kobe university repository : thesis · have also been incorporated into hydrogels. for example,...

Kobe University Repository : Thesis

学位論文題目Tit le

Novel Polymer Systems for Controlled Drug Delivery(高分子を用いた薬物送達システムの研究)

氏名Author Inoue, Tadaaki

専攻分野Degree 博士（工学）

学位授与の日付Date of Degree 1998-03-11

資源タイプResource Type Thesis or Dissertat ion / 学位論文

報告番号Report Number 乙2223

権利Rights

JaLCDOI 10.11501/3141268

URL http://www.lib.kobe-u.ac.jp/handle_kernel/D2002223※当コンテンツは神戸大学の学術成果です。無断複製・不正使用等を禁じます。著作権法で認められている範囲内で、適切にご利用ください。

PDF issue: 2021-06-26

Novel Polymer Systems for

Controlled Drug Delivery

SfJ5X 1 o if. 1 ~

*J:: FaaBB

Novel Polymer Systems for

Controlled Drug Delivery (~~-T- ~ ffll, \ t~~~m~)i ~.A T bO)~1f~)

Novel Polymer Systems for Controlled Drug Delivery

Tadaaki Inoue

1998

CONTENTS

Page

GENERAL INTRODUCTION ........................................................... 1

REFERENCES ....................................................................... 10

PART I NOVEL HYDROPHOBICALLY -MODIFIED POLYMERIC

CARRIERS FOR CONTROLLED DRUG DELIVERY

Chapter 1 A Hydrophobically-Modified Bioadhesive Polyelectrolyte

Hydrogel for Drug Delivery

INTRODUCTION .................................................................... 29

MATERIAlS AND METHODS .................................................... 30

RESULTS AND DISCUSSION .................................................... 38

CONCLUSIONS ..................................................................... 48

REFERENCES ....................................................................... 50

Chapter 2 A Hydrophobically-Modified, Bioadhesive Polymeric

Carrier for Controlled Drug Delivery to Mucosal Surfaces

INTRODUCTION .................................................................... 55



CONCLUSIONS ..................................................................... 70

REFERENCES ....................................................................... 70

i

Chapter 3 An AB Block Copolymer of Oligo(methyl methacrylate)

and Poly(acrylic acid) for Micellar Delivery of

Hydrophobic Drugs

INTRODUCTION .................................................................... 73



CONCLUSIONS ..................................................................... 89

REFERENCES ............................................ , '" ....................... 89

PART II TEMPERATURE SENSITIVITY OF A HYDROGEL

NETWORK CONTAINING DIFFERENT LCST OLIGOMERS

GRAFTED TO THE HYDROGEL BACKBONE

Chapter 4 Temperature Sensitivity of a Hydrogel Network Containing

Different LCST Oligomers Grafted to the Hydrogel

Backbone

INTRODUCTION ........................... '" ...................................... 97


RESULTS AND DISCUSSION ................................................... 108

CONCLUSIONS .................................................................... 116

REFERENCES ...................................................................... 116

ii

PART III HYDROGELS CARRYING PENDANT PHOSPHATE

GROUPS FOR CONTROLLED DRUG DELIVERY

Chapter 5 Lysozyme Loading and Release from Hydrogels Carrying

Pendant Phosphate Groups

INTRODUCTION ................................................................... 125

MA TERiAlii AND METHODS ................................................... 126


CONCLUSIONS .................................................................... 141

REFERENCES ...................................................................... 141

Chapter 6 Stirn uli-Sensitive Protein Release from Hydrogels Carrying

Pendant Phosphate Groups

INTRODUCTION ................................................................... 145

MATERIAlii AND METHODS ................................................... 147


CONCLUSIONS .................................................................... 158

REFERENCES ...................................................................... 161

SUMMARy ................................................................................ . 165

LIST OF PUBLICATIONS .......................................................... .. 169

LIST OF OTHER PUBLICATIONS ............................................... . 171

DELIVERED PAPERS .................................................................. 172

iii

PATENT APPLICATIONS ............................................................ 175

ACKNOWLEDGMENTS .............................................................. . 177

iv

GENERAL INTRODUCTION

The concept that delivering a therapeutic drug to where it is needed, when it is needed,

or how much it is needed, is called "drug delivery system (DDS)". After the

administration of a conventional drug formulation, the blood level of the drug rises and

then declines. A drug has a certain therapeutic blood level above which it is toxic and

below which it is no effect. In order to curtail the toxic blood level and prolong

therapeutic blood level, numerous efforts have been done. In the past two decades, new

drugs such as bioactive proteins are becoming available through genetic engineering. As

the delivery of the bioactive proteins is usually more complicated than that of conventional

low molecular weight drugs, new delivery systems are required for the bioactive proteins.

Natural products have been extensively investigated as drug carriers because of their

good biocompatibility [1-9]. Among them, phospholipids have been paid much attention

because they are major components of the body tissues. Liposomes and lipid

micro spheres are composed of phospholipids and have been extensively investigated as

drug carriers [1-9]. When these carriers injected intravenously, they distributed specific

sites of the body such as the inflammed regions, and the atherosclerotic region of the vain

[10]. These characters of lipid microspheres and liposomes were applied to

pharmaceutical preparations. The lipid microsphere preparations of corticosteroid [7], a

non-steroidal anti-inflammatory drug [8], and prostaglandin E1 [9] have been developed

and launched on Japanese market. They are prescribed for rheumatoid arthritis, pains and

peripheral vascular diseases, respectively.

In the past few decades, synthetic polymers have been widely used in medical field,

such as catheters, orthopedic prostheses, sutures, tissue engineering, implants, contact

lens, intraocular lens, kidney dialyzer, diagnostics, biosensors so on, because of their

compositional diversity and good biocompatibility [11-20]. Synthetic polymers also have

1

General Introduction

been extensively investigated as drug carriers in DDS [11-20]. A copolymer of lactic acid

and glycolic acid (PLGA) was applied to a controlled release preparation of a lcuterized

hormone release hormone (LHRH) analogue, leuprolide [21]. In this application, when

the formulation is injected to the body subcutaneously, the drug released from the

copolymer at zero order kinetics over a month period, and a good therapeutic index is

obtained in prostate cancer [21]. Moreover, since PLGA is gradually hydrolyzed in the

body, after the drug is completely released, PLGA is dissappeared from the body_ It has

been already launched on the world-wide market. Synthetic polymers are also clinically

used in transdermal formulations [22] and oral formulations [23]. To delivering a drug

when it is needed and how much it is needed could be accomplished by release rate

controlled formulations and has been paid much attention, because of its practicability.

