MULTIFUNCTIONAL MEDICAL DEVICES BASED ON PH-SENSITIVE

HYDROGELS FOR CONTROLLED DRUG DELIVERY

DISSERTATION

Presented in Partial Fulfillment of the Requirements

for the Degree of Doctor of Philosophy in the

Graduate School of The Ohio State University

By

Hongyan He, M.S.

The Ohio State University

2006

Doctor’s Examination Committee: Approved by Professor L. James Lee, Adviser Professor Kurt W. Koelling ______________________________ Professor Robert J. Lee Adviser

Graduate Program in Chemical Engineering

ii

ABSTRACT

Hydrogels are a desired material for biomedical and pharmaceutical applications

due to their unique swelling properties and highly hydrated structure. To better control

the synthesized hydrogels for various applications, it is necessary to have a thorough

understanding of hydrogel structure and reaction mechanism. In this study, pH-sensitive

hydrogel networks consisting of methacrylic acid (MAA) crosslinked with tri(ethylene

glycol) dimethacrylate (TEGDMA) were synthesized by free-radical

photopolymerization in the water/ethanol mixture with different ratios under various light

intensity. Reaction rate was measured using Photo-Differential Scanning Calorimetry

(PhotoDSC) with a modified sample pan designed for handling volatile reagents. A

photo-rheometer and a dynamic light scattering (DLS) goniometer were used to follow

the changes in viscosity and molecule size of the resin system during

photopolymerization. It was found that the rate of polymerization increased and more

compact and less swelling gels would form with a higher water fraction in 50wt%

solvent/reactant mixture. This is because the weaker interaction between MAA and

solvent gives a higher opportunity for propagation and a higher reaction rate. The

hydrophobic TEGDMA and initiator tend to form aggregates in the solution with a higher

water content, contributing to the inhomogeneous microgel formation. It was also noted

that the rate of polymerization and the MAA conversion were enhanced as the light

intensity increased. However, at too high a light intensity, an adverse effect was observed

iii

and the final conversion of MAA decreased to 43% at 24 mw/cm2. The optimal light

intensity was about 2.0 mw/cm2 to get the PMAA gels with low residue monomers. The

use of the high light intensity significantly shortened the reaction time to reach the

macro-gelation and increased the swelling ratio of formed hydrogels, which can be

explained by the mechanism of intra- and intermolecular reaction.

By using the desired functional hydrogels, several drug delivery systems were

developed based on the selected integration of a number of micro-manufacturing modules

such as soft-lithography, micro-imprinting, and polymer self-folding, to achieve

multi-functionalities such as drug protection, self-regulated oscillatory release, enhanced

mucoadhesion, and targeted unidirectional release. To evaluate the device performance,

adhesion measurement, dynamic flow testing, and targeted unidirectional release were

conducted for trans-luminal delivery of two model drugs, acid orange 8 and bovine serum

albumin. The self-folding device first attached to the mucosal surface and then curled into

the mucus, leading to enhanced mucoadhesion in the mode of “grabbing”. Furthermore,

the folded layer served as a diffusion barrier, minimizing the drug leakage in the small

intestine. The resulting unidirectional release provides improved drug transport through

the mucosal epithelium due to localized high drug concentration. The functionalities of

the devices have been successfully demonstrated in vitro using a porcine small intestine.

The novel delivery devices will be of great benefit to the advancement of oral

administration of proteins and DNAs. Since the mucus layer covers many tissues at other

specific sites, the devices may be applied for ocular, buccal, vaginal and rectal

administrations. The polymer self-folding at the microscale can also be applied as probe

arrays for bio/chemical sensing, carriers in cell-based bioreactors, and tissue clamping.

iv

This dissertation is dedicated

to

my parents

v

ACKNOWLEDGMENTS

I would like to express my great appreciation to my adviser, Dr. L. James Lee,

for his inspiring guidance, encouragement, and support throughout this work. I would

also like to acknowledge with sincere gratitude to the members of my dissertation and

candidacy exam committee, Dr. Kurt W. Koelling, Dr. Robert J. Lee, and Dr. James F.

Rathman for their valuable suggestions and comments on my work.

My gratitude is also expressed to Dr. Paula Stevenson, Paul Green, Karl Scott,

and Leigh Evrard for their great help in my research work. Special thanks go to my

fellow colleagues Dr. Xia Cao, Dr. Jingjiao Guan, and all other polymer research group

members, for their invaluable help and technical support.

Finally, I would like to thank my parents for their forever support through the

years of my study and my husband, Zhaohui Ning, for his understanding, support, and

encouragement.

vi

VITA

January 2nd, 1974........................................................Born - Taiyuan, Shanxi, P. R. China

September 1992−July 1997..........................................B.S. Chemical Engineering Tsinghua University Beijing, P. R. China

September 1997−March 2000......................................M.S. Environmental Engineering Shanghai University Shanghai, P. R. China

September 2000−December 2004.............................…Graduate Research Associate The Ohio State University Columbus, OH

June 2005−present.............................…........................Presidential Fellow The Ohio State University Columbus, OH

PUBLICATIONS

1. H. He, L. Li and L. J. Lee, “Photopolymerization and structure formation of

methacrylic acid based hydrogels in water/ethanol mixture”, Polymers, 47, 1612-1619,

2006.

2. H. He, J. Guan, and L.J. Lee, “Oral Delivery Devices Based on Self-folding

Hydrogels”, Journal of Controlled Release, 110(2), 339-346, 2006.

3. J. Guan, H. He, D.J. Hansford and L. J. Lee, “Self-folding Hydrogel

Three-Dimensional Microstructures”, Journal of Physics Chemistry B, 109(49),

23134-23137, 2005.

vii

4. H. He, J. Guan, D.J. Hansford and L.J. Lee, “Hydrogel-Based Multifunctional

Delivery Devices for Oral Protein Administration”, Abstracts of Papers PMSE-016, 229th

ACS National Meeting, San Diego, CA, March 13-17, 2005.

5. H. He, J. Guan, D.J. Hansford and L.J. Lee, “Hydrogel-Based Multifunctional

Delivery Devices for Oral Protein Administration”, Polymeric Materials: Science and

Engineering, 92, 28-30, 2005.

6. H. He and L. J. Lee, “Poly(lactic-co-glycolic Acid) and Functional Hydrogels for

Drug Delivery Applications”, Proceedings of Society of Plastics Engineers, 62(3),

3356-3360, 2004.

7. H. He, X. Cao and L. J. Lee, “Design of a Novel Hydrogel-based Intelligent

System for Controlled Drug Release”, Journal of Controlled Release, 95, 391-402, 2004.

8. H. He and J. Wei, “Synthesis and Properties of Modified Melamine Resin”,

Shanghai Huanjing Kexue, 19(9), 432-433, 2000.

9. H. He, J. Wei and G. Zhang, “Synthesis of Modified Melamine-Formaldehyde

Resin and Property Investigation as a Flocculent”, Shanghai Daxue Xuebao, V3, 2000.

10. H. He, “Synthesis of Modified Melamine-Formaldehyde Resin and Property

Investigation”, Master Thesis, Shanghai University, China, 2000.

11. H. He, “The Extraction of Glycin from Proteins”, Bachelor Thesis, Tsinghua

University, China, 1997.

FIELDS OF STUDY

Major Field: Chemical Engineering

Minor Field: Polymer Engineering

viii

TABLE OF CONTENTS

Page

Abstract………………………………………………………………………………...... ii Acknowledgments…………………………………………….………………….…........ v Vita……………………………………………………………………………….……….vi Table of contents………………………………………………………………..……….viii List of tables……………………………………………………………….…………….xii List of figures…………………………………………………………………………....xiii Chapters:

1. Introduction and motivation ……………………………………….………………. 1

2. Literature review…………………………………………………………………….8

2.1 Overview of pH-sensitive hydrogels………………….………………...…..8

2.1.1 Anionic hydrogels …………………………..…………………….10

2.1.2 Cationic hydrogels ……………………………….……………….12

2.2 Temperature-sensitive hydrogels ……………………………………….…14

2.2.1 Negatively temperature-sensitive gels ……………………………15

2.2.2 Positively temperature-sensitive gels ……….…………………….19

2.3 Properties of hydrogels……………………………….……………………20

2.3.1 Swelling properties ……………………………………………….20

2.3.2 Network structure and characterization ………………………….22

ix

2.3.3 Mechanical properties …………………………………………….30

2.4 Application of hydrogels in drug delivery ………………………………...33

2.4.1 Peroral drug delivery ………………………………………..…….34

2.4.1.1 Buccal route…………………………………..…………34

2.4.1.2 Gastrointestinal route……………………………………36

2.4.2 Nasal route …………………………………………………….….42

2.4.3 Ocular route ……………………………………………………….43

2.4.4 Rectal and vaginal routes ………………………………………....44

2.4.5 Transdermal route …………………………………………….…...45

2.4.6 Trends and perspectives……………………………………….…...46

3. Photopolymerization and structure formation of PMAA hydrogels in water/ethanol

mixture. ……………………………………………………………………….…………49

3.1 Introduction………………………………………………………………...50

3.2 Experimental……………………………………………………………….52

3.2.1 Materials and sample preparation…………..…………………….52

3.2.2 Modification of DSC pans …………………….…………….…...53

3.2.3 PhotoDSC measurement …………………………………….…..55

3.2.4 Rheological measurement……………………………………..…55

3.2.5 Dynamic light scattering analysis……………………………..…56

3.2.6 Swelling studies…………………………...…………………..…57

3.2.7 Scanning electron microscopy characterization…...…………..…57

3.3 Results and discussions…………………………………...………………..58

3.3.1 Kinetics of MAA/TEGDMA photopolymerization …...……..….58

3.3.2 Viscosity measurement and molecule size analysis …………..…63

3.3.3 Mechanism for gelation ……………………………..……..……67

3.3.4 Swelling ratio and structural characterization………….………...72

3.4 Conclusions………………………………………………………..…………77

x

4. Photopolymerization and structure formation of PMAA hydrogels cured under

various light intensities…………………………………………………...……….…….78

4.1 Introduction ………………………………….……………………….……79

4.2 Experimental …………………………………………….……………...…81

4.2.1 Materials and sample preparation………………..……………….81

4.2.2 PhotoDSC measurement ……………………….…………….…..82

4.2.3 Rheological measurement……………………….……………..…83

4.2.4 Dynamic light scattering analysis……………….……………..…83

4.2.5 Swelling studies………………………………………………..…84

4.3. Results and discussion…………………………………………………......84

4.3.1 Kinetics of MAA/TEGDMA photopolymerization ...………..….84

4.3.2 Viscosity measurement ………………………...……………..…89

4.3.3 Kinetic parameters ………………………………………..……..92

4.3.4 Molecular size analysis ………………………………………….97

4.3.5 Integrated analysis…………………………………………….....99

4.4 Conclusions………………………………………….…………………...106

5. Design of smart devices based on the functional hydrogels…….………….…….107

5.1 Introduction …………………………………………...…………….……108

5.2 Experimental …………………………………………………...……...…111

5.2.1 Materials……………………………………………….………..111

5.2.2 Device design and drug loading ………………………….……..113

5.2.3 In vitro drug release ………………………….………….……..115

5.2.4 Diffusion studies .. ……………….………………….…………116

5.2.5 Targeted unidirectional release……………….………….……..116

5.3. Results and discussion………………………………………..………......119

5.3.1 Swelling properties of hydrogels ………….……………………119

5.3.2 Model drug release from entrapped devices ……………….…...121

xi

5.3.3 Diffusion studies ……………………………………………..…126

5.3.4 Model drug release from assembled devices………..………..…130

5.3.4.1 Drug protection……………………….………………130

5.3.4.2 Self-regulated oscillatory release …….………………137

5.3.4.3 Targeted unidirectional release ………………………137

5.4 Conclusions……………………………………………………..………...141

6. An oral delivery device based on the self-folding hydrogels…………..………...142

6.1 Introduction………………………………………………….……...……143

6.2 Experimental………………………………………………..………...…..144

6.2.1 Materials………………………………………...……………...144

6.2.2 Device design and fabrication ………………..……………..….145

6.2.3 Swelling and self-folding studies ………………………………150

6.2.4 Mucoadhesion measurement ……………...…………………....151

6.2.5 Delivery performance...........................……...............................153

6.3 Results and discussion……………………………….…………………...154

6.3.1 Swelling and self-folding studies ………………………...…..…154

6.3.2 Mucoadhesion measurement ……………………………………158

6.3.3 Delivery performance ………………………………………..…165

6.4 Conclusions…………………………………………………………….…169

7. Conclusions and recommendations…………………………………………….…170

7.1 Conclusions………………………………………………………….……170

7.2 Recommendations…………………………………………………….…..172

References………………………………………………………………………………177

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LIST OF TABLES

Table Page

5.1 Physical properties of model drugs.............................................….....................112

5.2 Permeability and diffusion coefficient

of model drugs through different membranes......................................................129

xiii

LIST OF FIGURES

Figure Page

1.1 The engineering process applied for pH-sensitive hydrogels ………………....….6

2.1 Structures of anionic pH-sensitive hydrogels …………………………….….….10

2.2 Structures of negative temperature-sensitive hydrogels ………………….….….15

3.1 (A) DSC pan treated with PDMS; (B) Seal of DSC pan…………………..…….54

3.2 Comparison of PhotoDSC measurements by using

a modified and an un-modified pan at UV intensity

of 2.0 mw/cm2 in the MAA/TEGDMA system

(1.0 mole%TEGDMA, 50 wt.% solvent

mixture of the 1/1 water/ethanol ratio) ………………………………………….60

3.3 (A) Reaction rate and (B) conversion

versus reaction time for the isothermal

photopolymerization of MAA/TEGDMA

mole%TEGDMA, 50 wt.% solvent)

with different solvent compositions

at 30ºC and UV intensity of 2.0 mW/cm2………………………………..…..…..62

xiv

3.4 (A) Reaction rate and viscosity

rise as a function of conversion of

MAA/TEGDMA (mole% TEGDMA,

50 wt.% solvent) with different solvent

compositions cured at UVintensity of

2.0 mW/cm2, (B) Gel time and gel conversion

versus water/ethanol ratio in the solvent mixture ……………………….…..…..65

3.5 The size distribution of MAA/TEGDMA

resin (1.0 %TEGDMA, 50 wt.% solvent)

with different solvent ratios of water/ethanol:

(A) 1/4 and (B) 9/1 cured

at light intensity of 2.0 mW/cm2…………………….……………………..….....66

3.6 The schematic diagram of structure formation

of MAA/TEGDMA with different solvent qualities …………………..………...68

3.7 The size distribution of MAA/TEGDMA monomer solution

(1.0 %TEGDMA, 50 wt.% solvent) with different compositions ………………70

3.8 Equilibrium swelling ratios of the PMAA

(1.0 mole% TEGDMA) hydrogels with

different solvent ratios as a function of pH values …………………...……..…..74

3.9 SEM micrograph of swollen PMAA hydrogels

(1.0 mole% TEGDMA, 50 wt.% solvent)

with different swelling ratios (SR) in pH=7.4

buffer solution: (A) 9/1 and (B) 1/4……………………………………….…..…75

3.10 SEM micrograph of swollen PMAA hydrogels

xv

(1.0 mole% TEGDMA, 50 wt.% solvent) with

the same swelling ratio (SR=4.3) in different

buffer solution: (A) 9/1 in pH=6.2 buffer (B) 1/4 in pH=3.0 buffer.…....….……76

4.1 Reaction rate vs. conversion of MAA/TEGDMA

in the presence of 1% Irgacure 651 with 50 and

100 wt.% monomer content cured under 5.0 mw/cm2.…………………….…….86

4.2 Effect of light intensity on the polymerization

of MAA/TEGDMA system in the presence of

1% Irgacure 651 (A) reaction rate, (B) conversion ……….……………….…….88

4.3 Reaction rate and relative viscosity rise

as a function of conversion of MAA/TEGDMA

(1.0 mole% TEGDMA, 50 wt.% solvent)

cured under different light intensity:

(A) 0.25 and 2.0 mW/cm2, (B) 24 mW/cm2 ……………………..…………..…..90

4.4 Gel conversion versus light intensity

for polymerization of MAA/TEGDMA system

(1.0 mole% TEGDMA, 50 wt.% solvent)

in the presence of 1% Irgacure 651………………………………..………..…....91

4.5 Conversion dependence of the rate constant

of propagation pk and termination tk

for the polymerization of MAA/TEGDMA system at 2.0 mw/cm2………...…...95

4.6 Conversion dependence of the rate constant

of propagation pk and termination tk for the polymerization

of MAA/TEGDMA system at 24 mw/cm2.……………………..…………..…...96

xvi

4.7 The molecular size distribution of the MAA/TEGDMA

system (1.0% TEGDMA, 50 wt.% solvent)

cured at (A) 2.0 mw/cm2 and (B) 24 mw/cm2.……………………………..……98

4.8 Changes of reaction rate, viscosity during

the photopolymerization of MAA/TEGDMA

at light intensity of 2.0 mw/cm2: I initiation;

II microgel formation; III cluster formation;

IV macro-gelation; V post-gelation……………………….………….…….…..100

4.9 Changes of reaction rate, viscosity during

the photopolymerization of MAA/TEGDMA

at light intensity of 24 mw/cm2: I initiation;

II microgel formation; III cluster formation;

IV macro-gelation; V post-gelation……………………….………….…….…..101

4.10 Dynamic swelling behavior of the PMAA hydrogels

with 1.0% TEGDMA cured at different light intensity

and immersed in the different pH buffer solutions…………………………..…105

5.1 Schematic of the assembled device ………………………………………….…118

5.2 Dynamic swelling behavior of hydrogels. Samples

were 5.0 mm in diameter and 0.8 mm in thickness:

( ) PMAA hydrogel in pH=7.3 buffer.

( ) PMAA hydrogel in pH=3.0 buffer.

( ) PHEMA hydrogel in pH=7.3 buffer.

( ) PHEMA hydrogel in pH=3.0 buffer………………………….………..…120

xvii

5.3 Acid Orange 8 release to pH 7.3 buffer solution

from the entrapped 5.0 mm PMAA samples at 25 °C.

The samples were 0.8 mm in thickness ………………………………..………123

5.4 BSA release to pH 7.3 buffer solution from

the entrapped 5.0 mm PMAA samples at 25 °C.

The samples were 0.8 mm in thickness ……………..………………...……….124

5.5 AO8 and BSA release from the entrapped

5.0 mm PMAA samples at 25 °C. The

samples were 0.8 mm in thickness:

( ) AO8 at pH=3.0.

( ) AO8 at pH=7.3.

( ) BSA at pH=7.3.……………………………..………..……….…….……125

5.6 Permeation of AO8 and BSA through different

swollen hydrogel membranes at pH 7.3 and 25 °C.

( ) AO8 through PMAA.

( ) BSA through PMAA.

( ) AO8 through PHEMA....………...………..…………………………..….128

5.7 AO8 release from the assembled device at

pH=7.3 and 25°C. The diameter of the device

is 5.0 mm. The thickness of bilayered gate is 60 µm

and the thickness of the drug reservoir is 1.0 mm.

(A) Dry assembled device. (B) Releasing at t= 40 minutes.

(C) Released at t= 80 minutes. (D) Schematic of AO8

release from assembled device…………….…………………..…………..……132

5.8 AO8 release from the 5.0 mm assembled devices

xviii

with different gates at pH=3.0 and 25°C. The gate

thickness is 60 µm and the reservoir thickness is 1.0mm.

( ) PMAA hydrogel gate.

( ) PHEMA and PMAA bilayered gate………..………………………...…..133

5.9 AO8 and BSA release from the 5.0 mm assembled

device at 25°C. The thickness of the bilayered gate

is 60 µm and the thickness of the drug reservoir is 1.0 mm.

( ) AO8 at pH=3.0.

( ) AO8 at pH=7.3.

( ) BSA at pH=7.3 ……………………………………………………..……135

5.10 Thickness effects of the bilayered gate and reservoir

on AO8 release behavior at pH=7.3 and 25 °C.

( ) The gate thickness is 60 µm and the reservoir thickness is 0.5 mm.

( ) The gate thickness is 60 µm and the reservoir thickness is 1.0 mm.

