems include gold, silver,
and iron oxide nanoparticles for drug delivery. In the design of pharmaceutically
viable drug carriers at doses suitable for long-term therapies, organic materials
including polymers, lipids, proteins and carbohydrates are likely to be safer compared to metallic nanosystems. However, along with a drug molecule, the clinical
viability of all delivery systems is ultimately determined by the risk:benefit analysis
in a target patient population.
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Fig. 11.1 Some examples of nanoparticles currently under investigation for use as ocular drug
delivery vehicles: (a) albumin nanoparticles, (b) chitosan nanoparticles, (c) polyamidoamine
(PAMAM) dendrimers, (d) poly (lactic-co-glycolic) acid (PLGA) nanoparticles, and (e) polyethylene glycol (PEG)-coated liposomes
11.2 Nanoparticles
Disparate nanoparticles from those that mimic cellular structures, lipid-based delivery
carriers, to engineered branched structures such as dendrimers have been designed
to enhance the efficacy of specific therapeutic agents. US FDA already approved
some nanosystems such as Doxil® (a liposomal formulation of doxorubicin) and
Abraxane® (albumin-bound paclitaxel) due to their enhanced drug retention at the
target sites, compared to drug alone. Depending on the drug properties, target properties, and the purpose of the carrier, different synthetic materials and techniques
are available for designing specific nanoparticles. The following sections describe
the synthetic techniques and uses for polymer nanoparticles, liposomes (lipid-based
carriers), micelles (amphiphilic molecule-based, self-assembling systems), carbohydrate (chitosan) nanoparticles, protein (albumin) nanoparticles, branched nanoparticles (dendrimers) and other nanosystems consisting of multiple nanoparticle
structures.
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11.2.1 Polymer Nanoparticles
Polymers are large molecular weight compounds consisting of systematic or random
repeat units; for example, polyethylene glycol (PEG), polylactic acid (PLA), polyglycolic acid (PGA), and polylactic-co-glycolic acid (PLGA) that are currently in
clinical use in the form of drug delivery systems or surgical sutures are all polymers
of repeating units. Polymers can be classified based on their structure (e.g., polyesters, polyanhydrides), stability (e.g., biodegradable, nondegradable), charge (e.g.,
cationic, anionic), lipophilicity (e.g., hydrophobic, hydrophilic, amphiphilic), origin
(e.g., synthetic, natural, semi-synthetic), architecture (e.g., linear, branched, crosslinked), and nature of repeating units (e.g., homopolymers, copolymers, block
copolymers, random copolymers). Further, polymers can be formed in various
supramolecular architectures (e.g., interpenetrating and noninterpenetrating networks, micelles) (Qiu and Bae 2006). Also, polymers can be designed for sensitivity to various stimuli including pH and temperature. Given the versatility of
polymer design, polymer-based nanosystems are expected to be the mainstay of
nanotechnology- based drug delivery systems. Depending on their chemical makeup
and architecture, polymeric delivery systems can be designed to have certain properties such as biodegradation, sustained release, increased gene transfection efficiency,
and controlled release by actuated or stimuli-sensitive physicochemical changes.
Polymers such as PLA and PLGA are widely used in the design of a delivery
system since these compounds are biodegradable, biocompatible, and well tested in
humans. PLA and PLGA particles are degraded by autocatalysis (Dunne et al.
2000). PLGA degradation can range from days to months depending on the molecu-
lar weight of the polymer, lactide:glycolide ratio, and the size and shape of the
delivery system. In 1987, HV Maulding characterized the release profiles of PLGA
microparticles (~45 kDa) ranging from 45 to 177 mm and determined that it takes 70
days for 100% loss of the molecular weight of PLGA (Maulding 1987).
Drug release from nanoparticles can be more rapid compared to microparticles,
leading to less prolonged release. Kompella et
ticles (3.6 mm) containing budesonide, a glucocorticoid used to treat inflammation,
were able to sustain drug levels within the retina, vitreous, cornea and lens at similar
levels between day 1 and 14 days, whereas tissue drug levels for PLA nanoparticles
(345 nm) decreased by several fold in 7 days, with the levels being below detection
limits by 14 days in retina, vitreous, cornea, and lens (Kompella et al. 2003). Thus,
in addition to the polymer nature, the size of the delivery system influences drug
release and hence, delivery in vivo.