For example, controlled release preparations maintains the desired drug level by a single

administration, which have been improving the quality of life of patients as mentioned

above. The purpose of this study is to develop the release rate controlled drug delivery

system using synthetic polymers.

Hydrogels are of special interest in biological environments because they have high

water contents similar to body tissues and have, in general, high biocompatibility [15,

24-26]. Polymeric hydrogels have been widely studied for use in medical implants,

diagnostics, biosensors, bioreactors, and bioseparators, and are being used as matrices

for DDS [15, 24-26]. Peppas and his colleagues have extensively investigated hydrogels

as drug release matrices [27-31]. Lee and his colleagues have also reported swelling-

controlled drug release from hydrogels [32-35]. Many researchers have studied drug

release properties of thermally-sensitive hydrogels [36-44]. Hydrophobic components

have also been incorporated into hydrogels. For example, Hoffman and his colleagues

prepared temperature-sensitive and both temperature- and pH-sensitive hydrogels

containing silicone domains [37, 44]. Others have studied random copolymers of

hydrophobic polyelectrolyte hydrogels [32, 45-48]. Siegel et al. [46] synthesized

2


hydrophobic polyelectrolyte random copolymer hydrogels hased on methyl methacrylate

(MMA) and N,N -dimethylaminoethyl methacrylate and reported pH-sensitive, swelling-

controlled release of caffeine. Kim and Lee [32] synthesized hydrophobic polyelectrolyte

random copolymer hydrogels using MMA and methacrylic acid (MAA) and reported pH-

sensitive, swelling-controlled release of oxprenolol hydrochloride. However, there is no

report in the literature of hydrophobically-modified pH-sensitive hydrogels based on

hydrophobic oligomers grafted to a polyelectrolyte hydrogel matrix. In Chapter 1, a

hydrophobically-modified bioadhesive polyelectrolyte hydrogel with a hydrophobic

domain structure was prepared and applied as a bioadhesive hydrogel for controlled

release of either cationic drugs or hydrophobic drugs [49]. The hydrophobic domains

may also strengthen the hydrogel and control the pore structure within the hydrogel [44,

50]. Poly(acrylic acid) (PAAc) was used as the bioadhesive polyelectrolyte. The

bioadhesive properties of PAAc [51] should enhance the utility of this graft copolymer as

a drug delivery vehicle. It is well known that PAAc has been used clinically as a lightly

crosslinked hydrogel known as "Polycarbophil" [52] or "Carbopol" [53] for drug

delivery. Oligo(methyl methacrylate) (oMMA) , a polymer which is well-tolerated as an

implant material [54, 55] was selected as the hydrophobic component to be grafted to the

backbone of a lightly crosslinked PAAc hydrogel. The swelling behavior of the hydrogel

and drug release profiles from the hydrogel were investigated using six model drugs.

Water soluble polymers are of great interest as drug carriers, since the polymer may

dissolve as drug is released [5~4]. For example, Kopecek and his colleagues have

synthesized and extensively investigated poly(hydroxypropyl methacrylamide)

(PHPMA) as a drug carrier [63, 64]. Kim and his colleagues reported on soluble pH-

and temperature-sensitive linear polymers for protein and peptide delivery [57, 58]. Lee

and his colleagues investigated soluble polymers for drug delivery vehicles [59, 60].

Hoffman and his colleagues have reported on the synthesis and properties of a soluble

carrier prepared by grafting a thermally-sensitive polymer to the backbone of a

3


bioadhesive anionic polyelectrolyte, PAAc [56]. The grafted side chains were designed

to phase separate at body temperature and form hydrophobic domains within the poly

anionic matrix network, which shows the dissolution of the PAAc ionized [56]. This can

show the release of a drug from the bioadhesive formulation. In Chapter 2, the study in

Chapter 1 was extended to a soluble polymer system [66]. The oMMA and co-oligomers

with hydroxyethyl methacrylate (HEMA) were grafted to the soluble P AAc backbone.

These individual components are known to be biocompatible polymer for certain

applications [51, 54,67]. PAAc is known to have good bioadhesive properties [51],

and the graft copolymer synthesized is expected to adhere to mucosal surface, where the

drug is delivered. P AAc has also been reported to protect cationic proteins such as

transforming growth factor-~l (TGF-~l) from inactivation due to complexation with

alginate [68]. A series of comb-like graft copolymers with the hydrophobic oligomers

grafted onto the P AAc backbone was synthesized. In aqueous media, the hydrophobic

graft chains should form a physical network having hydrophobic domains surrounded

by the bioadhesive polyelectrolyte (PAAc), resulting in a carrier that could be useful for

topical, bioadhesive drug delivery of hydrophobic drugs and cationic drugs [49, 56, 69-

72].

In Chapter 3, an AB block copolymer of oMMA and PAAc using a novel synthetic

method was synthesized. The AB block copolymer should form micellar structure in

which a hydrophobic core (oMMA) is surrounded by a bioadhesive polyelectrolyte

(PAAc) outer shell [69, 73]. As the micelles are formed by intermolecular association of

single polymer units, they should eventually be dissociated into a single polymer chains,

for excretion from the body. The advantage of polymer micelles over conventional

surfactant micelles is the good thermodynamic stability in physiological solutions of the

former, as indicated by their low critical micellar concentration (CMC) [73]. Similar block

copolymers with glassy hydrophobic segments, such as polystyrene, form stable

micelles, and the release rate of the single polymer chains from the micelle is expected to

4


be slowed by the glassy structure of the core [73]. The micelle should be useful as a drug

carrier for hydrophobic drugs [69-71, 74-86] especially if the drug acts as a plastisizer in

the micellar core. Furthermore, the technique permits to vary the composition of the

hydrophobic block, which controls the Tg of the core before drug loading. The drug itself

will affect the Tg of the core, and that is another variable that should be optimized in any

delivery system utilizing such micelles. Kataoka and his colleagues have extensively

investigated polymeric micelles as carriers of anti-cancer drugs [69-71, 74-83]. They

synthesized AB block copolymers of poly(ethylene oxide) and L-benzyl aspartate that

formed micelles with diameters in a range of several tens of nanometers, mimicking

viruses, in order to prevent reticuloendothelial systems (RES) recognition when

administered intravenously to the body [69]. They reported enhanced tumor

accumulation, prolonged circulation times [81), and reduced toxicity [77] of the AB block

copolymer micelle when the aspartate block was conjugated with pendant adriamycin

molecules instead of benzyl groups. In these studies, although adriamycin was covalently

coupled with in the micelle core in most cases, it could also be physically entrapped

within the hydrophobic core of the micelle [70]. Akashi and his colleagues have studied

graft copolymer nanoparticles having hydrophobic backbone and hydrophilic branches