( ) The gate thickness is 90 µm and the reservoir thickness is 0.5 mm……..136

5.11 The oscillatory release behavior of the assembled device.

The gate thickness is 50 µm and the thickness ratio for

PHEMA to PMAA layer is 4.………………...….……...….……...…....…...….138

5.12 The comparison of the targeted

uni-directional release with untargeted release:

(A) Targeted release. (B) Untargeted release …..……………….……………...140

6.1 Schematic of the 3-layer device from

(A) side view and (B) top view,

(C) folding on the small intestine surface,

(D) a capsule containing devices ……………………………………..…..……148

xix

6.2 Fabrication procedure of the miniature devices ………………………….…….149

6.3 Experimental setup for (A) flowing testing and

(B) the detachment force measurement…………………………………..…….152

6.4 Dynamic swelling behavior PMAA and PHEMA hydrogels…………..……….156

6.5 Optical graphs of a bilayered structure at dried state

(A) top view, (B) side view,

(C) a curled bilayered structure in a

buffer solution. Scale bars=2.0 mm…………………………………………….157

6.6 (A) Number of bound samples and

(B) residence time for different samples

attached to intestinal mucus in the flow test………………………..…….…….159

6.7 Dynamic processes for (A) folding behavior and

(B) enhanced mucoadhesion. Buffer pH=6.5 and 25°C…………………..……161

6.8 Compared attachments for the devices with different

contact sides in the flow test. Buffer pH=6.5 and 25°C………………..………163

6.9 The detachment force of different samples

on the small intestinal surface. Buffer pH=6.5

and 25 °C. Error bar = SD, n = 3……………………………………………….164

6.10 The fractional leakage of AO8 from the drug

reservoir with different protection layers

(thickness=20 µm) at pH=6.5 and 25°C.

xx

Error bar = SD, n = 3………………………………………………….….…….166

6.11 AO8 transport from different systems

across the mucosal epithelium at pH=6.5

and 25°C. Error bar = SD, n = 3……………………………………..…………167

6.12 BSA transport from different systems

across the mucosal epithelium at pH=6.5

and 25°C. Error bar = SD, n = 3……………………………………..…………168

7.1 Schematic of fabrication of self-foldable microdevices.…………….…………173

7.2 Schematic of the self-foldable microdevice

with enhanced nanotips………………………….…………………..….………174

1

CHAPTER 1

INTRODUCTION AND MOTIVATION

The U.S. market for advanced drug delivery technology exceeded $10 billion in

1996 and is increasing rapidly [Langer, 1998]. A primary driving force is the fact that

many protein- and DNA-based drugs exhibit high sensitivity to the surrounding

physiological conditions as a result of their delicate physicochemical characteristics and

the susceptibility to degradation by proteolytic enzymes in biological fluids. They need

to be properly protected during administration and their release needs to be precisely

targeted and controlled. Most conventional drug delivery systems are based on polymers

or lipid vesicles: diffusion of the drug species through a polymer membrane; a chemical

or enzymatic reaction leading to cleavage of the drug from the system, and solvent

activation through swelling or osmosis of the system. A major limitation of these

available delivery devices is that they cannot fully protect the drugs and release them at a

controllable rate over a long period of time. Certain disease states, such as diabetes,

heart disease, hormonal disorders, and cancer, require drug administration either

repeatedly when needed, at a high release rate during the life-threatening moment, or at a

constant release rate during a sustained period of time. Drug delivery technology can be

brought to the next level by the fabrication of ‘smart materials’ into ‘miniature devices’

that are responsive to the individual patient’s therapeutic requirements and able to deliver

- 2 -

a certain amount of a drug in response to a biological state. Such smart therapeutics

should possess one or more properties such as proper drug protection, local targeting,

precisely controlled release, self-regulated therapeutic action, permeation enhancing,

enzyme inhibiting, imaging, and reporting. This is clearly a highly challenging task and it

is difficult to add all of these functionalities in a single device. Currently, there are no

commercial products based on the miniaturized responsive drug delivery approach, and

only limited research. Such a system would also need to exhibit good biocompatibility as

drug delivery carriers [Beyssac et al., 1996; Cohen et al., 1997; Draye et al., 1997; Kitano

et al., 1998; McNeill et al., 1984; Miyazaki et al., 1998; Petelin et al., 1998].

Hydrogels are crosslinked polymeric networks that are insoluble in water but

swell to an equilibrium size in the presence of excess water or biological fluids

[Brannon-Peppas et al., 1990; Peppas et al., 1986]. Research on hydrogels started in the

1960s with a landmark paper on poly(hydroxyethyl methacrylate) [Wichterle et al., 1960].

Due to the unique swelling properties and the biocompatible structure, these materials have

been extensively studied for biomedical and pharmaceutical applications, such as contact

lenses, membranes for biosensors, linings for artificial hearts, materials for artificial skin

and drug delivery devices [Peppas et al., 1994; Walther et al., 1995; Peppas et al., 1997;

Peppas et al., 2000]. In nature, polymeric hydrogel is a three-dimensional network

comprising interconnected hydrophilic macromolecules, with an inner space partially

filled with water molecules. The highly hydrated, non-ionic and good biocompatibility

provide the ability of hydrogels to release drug in a regulated mode, which can be achieved

by controlling the synthesis conditions, such as the reactant composition, the ratio of

crosslinked density, the method of polymerization, and the external environment.

- 3 -

Hydrogels are often synthesized by UV photopolymerization [Lu et al., 1999;

Ward et al., 2001] or redox polymerization [Hassan et al., 1999]. Photopolymerization is

favored because hydrogels can be synthesized at temperatures and pH conditions near

physiological conditions and even in the presence of biologically active materials.

Furthermore, photopolymerization can be easily controlled by adjusting the dosage and

intensity of UV light, and the curing temperature. Photo-Differential Scanning

Calorimetry is the most widely used technique to characterize the photopolymerization

kinetics. A great deal of research has been carried out using this approach for

photocurable materials. However, the application of this technique for highly volatile

reagents is limited since uncovered sample pans lead to significant sample loss during

measurement. Some researchers applied unsealed polyethylene (PE) films over the

sample pan to reduce the sample loss [Ward et al., 2001], while others used the sample

weight after the reaction to correct for the measurement error resulting from reagent

evaporation [Jakubia, 2000]. The results from such treatments are doubtful because

sample loss during the reaction is a time-dependent process. When preparing the carriers

for drug delivery, solvents like water and ethanol are often used in the synthesis to control

the hydrogel structure. Evaporation of highly volatile solvents like ethanol makes it

impossible to study the reaction kinetics using the existing approaches. We have recently

developed a modified DSC sample pan [Li et al., 2005]. Sample loss during reaction is

minimized, and loaded samples are much more uniform over the sample surface. This

new method is applied in this study.

To better control the synthesized hydrogels for various applications, it is essential

to understand how the polymerization conditions, chemical structure of reactants and

- 4 -

their composition, and solvent type and concentration affect the reaction and the resulting

properties of hydrogels. A number of studies have reported that varying curing conditions

may achieve different gel structures and swelling properties [Lowman et al., 1997; Elliott

et al., 2002; Kwok et al., 2003], and the compatibility between the solvent and the resin

may affect inter-molecular and primary cyclization of multi-vinyl monomers during the

polymerization, and, consequently, the hydrogel properties. However, there lacks a

thorough understanding on the interactions of reaction kinetics, rheological changes, gel

formation, and hydrogel structures.

Oral delivery of peptides and proteins has become a challenging and attractive

task with the enormous market potential in resent years. Typically, the intramuscular or

intravenous injection is used for their administration. However, due to the disadvantages,

such as the pain, inconvenience and inconsistent pharmacokinetics for this administration,

lots of work has been done to pursue alternative administration methods other than the

conventional injection approach. Among various potential routes, oral administration

could be the most convenient and ideal route since it is known as the most desirable route

of drug administration.

Although being an ideal non-invasive route of drug administration, the peptides

and proteins delivery through the oral route is fraught with difficulties around low

bioavailability, which results from the pH fluctuation, proteolytic degradation, low

transport, and short residence time. Many possible solutions, such as the inclusion of

protection, protease inhibitors, enhancers/promoters, and/or specific adhesion, do help the

increased drug bioavailability through oral route. Typical oral delivery systems can be

summarized as two categories: conventional systems, such as tablet, capsules and syrup,

- 5 -

and advanced systems, such as micro/nanoparticles and intestinal patches. For

microparticles and nanoparticles, the loaded drugs can be released to all directions due to

their symmetric shape. Asymmetric intestinal patches and some microdevices can provide

protected unidirectional release. Dorkoosh and coworkers [Dorkoosh et al., 2001;

Dorkoosh et al., 2002] designed a novel drug delivery system for site-specific drug

delivery of peptide drugs in the intestinal tract using superporous hydrogels (SPH) and

SPH composite polymers, which swell very rapidly by absorption of gut fluids. The

system attached to the intestinal wall and provided a longer residence time for drug

release. However, only a slight decrease in blood glucose levels was observed in animal

studies. Shen et al. [2002] reported an intestinal patch design for oral delivery. A longer

residence time and unidirectional diffusion were achieved for helping drug diffusion

through the intestinal barrier by using a mucoadhesive layer of Carbopol/ pectin. Tao et al.

[2004] combined microfabrication techniques with the use of mucoadhesive plant lectins

to design a microdevice with a long residence time. iMEDD Inc. developed Oral-MEDDs

(microfabricated particles) technology [Cohen et al., 2003] which combined several oral

delivery approaches into a single drug delivery system to deliver peptides and proteins.

The mucoadhesion for these systems is through surface-to-surface contact. Due to the

continuous shedding of surface mucus, these systems have the limited residence time and

the drug bioavailability is low. To match the patients’ needs, further efforts and better

solutions are still needed.

In this work, we design multi-functional devices based on the hydrogels that can

bind to the targeted issue for self-regulated and sustained release. A common process

model for engineering is used to show how materials appear likely to break previous

- 6 -

barriers in the process that ultimately results in applications with potential benefits. This

process development can be conveniently represented by the schematic description of

pH-sensitive hydrogels for oral drug delivery systems and sensors (Figure 1.1).

Figure 1.1 The engineering process applied for pH-sensitive hydrogels.

Polymeric Materials

Device Design Modeling

Micro-fabrication

Swelling, Kinetics, Rheology

Material Characteristics

Animal Studies,Clinic Trials

Better protection, Long residence time,

High transport

Proper swelling High final conversion

Better mechanical properties

Material/Process Design to

Improve the Delivery Performance

In Vitro Release,Targeting,

Unidirectional Rel.

PolymericSelf-folding

Products/Applications(Oral DDS, Biosensor and

Bioreactors, Tissue Clamping)

Polymeric Materials

Device Design Modeling

Micro-fabrication

Swelling, Kinetics, Rheology

Material Characteristics

Animal Studies,Clinic Trials

Better protection, Long residence time,

High transport

Proper swelling High final conversion

Better mechanical properties

Material/Process Design to

Improve the Delivery Performance

In Vitro Release,Targeting,

Unidirectional Rel.

PolymericSelf-folding

Products/Applications(Oral DDS, Biosensor and

Bioreactors, Tissue Clamping)

- 7 -

Some issues need to be addressed are as follows:

(1) What factors play the important roles in the synthesis of hydrogels with desired

properties?

(2) How will the solvent ratio and light intensity affect the structure and properties of

hydrogels?

(3) How will a multi-functional DDS be designed to integrate all possible solutions to

achieve high bioavailability?

The objectives of the research are (1) to generate functional hydrogels with desired

properties, (2) to develop an intelligent DDS, which is effective for controlled release,

drug protection, targeted unidirectional release, high transport, long residence time, as

well as a quick response time, and (3) to investigate the relationship between the

hydrogel properties and the release performance, and then optimize the device design.

Thus, we can extend the functionalities of hydrogels by combining with the fabrication

technology to match physiological needs for various pharmaceutical applications.

- 8 -

CHAPTER 2

LITERATURE REVIEW

2.1 Overview of pH-Sensitive Hydrogels

Hydrogels can be classified as neutral or ionic based on the type of repeating units

or the nature of the side chains on the polymer backbone. They can be homopolymer or

copolymer networks based on the preparation approach. The most important property for

hydrogel is the stimuli-sensitivity depending on the external conditions, which include

pH, temperature, pressure, ionic strength, electromagnetic radiation, ultrasonic energy,

buffer composition, the concentration of glucose, stress and strain, and photo [Peppas,

1991]. These conditions dramatically affect the swelling behavior, network structure,

permeability and mechanical strength of hydrogels. Such intelligent materials open the

door for novel applications in the areas of nanotechnology (actuators, substrates), surgical

implants and tissue engineering, due to hydrogel’s unique ability to undergo phase

transitions under the influence of small stimuli.

The pH-sensitive hydrogels exhibit swelling or deswelling behavior with changes

of pH values in the surrounding medium. The swelling behavior may be due to one of the

following mechanisms: (1) changes in the hydrophobic-hydrophilic nature of chains; (2)

9

inter- and intramolecular complexation by hydrogen bonding, or (3) electrostatic

repulsion. All these mechanisms are closely related to the protonation phenomena of the

ionizable moieties on the polymer backbone or the side chains. In the first case,

ionization makes the hydrophobic polymer network more hydrophilic because the ionized

structure usually posses more hydrophilicity which can imbibed more water into the

matrix. In the second case, ionization results in the breaking up of the hydrogen bonds

that exist in the polymeric matrix in the unionized state, leading to the hydrogel swelling.

In the third case, the ionization provides the electrostatic repulsion among charges present

on the polymer chain to keep the chains apart and allow more water absorbing into the

loose structure. In all these cases, the kinetics of the swelling process and the equilibrium

extent of swelling are affected considerably by several factors, such as ionic strength of

the medium, buffer composition, presence of salts [Hariharan et al., 1996]. Other factors,

such as the crosslinking ratio, solvent quality, chemical structure of monomers, and

synthesized conditions also influence the structure formation and the swelling behavior of

hydrogels.

pH-sensitive hydrogels can be divided into anionic and cationic depending on

the nature of pendant groups in the networks, which show sudden or gradual changes in

their dynamic and equilibrium swelling behavior as a result of pH changes. Anionic gels

often contain carboxylic or sulfonic acid. When the pH value of surrounding medium

rises above its pKa, the ionized structure will provide increased electrostatic repulsion

between chains and the hydrophilicity of network. Under these conditions, hydrogels are

capable of uptaking large amounts of water and forming very loose structure. In contrast,

cationic hydrogels usually contain pendant group such as amines. As pH values lower

10

than the pKb, the amine groups change from NH2 to NH3+, resulting in the increased

hydrophilicity, strong electrostatic repulsion, and high swelling ratio.

2.1.1 Anionic hydrogels

Many researchers have studied the dynamic swelling of anionic pH-sensitive

hydrogels, which often contain carboxylic groups. Typical examples of such polymers

include poly(acrylic acid) (PAA) and poly(methacrylic acid) (PMAA). Copolymers of

PAA and PMAA with poly(ethylene glycol) (PEG), poly(vinyl alcohol) (PVA), and

poly(hydroxyethyl methacrylate) (PHEMA) also exhibit the pH sensitivity due to the

presence of carboxylic segment. Additionally, incorporating other sensitive groups into

the networks of PAA or PMAA will give gels more interesting properties. For example,

the copolymer of PAA and PMAA with PNIPAAm can provide the coupling

environmental sensitivity of pH and temperature [Tian et al., 2003; Zhang et al., 2000].

Recently, a series of smart biomaterials, such as poly(ethylacrylic acid) (PEAA) and

poly(propylacrylic acid) (PPAA), has opened new opportunities for the molecular

imaging field because of their sharp pH-sensitivity [Stayton et al., 2005] .

PAA PMAA PEAA PPAA PBAA

Figure 2.1 Structures of anionic pH-sensitive hydrogels.

11

Hydrogels made of PAA or PMAA can be used to develop formulations that

release drugs in a neutral pH environment [Brannon et al. 1990]. Some researchers

[Hassan et al., 1999] focused on the synthesis of anionic pH-sensitive hydrogels and the

swelling behavior studies. Of particular interest was the design of a self-regulated release

device based on the mechanism of the “molecule gate” system. An important example of

copolymer networks was represented [Lowman et al., 1995] to verify the complexation

and decomplexation mechanism. The authors not only explored the influence of factors,

such as the solution pH, graft chain molecular weight, and copolymer composition, on

network structure and dynamic property of p(MAA-g-EG) hydrogels, but also studied the

complexation dependent diffusion coefficients.

p(MAA-g-EG) is a promising candidate for oral delivery of peptide and protein

drugs through the gastrointestinal tract [Torres-Lugo et al., 2002; Robinson et al., 2002;

Kim et al., 2003; Ichikawa et al., 2003]. Peppas’ group prepared p(MAA-g-EG)

micro/nanospheres with relatively narrow size distributions. The effects of various

reaction parameters on the particle size and the distribution were investigated. The

enhancing effect of p(MAA-g-EG) micro- or nano-particles for salmon calcitonin

delivery through intestinal epithelial cells was also evaluated using Caco-2 cell

monolayer. Results revealed that the p(MAA-g-EG) hydrogel microparticles could be

used as a cytocompatible carrier possessing the transport-enhancing effect on the

intestinal epithelial cells. PMAA crosslinked with azoaromatic crosslinkers was

developed for colon-specific drug delivery [Ghandehari et al., 1997]. The drug release

from such hydrogels in the stomach was very minimal. As the gels passed down the

intestinal tract, the extent of swelling increased. But, the azoaromatic crosslinks of the

12

hydrogels were degraded by azoreductase produced by the microbial flora of the colon.

It is known that the transition between the swollen and the collapsed state with

changes in pH can be moved to higher pH values by increasing the hydrophobicity of the

monomers. Tirrell and coworkers [1992] first described the pH-dependent properties of

PEAA for membrane-disruptive applications. PEAA is inactive at physiological pH and

has a sharp transition around pH of 6.3. To obtain a series of shifted pH profiles,

Hoffman’s group [Murthy et al., 1999] investigated the pH transition change of sensitive

hydrogels by using different monomers with increased methylene units and applied their

membrane-disruptive properties in a blood cell hemolysis assay. The PPAA exhibited a

shift to the membrane active state at a higher pH and a surprising increase over PEAA in

hemolytic efficiency. Further addition of another methylene unit with poly(butyl acrylic

acid) shifted the pH profile to the physiological pH. This general shift in pH profiles is

consistent with the trend expected for making the alkyl group longer and more

hydrophobic [Mourad et al., 2001].

2.1.2 Cationic hydrogels

The synthesis and properties of cationic pH-sensitive hydrogles have also been

investigated over the past three decades. Hariharan and Peppas [1996] investigated the

swelling behavior of cationic hydrogels as carriers for drug delivery. Diethylaminoethyl

methacrylate (DEAEM) and diethylaminoethyl acrylate (DEAEA) were used as the

cationic monomers copolymerized with HEMA. The equilibrium water uptake was a

strong function of the ionic strength of the medium. Podual [1998] provided a cationic

hydrogel prepared by the copolymerization of DEAEM and poly(ethylene glycol)

13

monomethacrylate (PEGMA). Not only the effect of crosslinking ratio on the swelling

properties was studied, but also the structure of hydrogels and the diffusion coefficients

were determined. Traitel et al. [2000] studied the insulin controlled release system based

on the cationic hydrogel, PHEMA-co-N,N-dimethylaminoethyl methacrylate

(DMAEMA). The effects of polymer morphology and oxygen availability on hydrogel

swelling and insulin release kinetics were studied. Hydrogels without the crosslinking

agent were stable in water and their sensitivity to pH was higher than the chemically

crosslinked hydrogels.

A pH-sensitive hydrogel containing glucose oxidase enzyme is called

glucose-sensitive hydrogel due to its responsiveness to ambient glucose concentration

[Jung et al., 2000]. These systems are functionalized with enzymes by binding the

enzyme into a polymer network during polymerization. Glucose oxidase is probably the

most widely used enzyme in glucose sensitivity. It oxidizes glucose to gluconic acid,

resulting in a pH change of the medium. Horbett’s group [Albin et al., 1985, Kost et al.,

1985; Klumb et al., 1992; Klumb et al., 1993] was the first to study systems consisting of

immobilized glucose oxidase in a pH responsive polymeric hydrogel, enclosing a

saturated insulin solution. Glucose oxidase has been successfully immobilized on a

wide variety of polymers, such as poly(MAA-g-EG) [Hassan et al., 1999],

poly(HEMA-DMAEMA) [Traitel et al., 2000], poly(DEAEM-g-EG) [Podual, 1998],

poly(HPMA-co-DMAEMA) [Jung et al., 2000], polyacrylates [Turmanova et al., 1993],

polyethylene [Hsiu et al., 1990], poly(vinyl alcohol) [Kozhukharova et al., 1988].

pH-sensitive hydrogels can serve as drug delivery carriers for oral, buccal, rectal,

vaginal, ocular, epidermal and subcutaneous applications. However, hydrogels made of

14

non-biodegradable polymers has to be removed from the body after use. The

non-biodegradability is not a problem for oral drug delivery, but it becomes a serious

limitation in other applications, such as the development of implantable drug delivery

carriers or implantable biosensors. Thus, much attention has been focused on the

development of biodegradable pH-sensitive hydrogels. Various formulations were

developed to obtain the biodegradable pH-sensitive hydrogels with appropriate properties,

such as dextran [Chiu et al., 1999, Franssen et al., 1999], semi-interpenetrating

Chitosan-PVA [Wang et al., 2004], PVA-gelatin [Wang et al., 2004], and poly(lactic

acid)-poly(ethylene glycol)-poly(lactic acid) hydrogels[Mason et al., 2001].