Polymers also have great potential as nonviral vectors due to their relative safety
compared to viral vectors. PLA and PLGA nanoparticles have been used for oligonucleotide or gene delivery to retinal pigment epithelial (RPE) cells and were proven
to be more effective than traditional transfection methods due to their ability to
transfect the cells without adverse effects such as cell toxicity (Aukunuru et al. 2003;
Bejjani et al. 2005). These nanoparticles have also been shown to protect the encapsulated plasmids from degradation by lysosomal nucleases (Hedley et al. 1998).
Other polymers have demonstrated preferential binding to genetic components by
al. demonstrated that PLA micropar-
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electrostatic interactions. The positively charged polymer, polyethylenimine, is able
to electrostatically interact with nucleic acid drugs to create a polymer–drug complex (Boussif et al. 1995). Due to the composition of the polymer, it acts as a proton
sponge in the presence of highly acidic environments such as the lysosomes. This
allows for the nucleic acid drugs to stay active while in the endosome after endocytosis and for the drug to be released from the polymer complex after the polymer
accepts protons. The versatile nature of polymers is largely attributed to the wide
range of materials that can be used to synthesize polymer structures. Using polymers
as gene transfection agents demonstrates only one of the many possible applications
for polymers.
Polymers can be designed to undergo structural changes upon activation to
release their drug components. Polymers containing cinnamic acid groups spontaneously form into linear, spiral, tube, or corkscrew structures depending on the UV
wavelength applied and whether the polymer is irradiated on both sides or only one
side (Lendlein et al. 2005). For instance, irradiation of an elongated polymer at
>260 nm for 60 min on one side of the polymer spontaneously forms to a corkscrew
structure. However, if the polymer is subjected to irradiation at >260 nm for 60 min
on both sides, a spiral conformation is obtained. This technique may be applied to
the development of drug delivery systems to allow for actuated release from polymer structures or formation of unique drug delivery systems upon irradiation.
Thermosensitive materials that exhibit temperature-dependent physicochemical
properties offer attractive opportunities in designing delivery systems. Such materials are useful in preparing delivery systems that can be injected as solutions to form
gels in the body (Zhang et
al. 2002). Synthesis of a two-component thermosensitive
polymer was reported by Lendlein et al. (Lendlein et al. 2005). The first component
is a molecular switch that conforms to a temporary shape at a given temperature.
This polymer is grafted on a permanent polymer network to form a particular architecture. Polymer grafting will vary depending on the network composition desired.
For example, one could copolymerize n-butylacrylate (BA), hydroxyethyl methacrylate (HEMA), and ethyleneglycol-1-acrylate-2-CA (HEA-CA) with the
crosslinker poly(propylene glycol)-dimethylcrylate. The elasticity of the polymer
is dependent on the amount of ethyleneglycol-1-acrylate-2-CA (HEA-CA) added.
The second component is synthesized by making a permanent network of
n-butylacrylate (BA) with the crosslinker, poly(propylene glycol)-dimethacrylate.
Star-poly(ethylene glycol) containing cinnamylidene acetic acid (CAA) in 10%
chloroform solution is then added to the network to functionalize the polymer
with cinnamic acid (CA) groups. Such polymers can potentially be transformed into
different architectures.