[87]. Akiyoshi, Sunamoto and colleagues reported cholesterol-conjugated polysaccharide

aggregates for use as a drug carrier [72, 88-90]. When the drug was covalently coupled

within the hydrophobic core of the micelle, it was difficult to control the cleavage rate of

the drug linkage. Furthermore a covalent coupling method is not suitable for drugs which

have no convenient functional group. For these reasons, in this study, doxorubicin

hydrochloride was physically entrapped in hydrophobic oMMA core of the micelle. From

a practical point of view, a micellar formulation has many advantages; for example, it can

be administered not only via oral, topical and percutaneous routes but also via the

parenteral route. Moreover it has potential for site specific drug delivery by conjugating

targeting ligands onto the P AAc blocks. Furthermore, it may be bioadhesive due to the

5


PAAc block which forms the outer shell of the micelle, and therefore especially suitable

for topical or oral delivery. The synthesis, characterization and in vitro drug release

behavior of the amphiphilic oMMA-b-PAAc block copolymer as a drug carrier was

studied [91]. The AB block copolymer formed micelle, which was studied by a

fluorescence probe technique using pyrene [92, 93].

Poly(N-isopropylacrylamide) (PNIPAAm) has a lower critical solution temperature

(LCST), and exhibits a discontinuous volume phase transition in response to a small

temperature change [94-100]. PNIPAAm and its copolymers have been widely

investigated in the fields of DDS [36, 38, 40-44, 57, 101-105], diagnostics [106, 107],

bioreactors [108-118], cell culture [119, 120] and bioseparations [121-123]. Oligo(N-

vinylcaprolactam) (oVCL) also has an LCST between 30-40" C depending on the

molecular weight of the oligomer [96]. PNIPAAm has one of the sharpest volume phase

transitions among many LCST polymers [124]. Okano and his colleagues reported a

hydrogel having grafted oligomers of NIP AAm on a crosslinked PNIP AAm backbone,

and reported enhanced rate of de-swelling of the hydrogel in response to temperature

increase above the LCST [125]. Hoffman and his colleagues reported a hydrogel based

on oligomers of NIPAAm grafted to a crosslinked poly(acrylic acid) backbone and

reported that the hydrogel exhibited phase transitions over a wide range of pHs [101]. A

hydrogel having two different grafted LCST oligomers is expected to undergo two

different volume phase transition temperatures. Above the LCST of anyone of the grafted

oligomers, the collapsed chains may interact hydrophobically with each other, which may

induce shrinkage of the hydrogel [125]. Such hydrogels having two different grafted

LCST oligomers may have potential applications as transducers for information storage

and retrieval. In Chapter 4, synthesis and properties of hydrogels having grafted LCST

oligomers with two different LCSTs were studied. The hydrogels were synthesized using

macromonomers of LCST oligomers. Oligo NIP AAm (LCST = 32" C), 0 VCL (LCST =

32-40°C), and a random co-oligomer of NIPAAm and acrylamide (AAm) (o(NIPAAm-

6

Genera/Introduction

co-AAm» (LCST > 32°C) were selected as the LCST oligomers. Random copolymers of

an LCST monomer with other monomers will have an LeST depending on the

composition and environmental conditions [96]. This system also can be applicable as a

temperature-sensitive biological or chemical agent delivery system. As temperature is

raised to the first LCST, a part of the loaded agent may be released from the hydrogel,

and at the second LCST, the remaining agent may be released. This could be especially

effective, if the backbone also exhibits its own distinct LCST. One possible use could be

during hyperthermia in cancer chemotherapies [126]. In this use, the amount of anti-

cancer drug released can be controlled as temperature is increased. Macromonomers were

synthesized by reaction of carboxy-terminated LCST oligomers with glycidyl

methacrylate (GMA) [127]. To synthesize the dual LCST hydrogels, two different LCST

macromonomers were copolymerized with AAm and a crosslinker and a hydrogel having

dual LCST oligomers was obtained. The volume phase transitions of the hydrogel in

response to environmental temperature changes were investigated [128].

Incorporation of physiologically active compounds with hydrogels or immobilization

onto polymer surfaces is a critical step in a variety of applications, including controlled

drug release [37, 50, 68, 129-131], biocompatible materials [132], biosensors [133),

bioreactors [108], clinical diagnostics [100], bioseparations [134), and tissue engineering

[20]. In such applications, it is crucial to maintain the biological activity of the

immobilized compounds during processing, packaging, sterilization and end use. The

conventional processes which have been developed for combining polymeric matrices

with physiologically active compounds of low molecular weight cannot usually be applied

for immobilizing high-molecular-weight biomolecules such as enzymes, antibodies, and

growth factors, especially if the process involves a treatment with elevated temperatures

or organic solvent, which may cause denaturation of biomolecules such as proteins.

Several methods have been proposed for immobilizing protein and retaining their

biological activity. They include utilization of electrostatic attraction between the protein

7


and the polymer carrier, where they have opposite, net charge. Most physiologically

active biomolecules carry charged functional groups in their sequences, and many

researchers have employed polyion complex formation for immobilizing proteins in ionic

hydrogels [37, 50, 68, 129-131], and nucleic acids [135] or heparin [136] onto

positively charged polymer surfaces. The topic of Chapter 5 is loading of lysozyme, a

cationic protein in polyanionic phosphate hydrogels through poly ion complexation and its

ion exchange release. The properties of lysozyme were extensively investigated [137,

138], and it was studied as a model protein in a drug delivery [130, 139]. The hydrogels

were synthesized using a methacrylate monomer carrying a pendant ethylene glycol chain

terminated with a dihydrogen phosphate. Most theoretical and applied studies on anionic

hydrogels have been made using networks which carry carboxylic or sulfonic anionic

charges, whereas less attention has been paid so far to phosphate anioins, which are a

bivalent and have two pKs (2.8 and 6.4) [140, 141], a group with a higher dissociation

constant than carboxylic acid in general. Nakamae and his colleagues have reported that

phosphate-carrying hydrogels exhibit large changes in their swelling ratio with small

changes in environmental factors such as pH and salt concentration. This behavior is

characteristic of a class of materials, called "intelligent-materials" [140, 141]. Chemically-

crosslinked hydrogels were synthesized by copolymerizing the phosphate monomer with

NIPAAm and N,N'-methylene-bis-acrylamide (MBA). In addition to the pH-sensitivity,

incorporation of a large fraction of NIP AAm provides hydrogels with exhibit a

discontinuous volume phase transition in response to a small change in temperature [36,

38, 102, 103, 108, 109, 111, 121]. Such dual stimuli-sensitive polymers have been

attracted much attention [57]. In Chapter 5, however, the efficient loading of a cationic

protein, lysozyme in the anionic phosphate hydrogels and its release from the hydro gels

were focused on.