2.2 Temperature-Sensitive Hydrogels

Temperature-sensitive hydrogels have received considerable attention for uses in

bioseparations, drug delivery, and diagnostics due to the ability of hydrogels to swell or

shrink as a result of temperature change in the surrounding fluid [Peppas et al., 2000].

Based on the transition mechanism, these hydrogels can be classified into three categories:

negatively temperature-sensitive gels, positively temperature-sensitive gels, and

thermo-reversible gels [Qiu et al., 2001].

Positive hydrogels have an upper critical solution temperature (UCST). If the

temperature is below UCST, the hydrogels contract and release solvent from the matrix.

In contrast, the swelling behavior of negative hydrogels is attributed to the lower critical

solution temperature (LCST). A temperature above LCST results in a collapsed structure

for hydrogels. For the thermo-reversible gels, the polymer chains are not covalently

crosslinked and the gels may undergo sol-gel phase transitions, instead of

15

swelling-shrinking transitions.

2.2.1 Negatively temperature-sensitive gels

For most polymers, the water solubility increases with the increasing temperature.

Negatively temperature-sensitive gels, however, have a critical parameter LCST. That

means these gels shrink as the temperature increases above the LCST and swell at the

lower temperature. The structures of some of those polymers are shown in Figure 2.2.

PNIPAAm PDEAAm P(NIAAm-co-AA)

Figure 2.2 Structures of negatively temperature-sensitive hydrogels.

Some of the earliest work with negatively temperature-sensitive hydrogels was

done by the Tanaka’s group [1978]. Poly(N-isopropylacrylamide) (PNIPAAm) is the best

example of a negatively temperature-sensitive hydrogel, which is made of polymer chains

containing a mixture of hydrophobic and hydrophilic segments. At lower temperatures,

water interacts with the side chains through the hydrogen bonds between water molecules

and the hydrophilic parts, –CONH–. These hydrogen bonds lead to enhanced dissolution

and well swelling in water [Zhang et al., 2003]. As the temperature is increased to higher

16

than LCST, the hydrophobic interactions among hydrophobic segments, –CH(CH3)2,

become stronger, while hydrogen bonds become weaker. These interactions result in the

shrinking of the hydrogels due to inter-polymer chain association [Qiu et al., 2001].

Hitotsu et al. [1987] worked with crosslinked PNIPAAm and determined that the LCST

of PNIPAAm gel was 34.38C. However, the response rate to external temperature

changes of typical PNIPAAm hydrogel is low, which limits its applications. Kabra et al.

[1991] synthesized fast temperature-response PNIPAAm gels by using a phase separation

technique. Preparation of gels at temperatures above LCST [Wu et al., 1992] or below the

freezing point [Zhang et al., 1999] results in an enhanced shrinking rate. Gas blowing

[Nakamoto et al., 2001] and radiation [Chen et al., 1999] may produce porous structures

leading to fast response. Other successful approaches to achieve a high

temperature-response rate involve using poly(ethylene glycol)s as pore-forming agents

[Zhang et al., 2000], interpenetrating poly(vinyl alcohol) within the hydrogel network,

using aqueous sodium chloride solution as the reaction medium for gel preparation, and

carrying out polymerizations in mixed sucrose solutions. These approaches could

significantly increase the response rate since these reaction mediums induce the phase

separation of gel system. For example, the fully deswelling time could be reduced to 2

min when the hydrogel polymerizations were carried out in aqueous glucose solutions

[Zhang et al., 2003].

LCST is a very important parameter for negatively thermo-sensitive gels. LCST

could be increased by mixing a small amount of ionic copolymers in the gels [Yu et al.,

1993] or by changing the solvent composition [Suzuki et al., 1996]. In general, as the

polymer chains contain more hydrophobic constituents, LCST moves to a lower

17

temperature. Thus, incorporating a hydrophilic monomer, acrylic acid (AAc), into the

PNIPAAm backbone is a good approach to modulate the properties of PNIPAAm gels

[Zhang et al., 2002]. Copolymerization of NIPAAm with different monomers results in

hydrogels with versatile properties. However, an increased hydrophilic content in the

copolymer network can reduce its temperature sensitivity [Beltran et al., 1991; Feil et al.,

1993]. In order to improve the temperature sensitivity of copolymers, several researchers

have prepared PNIPAAm-based copolymers. Okano and coworkers [Kaneko et al., 1998]

developed an exquisite method to prepare graft hydrogels of PNIPAAm. Small

PNIPAAm molecules were grafted with the main chain of the crosslinked PNIPAAm.

Above the LCST, hydrophobic regions in the network structure made the gels dehydrate

to a collapse state. At temperatures below LCST, the gels could transform into a fully

swollen conformation in less than 20 min, which was much faster than that of comparable

gels without graft chains. This group also proposed an incorporating carboxylate method

to promote gel shrinking [Ebara et al., 2001]. 2-carboxyisopropylacrylamide (CIPAAm)

was incorporated into PNIPAAm gels to induce rapid shrinking in response to small

temperature increases. In contrast, P(NIPAAm-co-AAc) copolymer gels lose their

temperature sensitivity with the introduction of only a few mole percent of AAc. Zhang et

al. [2002] synthesized P(NIPAAm-co-AAc) gels in an alkaline solution to achieve the

improved oscillating swelling properties. There has also been significant interest in the

synthesis of PNIPAAm-based hydrogels by other methods such as graft-, block- or comb-

copolymerization. Such systems show promise for rapid and abrupt or oscillatory release

of drugs, peptides, or proteins, because their swelling or syneresis process can occur

relatively fast [Yoshida et al., 1995; Kaneko et al., 1996; Inoue et al., 1997].

18

Recently, efforts have been made to prepare multifunctional hydrogels responding

to more than two stimuli, such as the pH and temperature sensitive hydrogels. Chen and

Hoffman [1995] prepared p(NIPAAm-g-AA) gels, which exhibited temperature- and

pH-sensitive behavior. These gels were able to respond rapidly to both temperature and

pH changes. The temperature- and pH-dependent swelling behaviors were better defined

in the graft copolymers than in random copolymers containing similar amounts of

components. Tian et al. [2003] developed a hydrogel of p(NIPAAm-co-AAc) modified by

a small amount of hydrophobic comonomers in tert-butanol solutions. The hydrogels

with a suitable 2-(N-ethylperfuorooctanesulfoamido) ethyl acrylate content showed good

pH and temperature sensitivity. Similar work was done by the Peppas group [Zhang et al.,

2000]. The interpenetrating gels of PNIPAAm and PMAA exhibited the ability of

responding to temperature and pH conditions. Additionally, the transition conditions were

determined at a pH value of approximately 5.5 and a temperature range of 31-32°C.

Negatively temperature-sensitive hydrogels have been studied extensively and

these materials can be used in a variety of applications, including controlled drug delivery,

immobilized-enzyme reactors, separation process, and biochips. In a monolithic device,

an on–off drug release profile could be obtained based on the reversible

thermo-sensitivity of hydrogels [Bae et al., 1990; Okano et al., 1990], which involve

crosslinked p(NIPAAm-co-BMA), and inter-penetrating PNIPAAm and

poly(tetramethyleneether glycol) (PTMEG). In order to increase the mechanical strength

of hydrogels, Okano and coworkers incorporated a hydrophobic comonomer, BMA into

NIPAAm gels and investigated the on–off release profile of indomethacin from the

matrices in response to a stepwise changing temperature. The hydrophobicity of the

19

comonomer influenced the shrinking process and thus controlled the release behavior of

the therapeutic agent dispersed in the matrix [Yoshida et al., 1991]. Negatively

temperature-sensitive gels are also utilized for controlled delivery of highly sensitive

therapeutic agents, such as peptides and proteins. Peppas et al. [1996] developed a

hydrogel of inter-penetrating PNIPAAm and PMAA and studied the release kinetics of

bioactive streptokinase. Kim et al. [1996] used an inter-penetrating hydrogel of

PNIPAAm and PAA to effectively release the protein drug, calcitonin, in response to

changing temperature and pH.

2.2.2 Positively temperature-sensitive gels

Certain hydrogels formed by IPNs show positive thermosensitivity. IPNs of PAA

and polyacrylamide (PAAm) or P(AAm–co-BMA) have positive temperature dependence.

IPNs composed of PAA and PAAm may shrink at low temperatures because of the

interpolymer complexes formed by hydrogen bonding. The complexes dissociate at

higher temperatures due to breaking of hydrogen bonds, and the gels rapidly swell above

the UCST [Klenina et al., 1981]. Katono et al. [1991] compared the temperature

dependent swelling behavior of poly(AAm-co-BMA), the IPNs of poly(AAm-co-BMA)

with PAA, and the random copolymer gel poly(AA-co-AAm-co-BMA). The IPNs and the

random gels showed the distinctly different profiles of temperature dependence, although

both had the positive temperature dependence. Only the IPNs showed a sigmoidal

alteration with a transition zone. The swelling of those hydrogels was reversible,

responding to stepwise temperature changes. This resulted in reversible changes in the

release rate of a model drug, ketoprofen, from a monolithic device.

20

Clinical applications of thermosensitive hydrogels based on NIPAAm and its

derivatives are limited due to the non-biocompatibility of the monomers and crosslinkers

and non-biodegradability of NIPAAm polymers and their derivatives. Further

development of new, biocompatible and biodegradable thermoreversible gels, such as

PEO-PLA block copolymers, is necessary to exploit the useful properties of

thermoreversible gles.

2.3 Properties of Hydrogels

2.3.1 Swelling properties

The swelling behavior of hydrogels is an important property for a variety of

applications. Generally, the swelling property of polymers is reflected by the

weight-swelling ratio, the ratio of the weight of the swollen sample to the weight of the

dry matrix. Factors affecting the swelling ratio mainly involve the crosslinking ratio, the

solvent concentration and quality, the chemical structure, and the specific stimuli.

The crosslinking ratio, the ratio of moles of crosslinking agent to the moles of

polymer repeating units, has a dominated effect. The higher the crosslinking ratio, the

more crosslinking agent is incorporated in the hydrogel structure. Highly crosslinked

hydrogels have a tighter structure, and will swell less compared to the same hydrogels

with a lower crosslinking ratio.

In many cases the influence of solvent is small. However, it is becoming

increasingly evident that solvent effects can be used to control the free radical

polymerization of hydrogels, both at the macroscopic and at the molecular levels. The

solvent concentration during the polymerization affects the material properties of the

21

polymer by increasing the rate of primary cyclization of multivinyl monomers during the

polymerization [Anseth et al., 1996; Elliott et al., 2001]. A primary cycle differs from a

crosslink in that the propagating free radical reacts intramolecularly with its own pendant

double bonds, which then loses the opportunity to crosslink. The greater the extent of

primary cyclization, the less crosslinked the polymer will be and the larger the mesh size.

This leads to the increased equilibrium swelling and reduced mechanical strength with

the increasing solvent concentration during the polymerization. The effects of solvent

concentration on the rate of primary cyclization and gel network formation can be

explained by the local dynamics of the propagating radical. For lower solvent

concentrations, the double bond concentration surrounding the free radical is relatively

high, leading to a faster rate of propagation and less opportunity for the free radical to

cycle by reacting with its own pendant double bonds. In addition to solvent concentration,

solvent quality also affects the three-dimensional network structure created during the

polymerization. For a better solvent, the propagating chain is less likely to cycle and thus

has a compact structure. However, the propagating chain is more likely to cycle for a

poor solvent, and the rate of primary cyclization is high, leading to a loose network

structure [Elliott et al., 2002].

The chemical structure of the polymer may also affect the swelling ratio.

Hydrogels containing hydrophilic groups swell to a higher degree compared to those

containing hydrophobic groups. Hydrophobic groups collapse in the presence of water,

thus minimizing their exposure to the water molecule. As a result, the hydrogels will

swell much less compared to hydrogels containing hydrophilic groups. Swelling of

environmentally sensitive hydrogels can be affected by specific stimuli. For example,

22

temperature and pH affect the swelling of temperature- and pH-sensitive hydrogels,

respectively. There are many other specific stimuli that can affect the gel swelling.

2.3.2 Network structure and characterization

The effect of chemical structure on polymer properties is without doubt the most

important aspect of polymer chemistry. Extensive uses of hydrogels in drug delivery

systems depend to a large extent on their structures in buffer solution. Based on the work

done by many researchers, the most important parameters used to characterize the

network structure of hydrogels are the polymer volume fraction in the swollen state (v2,s),

molecular weight of the polymer chain between two neighboring crosslinking points (Mc),

and the corresponding length or mesh size (ξ). In order to elucidate the structure of

hydrogels, the equilibrium swelling theory and the rubber elasticity theory are utilized

[Peppas et al., 2000].

The polymer volume fraction in the swollen state is a measure of the amount of

fluid imbibed and retained by the hydrogel. The molecular weight between two

consecutive junctions is a measure of the degree of crosslinking of the polymer. These

junctions may be chemical crosslinks, physical entanglement, crystalline regions, or even

polymer complex. It is important to note that only average values of Mc can be calculated

due to the random nature of the polymerization process. The correlation distance between

two adjacent crosslinks (ξ) provides a measure of the space available between the

macromolecular chains for drug diffusion. Also, it can be reported only as an average

value. These parameters can be determined theoretically or through a variety of

experimental techniques.

23

A Theoretical approaches

Several theories have been proposed to calculate the molecular weight between

crosslinks in a hydrogel matrix. Two theoretical methods, which are prominent among the

growing techniques utilized to elucidate the structure of hydrogels, are the equilibrium

swelling theory and the rubber elasticity theory.

The structure of hydrogels that contain ionic moieties was analyzed by Peppas

and Merrill [1977] based on the Flory-Rehner theory [Flory et al., 1943]. This

thermodynamic theory states that a crosslinked polymer gel, which is immersed in a fluid

and allowed to reach equilibrium with its surroundings, is subject only to three opposing

forces, the thermodynamic force of mixing, the retractive force of the polymer chains,

and the ionic force. At equilibrium, these forces are equal. Eq. (1) describes the physical

situation in terms of the Gibbs free energy.

ionicmixingelastictotal GGGG ∆+∆+∆=∆

Here, elasticG∆ is the contribution due to the elastic retractive forces developed

inside the gel, mixingG∆ is the result of the spontaneous mixing of the fluid molecules

with the polymer chains, and ionicG∆ is the contribution due to the ionic nature of the

polymer network. Eqations (2) and (3) are expressions that have been derived for the

swelling of anionic and cationic hydrogels prepared in the presence of a solvent.

)]2

()[()2

1)((])1[ln()10

)((4 ,2

,23/1

,2

,2,2

12,21,2,2

22,21

r

s

r

sr

n

c

csss

apH

as

v

v

v

vv

M

M

Mv

Vvvv

K

K

v

v

I

V −−+++−=−− χ

)]2

()[()2

1)((])1[ln()10

)((4 ,2

,23/1

,2

,2,2

12,21,2,2

214

2,21

r

s

r

sr

n

c

csss

apH

bs

v

v

v

vv

M

M

Mv

Vvvv

K

K

v

v

I

V −−+++−=−− χ

(1)

(2)

(3)

24

In these expressions, I is the ionic strength, and Ka and Kb are the dissociation constants

for the acid and base, respectively.

Hydrogels resemble natural rubbers in their remarkable property to elastically

respond to applied stresses. The elastic behavior of hydrogels can be used to elucidate

their structure by utilizing the rubber elasticity theory originally developed by Treloar

[1958] and Flory [1949]. However, the original theory or rubber elasticity does not apply

to hydrogels prepared in the presence of a solvent. Silliman [1972] and Peppas et al.

[1977] developed the expressions to analyze the structure of hydrogels prepared in the

presence of a solvent.

In Eq. (4), τ is the stress applied to the polymer sample, ρ is the density of the

polymer. The rubber elasticity theory has been used to analyze chemical and physical

crosslinked hydrogels [Mark, 1982; Anseth et al., 1996], as well as hydrogels exhibiting

temporary crosslinks due to hydrogen bonding [Lowman et al., 1997].

The primary mechanism of drug release from a hydrogel matrix is diffusion,

occurring through the space available between macromolecular chains in aqueous media

as a result of environmental stimuli. This space is often regarded as the pore. Depending

on the size of these pores, hydrogels can be conveniently classified as macro-porous,

micro-porous and non-porous. A structural parameter that is often used in describing the

size of the pores is the correlation length (ξ) which is defined as the linear distance

between two adjacent crosslinks, and can be calculated using the following equation

[Torres-Lugo et al., 1999],

(4) 31

))(1

)(2

1(,2

,22

r

s

n

c

c M

M

M

RT

υυ

ααρτ −−=

25

2/120

3/1 )(rQr=ξ

where Qr is the volume swelling ratio of the swollen polymer at equilibrium to the dry

polymer, and is the end-to-end distance in the unperturbed state, which can be

calculated by the following equation,

lM

MCr

r

cn 2/12/120 )

2()( =

where Cn is the polymer characteristic ratio (14.4 in case of a methacrylate chain), Mc is

the molecular weight between crosslinks, l is the carbon-carbon bond length (1.54 Å), ,

and Mr is the molecular weight of the repeating unit.

The average values of Mc and ξ are related to each other. The molecular weight

between crosslinks can be obtained by the following equation,

where Tg is the transition temperature of the crosslinking polymer, Tg0 is the transition

temperature of the uncrosslinking polymer. According to these equations, the parameters

characterizing the network structure of hydrogels can be experimentally obtained.

Although theoretical characterizations of the network structure are complicated,

they and the diffusion studies of model drugs provide an invaluable insight into the very

complex structure of gel networks and help in the design for drug delivery carriers

[Narasimhan et al., 1997].

B Experimental approaches

Besides theoretical approaches, simpler and more intuitionistic approaches can be

2/120 )(r

(5)

(6)

0

4109.3

ggc TT

M−×= (7)

26

used to investigate the hydrogel structure in buffer solutions. This section presents a brief

description of each specific approach followed by its advantages and limitations.

B.1 Scanning electron microscopy (SEM)

To visually examine the surface and interior morphology of a hydrogel in the

swollen state, scanning electron microscopy is commonly used to analyze the pore

structure and to observe the three dimensional structure. For example, Kim et al. [2000]

reported the visual observation of an unique 3D honeycomb-like network structure in the

interior of a swollen dextran-methacrylate hydrogel by SEM. Investigations of the

hydrogel structure by SEM lead to valuable results: 3D structure, pore shapes, and the

approximate pore size. However, this SEM-based technique suffers from a sever

disadvantage because the native state of hydrogels is characterized by the presence of

water and the need of dehydration and/or fixation procedures prior to SEM examination

inevitably affects the morphology of a hydrogel. The preparation of a hydrogel sample

for SEM examination involves critical-point drying and vacuum drying methods. Both

drying techniques result in volume shrinkage and significantly morphological alterations

of the gels. Other techniques, such as cryofixation, cryofracturing, and freeze-drying,

have been used to examine the interior structure of hydrogels because solvent (e.g., water)

inside can be easily removed by sublimation with minimal disturbance of structure [Hong

et al., 1998; Yang et al., 1983].

B.2 Environmental scanning electron microscopy (ESEM)

Although the dehydration and/or fixation procedures aid the swollen gel in SEM

testing, some reports indicated a discrepancy between the original hydrogel structure and

the deduced images from SEM. ESEM represents an important advance in conventional

27

SEM for hydrogel characterization. Whereas conventional SEM requires a relatively

high vacuum in the specimen chamber to prevent atmospheric interference with primary

or secondary electrons, an ESEM may be operated with a poor vacuum (up to 10 Torr of

vapor pressure, or one seventy-sixth of an atmosphere) in the specimen chamber. In such

"wet mode" imaging, the specimen chamber is isolated from the rest of the vacuum

system. Water is the most common imaging gas, and a separate vacuum pump permits

fine control of its vapor pressure in the specimen chamber. Due to the effect of the

electron beam, the water molecules are positively ionized, and thus they are

forced/attracted toward the hydrogel samples, serving to neutralize the negative charge

produced by the primary electron beam. In order to preserve the sample from

dehydration, the water vapor is kept at the saturation condition within the microscope

chamber by using a Pertiler cold stage. The field-emission gun produces a brighter

filament image and its accelerating voltage may be lowered significantly, permitting

nondestructive imaging of fragile specimens, such as swollen gels

[www.itg.uiuc.edu]. Because of these technical improvements, ESEM provides more

advantages for characterization of the hydrogel network. No additional sample treatment

is performed to avoid any introduction of possible specimen-coating artifacts, or

problems involved with either changing samples to a vacuum-friendly state or creating

their former replica. The controlled environment in the specimen chamber retains the

stable structure. Some accessory and the control valves could extend ESEM applications

for dynamic experiments. For instance, the morphology change of the pH sensitive

poly(AA-co-AAm) hydrogel swollen in different buffers was studied, and the 3D

network and the pore size were clearly observed from ESEM images [Zhou et al., 2003].