The complexity of polymer particle preparation is entirely dependent on the
properties desired and materials to be used. Solid polymeric particles containing a
specific drug can be synthesized using an oil-in-water (o/w) emulsion technique
(Fig. 11.2) (Kompella et al. 2001, 2003). The polymer and drug contents (e.g.,
PLGA and budesonide) are dissolved in an appropriate organic solvent (e.g., dichloromethane) and then emulsified in an aqueous medium containing an emulsifier
(e.g., polyvinyl alcohol, PVA). The mixture is further agitated with a probe sonicator
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S.A. Durazo and U.B. Kompella
Aqueous Phase Organic Phase
PVA (0.5 or 2% w/v) solution with
or without budesonide
Mix and sonicate at 50 W
for 3 min to obtain an
O/W emulsion
While stirring add drop-wise to an aqueous
PVA (0.5 or 2% w/v) solution with or without
budesonide. Continue to stir at 200 rpm for 4-
6 hours to obtain a nanosuspension
Ultracentrifuge to obtain
nanoparticles pellet
Wash the pellet twice with de-ionized
water and reconstitute the pellet in
deionized water
Lyophilize the reconstituted
contents for 48 hours to obtain a
dry product
Polymer+Budesonide+
Methylene chloride
Fig. 11.2 Scheme describing the emulsion-solvent evaporation method for preparing budesonideloaded polylactide (PLA) nanoparticles (Kompella et al. 2001)
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to obtain an o/w emulsion with smaller droplet size. The emulsion is then added
drop-wise to excess of aqueous medium containing PVA, while stirring overnight at
room temperature to allow the evaporation of the organic solvent. Upon removal of
organic solvent, the polymer precipitates along with the drug from emulsion droplets, resulting in fine nanoparticles. The final preparation may contain nanoparticles
as well as microparticles, depending on the materials and energy used. Particles can
be further segregated in different size ranges using ultracentrifugation at different
speeds. An alternative approach to prepare PLGA nanoparticles is supercritical fluid
extraction of emulsions (Mayo et al. 2010). This technique allows high encapsulation efficiencies for hydrophilic drugs in a polymer matrix. Further, it reduces organic
solvent content in polymeric particles to detection limits or a few parts per million.
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Fig. 11.3 Scheme depicting the dialysis method for analyzing sustained in vitro drug release from
budesonide-loaded polylactide nanoparticles (Kompella et al. 2001)
Drug release from any nanosystem or nanoparticles prepared as above can be
performed using a membrane-free system or a membrane-containing approach. In
both approaches, an aqueous medium with a preservative such as sodium azide is
used for prolonged drug release studies. In membrane-free systems, nanoparticle suspension is centrifuged periodically and the supernatants are quantified for the amount
of drug released. In membrane-containing systems (Fig. 11.3), a dialysis membrane
is used to separate the particles from the bulk of the release medium (Kompella
et al. 2001). The dialysis membrane has a cutoff size greater than drug molecule
size, but less than the particle size. At periodic intervals, aliquots are removed from
the bulk of the release medium to quantify drug release. In both approaches, the
release medium should provide sink conditions for continuous drug release.
Polymer-based nanoparticles are very attractive due to the availability of a variety
of polymers and architectures, possibility of structure manipulation and maintenance,
which can potentially translate into drug delivery systems with controlled properties.
11.2.2 Liposomes and Lipid Nanoparticles
When hydrated, lipids such as dipalmitoylphoshatidylcholine (DPPC) spontaneously form closed spherical lipid bilayers, called liposomes or vesicles (see Fig. 11.1)
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(Walker et al. 1997). The content within the spherical compartment of the liposome
is largely dictated by the contents of the medium in which the liposomes are formed.
The liposomal structure can be small, large or giant, and can be unilamellar (comprised of one lipid bilayer) or multilamellar (comprised of layers of lipid bilayers)
depending on the method of preparation. Acronyms such as SUV (small unilamellar
vesicles), LUV (large unilamellar vesicles), and MLV (multilamellar vesicles) are
used to describe the nature of the vesicle. Liposomes can be synthesized using a
variety of methods. Typically, the lipids obtained from a commercial source are
dried under N
followed by rehydration in a buffer (Wu et al. 2008). The buffer can
2
consist of the drug desired to be encapsulated in the liposome or it can consist of
other small molecule agents. Rehydration is performed under heating and vigorous
vortexing in a buffer over approximately 4–16 h to allow for the lipids to enter solution. The temperature is set above the phase transition temperature of the lipid from
the solid crystalline phase to the fluid crystalline phase. The size of the lipid can be
greatly reduced by sonication after rehydration or by extrusion through a porous
membrane with the pore diameter close to the desired liposome diameter. Thus,
liposomal preparation can be a simple process whereby liposomes are ready within
a day assuming the complexity of the liposome is minimal. For liposomes with
targeting ligands or multiple compartments, synthesis will likely be much more
complex and time-consuming.