In Chapter 6, pH-sensitivity of phosphate groups of Phosmer M was focused on. The

dependence of the complex formation and dissociation on stimuli were applied to DDS.

8


On/off control of lysozyme release from the hydrogel in response to pulsatile pH or ionic

strength change was investigated. Stimuli sensitive hydrogels are designed to exhibit

reversible phase transition in response to small changes in the environmental factors,

such as temperature, pH, ionic strength, photo irradiation, concentration of organic

compounds, and the electric field [18]. Recently, many efforts have been devoted to

improve sensitivity of these hydrogels. It was shown that PNIP AAm hydrogels which

exhibit volume phase transition near physiologic temperature based on its LCST [94],

undergo more abrupt swelling when hydrophilic poly(ethylene glycol) chains are tethered

to a PNIPAAm backbone [125]. Other direction of studies on stimuli-sensitive hydrogels

involves more smart performance such as multi-sensitivity that more than two different

physicochemical parameters bring synergistically about changes in the state of hydrogels

[37,57,58, 142]. Such studies mostly employed polyelectrolytes, such as polyacrylates,

as an ionic component of the hydrogels. In contrast, less attention has been paid so far to

polyphosphates as an anionic component of networks, which is also expected to provide

pH-sensitivity with hydrogels. This is probably because easily polymerizable monomers

carrying a phosphate moiety have not been available. Since phosphates are one of the

most abundant functional group in a living body and may play important roles in the

physiology, its use in biomedical applications seems interesting. In Chapter 6, the study

in Chapter 5 [143] was extended to on/off regulated drug release. The hydrogel-lysozyme

complex is subjected to an in vitro batch release test under different pHs. The external

solution pH is chosen to be 1.4 and 7.4, according to the consideration that the pH of the

stomach varies from 1.0 to 3.5, while the intestine contents from 3.8 to 8.0 [23]. The

feasibility of the system for enteric delivery of protein drugs was discussed. The testing

mode that applies pulsatile stimuli is further employed to examine sensitivity and

reproducibility of their response. On-off regulated drug release in response to stimuli has

been paid much attention, as drugs can be delivered when it is needed and stored in the

support when it is not needed by changing environmental conditions such as temperature,

9


pH, so on [38-41, 102, 103, 105, 144]. The dependence of the complex formation and

dissociation in response to step-wise pH and ionic strength changes was applied to on-off

regulated lysozyme release [145].

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submitted.

25

PART I

NOVEL HYDROPHOBICALLY-MODIFIED POLYMERIC

CARRIERS FOR CONTROLLED DRUG DELIVERY

Chapter 1

A Hydrophobically-Modified Bioadhesive Polyelectrolyte

Hydrogel for Drug Delivery

INTRODUCTION

Hydrogels are of special interest in biological environments because they have high

water contents similar to body tissues and have, in general, high biocompatibility [1-3].

Hydrogels have been widely studied for use in medical implants, diagnostics, biosensors,

bioreactors, and bioseparators, and are being used as matrices for drug delivery systems

(DDS) [1-4]. Peppas and his colleagues have extensively investigated hydrogels as drug

release matrices [5-9]. Lee and his colleagues have also reported swelling-controlled drug

release from hydrogels [10-13]. Many researchers have studied drug release properties of

thermally-sensitive hydrogels [14-20].

Hydrophobic components have also been incorporated into hydrogels. For example,

Hoffman and his colleagues have prepared temperature-sensitive and both temperature-

and pH-sensitive hydrogels containing silicone domains [21, 22]. Others have studied

random copolymers of hydrophobic polyelectrolyte hydrogels [10, 23-26]. Siegel et al.

[24] synthesized hydrophobic polyelectrolyte random copolymer hydrogels based on

methyl methacrylate (MMA) andN,N-dimethylaminoethyl methacrylate and reported pH-

sensitive, swelling-controlled release of caffeine. Kim and Lee [10] synthesized

hydrophobic polyelectrolyte random copolymer hydrogels using MMA and methacrylic

acid and reported pH-sensitive, swelling-controlled release of oxprenolol hydrochloride.

However, there is no report in the literature of hydrophobically-modified pH-sensitive

hydrogels based on hydrophobic oligomers grafted to a polyelectrolyte hydrogel matrix.

29

Chapter 1

A hydrophobically-modified bioadhesive polyelectrolyte hydrogel with a hydrophobic

domain structure as shown in Fig. 1 was synthesized, and was applied as a bioadhesive

hydrogel for controlled release of either cationic drugs or hydrophobic drugs. The

hydrophobic domains may also strengthen the hydrogel and control the pore structure

within the hydrogel [22,27].

In this study poly(acrylic acid) (PAAc) was used as the bioadhesive polyelectrolyte.

The bioadhesive properties of P AAc [28] should enhance the utility of this graft

copolymer as a drug delivery vehicle. It is well known that P AAc has been used clinically

as a lightly crosslinked hydrogel known as "Polycarbophil" [29] or "Carbopol" [30], for

drug delivery. Oligo(methyl methacrylate) (oMMA), a polymer which is well-tolerated as

an implant material [31, 32], was selected as the hydrophobic component to be grafted to

the backbone of a lightly crosslinked PAAc hydrogel. PMMA has been used as contact

lenses, intraocular lenses [31], and bone cements [32]. The swelling behavior of the

hydrogel and drug release profiles from the hydrogel were investigated using six model

drugs.

MATERIALS AND METHODS

Materials

Methyl methacrylate (MMA), acrylic acid (AAc) , and N,N -dimethylformamide

(DMF), ACS reagent grade, were purchased from Aldrich Chemical Company, Inc.,

Milwaukee, WI, and were used after distillation under reduced pressure. 2,2'-

Azobisisobutyronitrile (AIBN) and 2-amino ethanethiol hydrochloride (AE1) were

purchased from Aldrich Chemical Company, Inc., and were used after recrystallization

with methanol. Ethyleneglycol dimethacrylate (EGDMA), dicyclohexyl carbodiimide

(DCC), and ammonium persulfate (APS) were purchased from Aldrich Chemical

30

A HydrophobicaUy-Modified Polyelectrolyte Hydrogel

hydrophobic domains (oligo(methyl methacrylate), oMMA)

bioadhesive polyelectrolyte hydrogel (poly(acrylic acid), P AAc)

Fig. 1. The schematic structure of a hydrophobically-modified polyelectrolyte hydrogel

in aqueous media.