28

B.3 Confocal laser scanning microscopy (CLSM)

Confocal laser scanning microscopy (CLSM) is another valuable imaging

technique for direct observation of the hydrogel structure as it allows us to observe

non-destructive samples under the mild conditions. CLSM has been successfully applied

in biological, medical, and geological studies as an alternative investigative method for

hydrogel structure analysis. For CLSM, the hydrogel can remain in the aqueous

environment, thus avoiding the hazardous dehydration. The only sample treatment prior

to examination is the conjugation of a fluorescent dye to the polymer segments in a

hydrogel. However, the actual dye concentration can be very low so that the disturbance

of biological systems is kept to a minimum. Subsequently, images of the bulk structure

can be obtained at different locations without cutting or fracturing the hydrogels and

magnified images of any area of interest can be obtained [Fergg et al., 2001]. With the aid

of application software, the 3D nature of the hydrogel can be calculated from a series of

successful images taken at defined intervals and observed as a movie or as a single stereo

pair image.

There are a great number of advantages to using CLSM compared to conventional

fluorescence microscopy (FM). The two most important ones are the ability to eliminate

the out-of focus noise and the greatly increased sensitivity of the machine

[www.bioteach.ubc.ca]. CLSM also has some limitations. The first limitation of CLSM is

that the resolution is limited by the wavelength of light. Photo-damage is also a limitation

in the use of CLSM. The good axial resolution (between two focal planes) is obtained by

using two-photon fluorescence microscopy, which provides the possibility to obtain

biochemical information about cells or tissues and causes minimal photo-damage due to

29

its inherent 3D resolution and long penetration depth.

B.4 Porosimetry

The use of SEM, ESEM and CLSM approaches provide morphological details of

the interior and surface structure of hydrogels. However, there are needs to examine the

structure of hydrogels in a quantitative manner, because pore size, volume, and structure

of hydrogels are critical factors to control swelling, drug release behavior, and biological

interactions inside the body. The quantitative information of the pore structure of

hydrogels under a swollen condition could be obtained by nitrogen absorption and

mercury intrusion porosimetry.

Mercury intrusion porosimetry (MIP) has provided valuable information about

various aspects of pore structure characterization for porous media and powders [Mikijelj

et al., 1991; Liu et al., 2000; Gemeinhart et al., 2000]. The theory of all mercury

porosimeters is based on the physical principle that a non-reactive, non-wetting liquid

will not penetrate pores until sufficient pressure is applied to force its entrance. The

relationship between the applied pressure and the pore size is given by the Washburn

equation:

P

DΘ−= cos4γ

Where P is the applied pressure, D is the diameter, γ is the surface tension of

mercury (480 dyne cm-1) and Θ is the contact angle between mercury and the pore wall,

usually near 150○. As pressure increases, the instrument senses the intrusion volume of

mercury. As the mercury column shortens, the pressure and volume data are continuously

acquired and displayed by an attached personal computer.

(8)

30

As an analytical instrument, MIP can measure pores of diameters ranging from

3.6 nm to 200 mm. and give the porosity data from the intruded volume. Therefore, MIP

would be a good method to quantify pore size and volume of swollen hydrogel. However,

like SEM-based techniques, MIP also needs the dehydration and/or fixation procedures

prior to the examination, inevitably affecting the morphology and pores structure of a

hydrogel.

In this section, the theoretical approaches and experimental approaches to

characterize the network structure of swollen hydrogels are described. Decisions as to

which approach is most appropriate for the loose structure must consider the complexity

of sample preparation, gel deformation due to water loss, sample preparation, the

instrument operation, and the gel applications

2.3.3 Mechanical properties

Mechanical properties of hydrogels are extremely important in selecting a

material that is suitable for a specific pharmaceutical application. The theories of rubber

elasticity and viscoelasticity are used to understand the mechanical behavior of hydrogels.

These theories are based on time-independent and time-dependent recovery of the chain

orientation and structure, respectively. The use of these theories makes it possible to

analyze the polymer structure and determine the effective molecular weight between

crosslinks.

Anseth et al.[1996] summarized the dependence of the mechanical properties on

various parameters, which mainly include monomer composition, the crosslink density,

the degree of swelling, and medium conditions. Altering the composition of comonomers

31

used in preparing hydrogels is the simplest parameter to control the mechanical properties

of hydrogels. If the hydrogel is not a homopolymer, increasing the relative amount of

physically stronger components leads to an increased mechanical strength of the gels. For

instance, replacing acrylates with methacrylates causes the increased stiffness of the

polymeric backbone and increased mechanical strength. The change of hydrophilicity of

the polymer also alters the mechanical strength of the gels. Some results were reported

that the addition of N-vinyl-2-pyrrolidone (NVP) in the copolymer system of HEMA and

MMA resulted in a significantly decreased Young’s modulus since the hydrophilic NVP

alters the swelling properties of the hydrogel [Lustig et al., 1991; Davis et al., 1989].

Changing the crosslinking density has been utilized to achieve the desired

mechanical property of the hydrogel. The higher crosslinking density of the system will

result in a stronger gel. However, a higher degree of crosslinking creates a more brittle

structure and a lower swelling ratio. Hence, there is an optimum degree of crosslinking to

achieve a relatively strong and yet elastic hydrogel [Peppas et al., 2000].

The reaction conditions have the profound effects on the mechanical properties of

formed hydrogels. These conditions are summarized as reaction time, temperature, light

intensity, and amount and type of solvent. Of most importance are the amount and type of

solvent. If a large amount of solvent is used in polymerization, the crosslinking agent

prefers to intra-crosslinking than inter- crosslinking, which results in the loose network

and the low mechanical strength. The type of solvent or the nature of solvent is also used

as the controllable variable for mechanical properties. For example, the ionic strength and

pH values alter the reactivity of the monomers, leading to changed mechanical strength.

Usually, a highly ionic strength reduces the reactivity of monomers [Baker et al., 1994].

32

Other reaction conditions including reaction time and temperature can be changed to get

varied properties. For the photopolymerization, the light intensity and dosage influence

the network structure of gels and the mechanical properties [Crump, 2001]. Post-reaction

treatment can also work as a variable in manipulating the material strength. Techniques

such as the addition of a compound [Philippova et al., 1994] and thermal recycle [Cha et

al., 1993] can also be used to change the gel strength.

Most previously introduced variables such as monomer composition, crosslinking

density, and the reaction conditions, are designed to change the degree of swelling of the

hydrogels and thus modulate the mechanical properties. Typically, when the polymers

swell in a plasticizing solvent, the glass transition temperature of the mixture decreases

and the material becomes weaker. In most hydrogel applications, the swelling conditions

are usually predetermined according to the application. If not, the external conditions,

such as pH, temperature, ionic strength, pressure, or other swelling moduli, can be

controlled to get the desired mechanical properties for specific applications.

Common approaches for measuring mechanical properties of hydrogels involve

tensile or dynamic mechanical analysis. For most uniaxial tensile testing,

dumbbell-shaped samples are placed between two clamps and one end of the material is

pulled away from the other at varying loads and rate of extension. For most cases,

hydrogel samples are cut in their equilibrium-swollen state and the sample dimensions

must be measured in this swollen state. For tensile testing, hydrogel samples should be

immersed in a waterbath that is thermally regulated during the testing. To perform

dynamic mechanical testing, the samples are usually prepared in thin strips with square

edges and a uniform cross-sectional area through the sample length. Dumbbell-shaped

33

samples are no longer an optimal sample shape. The optimal cross-sectional area of the

sample is related to the modulus of the materials.

For hydrogel samples, the water loss during the experiment significantly

influences the mechanical behavior. With the increase of temperature, water loss becomes

more prominent and leads to increased moduli. Water loss can be minimized by coating

the hydrogel samples with petroleum gel (effective up to 45C) or silicon vacumm grease

(effective up to 85C) [Lustig et al., 1991]. Water loss limits the temperature range for

dynamic mechanical testing.

2.4 Applications of Hydrogels in Drug Delivery

Hydrogels, as a desired material, have been extremely useful in biomedical and

pharmaceutical applications due to their unique swelling properties and structures. Based

on the hydrogel functionalities, these biomaterials can be an excellent candidate for

controlled release devices, bioadhesive or targetable devices, and self-regulated release

devices. According to the delivery administration, hydrogel-based devices can be used for

oral, nasal, ocular, rectal, vaginal, epidermal and subcutaneous applications [Peppas et al.,

2000]. This section first summarizes applications of hydrogels for different

administrations, including its challenges and current status of development. Hydrogels for

gastrointestinal administration are introduced in detail because of their close relationship

with the work in this dissertation. This is followed by the trends and perspectives for drug

delivery.

34

2.4.1 Peroral drug delivery

Oral drug delivery is the most desirable and preferred method of administering

therapeutic agents for their systemic effects. In addition, the oral medication is

generally considered as the first avenue investigated in the discovery and development

of new drug entities and pharmaceutical formulations, mainly because of patient

acceptance, convenience in administration, and cost-effective manufacturing process.

Because of its enormous market potential, oral drug delivery using controllable hydrogels

has attracted considerable attention in the past 20 years.

In peroral administration, hydrogels can deliver drugs to four major specific sites:

mouth, stomach, small intestine and colon. By controlling their swelling properties or

bioadhesive characteristics in the presence of a biological fluid, hydrogels can be a useful

carrier for releasing drugs in a controlled manner at these desired sites. Furthermore, the

mucoadhesive hydrogels offer an attractive property for drug targeting at certain specific

regions, leading to a locally increased drug concentration, and thus, enhancing the drug

absorption at the release site.

2.4.1.1 Buccal route

Drug delivery to the oral cavity has versatile applications in the local treatment of

diseases of the mouth, such as periodontal disease, fungal and viral infections, and oral

cavity cancers. To ensure the long-term adhesion of the delivery carrier at specific site

and to improve the drug absorption, many types of bioadhesive hydrogels have been

considered in the device design since the early 1980s. The typical delivery carrier for

buccal route comprises tablets, patches, and ointment. Some of these are already on the

35

market. For example, a double-layered tablet with a bioadhesive layer made of

hydroxypropyl cellulose/PAA and a lactose non-adhesive backing layer was introduced in

the market by Nagai et al. [1999] for the treatment of aphthous stomatitis. Bouckaert et al.

[1993] tested the buccal tablets of miconazole based on a modified starch-PAA mixture.

Although these tables showed different mucoadhesion properties, there was no significant

difference in the salivary content of miconazole for human volunteers. Nair and Chien

[1996] compared patches and tablets of different polymers and different released drugs.

Sustained release of all four compounds from mucoadhesive tablets was observed.

For systemic drug administration, new buccal bilayered tablets, comprising two

layers–a drug-containing mucoadhesive layer of chitosan with polycarbophil and a

backing layer of ethylcellulose, were developed by direct compression. The

double-layered structure design provided a unidirectional drug delivery towards the

mucosa, and minimized the drug leakage. The striking feature of this device would be the

utilization of an in-situ crosslinking reaction between the cationic chitosan and the

anionic polycarbophil, leading to the controlled swelling, prolonged drug release, and an

adequate adhesiveness [Remunan-Lopez et al., 1998].

A hydrogel-based ointment can also be utilized as a drug delivery device or a

liposome delivery vehicle for the topical treatment of certain diseases in the oral cavity.

Compared with the conventional ointment-drug formulations, liposomal formulations

within ointment may provide more desirable properties for topical use, such as the

reduction of uncontrolled release of drugs into the blood circulation and certain

undesirable side effects. Petelin et al. [1998] investigated the pharmaceutical performance

of three different hydrogel-based ointments as possible vehicles for liposome delivery

36

into the oral cavity tissues by electron paramagnetic resonance (EPR). Liposome

containing mucoadhesive ointments were prepared by simply mixing multilamellar

liposomes with each ointment pre-diluted with phosphate-buffered saline of pH 7.4. An

EPR study showed that p(MAA-co-MMA) was the most appropriate ointment in terms of

liposomal stability in the ointment, transport of liposome-entrapped molecules from the

ointment into the oral soft tissues, and washing-out time from oral mucosa or gingvia.

2.4.1.2 Gastrointestinal route

The peroral route represents the most convenient route of drug administration,

being characterized by high patient compliance. The mucosal epithelium along the

gastrointestinal tract varies. In the stomach, the surface epithelium consists of a single

layer of columnar cells. A thick layer of mucus covers the surface to protect against

aggressive luminal content. This specific site is of minor interest for drug delivery since

the low pH and the presence of proteolytic enzymes make the stomach a rather harsh

environment. However, there are examples of hydrogel-based devices specially designed

to delivery in the stomach.

Patel and Amiji [1996] developed stomach-specific antibiotic drug delivery

systems for the treatment of peptic ulcer disease using pH-sensitive cationic hydrogels.

The hydrogels were composed of freeze-dried chitosan-poly(ethylene oxide)

interpenetrating network. pH-dependent swelling properties and the release of two

common antibiotics, amoxicillin and metronidazole were evaluated in an enzyme-free

simulated gastric fluid (pH=1.2) and a simulated intestinal fluid (pH=7.2). The rapid

swelling and drug release demonstrated by these hydrogel formulations in the lower pH

37

fluid may be beneficial for site-specific antibiotic delivery in the stomach. Amiji et al.

[1997] also reported enzymatically degradable gelatin-PEO semi-IPN with pH-sensitive

swelling properties for oral drug delivery. The incorporation of gelatin in the IPN made it

possible to swell in the acidic pH of the gastric fluid due to the ionization of the basic

amino acid residues of gelatin.

The small intestine is characterized by an enormous surface area available for the

absorption of nutrients and drugs. The most important structural aspect of small intestine

is the means by which it greatly increases its effective luminal surface area by folds of

mucosa, fingerlike villi, and microvilli. The microvilli region has been referred as the

specialized location since regions of the device can be surfaced-modified to incorporate

cell-targeting mechanism that localize the vehicles at the specific site of reaction to

ensure that the drug diffuses the shortest distance in one direction towards the intestinal

epithelium. At the terminal ileum, the Peyer's patches, a particular specialization of the

gut-immune system, contain the M cells, which are specialized in endocytosis and

processing luminal antigens. The large intestine (colon) has the same cell populations as

the small intestine, and its main function is the absorption of water and electrolytes.

Aside from being an ideal non-invasive route of drug administration, the peptide

and protein delivery through the GI tract is fraught with difficulties around low

bioavailability, which results from the pH fluctuation, proteolytic degradation, low

transport, and short residence time. The pH fluctuation greatly influences the drug

integrity. For example, the high acidity of the stomach fluid can preclude the stability of

proteins. And the bile salt secreted from the gall bladder into the small intestine can

compromise the protein stability. Therefore, proper protection is required during oral

38

administration of bioactive molecules. Enteric-coated systems have been used in

commercial applications for releasing drugs through oral administration [Brogmann et al.,

2001]. The encapsulation of drugs within lipid vesicles also has the potential advantage

of protection and high drug loading [Park et al., 1997; Gregoiraidis, 1995]. However, a

major limitation is that these systems cannot fully protect the drugs and release them at a

targeted area with a precisely controllable rate over a long period of time. The use of

microspheres or nanoparticles made of pH-responsive complexation hydrogels to protect

drugs for site-specific delivery has been of interest. [Lowman et al.,1999; Morishita et al.,

2002]. Lowman’s group prepared crosslinked copolymer gels of PMAA with graft chains

of polyethylene glycol to protect the insulin in the harsh, acidic environment of the

stomach before releasing the drug in the small intestine. The insulin-containing

p(MAA-g-EG) microparticles demonstrated strong dose-dependent hypoglycemic effects

in in-vivo oral administration studies using both healthy and diabetic rats.

For a bioactive macromolecule, it is quickly denatured and degraded by

proteolytic enzymes in the GI tract. Much work has been done to protect against

enzymatic activity by adding protease inhibitors or coating the drug with liposomes and

polymeric film. Carbopol 934P and chitosan gels were tested in vivo for their ability to

increase the absorption of the peptide when administered intraduodenally in rats [Luellen

et al., 1996]. Both polymers increased the absorption of the peptide significantly,

probably due to both permeation enhancing and enzyme-inhibition properties. Akiyama

et al. [1996] reported novel peroral dosage forms of hydrogel formulations with protease

inhibitory activities using Carbopolw (C934P), which has been shown to have an

inhibitory effect on the hydrolytic activity of trypsin, and its neutralized freeze-dried

39

modification (FNaC934P). They demonstrated that two-phase formulations had the most

profound effect on trypsin activity inhibition. Ramdas et al. [2000] developed an oral

formulation for insulin delivery based on liposome encapsulated alginate-chitosan gel

capsules to increase the encapsulation efficiency to preserve the insulin stability through

the acidic media in the stomach and the enzyme-actively intestinal barrier. In animal

studies, it was reported that variable reductions in blood glucose were dependent on

factors including the lipid composition, size, surface charge and the physical state of the

phospholipid bilayer employed [Choudhari et al., 1994; Kisel et al., 2001]. Besides

liposomal approach, coating insulin with a pH-dependent acrylic based biodegradable

polymer and its encapsulation in enteric-coated microspheres has also been tried

[Musabayane et al., 2000]. Oral administration of insulin encapsulated in biocompatible

self-assembled ‘nanocubicles’ also appears to be effective in animal studies [Chung et

al., 2002].

The drug release at specific sites has received much attention. Based on the

surface receptors, various targeting molecules are utilized to achieve the local targeting.

For instance, a polymer-drug conjugate with an antibody can be recognized by the cell

surface antigen for cancer diagnostics and therapeutics [Jelinkova et al., 1999]. For

peptides or proteins through GI tract, the drug delivery system (DDS) can bind

specifically to the mucosal layer or cell surface to increase the residence time and

improve the drug bioavailability. Residence time is an important factor influencing drug

transport through the GI barrier. Several groups developed DDS with site-specific

delivery for peptides and proteins by the choice of material characteristics and the

combination of advanced manufacturing techniques. Dorkoosh et al. [2001] designed a

40

novel DDS for site-specific drug delivery of peptide drugs in the intestinal tract using

superporous hydrogels (SPH) and SPH composite polymers, which swell very rapidly by

absorption of gut fluids. Thus, the system attached to the intestinal wall and provided a

longer residence time for drug release. Shen et al. [2002] reported an intestinal patch

design for oral delivery. A longer residence time and unidirectional diffusion were

achieved for better drug diffusion through the intestinal barrier by using a mucoadhesive

layer of Carbopol/ pectin. Tao et al. [2003] combined microfabrication techniques with

the use of mucoadhesive plant lectins to design a microdevice with a long residence time.

These mucoadhesive drug delivery systems (MDDSs) have attracted considerable interest

because of their sustained drug release profile at the absorption site and increased drug

bioavailability due to the intimate contact with the absorbing tissue. However, a major

physiological condition, continuous shedding of the mucus, leads to the limited retention

of these conventional mucoadhesive devices that can only attach to the surface layer of

mucus due to their relatively large sizes [Ponchel et al., 1998]. It is also generally known

that gastrointestinal mucus renews completely within a few hours [Rubinstein et al.,

1994], which apparently sets an upper limit on the retention time of a mucoadhesive

system. In addition, the mucus layer can hinder the diffusion of drugs or drug carriers

from the device to the absorption site [Meaney et al., 1999; Khanvilkar et al., 2001]. The

GI mucus is a bilayered structure. One of the two layers is on the lumen side and called

the loosely-adherent layer because it can be easily sucked away. The other is on the

epithelium side and called the firmly-adherent layer since it is tightly attached to the

epithelial cells and is resistant to suction. It was reported that the mucus that experiences

full renewal in the generally-regarded turnover time might solely be the loosely-adherent

41

layer, and the firmly-adherent mucus probably has a longer turnover time [Stuma et al.,

2001; Brownlee et al., 2003]. As a result, longer retention than a few hours may be

achieved if a device can penetrate the loosely-adherent layer and adhere to the

firmly-adherent mucus layer.

Another typical approach to extend the duration time is to reduce the delivery

device to micron-sized or smaller. The microvilli region has been referred as the

specialized location. Currently, advanced DDS contain components on the micro- and

nanoscale, but the devices themselves remain in the macroscale (>1mm). As the scale

decreases, micro-fabricated DDS may be delivered by ingestion (<1mm), injected into

tissue (<200 µm), inhaled (<100µm), or released into the systemic circulation (<10nm).

To directly deliver the devices into the microvilli extending the residence time, the device

scale is required to be 5 µm or less. For hydrophilic and macromolecular compounds

such as peptides and proteins, which have to be absorbed preferably through the

paracellular route, the tightness of the intercellular junctions of the mucosal epithelia

forms a very strong absorption barrier [Luessen et al., 1997]. In an effort to increase

intestinal absorption of various macromolecules, permeation enhancers have been found

to reversibly open epithelial tight junctions. To date, numerous compounds have been

reported to have absorption-promoting activity and many researchers have tried to

elucidate the mechanisms by which the absorption can be enhanced [Yeh et al., 1994;

Lindmark et al., 1998; Kotze et al., 1999]. Nevertheless, the potential local toxicity of the

enhancers themselves has made it difficult to apply them to practical use. Only sodium

caprate is used as an absorption-enhancing adjuvant in drug products. Another major

disadvantage of permeation enhancers is their lack of specificity, opening the possibility

42

that food-borne pathogens and toxins migrate along with therapeutic compounds [Foraker

et al., 2003].