Evading the reticuloendothelial system (RES) is a major hurdle in the development of nano-drug delivery systems. RES is responsible for the detection and elimination of foreign objects including nano-drug delivery systems by the phagocytic
cells, macrophages, and primary monocytes. Several investigators have focused on
determining the optimal particle size and composition of delivery devices that
bypass the RES. Litzinger et al. in 1994 demonstrated that PEG-coated liposomes
with a diameter of 150–200 nm have optimal blood retention levels, whereas liposomes within the range of 200–300 nm experienced highest uptake by the RES and
were found in the spleen (Litzinger et al. 1994).
STEALTH liposomes, such as liposomes coated with PEG that are capable of
evading RES at least in part, are FDA approved and have been implicated to improve
target tissue availability of cytotoxic drugs (e.g., doxorubicin, Doxil®), while minimizing their nontarget tissue delivery. The PEG moiety provides for increased retention in the circulation as well as tissue compartments by preventing premature lipid
degradation by lipsases (e.g., phospholipase) and providing a steric barrier that prevents liposomes from interacting with opsonins and macrophages that are involved
in liposomal clearance (Papahadjopoulos et al. 1991). However, the precise mechanism by which PEG increases circulation time and tissue drug retention is highly
debated among scientists in this field. Interestingly, PEG-coated 100 nm liposomes
(DoxilTM) can sustain the release of doxorubicin across sclera when compared to the
plain drug (Kim et al. 2009a).
Several investigators have realized the potential for liposomes as nonviral vectors. As early as 1998, it was discovered that positively charged lipids are capable
of binding to deoxyribonucleic acid (DNA) and releasing them into the cell for
gene expression (Koltover et al. 1998). It was later discovered that PEG-coated
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lipid–DNA complexes had extended circulation times and tumor-selective gene
expression due to increased blood circulation time to reach the tumor site (Ambegia
et al. 2005). Most recently, several of the primary investigators in this field including Drs. Ian MacLachlan and Pieter Cullis have developed a rational design approach
to develop lipids for siRNA delivery (Semple et al. 2010). The use of liposomes for
gene delivery and transfection is a prominent field and although there have been
several breakthroughs, the field is still in its infancy and little is known regarding the
mechanism of transfection and cellular uptake. The potential for liposomes to be
marketed is largely hindered by this “black box” in which controlling the retention
and cell uptake properties are poorly understood.
Liposomal delivery devices for ocular drug delivery have been under investigation at least since 1981 and various routes of administrations have been used in the
comparison of liposomal formulations vs. traditional therapeutic formulations
(Meisner and Mezei
1995). Smolin et al. in 1981 found that topical administration
of idoxuridine in the liposomal formulation was much more effective than in its
marketed formulation in the treatment of acute and chronic herpetic keratitis in rabbit
(Smolin et al. 1981). Peyman et al. further validated this study in 1986 and confirmed that liposomal formulations of idoxuridine were able to penetrate the cornea
much better than the marketed formulation (Dharma et al. 1986). The administration of liposomal formulations by subconjunctival (SC) injection in rabbit was also
evaluated by other investigators in 1991 (Hirnle et al. 1991). Liposomal formulations made with phosphatidylcholine (PC) cholesterol (Chol) dicethyl phosphate
(DP) lipids encapsulating carboxyfluorescein (CF) were injected subconjunctivally
and 30 min after injection, ~87.4, 5.8 and 2.3% of CF was distributed to the sclera,
retina and choroid tissues totaling in 95.5% of total CF distribution in the intraocular tissues of the posterior segment. The following data describe the concentrations
of CF in each tissue after administration of the liposomal formulation. After 1 h, the
amount of CF at the site of injection dropped from ~600 to 275 mg/g. In the sclera,
the amount of CF at 1 h was ~40 mg/g and at 6 h reduced to ~15 mg/g and in the
retina-choroid after 1 h ~8 mg/g of CF was present and after 6 h, ~6 mg/g was present. However, CF levels as high as 5 mg/g persisted in the retina-choroid for at least
7 days. At 1 h only ~7.5 mg/g of CF was found in the cornea and decreased to
~2 mg/g at 6 h. In the iris-ciliary body, at 1 h ~8 mg/g of CF was present and decreased
to ~1.5 mg/g at 6 h. Less than ~2 mg/g of CF was in the vitreous and less than
~2.5 mg/g of CF was in the anterior chamber over the entire time period 0–6 h.