31

Chapter 1

Company, Inc., and were used as received. Tetrahydrofuran, HPLC grade was

purchased from Aldrich Chemical Company, Inc., and was used as received.

Poly(methyl methacrylate) standards were purchased from American Polymer Standards

Corporation, Mentor, OH, and used as received. Theophylline and indomethacin were

purchased from Fluka Chemika-BioChemika, Buchs, Switzerland, and used as received.

Caffeine, propranolol hydrochloride, and ~-estradiol were purchased from Aldrich

Chemical Company and used as received. Lysozyme (from chicken egg white, L-6876)

was purchased from Sigma Chemical Company, St. Louis, MO, and used as received.

All other chemicals were of reagent grade and were used as received.

Preparation of amino-terminated oMMA

Amino-terminated methyl methacrylate oligomer (oMMA) was prepared by free radical

polymerization of MMA using AIBN as an initiator and AET as a chain transfer reagent,

respectively [33]. The reaction was performed in a sealed ampoule at 60°C in DMF

solution. Before the reaction, the reaction mixture was subjected to repeated freeze thaw

cycles using liquid nitrogen as a coolant. When the reaction mixture was frozen, the

ampoule was degassed using a vacuum pump. Finally, the reaction was carried out under

vacuum. After the reaction, the oMMA was precipitated by water. The precipitate was

filtered and washed with water, then dried in vacuum. Oligo MMA with different

molecular weights were prepared by changing the mole ratio of the chain transfer agent to

the monomer from 0.02 to 0.08.

Characterization of amino-terminated oMMA

The molecular weight of the amino-terminated oMMA was determined by gel

permeation chromatography (GPC) using poly(methyl methacrylate)s standards. GPC

32

A Hydrophobically-Modified Polyelectrolyte Hydrogel

was carried out using a Waters 501 type pump, Millipore Corporation, Milford, MA, at a

flow rate of 0.7 ml/min at room temperature with a Waters Ultrastyragcl 1 O~ A column

(7.8 x 300 mm), a Waters Ultrastyragel 10-' A column (7.8 x 300 mm) and a Waters

Ultrastyragel500 A column (7.8 x 300 mm) connected in series, in tetrahydrofuran. The

polymers were detected by refraction index with a Waters 410 Differential Refractometer.

The sample solution of 50 f..tl was injected to the GPC system using a Waters U-6K

Universal Injector, Millipore Corporation.

Preparation of PAAc hydrogel

PAAc hydrogel was prepared by co-polymerization of AAc with added EGDMA as a

crosslinking reagent (0.5w/w% to the monomer) and APS as an initiator (O.4w/w% to the

monomer) in a 1.5 mm thickness glass chamber with a 30w/v% aqueous solution. The

reaction mixture was deoxygenated by bubbling nitrogen gas. The reaction was

performed in the sealed chamber at 60°C for 18 hours, followed by cooling to room

temperature. The hydrogel was separated from the glass chamber and immersed in

ethanol for three days in order to remove impurities, and was cut into 15 mm diameter

discs using a cork borer. The disc was air dried for three days and completely dried in

vacuum at 40°C.

Preparation of oMMA-grafted-PAAc hydrogel

The oMMA-grafted-PAAc (MMA-g-AAc) hydrogel was prepared by coupling the

amino-terminated oMMA onto the P AAc hydrogel backbone through the reaction of the

amino group with an activated carboxyl group in the P AAc hydrogel using DCC as an

activation reagent [34]. The schematic reaction of PAAc hydrogel and amino-terminated

oMMA is shown in Fig. 2. The dried PAAc hydrogel was put into a screw-capped glass

33

Chapter 1

H H I I

___ .1--__ (C- C) ----

I I H C=O

I OH

PAAc hydrogel

.. H H I I

----'---(C- C) -----I I H c=o

I H-7 ~ yH 3

+ ~N=C=N- CH-CH -S-(C-C)-H

2 2 I I m H C=O

I OCH3

CH3H I I

H-(C-C)-S-CH -CH "NH I 1 m2 2 2

O=C H I

OCH3

amino-terminated oMMA

dicyclobexylcarbodiimide

MMA-g-AAc

Fig. 2. Synthesis of MMA-g-AAc hydrogel through the reaction of amino-terminated

oMMA with crosslinked P AAc.

34


vial, then a DMF solution of the amino-terminated oMMA was added and shaken for 24

hours. After the equilibrium, excess solution was removed from the vial and a DMF

solution of Dee was added. The solution was reacted for 24 hours at room temperature.

After the reaction, the solution was removed from the vial and the hydrogel was washed

with DMF followed by an ethanol/water = 80/20 (v/v) mixture. The hydrogel was air-dried for three days and then completely dried in vacuum at 40°C. The oMMA graft level

was varied from 10 to 50w/w%. All hydrogels were opaque and opacity increased with

the content of the oMMA, in either the hydrated or dehydrated state. In the hydrated state,

the strength of the hydrogels increased with the content of oMMA.

Swelling behavior of PAAc andMMA-g-AAc hydrogels

In order to investigate the swelling behavior of the hydrogels at different pHs, each

dried hydrogel was put into a 20 ml screw-capped glass vial, to which was added 10 ml

of one of the following solutions: (a) distilled water, (b) 50 mM phosphate buffered

saline, pH 7.4 (PBS), (c) 50 mM acetate buffered (potassium phosphate, dibasic and

sodium acetate) saline, pH 4.0, and (d) 1 M hydrochloric acid containing 0.15 M sodium

chloride, pH 1.0, respectively, and shaken at room temperature for 48 hours. The

hydrogels were removed from the solution periodically and blotted by wax paper and

weighed. "Swelling Ratio" was calculated by:

Swelling Ratio (w/w%) = WWel/Wdry'

where Wand Wd are the hydrogel weights in the wet and dry states, respectively. The weI ry

Swelling Ratio data are expressed as the mean value of the three determinations ±

standard deviations.