2.4.2 Nasal route

The nasal route of drug administration is the most suitable alternative of delivery

for poorly absorbable compounds such as peptide or protein drugs. The nasal epithelium

exhibits relatively high permeability, and only two cell layers separate the nasal lumen

from the dense blood-vessel network in the lamina propria. The respiratory epithelium

covered by a mucus layer is the major lining of the human nasal cavity and is essential in

the clearance of mucus by the mucociliary system.

Various structurally different mucoadhesive polymers were tested for their ability

to retard the nasal mucociliary clearance in rats [Zhou et al., 1996]. The clearance was

measured using microspheres labeled with a fluorescent marker incorporated into the

formulation. The clearance rate of each polymer gel was found to be lower than that of a

control microsphere suspension, resulting in an increased residence time of the gel

formulations in the nasal cavity. Ilium et al. [1994] evaluated chitosan solutions as

delivery platforms for nasal administration of insulin to rats and sheep. They reported a

concentration-dependent absorption-enhancing effect with minimal histological changes

of the nasal mucosa. Oechslein et al. [1996] studied various powder formulations of

mucoadhesive polymers for their efficacy to increase the nasal absorption of octreotide in

rats. The chitosan delivery systems can reduce the rate of clearance from the nasal cavity,

thereby increasing the contact time of the delivery system with the nasal mucosa and

providing the potential for raising the bioavailability of drugs incorporated into these

43

systems. Nakamura et al. [1999] described a microparticulate dosage form of budesonide,

consisting of bioadhesive and pH-dependent graft copolymers of PMAA and PEG,

resulting in elevated and constant plasma levels of budesonide for 8 h after nasal

administration in rabbits.

2.4.3 Ocular route

The ocular route is mainly used for the local treatment of eye pathologies. Many

physiological constraints prevent a desired drug delivery to the eye due to its protective

mechanisms, such as effective tear drainage, blinking and low permeability of the cornea.

Therefore, conventional eyedrops containing a drug solution tend to be eliminated rapidly

from the eye, and the drugs administered exhibit limited absorption, leading to poor

ophthalmic bioavailability (2-10%). Additionally, their short retention often results in a

frequent dosing regimen to achieve the therapeutic efficacy for a sufficiently long

duration. These challenges have motivated researchers to develop drug delivery systems

to provide a prolonged ocular residence time of drugs.

The following types of mucoadhesive formulations have been evaluated for ocular

drug delivery: viscous liquids (suspensions and ointments), hydrogels, and solids

(inserts). Certain dosage forms, such as suspensions and ointments, can be retained in the

eye, although these sometimes give patients an unpleasant feeling because of the

characteristics of solids and semi-solids. Due to their elastic properties, hydrogels can

also represent an ocular drainage-resistant device. In particular, in-situ hydrogels are

attractive as an ocular drug delivery system because of their facility in dosing as a liquid,

and their long-term retention property as a gel after dosing. Hui and Robinson [1985]

44

introduced hydrogels consisting of crosslinked PAA for ocular delivery of progesterone

in rabbits. These preparations increased progesterone concentrations in the aqueous

humor four times over aqueous suspensions. Cohen et al. [1997] developed an in situ gel

system of alginate with high guluronic acid contents for the ophthalmic delivery of

pilocarpine. This system significantly extended the duration of the pressure-reducing

effect of pilocarpine. Carlfors et al. [1998] investigated the rheological properties of the

deacetylated gellan gum, which gels upon instillation in the eye due to the presence of

cations, and indicated that a high rate of the sol/gel transition of in-situ gels results in

long precorneal contact times. An approach to ocular inserts was presented by Chetoni et

al. [1998] in a study of cylindrical devices for oxytetracycline, made from mixtures of

silicone clastomer and grafted on the surface of the inserts with an interpenetrating

mucoadhesive polymeric network of PAA or PMAA. The ocular retention of IPN-grafted

inserts was significantly higher than the ungrafted ones. An in-vivo study using rabbits

showed a prolonged release of oxytetracycline from the inserts for several days.

2.4.4 Rectal and vaginal routes

The rectal and vaginal routes are considered to be suitable for the local

application and absorption of therapeutics, although patient acceptability is a variable due

to the discomfort arising from administered dosage forms. The drugs are absorbed from

these specific sites and into the circulation directly. Thus, the rectal and vaginal routes are

useful for drugs suffering heavy first-pass metabolism. Conventional delivery systems at

both sites include tablets, foam gels, suppositories. Typical suppositories hitherto adapted

as dosage forms are solids at room temperature, and melt or soften at body temperature.

45

However, an uncontrolled release pattern of drugs leads to short residence time at the

specific position, and a variation of the bioavailability of certain drugs.

Mucoahesive hydrogels may offer a valuable way to overcome the problem in

conventional suppositories, providing a sufficient bioadhesive property. Mucoadhesive

gel formulations based on polycarbophil have been reported to remain 3–4 days at the

vaginal tissue, providing an excellent vehicle for the delivery of progesterone and

nonoxynol-9 [Robinson et al., 1994]. To improve the propranolol bioavailability, Ryu et

al. [1999] added certain mucoadhesive polycarbophil and sodium alginate to

poloxamer-based thermally gelling suppositories. The largest mucoadhesive force and the

smallest intrarectal migration for the suppositories resulted in the largest bioavailability

of propranolol. Miyazaki et al. [1998] investigated the potential application of xyloglucan

gels with a sol-gel transition temperature of around 22-27C as vehicles for rectal drug

delivery. This thermal gelling property provided easy administration at room temperature

and a gel status at body temperature. In-vivo rectal administration of indomethacin

showed a well-controlled drug plasma concentration-time profile without reduced

bioavailability.

2.4.5 Transdermal route

A transdermal route has been considered as a possible site for the systemic

delivery of drugs. The possible benefits of transdermal drug delivery include ease of

access, applying, and easing the delivery, sustained and steady drug release, reduced

systemic side effects, avoidance of drug degradation in the GI tract and first-pass hepatic

metabolism. Furthermore, swollen hydrogels with a high water content can provide a

46

better feeling for the skin in comparison to conventional ointments and patches. Versatile

hydrogel-based devices for transdermal delivery have been proposed. Sun et al. [1997]

prepared composite membranes comprising of crosslinked PHEMA with a non-woven

polyester support. Depending on the preparation conditions, the composite membranes

could be tailored to give a permeation flux ranging from 4 to 68 mg/cm2 per h for

nitroglycerin. Gayet and Fortier [1996] reported the use of the BSA-PEG hydrogels

containing high water content over 96% as controlled release devices in the field of

wound dressing. However, the skin functions naturally as a barrier to foreign substances,

preventing the entrance of the majority of drugs. Therefore, researchers are developing

various electrically assisted methods to enhance the drug permeation across the skin. The

notable technologies include electroporation, ionophoresis, sonophoresis, and laser

irradiation [Bellhouse et al., 2003; Mehier-Humbert et al., 2005; Prausnitz et al., 2004].

Several hydrogel-based formulations are being investigated as vehicles for transdermal

iontophoresis to obtain the enhanced permeation of hormone [Chen et al., 1996] and

enoxacin [Fang et al., 1999]. A methyl cellulose-based hydrogel was used as a viscous

ultrasonic coupling medium for transdermal sonophoresis assisted with an AC current,

resulting in an enhanced permeation of insulin and vasopressin across human skin in vitro

[Zhang et al., 1996].

2.4.6 Trends and perspectives

In this chapter, a number of sensitive hydrogels with various applications have

been described as novel drug delivery platforms. These polymers, as useful drug carriers,

or as safe absorption enhancers, or as improved mucoadhesive hydrogels, are the recent

47

developments in drug delivery platforms for intestinal absorption of drugs. Another trend

observed during the past few years is the new methods of preparation of hydrogels with

desirable functional groups that may be used in the future for drug delivery applications.

For example, novel biodegradable polymers include polyrotaxanes, which are considered

potentially useful for molecular assemblies for drug delivery. In the synthesis, choice of

new functional monomers and adjusting of hydrophobicity/hydrophilicity of copolymers

can be used to better control the swelling/deswelling behavior of novel gels. Moreover,

graft, block, and comb-like copolymerizations offer better advantages and the produced

novel gels have the interesting applications for treatment of diabetes, osteoporosis, cancer

or thrombosis.

Besides the development of novel materials for drug delivery, applications of

functional hydrogels as the promising materials have been extended in biomedical and

pharmaceutical fields when combined with the advanced manufacturing techniques, such

as micro- and nanoscale machining techniques. Drug delivery technology can be brought

to the next level by the fabrication of ‘smart materials’ into ‘miniature devices’ that are

‘responsive’ to the individual patient’s therapeutic requirements and able to deliver a

certain amount of a drug in response to a biological state. Bures and Peppas [2001]

have prepared gels of controlled structure and large biological functionality by irradiation

of PEO star polymers. Combined with the techniques of molecular imprinting. Such

highly crosslinked gels with the sending/activation mechanism may lead to a variety of

new, and robust biomolecular sensing hydrogel networks for drug delivery.

Gene therapy with the broad potential has been heavily investigated during last 15

years. Many types of polymers are specifically designed for gene delivery. Gene therapy

48

requires the identification of a therapeutic gene and the transfer of the gene to target cells

with high efficiency and without hazard for the patients. Hydrogels are designed to

address a specific intracellular barrier based on their stability/degradability,

biocompatibility, and sensitivity. Hoffman’s group [Pack et al., 2005] developed one class

of hydrogel carriers to reversibly control membrane stability in response to sharp pH

changes for delivering proteins and nucleic acids to intracellular compartments in gene

delivery. To enhance the transfection efficiency of gene into mammalian cells, a new

system of plasmid DNA release with a biodegradable hydrogel is described while the

biological activity of a plasmid DNA of hepatocyte growth factor is augmented by the use

of the release system [Kushibiki et al., 2004]. All these promising applications

demonstrate that the use of functional hydrogels is a powerful strategy to improve the

controlled drug delivery and may benefit the human being.

49

CHAPTER 3

PHOTOPOLYMERIZATION AND STRUCTURE FORMATION OF PMAA

HYDROGELS IN WATER/ETHANOL MIXTURE

SYNOPSIS

Hydrogels are a desired material for biomedical and pharmaceutical applications.

To better control the synthesized hydrogels for various applications, it is necessary to

have a thorough understanding of hydrogel structure and reaction mechanism. In this

study, pH-sensitive hydrogel networks consisting of methacrylic acid (MAA) crosslinked

with tri(ethylene glycol) dimethacrylate (TEGDMA) were synthesized by free-radical

photopolymerization in the water/ethanol mixture. Reaction rate was measured using

Photo-Differential Scanning Calorimetry (PhotoDSC) with a modified sample pan

designed for handling volatile reagents. A photo-rheometer and a dynamic light scattering

(DLS) goniometer were used to follow the changes in viscosity and molecule size of the

resin system during photopolymerization. It was found that the rate of polymerization

increased and more compact and less swelling gels would form with a higher water

fraction in the 50wt% solvent/reactant mixture. This is because the weaker interaction

50

between monomer and solvent gives a higher opportunity for propagation and a higher

reaction rate. And the hydrophobic TEGDMA and initiator tend to form aggregates in the

higher water solution, contributing to the inhomogeneous microgel formation. This

mechanism is conformed by viscosity measurement, DLS analysis, scanning electron

microscopy (SEM) observation, and kinetics analysis.

3.1 Introduction

Hydrogels are a desired material for biomedical and pharmaceutical applications

due to their unique swelling properties and structures. The highly hydrated structure and

good biocompatibility make them suitable for contact lenses, biosensors, artificial organs,

and drug delivery devices [Peppas, 1997; Peppas et al., 2000]. In drug delivery,

functional hydrogels may release drugs in an aqueous median at regulated rate by

controlling the synthesis conditions such as the method of polymerization, the

crosslinking ratio, and the solvent composition.

Hydrogels are often synthesized by UV photopolymerization [Lu et al., 1999;

Ward et al., 2001] and redox polymerization [Hassan et al., 1999]. Photopolymerization is

favored because hydrogels can be synthesized at temperatures and pH conditions near

physiological conditions and even in the presence of biologically active materials.

Furthermore, photopolymerization can be easily controlled by adjusting the dosage and

intensity of UV light, and the curing temperature [Crump, 2001]. Photo-Differential

Scanning Calorimetry (PhotoDSC) is the most widely used technique to characterize the

photopolymerization kinetics. A great deal of research has been carried out using this

approach for photocurable materials. However, the application of this technique for

51

highly volatile reagents is limited since uncovered sample pans lead to significant sample

loss during measurement. Some researchers applied unsealed polyethylene (PE) films

over the sample pan to reduce the sample loss [Ward et al., 2001], while others used the

sample weight after the reaction to correct for the measurement error resulting from

reagent evaporation [Jakubiak et al., 2000]. The results from such treatments are doubtful

because sample loss during the reaction is a time-dependent process. When preparing the

carriers for drug delivery, solvents like water and ethanol are often used in the synthesis

to control the hydrogel structure. Evaporation of highly volatile solvents like ethanol

makes it impossible to study the reaction kinetics using the existing approaches. We have

recently developed a modified DSC sample pan [Li et al., 2005]. Sample loss during

reaction is minimized, and loaded samples are much more uniform over the sample

surface. This new method is applied in this study.

To better control the synthesized hydrogels for various applications, it is essential

to understand how the polymerization conditions, chemical structure of reactants and

their composition, and solvent type and concentration affect the reaction and the resulting

properties of hydrogels. A number of studies have reported that varying curing conditions

may achieve different gel structures and swelling properties [Lowman et al., 1997;

Anseth et al., 1996; Peppas et al., 1991], and the compatibility between the solvent and

the resin may affect inter-molecular and primary cyclization of multi-vinyl monomers

during the polymerization, and, consequently, the hydrogel properties [Kwok et al., 2003;

Elliott et al., 2002; Elliott et al., 2001]. However, there lacks a thorough understanding on

the interactions of reaction kinetics, rheological changes, hydrogel structures, and

solvent-resin compatibility. In this chapter, PMAA gels synthesized in a water/ethanol

52

mixture were investigated by using a series of analytical tools including PhotoDSC,

photo-rheometry, dynamic light scattering goniometry, and scanning electron microscopy

of freeze-dried hydrogels.

3.2 Experimental

3.2.1 Materials and sample preparation

The monomer, MAA (Sigma-Aldrich) and the crosslinking agent, TEGDMA

(Sigma-Aldrich) were used to prepare pH-sensitive hydrogels. For all reactions, the

crosslinking agent was presented at a level of 1.0 mole% based on the total mole of

monomers. A photoinitiator, 2,2-dimethoxy-2-phenylacetophenone (Irgacure 651, Ciba

Specificity Chemicals), was used at 1.0 wt% of the monomer mixture. The free-radical

photopolymerization was carried out in a mixed solvent of distilled water and ethanol

with varying ratios. The ratio of monomer to solvent was kept at 50:50 (w/w). All

reagents, unless specified, were of anylytical grade and were used without further

purification.

To prepare hydrogel films for the swelling test and structure analysis, 5.0 grams

of MAA were mixed with a proper amount of TEGDMA and initiator. An equal weight of

solvent mixture was then added. The solution was transferred to a glove box where it was

kept under a nitrogen atmosphere. Nitrogen was bubbled through the solution for 20

minutes. Then the mixture was pipetted between two glass slides separated by a Teflon

spacer. The thickness of the spacers was 0.3 mm. The setup was then placed under a UV

light for photopolymerization at 2.0 mw/cm2. The cured hydrogels were then rinsed in

double deionized water for 5 days to remove unreacted monomer, initiator and sol

53

fraction. Subsequently, the monomer-free films were cut into samples with a 5.0 mm

diameter for swelling test.

3.2.2 Modification of DSC pans

A poly(dimethyl siloxane) (PDMS) curing kit (Sylgard®184 silicone kit, Essex

Group Inc.) was prepared and dissolved in hexane to form a 0.05 g/ml PDMS solution.

About 10 µl PDMS solution was placed in the DSC pan, which quickly spread to the

inner corner of the pan by capillary forces. After solvent evaporation, the pan was heated

at 60oC for 4 hours to cure the PDMS resin. The cured PDMS formed a thin layer of

O-ring-like hydrophobic film inside the pan, as shown in Figure 3.1(a). This PDMS ring

can prevent the hydrophilic sample from flowing towards the inner corner during sample

loading. Through this treatment, the loaded resin sample can form a thin film with

uniform thickness, essential for consistent UV irradiation.

To minimize the sample weight loss during measurements, the sample pan was

further modified as shown in Figure 3.1(b). The PhotoDSC pan was placed face-down

and adhere to a layer of photo-safe, double-sided Scotch tape. A small amount of

partially-cured HEMA/DEGDMA/PI solution was applied around the outside edge of the

pan, which was then completely cured under the UV light. The cured poly(HEMA)

formed an edge around the open pan. The Scotch tape in the center above the original pan

was removed by a razor. After loading the sample, the pan was covered with a layer of

polyethylene (PE) film and sealed by the double-sided Scotch tape along the edge area.

54

(a)

(b)

Figure 3.1 (A) DSC pan treated with PDMS; (B) Seal of DSC pan [Li et al., 2005].

A layer of cured PDMS

DSC pan

HEMA/DEGDMA Double-sided Scotch tape

hv hv (ii) photocure the edge

(i) apply photocurable material around the pan

(iii) remove the Scotch tape in the center

(iv) pan sealed by PE film

PE film to seal the pan cover

Monomer solution

55

3.2.3 PhotoDSC measurement

The reaction kinetics and heat of reaction of PMAA gels were measured using a

PhotoDSC (TA 2920, TA Instruments). A UV light source (Novacure, 100W Hg short-arc

lamp, EXFO, Mississaugua, Ont., Canada) was used to cure the samples. In order to

prevent the weight loss of volatile MAA and ethanol, the DSC pans were physically and

chemically modified by using the technique described elsewhere [Li et al., 2005]. We

compared the performance of modified sample pans vs. the ones covered with a layer of

PE film. A micropipette was used for PhotoDSC sampling (5~7 µl), which controlled the

sample weight for each test. All measurements were carried out at 30oC and the light

intensity was kept at 2.0 mw/cm2. Each run was conducted by purging the sample with

nitrogen gas until reaching equilibrium (around 2 minutes), and then UV irradiation was

applied to induce the free-radical polymerization.

The DPC measured the heat flow per unit mass as a function of time. The rate of

polymerization, Rp, was calculated by dividing the measured heat flow per unit mass by

the theoretical enthalpy. The units of Rp were fractional double bond conversion per

second. Integration of Rp curve versus time provided the conversion as a function of time.

It is assumed that in the polymerization of two monomers, the functional groups have

equal reactivity. In other words, the theoretical enthalpy derived for a comonomer

mixture is an average of the enthalpies of the individual monomers.

3.2.4 Rheological measurement

A photo stress rheometer MCR 300 (Physica, Anton Paar) was used to follow the

viscosity change during the isothermal photopolymerization. A UV cell, including a top

56

steel plate with a diameter of 50 mm and a bottom plate made of quartz glass, was

utilized in this test. The UV light source (Acticure 4000, EXFO, Canada) was illuminated

from the bottom. The light intensity on the sample surface was kept at 2.0 mw/cm2. The

gap between the two plates was set at 1.0 mm and the shear rate used was 0.1s-1. The gel

point was assumed when the relative viscosity, i.e. viscosity of the reactive resin vs. its

initial viscosity, reached 104.

3.2.5 Dynamic light scattering analysis

Dynamic light scattering (DLS) measurements at 30°C were carried out to

determine the molecule size and size distribution before gelation during

photopolymerization by using a BI-DNDC Differential Refractometer (Brookhaven

Instruments) with a 10 mW He-Ne laser beam at a wavelength of 633 nm. A scattering

angle was held constant at 90°in the measurement. Before the DLS analysis, the

partially reacted sample (around 0.3 ml) was dispersed in 3 ml of ethanol, and the diluted

solution was then filtrated through a filtration unit with 0.45-micron pore size (Whatman

Puradisc 25TF). Count rates between 10 to 200 kilocounts per second were used to obtain

meaningful results by changing the sample concentration and adjusting the laser power.

Autocorrelation of the intensity was carried out by the method of cumulate analysis to

obtain an average diameter of the molecules and the polydispersity. The molecule size

distribution was obtained from the correction function by CONTIN analysis using the

standard software BI-DNDCW.

57

3.2.6 Swelling studyies

The swelling tests were performed at various pH values ranging from 2.6 to 7.4 to

characterize the swelling behavior for synthesized pH-sensitive hydrogels. The buffer

solutions with different pH values were prepared by mixing the citric acid with

appropriate amounts of sodium phosphate solution. Sodium chloride was used to adjust

the ionic strength of all solutions to I=0.1M, which is the near-physiological condition.