Concentrations of CF after administration of aqueous CF (nonliposomal) were
much less in all ocular tissues. After subconjunctival injection of CF, the concentration dropped to ~5 mg/g after 1 h; at 1 h, the concentration in the sclera, retinachoroid, cornea, iris-ciliary body, anterior chamber and vitreous was ~8, 2, 0.2, 2,
0.2, and 0.2 mg/g, respectively. The liposomal formulation of CF was found to have
higher drug retention in all ocular tissues compared to aqueous CF, but the concentrations within the posterior segment of the eye was highest.
Liposomal formulations were shown to be much more effective than unencapsulated drugs as early as 1986 and continue to show superiority over small molecule
ocular therapeutics. Topical administration of positively charged and neutral
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acetazolamide MLVs consisting of phosphatidylcholine:cholesterol:stearylamine in
a molar ratio of 7:4:1 were much more effective in lowering intraocular pressure
(IOP) than negatively charged MLVs of the same composition and aqueous acetazolamide (Hathout et al. 2007). In 2009, another study investigated the potential use
of a chitosan-coated liposomes encapsulated with diclofenac sodium for sustaining
drug release (Li et al. 2009). The authors used 0.25 and 0.5 mol% 540 kDa chitosan
and the diameter was 82.4 ± 2.2 and 84.0 ± 4.7 nm, respectively. Both liposomal
formulations released only 50% of the drug over 24 h, whereas the nonchitosancoated diclofenac liposomes and aqueous diclofenac released 60% and 90%, respectively, over 24
360 min for chitosan-coated diclofenac liposomes (both 0.25 and 0.5 mol% chitosan), noncoated diclofenac liposomes, and aqueous diclofenac dropped from 16 to
6 mg/ml, 15 to 2 mg/ml, and 13 to 1.5 mg/ml, respectively. The cumulative penetration for the 0.25 mol% chitosan-coated diclofenac liposomes, 0.5 mol% chitosancoated diclofenac liposomes, noncoated diclofenac liposomes, and aqueous
diclofenac was 35, 32, 22 and 25 mg/cm2, respectively. From these studies, the
importance of liposomal surface charge and composition in determining tissue
penetration, drug retention, and drug concentration in ocular tissues is apparent.
Similar to polymeric systems, laser irradiation-triggered release is also feasible
with liposomal systems. Zasadzinski et al. have showed that liposomal release
can be controlled by triggering the release by activating gold nanoparticles
encapsulated within the liposome (Wu et al. 2008). Further, by modifying lipids
with targeting elements, targeted liposome delivery systems can be prepared. Thus,
liposomes offer a commercially viable approach for encapsulation and sustained
drug delivery.