35

Chapter 1

Drug loading on PAAc and MMA-g-AAc hydrogels

The model drugs used for the release experiments were theophylline, caffeine,

indomethacin, propranolol hydrochloride, and ~-estradiol. Lysozyme was selected as a

model cationic protein drug. The chemical structures of theophylline, caffeine,

indomethacin, propranolol hydrochloride, and ~-estradiol are shown in Fig. 3. The dried

hydrogels were soaked in ethanol/water = 80/20 (v/v) solutions (theophylline, caffeine,

indomethacin, propranolol hydrochloride, or ~-estradiol) or 50 mM phosphate buffer, pH

7.4 solution (lysozyme) of the drugs and equilibrated. In the case of lysozyme, the dried

hydrogel was pre-equilibrated in 50 mM phosphate buffer, pH 7.4 prior to loading [35].

After the equilibrium, the hydrogels were removed from the vial and lyophilized at -20oe

in vacuum. The amount of the drug loaded in the hydrogels was determined by measuring

the drug concentration before and after soaking in these solutions using a

spectrophotometer (Spectronic 1001, Bausch & Lomb, Rochester, NY) at 274 nm for

theophylline and caffeine, at 270 nm for indomethacin, at 290 nm for propranolol

hydrochloride, and at 280 nm for ~-estradiol and lysozyme.

Drug release from PAAc and MMA-g-AAc hydrogels

The drug release studies were carried out in 20 rnl screw-capped glass vials. Each

dried, drug-loaded hydrogel was put into a separate vial, to which was added 10 rnl of

PBS or distilled water, and shaken by an Orbit Shaker (model 3540, Lab-line

instruments, Inc., Melrose Park, IL) at a shaking speed of 100rpm, at room temperature.

0.5 ml of the solution was collected from the vials periodically and the released drug was

determined spectrophotometrically. The volume of the release medium in the vials was

held constant by adding 0.5 ml of the fresh release medium after each sampling. The

fractional release of the drug was calculated as a function of time. All the data points are

36


theophylline o H

H3~(X~) 00 N N

I

caffeine ~CH3

H3

C'J, 1-~ 0 N N o I

CH3 CH3

propranolol hydrochloride HO HCI I

0-CH2- CH-CH -NH-CH-CH 2 I 3 ~ CH3

indomethacin

CO-O-~ CI I -

(3 -estradiol

N CH3

HO

Fig. 3. Chemical structures of theophylline, caffeine, indomethacin, propranolol

hydrochloride, and ~-estradiol.

37

Chapter 1

averages of three determinations ± standard deviations.

RESULTS AND DISCUSSION

Characterization of the amino-terminated oMMA

The molecular weights of the amino-terminated oMMAs, as prepared by the chain

transfer free radical polymerization, were determined by GPC using poly(methyl

methacrylate)s as standards. The data are shown in Table 1. Three amino-terminated

oMMA's having molecular weights ranging fro 4,700 to 11,000 were synthesized. It

was found that as the ratio of the chain transfer reagent increased, molecular weight of

amino-terminated oMMA decreased, as expected [36].

Table 1 Molecular weight of amino-terminated oMMA in different method1

MMA2 : AIBN3 : A.Er pol ymerization yield

(in mole) time (hours) (w/w%)

100: 0.5 : 2 4 55

100: 0.5 : 4 6 78

100: 0.5 : 8 14 56

1 Polymerization was carried out in DMF at 60·C, [M] = 30w/v%.

2 methyl methacrylate

3 2,2' -azobisisobutyronitrile

4 2-amino ethanethiol hydrochloride

5 Determined by GPC, PMMAs as standards

38

11,000 2.4

6,600 2.0

4,700 1.8


Swelling behavior of the PAAc and the MMA-g-AAc hydrogeL 'I

The swelling behavior of the P AAc and the MMA-g-AAc hydrogels of different graft

levels was investigated in (a) distilled water, (b) PBS (pH 7.4), (c) 50 mM acetate

buffered saline (pH 4.0), and (d) 1 M hydrochloric acid containing 0.15 M sodium

chloride (pH 1.0). The swelling ratios of the hydrogels as a function of time are shown in

Fig. 4. The molecular weight of the grafted oMMA used in these hydrogels was 6,600.

The P AAc hydrogels after the oMMA modification showed different swelling behaviors.

It was found that the greater the graft level, the slower the swelling as well as the lower

the extent of swelling in distilled water, which is expected, due to the higher amount of

hydrophobic component being introduced into the hydrogel. The same phenomenon was

found at the different pHs, as expected, since the hydrophobic interactions should be

relatively independent of pH, at either high or low pHs.

The swelling behavior of the 50w/w% MMA-g-AAc hydrogel of different oMMA

molecular weights was investigated in the same four solutions. The swelling ratios of the

hydrogels as a function of time are shown in Fig. 5. With the same graft level, more rapid

swelling and higher extents of swelling were found for the hydrogels modified with

higher molecular weights of oMMA, which might be due to fewer hydrophobic domains

in the hydrogel. The same phenomenon was found for the hydrogels at either high or low

pHs. From these results, it would be expected that the MMA-g-AAc hydrogels could be

used as matrices for delivery of either cationic drugs or hydrophobic drugs, and the rate

of drug release from the hydrogels could be controlled by controlling the rate of swelling.

This rate of swelling should be directly related to both the graft level and the molecular

weight of the oMMA being grafted.

39

Chapter 1

(a) distilled water (b) PBS (pH 7.4)

_300 _100 ... C' -~250 ~ 80 ~ ~

~200 ~ '-' Q .S 60 .~ 150 .... ~ cr:: cr:: 40 ~100 ~ :5 a:i a:i ~ ~

00 00

0 10 20 30 40 50 0 10 20 30 40 Time (hours) Time (hours)

(c) McIlvaine butTer (pH 4.0) (d) IN hydrochloric acid

~ 50 _ 10 C' ... -~ 40 ~

~ ~

8 ~ ~

~ 30 '-'

6 Q .-.... .... ~ ~

cr:: 20 cr:: ~ ~

:S .5 -a:i a:i ~ ~

00 00 0

0 10 20 30 40 50 0 10 20 30 40 Time (hours) Time (hours)

Fig. 4. The swelling behavior of PAAc and MMA-g-AAc hydrogels of different graft

level. The graft levels were Ow/w% (0), lOw/w% (e), 30w/w% (.&.), and 50w/w% (II).

The molecular weight of the oMMA was 6,600. The dried hydrogels were put into 20 ml

glass vial and added 10 ml of (a) distilled water, (b) 50 mM phosphate buffered saline,

pH 7.4 (PBS), (c) 50 mM acetate buffered saline, pH 4.0, and (d) 1 M hydrochloric acid

containing 0.15 M sodium chloride, pH 1.0, and shaken for 48 hours. The weight of the

hydrogels were measured periodically after blotting. The data are expressed as the mean

value of the three samples ± standard deviations.