The dried hydrogel samples were weighed and placed in the buffer solution at room

temperature (25°C). The samples were taken out of the solution at pre-selected time

intervals. After the extra water on the surface was removed by laboratory tissue, the

weight of the wet hydrogels was measured. The weight-swelling ratio was calculated by

the weight of the swollen sample to the weight of the dried sample. The samples were

blotted and weighed until the weight change is less than 0.1 mg over a 24-hour period.

3.2.7 Scanning electron microscopy characterization

To visually examine the surface and interior morphology of hydrogels in the swollen

state, a Hitachi Model S-4300 SEM was used to analyze the pore structure. The samples

cured under UV radiation were first swollen to reach equilibrium in buffer solutions for

24 hours, and then quickly frozen below its freezing point using liquid nitrogen. The

sample containers were transferred to a freeze dryer (Labconco 75150, Labconco Inc.

Kansas City, MI) and freeze-dried for 48 hours until all solvent was sublimed. The

freeze-dried samples were loaded on the surface of an aluminum SEM specimen holder

and sputter coated with gold for 40 s (Pelco Model 3 Sputter Coater) before observation.

A working distance about 8-10 mm, an accelerating voltage of 10 KV, and a chamber

58

pressure of 10-8 Torr were found to be suitable for obtaining high-resolution images of

hydrogel samples. The magnification in this study varied from 2000× to 20,000×

depending on the network structure.

3.3 Results and Discussions

An important feature of this curing system was the formation of heterogeneous

structure in different solvent compositions, which influenced not only the reaction

kinetics and rheological changes of the resin, but also the swelling behavior and network

structure of the formed gels.

3.3.1 Kinetics of MAA/TEGDMA photopolymerization

To minimize the sample weight loss during DSC measurements, the sample pan

was physically and chemically modified. The advantage of such treatment was

demonstrated via the photopolymerization of the MAA/TEGDMA system. The measured

heat flow by using both modified and un-modified pans is shown in Figure 3.2. With a

modified sample pan, an equilibrium state was reached in about 1-2 minutes, and the

measurement started at a level close to the “zero” heat flux. While, with a regular sample

pan covered with a layer of PE film, there was a continuous endotherm due to the

evaporation of monomers and solvents, leading to a negative starting point for heat flux.

Additionally, a longer time was needed to reach equilibrium, which would inevitably

cause more weight loss. For systems containing highly volatile MAA and ethanol, a

strong competition occurred between sample evaporation and chemical reaction.

Consequently, a complete change in the reaction rate profile was observed with the use of

59

an un-modified DSC pan. The sample weights before and after the test showed that there

was less than 5% weight loss using a modifies pan, compared to about 40% loss using an

un-modifies pan (the data represents the mean of six samples). It is clear that the

modified pans have to be used in the DSC kinetic analysis of volatile monomers and

solvents.

Using the modified pans, the effect of solvent composition on the reaction

kinetics of MAA/TEGDMA was investigated. Figure 3.3(A) illustrates the reaction rate

versus reaction time for the isothermal photopolymerization of MAA/TEGDMA (1.0

mole%TEGDMA, 50 wt.% solvent) with different solvent compositions at 30 ºC and a

UV intensity of 2.0 mW/cm2. As can be seen, the solvent composition had a great

influence on the reaction kinetics of the photocurable MAA/TEGDMA system. With an

increase of the ethanol content in the solvent mixture, the polymerization rate decreased

correspondingly, and multiple exothermic peaks were observed on the reaction rate

profiles for all cases. A peak occurred at the very early stage of polymerization, followed

with a stronger second peak. A higher ethanol content delayed and broadened the first

peak and substantially reduced the second peak. It is also noted from the conversion

profiles shown in Figure 3.3(B) that the higher ethanol content delayed the time to

achieve a high conversion.

60

Figure 3.2 Comparison of PhotoDSC measurements by using a modified and an un-modified pan at UV intensity of 2.0 mw/cm2 in the MAA/TEGDMA system (1.0

mole%TEGDMA, 50 wt.% solvent mixture of the 1/1 water/ethanol ratio).

-4

-2

0

2

4

0 5 10 15 20Time (min)

Hea

t F

low

(m

w)

un-modified

modified

61

The multiple peaks observed in the free radical crosslinking polymerization have

been reported for several mono- and divinyl monomers [Jakubiak, 2000; Li et al., 2005;

Lai et al., 1997; Horie et al. 1975; Cook, 1993; Anseth et al., 1994]. Horie and coworkers

postulated that the double maxima in the reaction rate of MMA/EGDM systems were

caused by microgel formation. They attributed the first peak to the Trommsdorff effect in

the bulk material while the resin mixture was still homogeneous, and the second one to

the Trommsdorff rate acceleration in the microgels. As the polymerization proceeded

further, the system viscosity limited propagation and the autodeceleration in the reaction

rate occurred, as monomer could not diffuse to the relatively immobile radicals. Such

hypothesis has also been used to interpret the occurrence of multiple reaction peaks in the

acrylic acid (and N-vinylpyrrolidone) copolymerization with TEGDMA [Jakubiak,

2000], in the photopolymerization of HEMA/glycerin [Horie et al. 1975], in the

photopolymerization of a series of oligo(methylene) oxide and oligo (ethylene oxide)

dimethacrylates [Cook, 1993], and in the reaction between multifunctional methacrylate

and acrylate monomers [Anseth et al., 1994]. Although our kinetics results show a

similar trend, the viscosity and molecule size analysis presented in the next section,

however, show a different mechanism.

62

Figure 3.3 (A) Reaction rate and (B) conversion versus reaction time for the isothermal photopolymerization of MAA/TEGDMA (1.0 mole%TEGDMA, 50 wt.% solvent) with

different solvent compositions at 30ºC and UV intensity of 2.0 mW/cm2.

0

0.001

0.002

0.003

0.004

0.005

0.006

0 5 10 15 20Tim e(m in)

Rea

ctio

n R

ate(

1/s)

9/1

4/1

1/1

1/4

a

b

c

a’ b’

c’

0

0.001

0.002

0.003

0.004

0.005

0.006

0 5 10 15 20Tim e(m in)

Rea

ctio

n R

ate(

1/s)

9/1

4/1

1/1

1/4

a

b

c

a’ b’

c’

(A)

Water to ethanol weight ratio:

0

0.2

0.4

0.6

0.8

1

0 5 10 15 20

Time(min)

Con

vers

ion

9/1

4/1

1/1

1/4

(B)

Water to ethanol weight ratio:

63

3.3.2 Viscosity measurement and molecule size analysis

In order to evaluate the effect of solvent composition on the polymeric structure

formation, rheological and DLS measurements were carried out to follow the viscosity

change and the growth of molecule size during photopolymerization. Figure 3.4(A)

displays both the relative viscosity and reaction rate as a function of double bond

conversion for PMAA gels with different solvent compositions. Approaching the gel

point, there was the steep increase of relative viscosity (104). For the gels with the

water/ethanol ratio of 1/4, the macrogelation occurred at 9 minutes or around a

conversion of 78%. With an increase of water content, the curves of relative viscosity

shifted to a higher conversion. Figure 3.4(B) presents the corresponding gel time and gel

conversion versus water content based on the weight of solvent mixture. The gelation

time was linearly decreased and the gel conversion was increased with the increasing

water content. For the system with the highest water content (90 wt.%), it only took

around 5.5 minutes to reach the gel point. However, its gel conversion could reach 88%.

Figures 3.5(A) and (B) summarize the size distribution of polymers formed during

the photopolymerization of MAA/TEGDMA in ethanol. For MAA/TEGDMA with the

1/4 solvent ratio, the double bond conversion was around 78% at the gel point. The

macromolecules formed at a conversion of 23% (point ‘a’, the first maxima of reaction

rate in Figures 3.3A and 3.4A) exhibited a narrow unimodal size distribution, ranging

from 5 to 80 nm. The intensity reached the maximum value at 17.5nm. With the reaction

progressed to a conversion of 45% (point ‘b’, onset of the second autoacceleration in

Figures 3.3A and 3.4A), the peak was shifted to 64 nm. In addition, a bimodal size

distribution occurred, which contains a very narrow peak (13-32 nm) with the same

64

maximum value at 17.5 nm and a broader and larger size distribution (40-164 nm). A

further increase in the conversion to 76% (point ‘c’, before macrogelation) showed very

large clusters with the size distribution from 83 to 223 nm, while the intensity ratio of

smaller molecules decreased significantly. Apparently, most small molecules had

converted into larger clusters.

Compared with the system with the 1/4 solvent ratio, the size distribution curves

for the system with the 9/1 solvent ratio exhibited a similar shape and trend. Increasing

the water content in the solvent mixture shifted the polymer size distribution to a larger

size. For example, the formed polymer showed a unimodal size distribution at the same

conversion of 23%, point a’, and a bimodal size distribution around the onset of the

second autoacceleration, point b’, except that the molecule clusters were large. At a

conversion of 86%, point c’ which was close to the gel conversion, the peak for larger

molecules moved to 204 nm and the width of the distribution spread from 136 to 304 nm.

Obviously, the resin system with a higher water/ethanol ratio formed larger polymer

clusters under the same UV radiation when the reaction approached macrogelation.

65

Figure 3.4 (A) Reaction rate and viscosity rise as a function of conversion of MAA/TEGDMA (1.0 mole% TEGDMA, 50 wt.% solvent) with different solvent

compositions cured at UV intensity of 2.0 mW/cm2, (B) Gel time and gel conversion versus water/ethanol ratio in the solvent mixture.

0

0.002

0.004

0.006

0.008

0.01

0 0.2 0.4 0.6 0.8 1

Conversion

Rea

ctio

n R

ate(

1/s)

0

3000

6000

9000

12000

Rel

ativ

e V

isco

sity

II IIII IIIIII IVIV VV

a

bc

a’ b’ c’

0

0.002

0.004

0.006

0.008

0.01

0 0.2 0.4 0.6 0.8 1

Conversion

Rea

ctio

n R

ate(

1/s)

0

3000

6000

9000

12000

Rel

ativ

e V

isco

sity

II IIII IIIIII IVIV VV

a

bc

a’ b’ c’

II IIII IIIIII IVIV VV

a

bc

a’ b’ c’

(A)

◊ 9/1 ○ 4/1 □ 1/1 ∆ 1/4

Water to ethanol ratio:

(B)

2

4

6

8

10

W ater to Ethanol Ratio

Gel

Tim

e (m

in)

70

80

90

100

Gel

Con

vers

ion

(%)

1/4 1/1 4/1 9/12

4

6

8

10

W ater to Ethanol Ratio

Gel

Tim

e (m

in)

70

80

90

100

Gel

Con

vers

ion

(%)

1/4 1/1 4/1 9/1

66

Figure 3.5 The size distribution of MAA/TEGDMA resin (1.0 %TEGDMA, 50 wt.% solvent) with different solvent ratios of water/ethanol: (A) 1/4 and (B) 9/1 cured at light

intensity of 2.0 mW/cm2.

(A)

0

30

60

90

120

0 100 200 300 400

Diameter(nm)

Inte

nsi

ty

3.00min, 23%

5.50min, 45%

8.80min, 76%

(a’) (b’) (c’)

(B)

0

30

60

90

120

0 100 200 300 400Diameter(nm)

Inte

nsi

ty

2.09min, 23%4.00min, 51%5.40min, 86%

(a) (b) (c)

67

3.3.3 Mechanism for gelation

It is well known that free-radical polymerization of multifunctional monomers

forms heterogeneous polymer networks, leading to microgel formation [Hsu et al., 1993;

Chiu et al., 1995; Sun et al., 1997]. Such entities are a result of strong intramolecular

crosslinking of the growing macroradicals. Eventually, intermolecular reactions among

microgels form the network structure. The relative rates of intra- and intermolecular

reactions depend on the initial monomer composition, as well as other reaction conditions.

The solvent composition is a major factor influencing the gelation kinetics. According to

the experimental results shown in the previous section, the photopolymerzation of

MAA/TEGDMA system can be described in five stages: initiation, microgel formation,

cluster formation, macro-gelation, and post-gelation. The schematic diagram of structure

formation in the MAA/TEGDMA photopolymerization describing the first four stages is

given in Figure 3.6.

In the first stage, all reactants are mixed together and UV radiation initiates

initiator decomposition to form radicals (shown as filled dots). In the MAA/TEGDMA

system with a good solvent, such as the one with a high-ethanol content (ethanol is a

good solvent for both hydrophilic MAA and hydrophobic TEGDMA and Irgacure 651

due to its participation in both hydrogen bonding and hydrophobic interactions), a

homogeneous solution is formed with uniform distribution of all reactants. While in a

poor solvent with a high water content, TEGDMA tends to form a micelle-like structure

due to the amphiphilic properties. Its hydrophilic ends prefer to be in contact with the

water phase by hydrogen bonding while the hydrophobic area is located in the center,

68

Figure 3.6 The schematic diagram of structure formation of MAA/TEGDMA with different solvent qualities.

I) Initiation

II) Microgel formation

III) Cluster formation

IV) Macrogelation

=

=

=

=

=

=

=

=

Good solvent (high ethanol content)

Poor solvent (high water content)

=

=

=

=

= ==

=

=

=

= =

=

=

=

=

=

=O

C–C=CC

C=C–C

=OC

=O

C–C=CC

C=C–C

=OC

C–C=CC

C=C–C

=OC

=O

C–C=CC

C=C–C

=O

C

=O

C–C=CC

C=C–C

=O

CC–C=C

CC=C–C

=O

C

=O

C–C=CC

C=C–C

=O

C

=O

C–C=CC

C=C–C

=O

CC–C=C

CC=C–C

=O

C

=O

C–C=CC

C=C–C

=O

C

=O

C–C=CC

C=C–C

=O

CC–C=C

CC=C–C

=O

C

=O

C–C=CC

C=C–C

=O

C

=O

C–C=CC

C=C–C

=O

CC–C=C

CC=C–C

=O

C

=O

C–C=CC

C=C–C

=O

C

=O

C–C=CC

C=C–C

=O

CC–C=C

CC=C–C

=O

C

= =

= =

=

= =

=

=

C–C=C

=O C

C=C

–C =O

C

C–C=C=O

C

C=C–C

=O

C

C–C

=C

=O

C

C=C–C

=O

CC–C=C

=O C

C=C

–C =O

C

C–C=C

=O CC–C=C

=O C

C=C

–C =O

CC

=C–C =O

C=C

–C =O

C

C–C=C=O

C

C=C–C

=O

C

C–C=C=O

C

C–C=C=O

C

C=C–C

=O

CC=C–C

=O

C=C–C

=O

C

C–C

=C

=O

C

C=C–C

=O

C

C–C

=C

=O

CC

–C=C

=O

C

C=C–C

=O

CC=C–C

=O

C=C–C

=O

C

C–C=C

=O C

C=C

–C =O

C

C–C=C=O

C

C=C–C

=O

C

C–C

=C

=O

C

C=C–C

=O

CC–C=C

=O C

C=C

–C =O

C

C–C=C

=O CC–C=C

=O C

C=C

–C =O

CC

=C–C =O

C=C

–C =O

C

C–C=C=O

C

C=C–C

=O

C

C–C=C=O

C

C–C=C=O

C

C=C–C

=O

CC=C–C

=O

C=C–C

=O

C

C–C

=C

=O

C

C=C–C

=O

C

C–C

=C

=O

CC

–C=C

=O

C

C=C–C

=O

CC=C–C

=O

C=C–C

=O

C

2nm

=

=

=

= = = = =

= = =

20nm

90nm

=

100nm

MAA = Free radical

=O

C–C=CC

C=C–C

=O

C

=O

C–C=CC

C=C–C

=O

CC–C=C

CC=C–C

=O

CTEGDMA

69

where most Irgacure 651 molecules are located. This initial structure is verified by the

DLS measurement of MAA/TEGDMA mixtures without UV radiation shown in Figure

3.7. In the MAA/TEGDMA system with the 1/4 solvent ratio, no “particles” were

observed in the DLS analysis. On the other hand, in the system with the 9/1 solvent ratio,

a peak about 6 nm was observed with or without Irgacure 651, supporting the complex

formation by amphiphilic TEGDMA.

After initiation, radicals react with monomers to produce monomeric radicals.

Because of the presence of multifunctional monomers, the monomeric radicals have

chances to link with these molecules to form the growing macroradicals with pendant

double bonds, leading to the cyclization or ring formation through intramolecular

reactions. The intramolecular reactions consume vinyl groups, but do not contribute to

the increase of molecule weight and macroscopic network formation. This internal

crosslinking on the primary polymer chains leads to the formation of “microgels” [Dusek

et al., 1980]. Inside the microgels, the Trommsdorff effect may occur because termination

is largely hindered due to immobilized polymerical radicals, while the propagation rate is

less affected since small MAA monomers are still mobile. This leads to a small peak or

shoulder in the early stage of the reaction profiles. However, the relative viscosity

remains nearly unchanged. The greater extent of intramolecular cyclization means less

intermolecular crosslinking. This leads to larger mesh size in formed hydrogels, and the

weaker mechanical properties. This mechanism of intramolecular cyclization has been

used to explain the network formation influenced by the light intensity [Li et al., 2005],

the solvent concentration [Elliott et al., 2001], the solvent quality [Kwok et al., 2003;

Elliott et al., 2002], and the curing temperature [Chiu et al., 1995].

70

Figure 3.7 The size distribution of MAA/TEGDMA monomer solution (1.0 %TEGDMA, 50 wt.% solvent) with different compositions.

0

30

60

90

120

0 20 40 60 80 100

Diameter(nm)

Inte

nsi

ty

MAA/TEGDMA (9/1, no Irgacure 651)

MAA/TEGDMA (9/1, Irgacure 651)

MAA/TEGDMA (1/4, Irgacure 651)

6nm

71

In the solvent mixture, it is favorable for ethanol to participate in the formation of

hydrogen bonding with MAA molecules. Thus, more ethanol indicates stronger

interaction with the MAA molecules. According to the theory of complex [Henrici-Olive

et al., 1965], the propagating macroradicals continually interacts with the surrounding

medium (i.e. monomer and solvent). The stronger the interaction between the MAA and

the solvent, the lower the overall rate of polymerization since the propagation can only

take place if the propagating macroradical is in the vicinity of the monomer molecules.

Therefore, the high ethanol content in good solvent system gives a less opportunity for

propagation and a lower reaction rate under the UV radiation. Additionally, the uniform

distribution of TEGDMA and radicals increase the distance between radicals and free

vinyls or pendant vinyls, resulting in a high extent of intramolecular cyclization and

smaller microgels with loose structure. On the other hand, there is a higher reaction rate

of adding monomers onto the growing radicals and a fast microgel formation in the poor

solvent. And the localized TEGDMA and radicals leads to a high extent of intermolecular

crosslinking and larger microgels with smaller mesh size. The solvent composition has

little effect on the solution viscosity at this stage since microgel formation does not

significantly affect bulk properties in the solution.

During the cluster formation stage (stage III), the reactive microgels with pendant

double bonds may react with free monomers and other microgels to form larger clusters,

resulting in a bimodal molecular size distribution. The Trommsdorff effect in the clusters

leads to the second autoacceleration in the reaction profiles. At the later part of this stage,

the presence of a larger number of clusters and the inter-connection of some clusters lead

to an increase of solution viscosity.

72

As a macroscopic polymeric network is formed by chemical or physical

crosslinking, the resin system reaches the gel point in stage IV. Approaching the gel point,

most small microgels have converted to the larger clusters and intermolecular reactions

among these clusters finally lead to macrogelation. For the transition from microgels to

macrogels, intermolecular crosslinking reactions require the displacement of neighboring

solvent molecules from the vicinity of the microgels. In the system with a higher water

content, the microgels can easily form larger aggregates at a higher reaction rate due to

the weaker interaction between the microgels and solvent mixture. Therefore, the

MAA/TEGDMA with the 9/1 solvent ratio exhibited the shortest gel time and the highest

gel conversion as shown in Figure 3(B). While the uniformly distributed smaller

microgels in a system with a higher ethanol content have less chance to connect with

each other, taking longer time to reach the gel point. As the system entered the

post-gelation stage (V), the reaction rate abruptly decreased since both propagation and

termination became diffusion limited.

3.3.4 Swelling ratio and structural characterization

Figure 3.8 compares the equilibrium swelling ratio (SR) in different pH buffer

solutions for hydrogels synthesized with various solvent compositions. In all cases, the

hydrogel samples swelled more at higher pH due to the electrostatic repulsion between

the ionized forms of the carboxylic segments, as well as the dissociation of hydrogen

bonds between the carboxylic acid groups of MAA and the oxygen of the ether groups of

TEGDMA and the hydrophilicity of ionized molecules. Below a pH of 6.0, the swelling

ratio drastically decreased, indicating the hydrogel was in a relatively collapsed state

73

mainly due to the formation of hydrogen bonding. It is also interesting to note that the

gels with the highest ethanol content had the highest swelling ratio for a specific pH

value and its value reached approximately 33 at a pH of 7.3.