Alternative lipid-based nanosystems include solid lipid nanoparticles and nanostructured lipid carriers (Pardeike et al. 2009). In this case, similar to PLGA- and
PLA-based nanoparticles, lipids are formulated into solid nanoparticles or nanoparticles with defined nanostructure, without the core (liquid) and coat (lipid) structure
of liposomes. Similar to polymeric systems as well as liposomes, solid lipid nanoparticles and nanostructured lipid carriers can also be surface modified to improve
circulation time or to enhance cell uptake.
h. The concentrations of diclofenac in precorneal regions after
11.2.3 Micelles
Similar to liposomes, micelles are formed by self-assembly of amphiphilic molecules; however, micelles exhibit a slightly different higher-ordered structure
(Israelachvili et al. 1975). Micelles may assemble in three different distinct structures (Trivedi and Kompella 2010): (1) standard micelles, where the hydrophilic
portion is oriented toward the solvent to form the shell of the micelle and the hydrophobic portion is oriented away from the solvent to form the core of the micelle,
(2) inverse micelles, where the hydrophobic parts form the shell of the micelle and
the hydrophilic parts form the core of the micelle, and (3) unimolecular micelles,
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where one molecule with hydrophilic and hydrophobic blocks forms the entire
micelle. On exposure to aqueous solvent, the hydrophobic parts are clustered in the
center of the structure to form the core of the micelle and the hydrophilic parts are
primarily on the outside of the structure to form the shell of the micelle. Depending
on the solvent and type of amphiphilic molecules used, different structures will
predominate. For example, in hydrophobic solvents, the inverted micelle will predominate so that the hydrophobic interactions between the hydrophobic parts of the
micelle and the solvent are maximized while the hydrophilic portions of the molecule are oriented away from the solvent.
The propensity of an amphiphilic molecule to form a micelle will largely depend
on the characteristics of the molecule (Lukyanov and Torchilin
2004).
Polyethyleneglycol (PEG) is a polymer commonly used in the synthesis of lipid
drug carriers due to its ability to reduce particle aggregation and prolong systemic
exposure time. PEG between the molecular weights of 750 and 5,000 when ligated
to phosphatidylethanolamine (PE) lipids can spontaneously form micelles at a concentration above the critical micellar concentration (CMC). For example, PEG(750)distearoyl-phosphatidyl-ethanolamine (DSPE) has a CMC of 0.1 mM and forms
7–15-nm-sized micelles. These micelles assemble spontaneously when PEG–DSPE
dry lipid is hydrated for several hours (Lukyanov and Torchilin 2004). Methods
such as sonication or extrusion can be performed to further control the size of the
micelles.
Several authors (Kataoka et al. 2000; Liu et al. 2000; Yoo and Park 2001) have
demonstrated the ability of different micelle structures to sustain the release of
drug molecules over a few days. A prodrug of paclitaxel loaded in PEG-b-poly
(e-caprolactone) micelles (PEG-b-PCL) of 27–44 nm diameter resulted in ~40%
cumulative release of the prodrug in 14 days (Forrest et al. 2008). These paclitaxel
prodrug-loaded micelles were also shown to have a slightly higher anticancer activity than free paclitaxel as well as prodrug cremophor formulation in breast cancer
cell lines. In serum, the total concentration of paclitaxel was reduced threefold in
~25 h for the cremophor formulation of paclitaxel prodrug and in ~50 h for the
PEG-b-PCL micellar formulation of paclitaxel prodrug. These paclitaxel prodrugloaded PEG-b-PCL micelles were fabricated by dissolving PEG-b-PCL and paclitaxel with a minimal volume of acetone and adding it dropwise with the help of a
syringe pump to vigorously stirred distilled water. The organic solvent was removed
by stirring under air purge and the micelles were purified by extruding through a
membrane filter with a 0.22-mm pore size.
Indomethacin-loaded micelles using methoxy-PEG/e-caprolactone (e-CL) block
copolymers (MePEG/e-CL) have been shown to be superior over plain drug in sustaining drug release (Kim et al. 1998). The indomethacin-loaded MePEG(5000)/e-CL
micelles were synthesized by dissolving the MePEG/e-CL block copolymer in an
organic solvent and then indomethacin was added to the mixture and stirred at room
temperature. The micelles were dialyzed and then centrifuged to remove the unencapsulated drug. The size of the resulting micelles ranged from 54 to 180 nm and
the size depended on the molecular weight of the copolymer and the solvent in
which it was dissolved to form micelles. As the molecular weight of the copolymer
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