40

50

50


(a) distilled water (b) PBS (pH 7.4)

~ 50 ~ 30 ~ ~ .. .. ~ 40 "0 ~ Qj Qj

~ ~ 20 :s 30 0 .... ;: = = =: 20 =: OIJ .5 10 = .- --~ ~ ~ ~ ~ ~

0 10 20 30 40 50 0 0 10 20 30 40 50

Time (hours) Time (hours)

(c) McIlvaine butTer (pH 4.0) (d) IN hydrochloric acid

~ 20 ~ 5 ~ ~ .. ..

"0 ~ ~ 15 4 Qj

~ ~ '-' '-' 0 3 .S .- a ... -~1O .... = ... --IIL & =: =: 2 fiT T T T T OIJ OIJ .5 :§ - Ii ~ ~ ~ ~ ~ ~

0 10 20 30 40 50 0 10 20 30 40

Time (hours) Time (hours)

Fig. 5. The swelling behavior of MMA-g-AAc hydrogels of different molecular weight

of oMMA. The molecular weights were 11,000 (e), 6,600 (A), and 4,700 (->. The graft

level of the oMMA was 50w/w%. The dried hydrogel were put into 20 ml glass vial and

added 10 ml of (a) distilled water, (b) 50 mM phosphate buffered saline, pH 7.4 (PBS),

(c) 50 mM acetate buffered saline, pH 4.0, and (d) 1 M hydrochloric acid containing 0.15

M sodium chloride, pH 1.0, and shaken for 48 hours. The weight of the hydrogels were

measured periodically after blotting. The data are expressed as the mean value of the three

samples ± standard deviations.

41

50

Chapter 1

Drug release from the PAAc and the MMA-g-AAc hydrogels in PBS

The drug release profiles in PBS at room temperature from the P AAc and the SOw Iw%

MMA-g-AAc hydrogels were investigated using theophylline, caffeine, propranolol

hydrochloride, indomethacin, and l3-estradiol as model drugs. The fractional release of

these model drugs from the hydrogels as a function of time is shown in Fig. 6. In the

case of hydrophilic drugs, such as theophylline, the drug release rate was enhanced by

oMMA grafting. The formation of hydrophobic domains could create large pore sizes in

the hydrophilic portions of the hydrogel, permitting hydrophilic drug to diffuse out more

rapidly [22, 27]. Caffeine showed a similar release profile to theophylline. In contrast,

the release rate of the moderately hydrophobic drug, propranolol hydrochloride, was

slowed down by the oMMA modification. This may be due to the hydrophobic

interactions between the oMMA domains and the drug. Indomethacin showed a similar

release profile to propranolol hydrochloride. In the case of a very hydrophobic drug, such

as l3-estradiol, no drug was released from either the P AAc or the MMA-g-AAc hydrogel,

probably due to a combination of its low solubility in the release medium and its high

affinity for the hydrophobic domains which are presumably too few to form an

interconnected network within the hydrogel. Based on these results, it is concluded that

drugs with a combination of moderate hydrophilicity and hydrophobicity are most

suitable for delivery from this "hybrid" hydrogel system. Therefore propranolol

hydrochloride was selected for further study.

The effect of the graft level on the release of propranolol hydrochloride from the

hydrogel was investigated. The fractional release of propranolol hydrochloride from the

hydrogels as a function of time is shown in Fig. 7. As the graft level of the oMMA

increases from 0 to 30w/w%, the total amount of propranolol hydrochloride released at

long times decreased. From this result, it seems that the hydrophobicity of the hydrogel

dominates the release of the propranolol hydrochloride.

42

1.0

0.8 8 ~ 0.6 S ~ 0.4

0.2

o o 5

A Hydrophobically-Modijied Polyelectrolyte Hydrogel

10 15 20 25 Time (hours)

Fig. 6. The drug release profile from PAAc (open symbols) and MMA-g-AAc (50w/w%

graft level) (closed symbols) hydrogels in 50 mM phosphate buffered saline, pH 7.4

(PBS). The drugs were theophylline (0, e), propranolol hydrochloride (.60, A), and ~estradiol (0, ~. The molecular weight of the oMMA was 6,600. The dried, drug-loaded

hydrogels were put into 20 rnl glass vials, 10 rnl of PBS was added and the vial was

shaken for 24 hours. The drug released was measured spectrophotometrically at 274 nm

for theophylline, at 290 nm for propranolol hydrochloride, and at 280 nm for ~-estradiol.

The data are expressed as the mean value of the three samples ± standard deviations.

43

Chapter 1

1.0

0.8 8 ~ 0.6 .... ~ 0.4

0.2

o o 5 10 15 20 25

Time (hours)

Fig. 7. Release of propranolol hydrochloride from PAAc and MMA-g-AAc hydrogels in

50 mM phosphate buffered saline, pH 7.4 (PBS). The graft levels were Ow/w% (0),

lOw/w% (e), 30w/w% (A), and 50w/w% (II). The molecular weight of oMMA was

6,600. The propranolol hydrochloride loaded hydrogels were put into 20 m1 glass vials.

10 ml of PBS was added to each and the vials were shaken for 24 hours. Released

propranolol hydrochloride was measured spectrophotometrically at 290 nm. The data are

expressed as the mean value of the three samples ± standard deviations.

44


Loading of lysozyme on the PAAc and the MMA-g-AAc hydrogels

This grafted hydrogel system was also applied for release of a model cationic protein.

Lysozyme was selected as the model protein [37]. Lysozyme has 18 basic amino acids

and its pI is ca. 11 [38], so at neutral pH, it is highly positively charged. The loading

level of lysozyme into the hydrogel as a function of the graft level is shown in Fig. 8.

The higher the graft level, the higher the loading level. This result suggests that in

addition to its ionic interactions with the negatively-charged PAAc backbone, lysozyme

also may have some affinity for the hydrophobic/hydrogel interfaces of the oMMA

domains and/or it may partition into larger pores caused by the oMMA hydrophobic

domains. Each of these effects might enhance lysozyme absorption into the hydrogel with

increasing graft level.