SEM technique is useful to reveal hydrogel structure, although the pre-treatment

of dehydration and/or fixation procedures for SEM examination may affect the

morphology of a hydrogel [Hong et al., 1998]. As shown in Figure 3.9, the pore structures

of the swollen interior of PMAA hydrogels are different depending on the solvent

composition. Figure 3.9(A) presents the SEM micrograph of PMAA hydrogel with the

9/1 solvent ratio. In a pH=7.4 buffer solution, this hydrogel (SR=10.0) exhibited mostly

circular and elliptical pores with smaller pores. Its pore size varies from very small to

very large pores, which may be a result of inhomogeneous reaction during

photopolymerization. On the other hand, the swollen gel with the 1/4 solvent ratio in the

same buffer solution showed larger and more uniform pores as shown in Figure 3.9(B).

Figure 3.10 shows the different morphology of swollen PMAA gels with the same

swelling ratio (SR=4.3) in the freeze-dried state. To obtain the same swelling ratio, the

gels with the solvent ratios of 1/4 and 9/1 were immersed in buffer solutions with the pH

values of 3.0 and 6.2, respectively. The gel with a higher water content displayed

smaller pores and much thicker pore walls at the same SR value.

These results are consistent with the solvent effect discussed in the previous

section. The localized reactants contribute to the formation of highly crosslinked network

structure in the poor solvent, leading to the smaller pores with thicker wall, while the

uniformly distributed reactants in a good solvent lead to a looser network structure,

forming larger pores with thinner wall.

74

Figure 3.8 Equilibrium swelling ratios of the PMAA (1.0 mole% TEGDMA) hydrogels

with different solvent ratios as a function of pH values.

0

5

10

15

20

25

30

35

2 3 4 5 6 7 8

pH

Wei

gh

t S

wel

ling

Rat

io (

g /

g)

1/41/14/19/1

Water to ethanol ratio:

75

Figure 3.9 SEM micrograph of swollen PMAA hydrogels (1.0 mole% TEGDMA, 50 wt.% solvent) with different swelling ratios (SR) in pH=7.4 buffer solution: (A) 9/1 and

(B) 1/4.

(A)

25 µm

SR=10.0

25 µm

SR=10.0

(B)

18–Jul – 05 s13 ×1.8k 25 um

25 µm

SR=33.0

18–Jul – 05 s13 ×1.8k 25 um

25 µm

SR=33.0

76

Figure 3.10 SEM micrograph of swollen PMAA hydrogels (1.0 mole% TEGDMA, 50 wt.% solvent) with the same swelling ratio (SR=4.3) in different buffer solution: (A) 9/1

in pH=6.2 buffer (B) 1/4 in pH=3.0 buffer.

5 µm

SR=4.3

5 µm5 µm

SR=4.3

(A)

(B)

5 µm

SR=4.3

5 µm5 µm

SR=4.3

77

3.4 Conclusion

This work clarified the role of the solvent composition in the photopolymerization

of hydrogels. The solvent composition has a great influence on the reaction kinetics of

photocurable MAA/TEGDMA system. With the increase of the ethanol content in the

solvent mixture, the photopolymerization rate and the gel conversion decreased, while the

gel time and the swelling ratio of PMAA hydrogels increased.

This can be explained by the solvent compatibility and interaction with the

reactants and the initiator. A less ethanol content indicated less compatibility of

TEGDMA and initiator and weaker interaction between MAA and solvent. This weaker

interaction led to a higher reaction rate and faster gel formation. The less compatibility

resulted in localized TEGDMA and initiator distribution. Since the localized TEGDMA

contributed to more highly crosslinked microgels, the resulting hydrogel had a lower

swelling ratio and less uniform pore distribution. This mechanism has been confirmed by

viscosity measurement, dynamic light scattering analysis, and SEM observation.

78

CHAPTER 4

PHOTOPOLYMERIZATION AND STRUCTURE FORMATION OF PMAA

HYDROGELS CURED AT VARIOUS LIGHT INTENSITIES

SYNOPSIS

Hydrogels are a desired material for biomedical and pharmaceutical applications

due to their unique swelling properties, the highly hydrated structure and good

biocompatibility. To better control the properties of synthesized hydrogels, it is necessary

to have a thorough understanding of hydrogel structure and reaction mechanism. The

solvent effect on the reaction kinetics and structure formation of pH-sensitive hydrogel

networks comprising a PMAA backbone crosslinked by TEGDMA has been discussed in

the previous chapter. In this chapter, the effect of light intensity on the reaction kinetics

and structure formation is addressed. A series of analytical tools including PhotoDSC,

photo-rheometry, and DLS goniometry were used for this study. The kinetics-gelation

mechanism based on the concept of microstructure formation is also discussed.

79

4.1 Introduction

Hydrogels with the highly hydrated structure and good biocompatibility have

been employed as contact lenses, artificial organs, and drug delivery devices [Peppas,

1997]. The volumetric shape memory capability makes hydrogels an ideal choice as

actuator, fluid pump, and valves in microfluidic devices [Osada et al., 1993; Seigel et al.,

1991]. In an aqueous environment, hydrogels will undergo a reversible phase

transformation that results in dramatic volumetric swelling and shrinking upon exposure

and removal of a stimulus, such as pH value. Typically, pH-sensitive hydrogels contain

carboxylic groups capable of uptaking a large amount of water above its pKa. Such

polymers mainly include poly(acrylic acid) (PAA) and poly(methacrylic acid) (PMAA).

Copolymers of PAA and PMAA with poly(ethylene glycol) (PEG), poly(vinyl alcohol)

(PVA), and PHEMA also exhibit the pH sensitivity due to the presence of carboxylic

segments. Additionally, incorporating other sensitive groups into the networks of PAA or

PMAA may give gels more interesting properties. For example, the copolymer of PAA

and PMAA with PNIPAAm can provide the environmental sensitivity of both pH and

temperature [Tian et al., 2003; Zhang et al., 2000]. Recently, a series of smart

biomaterials such as poly(ethyl acrylic acid) (PEAA) and poly(propyl acrylic acid)

(PPAA), has opened new opportunities for applications in the molecular imaging field

because of their sharp pH-sensitivity [Mourad et al., 2001].

PH-sensitive hydrogels exhibit swelling or deswelling behavior with changes of

pH values due to one of the following mechanisms: (1) changes in the

hydrophobic-hydrophilic nature of chains, (2) inter- and intramolecular complexation by

hydrogen bonding, or (3) electrostatic repulsion. All these mechanisms are closely related

80

to the protonation phenomena of the ionizable moieties on the polymer backbone or the

side chains. The kinetics of the swelling process and the equilibrium extent of swelling

are affected considerably by several factors, such as ionic strength of the medium, buffer

composition, and the presence of salts [Hariharan et al., 1996]. Other factors such as the

crosslinking ratio, solvent quality, chemical structure of monomers, and reaction

conditions during the photopolymerization also influence the structure formation and

hydrogel swelling properties.

Photopolymerization is a widely used technique to synthesize polymers and

hydrogels due to its distinct advantages of rapid cure, low curing temperature, in-line

production, low energy consumption, and easy process control. A great deal of research

has been carried out to investigate the effect of light intensity on the reaction kinetics of

UV-curable materials with the use of PhotoDSC, in which the hydrogel matrix is loaded

into an aluminum pan and then exposed to UV irradiation [Cook, 1993; Ward et al.,

2001; Li et al., 2005]. In the experiment, evaperation of volatile solvent or reactants may

cause significant measurement errors. Recently, Li et al. [2005] reported a technique of

modifying the DSC sample pan to minimize the sample loss and improve the accuracy for

volatile systems.

PH-sensitive hydrogels has the unique swelling/deswelling behavior with changes

of pH values in the surrounding medium. The structure formation and hydrogel swelling

properties are affected considerably by several factors, such as the properties of the

monomer solution (composition, solvent quality, chemical structure), synthesized

conditions during the photopolymerization, and the conditions of medium (ionic strength,

composition, pH values). A number of studies have reported that varying curing

81

conditions may achieve different gel structures and swelling properties [Lowman et al.,

1997; Anseth et al., 1996; Peppas et al., 1991; Kwok et al., 2003; Elliott et al., 2001;

Elliott et al., 2002]. The UV light intensity is one of the most important factors that affect

the reaction kinetics of the resin systems and the properties of the formed gels. It was

reported that an increase of the intensity led to a higher maximum polymerization rate of

the acrylate resin systems. The maximum was achieved more rapidly after the start of the

reaction and the induction period slightly decreased [Lovell et al., 1999; Scherzer et al.,

1999]. The effect of light intensity on the hydrogel system becomes more complex due to

the solvent influence. There lacks a thorough understanding on the interactions of

reaction kinetics, rheological changes, gel formation, and hydrogel structures as a result

of the UV radiation with different light intensities. In this study, PMAA gels synthesized

in a water/ethanol mixture are investigated by using a series of analytical tools including

PhotoDSC, photo-rheometry, and dynamic light scattering goniometry. The effects of

light intensity on the reaction kinetics and structural properties are addressed.

4.2 Experimental

4.2.1 Materials and sample preparation

The monomer, MAA (Sigma-Aldrich) and the crosslinking agent, TEGDMA

(Sigma-Aldrich) were used to prepare pH-sensitive hydrogels. For all reactions, the

crosslinking agent was presented at the level of 1.0 mole% based on the total mole of

monomers. A photoinitiator, 2,2-dimethoxy-2-phenylacetophenone (Irgacure 651, Ciba

Specificity Chemicals), was used at 1.0 wt% of the monomer mixture. The free-radical

photopolymerization was carried out in a mixed solvent of distilled water and ethanol

82

with the 1/1 ratio. The ratio of monomer to solvent was kept at 50:50 (w/w). All reagents,

unless specified, were of anylytical grade and were used without further purification.

To prepare hydrogel films for the swelling test and structure analysis, 5.0 grams

of MAA were mixed with a proper amount of TEGDMA and initiator. An equal weight of

solvent mixture was then added. The solution was transferred to a glove box where it was

kept under a nitrogen atmosphere. Nitrogen was bubbled through the solution for 20

minutes. Then the mixture was pipetted between two glass slides separated by a Teflon

spacer. The thickness of the spacers was 0.3 mm. The setup was then placed under a UV

light for photopolymerization at 0.25~24 mw/cm2. The cured hydrogels were then rinsed

in double deionized water for 5 days to remove unreacted monomer, initiator and sol

fraction. Subsequently, the monomer-free films were cut into samples with 5.0 mm

diameter for swelling test.

4.2.2 PhotoDSC measurement

The reaction kinetics and heat of reaction of PMAA gels were measured using a

PhotoDSC (TA 2920, TA Instruments). A UV light source (Novacure, 100W Hg short-arc

lamp, EXFO, Mississaugua, Ont., Canada) was used to cure the samples. In order to

prevent the weight loss of volatile MAA and ethanol, the DSC pans were physically and

chemically modified by using the technique described elsewhere [Li et al., 2005]. A

micropipette was used for PhotoDSC sampling (5~7 µl), which controlled the sample

weight for each test. All measurements were carried out at 30oC and the light intensity

was varied from 0.25 to 24 mw/cm2. Each run was conducted by purging the sample with

nitrogen gas until reaching equilibrium (around 2 minutes), and then UV irradiation was

83

applied to induce the free-radical polymerization.

To obtain the kinetic parameters, a series of unsteady state polymerizations was

performed. At a given time, the light source was extinguished and the “dark”

polymerization was continuously monitored by the DSC. Along with an expression for

the steady state polymerization, an expression for the unsteady state polymerization was

used to determine the kinetic parameters as a function of conversion. The details of this

experimental technique are available in the literature [Lovell et al., 1999].

4.2.3 Rheological measurement

A photo stress rheometer MCR 300 (Physica, Anton Paar) was used to follow the

viscosity change during the isothermal photopolymerization. A UV cell, including a top

steel plate with a diameter of 50 mm and a bottom plate made of quartz glass, was

utilized in this test. The UV light source (Acticure 4000, EXFO, Canada) was illuminated

from the bottom. The light intensity on the sample surface was kept at 2.0 mw/cm2. The

gap between the two plates was set at 1.0 mm and the shear rate used was 0.1s-1. The gel

point was assumed when the relative viscosity, i.e. viscosity of the reactive resin vs. its

initial viscosity, reached 104.

4.2.4 Dynamic light scattering analysis

Dynamic light scattering (DLS) measurements at 30○C were carried out to

determine the molecule size and size distribution before gelation during

photopolymerization by using a BI-DNDC Differential Refractometer (Brookhaven

Instruments) with a 10 mW He-Ne laser beam at a wavelength of 633 nm. A scattering

84

angle was held constant at 90°in the measurement. Because the formed polymer swells

more in water than in ethanol, the ethanol (3ml) was used as a solvent to dilute the

partially reacted sample (around 0.3 ml). The diluted solution was then filtered through a

filtration unit with 0.45 micron pore size (Whatman Puradisc 25TF) before measurement.

Count rates between 10 to 200 kilocounts per second were used to obtain meaningful

results by changing the sample concentration and adjusting the laser power.

Autocorrelation of the intensity was carried out by the method of cumulate analysis to

obtain an average diameter of the molecules and the polydispersity. The molecule size

distribution was obtained from the correction function by CONTIN analysis using the

standard software BI-DNDCW.

4.2.5 Swelling study

The swelling tests were performed in a pH=4.2 (or 7.3) buffer solution to

characterize the swelling behavior of synthesized pH-sensitive hydrogels. The buffer

solutions with specfic pH values were prepared by mixing the citric acid with appropriate

amounts of sodium phosphate solution. Sodium chloride was used to adjust the ionic

strength of all solutions to I=0.1M, which is the near-physiological condition. The dried

hydrogel samples were weighed and placed in the buffer solution at room temperature

(25°C). The samples were taken out of the solution at pre-selected time intervals. After

the extra water on the surface was removed by laboratory tissue, the weight of the wet

hydrogels was measured. The weight-swelling ratio was calculated by the weight of the

swollen sample to the weight of the dried sample. The samples were blotted and weighed

until the weight change was less than 0.1 mg over a 24-hour period.

85

4.3 Results

4.3.1 Kinetics of MAA/TEGDMA photopolymerization

Using the modified DSC sample pan, the effects of monomer content and UV

irradiation intensity on the reaction kinetics of the MAA/TEGDMA resin system were

investigated. Figure 4.1(A) shows the polymerization rate versus conversion for

MAA/TEGDMA (100/1 mol.%) with 50 or 100 wt.% monomer cured at a light intensity

of 5.0 mw/cm2. As expected, decreasing the monomer content diluted the reactant

concentration, hence slowed down the polymerization rate. The addition of solvent in the

monomer solution significantly changed the reaction profiles. For the bulk resin system, a

large exothermic peak was observed, while the resin system with 50% solvent had

multiple exothermic peaks on the reaction profile. The first peak (or shoulder) occurred at

the very early stage of polymerization. Regardless of solvent addition, the reaction rate vs.

conversion profile followed nearly the same path in the beginning. In other words,

changing the solvent content had little influence on the early reaction. It was also noted

that the addition of solvent allowed the polymerization to achieve a higher final

conversion as compared to the bulk condition (conversion of 99% vs. 61%). This is

because the resin system with much higher monomer content reacted faster, leading to

more buried monomer and consequently lower double bond conversion.

To study the effect of light intensity on the reaction kinetics, isothermal reactions

were carried out at 30°C for MAA/TEGDMA (100/1 mol.%) with 50 wt.% solvent

mixtures. The light intensities varied from 0.25 to 24 mw/cm2. Results are shown in

Figures 4.2(A) and (B). As the light intensity was raised, the initiation rate and the

86

Figure 4.1 Reaction rate vs. conversion of MAA/TEGDMA in the presence of 1% Irgacure 651 with 50 and 100 wt.% monomer content cured under 5.0 mw/cm2.

0

0.002

0.004

0.006

0.008

0 0.2 0.4 0.6 0.8 1

Conversion

Rea

ctio

n R

ate(

1/s)

100%50%

87

polymerization rate increased. The light intensity significantly influenced the reaction

rate profiles (i.e. the size and shape of the exothermic peaks). Under a low light intensity,

the first peak was small. It gradually became larger and took place at an earlier time with

an increased light intensity. However, the second peak tended to become smaller at a

higher light intensity. When the sample was cured at a light intensity larger than 5.0

mw/cm2, the first peak dominated and the second one became a shoulder. A further

increase in the light intensity caused the size of the second peak to become even smaller.

From the conversion versus time curves presented in Figure 4.2(B), one can see that an

increase in the light intensity generally reduced the time required to achieve a high

conversion. For example, to reach a conversion of 40%, the time required was shortened

from 10.8 to 3.4 minutes when the light intensity increased from 0.25 to 5.0 mw/cm2.

However, if the sample was cured at a light intensity larger than 5.0 mw/cm2, a higher

reaction rate was observed at the early stage, but the reaction rate became lower later than

that at a low light intensity at a later time. Consequently, the time to reach 40%

conversion at a light intensity of 24 mw/cm2 was as long as 4.3 minutes. This indicates

that too high a light intensity has an adverse effect on the photopolymerization of the

resin system.

The multiple peaks observed in the free radical polymerization can be explained

by microgel formation, which may affect the onset of macrogelation and the curing

behavior. Horie et al. [1975] has postulated this hypothesis to explain the occurrence of

double maxima in the reaction rate of MMA/EGDM systems: the first peak attributes to

the Trommsdorff effect in the bulk material and the second one to the Trommsdorff effect

in the microgels.

88

Figure 4.2 Effect of light intensity on the polymerization of MAA/TEGDMA system in the presence of 1% Irgacure 651 (A) reaction rate, (B) conversion.

0

0.001

0.002

0.003

0.004

0 10 20 30

Time(min)

Rea

ctio

n R

ate(

1/s)

24mw/cm25.0mw/cm22.0mw/cm20.25mw/cm2

(A)

a b

c

a’

c’ b’

24 mw/cm2

5.0 mw/cm2

2.0 mw/cm2 0.25 mw/cm2

0

0.2

0.4

0.6

0.8

1

0 10 20 30

Time(min)

Co

nve

rsio

n

24mw/cm25.0mw/cm22.0mw/cm20.25mw/cm2

(B)

24 mw/cm2

5.0 mw/cm2

2.0 mw/cm2 0.25 mw/cm2

89

4.3.2 Viscosity measurement

In order to evaluate the effect of light intensity on the polymeric structure

formation, a rheometer equipped with a UV cell was used to follow the viscosity change

during the reaction. Figures 4.3(A) and (B) display both the relative viscosity and

reaction rate as a function of double bond conversion for MAA/TEGDMA (100/1 mol.%)

cured at 0.25, 2.0, and 24 mw/cm2. Approaching the gel point, there was a steep increase

of the relative viscosity. At a low light intensity of 0.25 mw/cm2, macrogelation occurred

before the maximum of the second peak. As the intensity increased to 2.0 mw/cm2, the

gelation point reached the maximum of the second peak. While for a high intensity of 24

mw/cm2, macrogelation occurred near the end of the first peak. Combining these two

figures shows that the on-set of macrogelation shifted to a higher conversion when the

light intensity increased from 0.25 to 2.0 mw/cm2. Approaching an optimal intensity, the

gel conversion reached the maximum. However, if the light intensity was larger than 2.0

mw/cm2, a decreased gel conversion was observed. Figure 4.4 presents the gel conversion

versus light intensity. The gel conversion was only 71% when cured at 0.25 mw/cm2, but

rose to around 80% at 2.0 mw/cm2, after which the gel conversion significantly decreased.

According to this figure, an optimal intensity (2.0mw/cm2) can be used for curing the

PMAA hydrogels as drug delivery carriers to minimize the negative effect of residue

monomers.

90

Figure 4.3 Reaction rate and relative viscosity rise as a function of conversion of MAA/TEGDMA (1.0 mole% TEGDMA, 50 wt.% solvent) cured at different light

intensity (A) 0.25 and 2.0 mW/cm2, (B) 24 mW/cm2.

0

0.001

0.002

0.003

0.004

0 0.2 0.4 0.6 0.8 1

Conversion

Rea

ctio

n R

ate(

1/s)

0

2000

4000

6000

8000

10000

Rel

ativ

e V

isco

sity

(A)

0.25 mw/cm2

2.0 mw/cm2

0

0.001

0.002

0.003

0.004

0.005

0 0.2 0.4 0.6 0.8 1

Conversion

Rea

ctio

n R

ate(

1/s)

0

3000

6000

9000

12000

Rel

ativ

e V

isco

sity

(B)

24 mw/cm2

91

Figure 4.4 Gel conversion versus light intensity for polymerization of MAA/TEGDMA system (1.0 mole% TEGDMA, 50 wt.% solvent) in the presence of 1% Irgacure 651.