Drug release from the PAAc and the MMA-g-AAc hydrogels in distilled water

The lysozyme or propranolol hydrochloride loaded and lyophilized hydrogels were put

into distilled water and the release profiles were investigated. Distilled water is interesting

as a possible storage and delivery medium for drug-loaded hydrogel systems. The

fractional release of these two model drugs from the hydrogels as a function of time is

shown in Fig. 9. In the case of lysozyme, none was released from either the non-grafted

and the grafted hydrogels. This is probably because the highly positively charged

lysozyme is associated with the negative charges of PAAc and cannot be released by ion

exchange in distilled water. Contrary to lysozyme, in the case of propranolol

hydrochloride, about 80% of the loaded drug comes out from the non-grafted hydrogel,

while only about 20% of the loaded drug comes out from the grafted hydrogel. Since

both propranolol samples lose 20% of the stored drug as a "burst", it is assumed that this

drug is probably free drug accumulated at the surface during drying. These results

45

Chapter 1

50 ,.-..

~ 40 '-' -~ ~ 30 ~ ~ 0.1)

= 20 .... "0 = 0 ~ 10

o------~----------------~--------~ o 10 20 30 40 50

Graft Level of oMMA (%)

Fig. 8. Effect of graft level of oMMA on loading ratio of lysozyme on P AAc and MMA-

g-AAc hydrogels. The molecular weight of the oMMA was 6,600. The hydrogels were

pre-swollen by using 50 mM phosphate buffer, pH 7.4 (PB) and put in PB solution of

lysozyme. The amount of the lysozyme loaded on the hydrogels was determined by

measuring the lysozyme concentration before and after soaking the hydrogels in the

solution by spectrophotometer at 280 nm. The data are expressed as the mean value of the

three samples ± standard deviations.

46

A Hydrophobically-Modijied Polyelectrolyte Hydrogel

100

,-- 80 ~ '-" OJ)

60 = "-~ ~ 0 40 ~ \IJ e'd ~

20 • • - • ~ =: 0 0 20 40 60 80

Time (bours)

Fig. 9. Drug release from PAAc (open symbols) and MMA-g-AAc (30w/w% graft level)

(closed symbols) hydrogels in distilled water. The drugs were propranolol hydrochloride

(0, e), and lysozyme (~, ~). The molecular weight of the oMMA was 6,600. The drug

loaded hydrogels were put into 20 rnl glass vial and added 10 rnl of distilled water and

shaken for 72 hours. The drug released were measured by spectrophotometer at 290 nm

for propranolol hydrochloride and at 280 nm for lysozyme. The data are expressed as the

mean value of the three samples ± standard deviations.

47

Chapter 1

suggest (1) that the ionic interaction of propranolol hydrochloride (which has one

secondary amine group) with the network is much weaker than that of the polycationic

lysozyme, as one would expect, and (2) that hydrophobic interactions seem to be the

major force retaining propranolol hydrochloride in the hydrogel in distilled water.

Lysozyme release from the PAAc and the MMA-g-AAc hydrogels in PBS

The dried lysozyme-loaded hydrogel was put into PBS, and the release of lysozyme

was investigated. The fractional release of lysozyme as a function of time is shown in

Fig. 10. Lysozyme is readily released from the hydrogels in PBS. Since it was not

released in distilled water (Fig. 9), it is probable that lysozyme release is via an ionic

exchange mechanism. It is also noteworthy that the higher the graft level of the oMMA,

the slower the release (Fig. 10). Therefore either the hydrophobic domain interfaces

and/or change of pore structure caused by the oMMA domain formation seem to slow

lysozyme release from the hydrogel by the exchange of Na+ ions with the lysozyme

cations.

Kim and his colleagues synthesized albumin-heparin microspheres and reported ion-

exchange lysozyme delivery from the microspheres [37]. They concluded that the release

was independent of diffusion, SInce the rate determining step was an

adsorption/desorption process, and that low release of the lysozyme from the

micro spheres was observed in distilled water, consistent with an ion-exchange release

mechanism.

CONCLUSIONS

New oMMA hydrophobic graft copolymers on a PAAc bioadhesive, polyelectrolyte

backbone have been synthesized. Swelling of MMA-g-AAc hydrogels depends on the

48

1.0

0.8 8 ~ 0.6

~ 0.4

o

A HydrophobicaUy-ModiJied Polyelectrolyte Hydrogel

20 40 Time (hours)

60 80

Fig. 10. Release of lysozyme from PAAc and MMA-g-AAc hydrogels in 50 mM

phosphate buffered saline, pH 7.4 (PBS). The graft levels were Ow/w% (0), lOw/w%

(e), 30w/w% (A), and 50w/w% (II). The molecular weight of oMMA was 6,600. The

lysozyme loaded hydrogels were put into 20 ml glass vial and added 10 ml of PBS and

shaken for 96 hours. Released lysozyme was measured by spectrophotometer at 280 nm.

The data are expressed as the mean value of the three samples ± standard deviations.

49

Chapter 1

graft level and molecular weight of the oMMA. Higher graft levels and lower molecular

weights of the grafted oMMA chains yield lower swelling ratios. Hydrophobic drugs and

a positively charged protein, lysozyme are more slowly released from the MMA-g-AAc

hydro gels compared to more rapid release from the ungrafted P AAc hydrogel. The

hydrophobic drugs may associate with the hydrophobic domains. Positively charged

lysozyme may associate both with the negative charges of AAc and the hydrophobic

domains. Therefore both of these drug types release slowly. On the other hand,

uncharged hydrophilic drugs release faster because of their lower hydrophobic or ionic

interactions within the hydrogel.

REFERENCES

[1] W. R. Gombotz and A. S. Hoffman, Immobilization of biomolecules and cells on

and within synthetic polymeric hydrogels, in: N. A. Peppas (Ed.), Hydrogels in

Medicine and Pharmacy, vol. 1, CRC Press, Boca Raton, FL, 1986, pp. 95-126.

[2] N. A. Peppas and R.W. Korsmeyer, Dynamically swelling hydrogels in controlled

release applications, in: N. A. Peppas (Ed.), Hydrogels in Medicine and Pharmacy,

vol. 3, CRC Press, Boca Raton, FL, 1987, pp. 109-135.

[3] A. S. Hoffman, Conventional and environmentally-sensitive hydrogels for

medical and industrial uses: a review paper, in: D. DeRoss, K. Kajiwara, Y. Osada

and A. Yamauchi (Eds.), Polymer Gels, Plenum Press, New York, NY, 1991, pp.

289-297.

[4] P. Colombo, P. L. Catellani, N. A. Peppas, L. Maggi and U. Conte, Swelling

characteristics of hydroph

kobe university repository : thesis · have also been incorporated into hydrogels. for example,...

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