20

40

60

80

100

0 6 12 18 24 30

Intensity (mw/cm2)

Gel

Co

nve

rsio

n (

%)

Optimal

Intensity (mw/cm2)

92

4.3.3 Kinetic parameters

Polymerization is often described as a chain reaction with a set of rate constants

of elementary reactions among which the most important ones are the rate constants of

propagation ( pk ) and termination ( tk ). Photoinitiation is a useful process for determining

the kinetic rate constants in free radical polymerization. By monitoring the rate of

polymerization during UV-exposure and afterwards in the dark, one can evaluate the rate

constants pk and tk . The ratio 5.0tp kk is calculated from rate measurements under

steady-state irradiation conditions using the following equation [Decker, 1998]:

Here, the rate of propagation ( pR ) is directly related to the incident light intensity ( 0I ),

sample thickness (l), the absorptivity (ε ), concentration of the photoinitiator ([PI]), and

the quantum yield of initiation (φ ) (number of initiating species produced per photon

absorbed).

During the dark polymerization, no more radicals are produced and the rate

equation becomes [Tryson et al., 1979]:

where i and t refer to the monomer concentration and the rate of polymerization at the

onset of the dark reaction and after a given time, respectively. The linear time

dependence allows one to evaluate the ratio ptb kk . Together with the 5.0tpp kk ratio, the

individual values of pk and tk can be determined.

[ ] [ ] [ ][ ]( ) 2/1

5.0 1 lPIo

tb

pp eIM

k

k

dt

MdR εφ −−=−=

i

i

p

tb

t

t

dtMd

Mt

k

k

dtMd

M

)/][(

][

)/][(

][

−+=

−

(1)

(2)

93

Figures 4.5 and 4.6 show the variation of pk and tk with the degree of

conversion for the MAA/TEGDMA resin system cured at 2.0 mw/cm2 and 24 mw/cm2,

respectively. Under a low light intensity, the propagation and termination processes were

reaction controlled at the very beginning of the polymerization in the solvent mixture, so

pk and tk remained relatively constant in Figure 4.5. Above a conversion of 10%, tk

started to decrease gradually. When the reaction reached a conversion of 46%, the

termination rate curve leveled off. In the corresponding process, the pk value kept

increasing. This phenomenon can be explained by the theory of complex [Henrici-Olive

et al., 1962 &1965]. The essence of the theory is the assumption that the propagating

macroradicals continually interact with the surrounding medium. In the solution

polymerization, the propagating macroradical is surrounded by monomers as well as the

solvent molecules. Since the propagation can only take place if the propagating

macroradical is in the vicinity of the monomer molecules, the local concentration of

monomer molecules influences the rate of solution polymerization and the rate constant

for propagation. This hypothesis has been used to explain the variation of the rate

constant for propagation in systems containing monomers, such as acrylamide,

methacrykaminde, acrylic acid, methyacrylic acid, and their derivatives. To explain the

variation of pk and tk at low light intensity, we also need consider the microgel

formation. Above the conversion of 10% in this case, the microgel entanglements started

to form, although the viscosity of the bulk system showed little change. Inside the

microgels, the motion of the macroradicals was restricted due to increased diffusional

limitations, leading to a decrease in the overall value of the termination kinetic constant.

94

This entanglement formation made it possible for propagating macroradicals in the bulk

system to be surrounded by more monomer molecules. Consequently, the propagation

constant gradually increased and the first autoacceleration of the polymerization rate

occurred in Figures 4.2(A) and 4.3(A). A further increase in the conversion close to

80% induced a dramatic increase of bulk viscosity. At this point, the propagation rate

dropped rapidly since pk also became controlled by diffusion due to the increasing

mobility restriction in bulk materials.

In contrast, a high light intensity provided more energy for initiator to activate,

leading to more free radicals. Because there were so many reactive molecules in the

system and the polymerization reacted so fast, both the propagation and termination

processes were controlled by the diffusion even at the very beginning of the

polymerization. The values of pk and tk dramatically reduced, although the bulk

viscosity maintained a relative constant. After the macrogelation (above a conversion of

42%), the values of pk and tk varied with the increasing monomer conversion.

95

Figure 4.5 Conversion dependence of the rate constants pk and tk for the

polymerization of MAA/TEGDMA system at 2.0 mw/cm2.

1

10

100

1000

0 0.2 0.4 0.6 0.8 1

Conversion

Kp

(Kt)

( l/

mol

's)

kt, 2.0mw/cm2kp, 2.0mw/cm2

2.0 mw/cm2 2.0 mw/cm2

96

Figure 4.6 Conversion dependence of the rate constants pk and tk for the

polymerization of MAA/TEGDMA system at 24 mw/cm2.

0.01

0.1

1

10

100

0 0.2 0.4 0.6 0.8 1

Conversion

Kp

(Kt)

( l/

mol

's)

kt, 24 mw/cm2kp, 24 mw/cm2

24 mw/cm2 24 mw/cm2

97

4.3.4 Molecular size analysis

Figure 4.7(A) summarize the molecular size and its distribution of polymers

formed during the photopolymerization of MAA/TEGDMA cured at 2.0 mw/cm2. Under

this condition, the gel conversion was around 80%. The macromolecules formed at a

conversion of 23% (point ‘a’, the first maximum of reaction rate in Figure 4.2A)

exhibited a narrow unimodal distribution, ranging from 6 to 45 nm. The intensity reached

the maximum value at 18 nm polymer diameter. With the reaction progressed to a

conversion of 39% (point ‘b’, onset of the second autoacceleration in Figure 4.2A), the

peak was shifted to 62 nm. In addition, a bimodal size distribution occurred, which

contained a relatively narrow peak (11~22 nm) and a larger size distribution (44~87nm).

A further increase in the conversion to 78% (point ‘c’, before macrogelation) induced a

broad size distribution from 116 to 303 nm, while the intensity ratio of smaller molecules

decreased significantly. This suggests that most small molecules have converted into

larger clusters. The growth of hydrogel particles under UV radiation of 24 mw/cm2 was

investigated and is shown in Figure 4.7(B). The size distribution curves exhibit similar

shape under this condition. Increasing the light intensity shifted the polymer size

distribution to a smaller size. For example, the formed particles showed a unimodal size

distribution at the conversion of 9% (point a’), and a bimodal size distribution at the

conversion of 40% (point b’), except that the molecule clusters were small. At a

conversion of 42% (point c’), which was close to the gel conversion, the peak for larger

molecules was at 123 nm and the width of the distribution was from 54 to 212 nm. The

resin system cured at a lower light intensity formed larger polymer clusters when the

reaction approached macrogelation.

98

Figure 4.7 The molecular size distribution of the MAA/TEGDMA system (1.0 mole% TEGDMA, 50 wt.% solvent) cured at (A) 2.0 mw/cm2 and (B) 24 mw/cm2.

(A) 2.0mw/cm2

0

30

60

90

120

0 70 140 210 280 350

Diameter(nm)

Inte

nsi

ty

2.50min, 22%4.23min, 39%7.80min, 78%

(a) (b) (c)

(B) 24mw/cm2

0

30

60

90

120

0 70 140 210 280 350

Diameter(nm)

Inte

nsi

ty

0.51min, 9%3.93min, 40%5.00min, 42%

(a’) (b’) (c’)

99

Batzilla and Funke [1987] used poly(4-vinyl styrene) monomer to synthesize

highly crosslinked microgels under different conditions. The viscosity of the reactive

system decreased and then increased during polymerization. The initial viscosity decrease

was due to the intramolecular cyclization in the beginning of the reaction. As the reaction

proceeded, the viscosity increased due to intermolecular crosslinking. Although our

overall reaction kinetics followed a similar trend, the viscosity, the kinetic parameters,

and molecule size analysis, showed a different mechanism.

4.3.5 Discussion

For the chain crosslinking polymerization, the existence of multifunctional

monomers leads to the formation of pendant double bonds on the growing macro-radicals.

The pendant double bonds can react with propagation radicals through intramolecular

reactions to form cycles, and may also react through intermolecular reactions to form

network structures. Therefore, the network formation may coexist with the microgel

formation during polymerization. Based to the reaction kinetics, the changes of viscosity,

and the corresponding particle formation discussed in the previous sections, the curing

process of MAA/TEGDMA system can be described in five stages: initiation, microgel

formation, cluster formation, macrogelation, and post-gelation. The schematic diagram of

the structure formation in the MAA/TEGDMA photopolymerization at different light

intensities is described in Figures 4.8 and 4.9 for the first four stages.

In the first stage, all reactants are mixed together and UV radiation initiates

initiator decomposition to form free radicals (shown as filled dots). In the

MAA/TEGDMA system with 50 wt.% solvent mixture, a homogeneous solution is

100

Figure 4.8 Changes of reaction rate, viscosity during the photopolymerization of MAA/TEGDMA at light intensity of 2.0 mw/cm2: I initiation; II microgel formation; III

cluster formation; IV macrogelation; V post-gelation.

0% 20% 40% 60% 80% Conversion

Rea

ctio

n R

ate(

1/s)

Rel

ativ

e V

isco

sity

ab

c

VΙ

MAA

Free radical

ΙΙ ΙΙΙ ΙV

0% 20% 40% 60% 80% Conversion

Rea

ctio

n R

ate(

1/s)

Rel

ativ

e V

isco

sity

ab

ca

b

c

VΙ

MAA

Free radical

MAA

Free radical

ΙΙ ΙΙΙ ΙV

Intermolecular crosslinks

101

Figure 4.9 Changes of reaction rate, viscosity during the photopolymerization of MAA/TEGDMA at light intensity of 24 mw/cm2: I initiation; II microgel formation; III

cluster formation; IV macro-gelation; V post-gelation.

0% 20% 40% 60% 80% Conversion

Rea

ctio

n R

ate(

1/s)

Rel

ativ

e V

isco

sity

a’

b’

c’

ΙΙ ΙΙΙ ΙV VΙ

MAA

Free radical

0% 20% 40% 60% 80% Conversion

Rea

ctio

n R

ate(

1/s)

Rel

ativ

e V

isco

sity

a’

b’

c’

ΙΙ ΙΙΙ ΙV VΙ

MAA

Free radical

MAA

Free radical

Intermolecular crosslinks

102

formed with uniform distribution of all reactants since ethanol is a good solvent for both

hydrophilic MAA and hydrophobic TEGDMA (Irgacure 651). This is verified by the DLS

measurement of MAA/TEGDMA mixtures without UV radiation. According to the

measurement, no “particles” were observed in the DLS analysis. The initiation step of the

radical polymerization may be divided into the radical formation and the addition of a

monomer to the radical. Since the rate constant for the addition of a monomer to the

radical is usually several orders of magnitude higher than the value for the radical

formation (primary radicals), the decisive step of the initiation process is the formation of

primary radicals. A high light intensity provides more energy for initiator to activate,

leading to more formed primary radicals in solution. Therefore, more filled dots are

distributed in the proposed diagram in the first stage (Figure 4.9).

After the formation of monomeric radicals, the monomeric radicals may link with

multifunctional monomers to form the growing macroradicals with pendant double bonds,

leading to the cyclization through intramolecular reactions. This internal crosslinking on

the primary polymer chains leads to the formation of “microgels” [Dusek et al., 1980].

Simultaneously, the pendent double bonds may react through intermolecular reaction to

form a network structure. The relative rates of the intra- and intermolecular reactions are

strongly affected by the monomer composition, solvent concentration and quality, and the

curing conditions, such as the temperature and the intensity of incident light. Here, we

focused on the influence of light intensity.

A high light intensity leads to a faster initiation, more radicals and more pendant

vinyls in the system. The concentration of active radicals is relatively high, leading to a

faster polymerization rate and a higher possibility for the polymeric radical to cycle by

103

reacting with its own pendant double bonds. Consequently, cyclization may dominate

from the beginning of the reaction. The greater extent of intramolecular cyclization

means less intermolecular crosslinking, resulting in larger mesh and smaller size of

formed particles (Figures 4.7B and 4.9), and the weaker mechanical properties. The

propagation rate decreased with the reaction progress due to the comsumption of bulk

monomers. However, the Trommsdorff effect inside the microgels may occur because

termination is largely hindered due to immobilized macroradicals, Therefore, a large peak

was shown in the early stage of the reaction profile (Figures 4.3B and 4.9).

On the other hand, at a low light intensity, less radicals are fromed and the

reaction rate is low at the beginning. Due to the microgel formation, the Trommsdorff

effect may occur becuase termination is diffusion controlled, while the propagating

process is still in the reaction-controlled stage in the bulk system. Thus, a small shoulder

was observed in the early stage of polymerization (Figures 4.2A and 4.8). In addition, the

active radicals prefer to intermolecularly react with the double bonds. Therefore, the

formed molecules are generally larger in size with a more compact structure.

During the cluster formation stage (III), the reactive microgels with pendant

double bonds may react with free monomers and other microgels to form larger clusters,

resulting in a bimodal molecular size distribution. At the later part of this stage, the

presence of a larger number of clusters and the inter-connection of some clusters lead to

an increased viscosity.

Approaching the gel point in stage IV, most small microgels have converted to the

larger clusters and intermolecular reactions among these clusters finally lead to

macrogelation. For the transition from microgels to macrogels, intermolecular

104

crosslinking reactions require the displacement of neighboring solvent molecules from

the vicinity of the microgels. In the system cured at a higher light intensity, the dominant

intramolecular reaction can form many microgel particles. These microgels can easily

form large aggregates and quickly reach the gel point. In contrast, the distributed

microgels in a system with a lower light intensity have less chance to connect with each

other, taking a longer time to reach the gel point. As the system enters the post-gelation

stage (V), the reaction rate abruptly decreases since both propagation and termination

become diffusion limited.

Obviously, the high light intensity facilitates the cyclization, thus playing a

significant role in the overall structure of formed gels. One of the most important

physical properties characterizing the hydrogels structure is the weight swelling ratio.

Figure 4.10 illustrates this property of PMAA hydrogels cured under different light

intensities and immersed in different pH buffer solutions. When the light intensity

increased from 2.0 to 24 mw/cm2, the swelling ratio of cured hydrogels only rose from

5.3 to about 5.7 after immersing in a pH=4.2 buffer for 4 hours. In a higher pH buffer

(pH=7.3), the difference of the swelling ratio became very significant and increased from

21.4 to 32.8. The structure difference of formed PMAA gels is more easily characterized

in higher pH buffer solutions due to the electrostatic repulsion between the ionized forms

of the carboxylic segments, as well as the dissociation of hydrogen bonds between the

carboxylic acid groups of MAA and the oxygen of the ether groups of TEGDMA. These

swelling results are consistent with the particle size and integrated analysis discussed in

the previous section.

105

Figure 4.10 Dynamic swelling behavior of the PMAA hydrogels with 1.0% TEGDMA cured at different light intensity and immersed in the different pH buffer solutions.

0

7

14

21

28

35

0 40 80 120 160 200 240

Time(min)

Wei

gh

t S

wel

ling

Rat

io(g

/g)

24mw/cm2, pH 7.3

24mw/cm2, pH 4.2

2.0mw/cm2, pH 7.3

2.0mw/cm2, pH 4.2

106

4.4 Conclusions

This work studied the effect of the light intensity in the photopolymerization of

hydrogels. The copolymerization of photocurable MAA/TEGDMA system was enhanced

as the light intensity increased, especially at the low light intensity range and low

conversion. At too high a light intensity, an adverse effect was observed and the final

conversion of MAA decreased to 43% at 24 mw/cm2. The optimal light intensity was

about 2.0 mw/cm2 to get the PMAA hydrogels with low residue monomers. The use of

the high light intensity significantly shortened the reaction time to reach macrogelation

and increased the swelling ratio of formed hydrogels, which can be explained by the

mechanism for the relative rates of intra- and intermolecular reactions. With a high light

intensity, more free radicals and more intramolecular reactions led to a higher reaction

rate and faster gel formation. Since the intramolecular reaction contributed to less

crosslinked microgels, the resulting hydrogels had a higher swelling ratio.

107

CHAPTER 5

DESIGN OF SMART DEVICES BASED ON THE FUNCTIONAL HYDROGELS

SYNOPSIS

This chapter focused on the design of an assembled drug delivery system (DDS)

to provide multifunctions, such as drug protection, self-regulated oscillatory release, and

targeted uni-directional delivery by a bilayered self-folding gate and simple surface

mucoadhesion. In this device, a pH-sensitive hydrogel together with a poly(hydroxyethyl

methacrylate) (HEMA) barrier was used as a gate to control drug release. In addition,

PHEMA coated with poly(ethylene oxide) / poly(propylene oxide) / poly(ethylene oxide)

(PEO-PPO-PEO) surfactant was utilized to enhance mucoadhesion on the device surface.

The release profiles of two model drugs, acid orange 8 (AO8) and bovine serum albumin

(BSA) were studied in this assembled system, which compared with the conventional

drug-entrapped carriers and enteric-coating systems. Furthermore, targeted unidirectional

release was demonstrated in a side-by-side diffusion cell. In conclusion, for such an

assembled device, the PHEMA layer not only affects the folding direction but also serves

108

as a barrier to protect the model drugs. The release time can be controlled by the

thickness of the bilayered gate and the drug reservoir. Due to the reversible swelling

behavior of PMAA gels, the bilayered gate can sense the environmental pH change and

achieve an oscillatory release pattern. Moreover, the local targeting and uni-directional

release have been successfully demonstrated in vitro.

5.1 Introduction

It would be most desirable for drug release to match a patient’s physiological

needs at the proper time and/or proper site. This is why there is a great interest in the

development of controlled delivery systems [Qiu et al., 2001]. Drug delivery technology

can be brought to the next level by the fabrication of smart materials into a single

assembled device that is responsive to the individual patient’s therapeutic requirements

and able to deliver a certain amount of drug in response to a biological state. Such smart

therapeutics should possess one or more properties such as proper drug protection, local

targeting, precisely controlled release, self-regulated therapeutic action, permeation

enhancing, enzyme inhibiting, imaging, and reporting. This is clearly a highly

challenging task and it is difficult to add all of these functionalities in a single device. The

objective of this study is to develop an intelligent system for drug protection,

self-regulated oscillatory release, and targeted uni-directional release based on hydrogels.

Such a system would need to exhibit [Park et al., 1993], serving as drug delivery

carriers for oral, buccal, rectal, vaginal, ocular, epidermal and subcutaneous applications

[Petelinet al., 1998; Kitano et al., 1998; Miyazaki et al., 1998; McNeill et al., 1984;

Cohen et al., 1997; Draye et al., 1997; Beyssac et al., 1996].

109

Proper protection is required during administration of bioactive molecules.

Enteric-coated systems have been used in commercial applications for releasing drugs

through oral administration [Brogmann et al., 2001]. The encapsulation of drugs within

lipid vesicles also has the potential advantage of protection and high drug-loading [Park

et al., 1997; Gregoiraidis 1995]. However, a major limitation is that these systems cannot

fully protect the drugs and release them at a targeted area with a precisely controllable

rate over a long period of time. The use of microspheres or nanoparticles to protect drugs

for site-specific delivery has been of interest [Lowman et al., 1999; Horak et al., 2001;

Morishita et al., 2002]. In order to avoid periodic insulin injection, Lowman et al [1999].

prepared p(MAA-g-EG) hydrogel microparticles containing insulin for in vivo oral

administration. The hydrogel protects the insulin in the acidic condition of the stomach.

However, protein instability resulting from exposure to an organic solvent during loading

is a major problem [Li et al., 2000; Sah et al., 1999]. The applications are also limited by

organic solvent residues, the complexity of the process, and the need to sterilize the

microspheres.

Besides proper protection, controlled release and self-regulation of drug delivery

are highly desirable in many applications. Self-regulated devices can be classified into

substrate-specific and environment-specific devices [Heller 1996]. Makino et al. [1990]

developed a sugar-insulin conjugate, which was complexed with the protein

Concanavalin A (Con A). Such a device could deliver insulin in response to a change in

blood glucose concentration. In order to adjust the release of insulin by a “molecule gate”

system, Hassan et al. [1999] synthesized glucose-oxidase containing gels to convert the

pH-sensitivity to glucose-sensitivity. These substrate-specific devices are still under

110

development. In environment-specific devices, devices can directly respond to changes in

pH, temperature, ion strength, electromagnetic radiation, ultrasound, and photo or

pressure stimulation [Neuberger 2002; Peppas 1991]. Using functional hydrogels as a

switch or gate for controlled drug delivery has been explored recently by several

researchers [Kaetsu et al., 1999; Cao et al., 2001]. However, these devices either had a

very long response time or could not completely stop drug diffusion in non-delivery

conditions.

The release of drugs at specific sites has received much attention lately. Based on

the surface receptors, various targeting molecules are utilized to achieve the local

targeting. For instance, a polymer-drug conjugate with an antibody can be recognized by

the cell surface antigen for cancer diagnostics and therapeutics [Jelinkova et al., 1999].

For peptides or proteins through the gastrointestinal (GI) tract, the DDS can bind

specifically to the mucosal layer or cell surface to increase the residence time and

improve the bioavailibity of drugs. Residence time is an important factor for drug

transport through the GI-tract barrier. Dorkoosh et al. [Dorkoosh et al., 2001] designed a

novel DDS for site-specific drug delivery of peptide drugs in the intestinal tract using

superporous hydrogels (SPH) and SPH composite polymers, which swell very rapidly by

absorption of gut fluids. Thus, the system attached to the intestinal wall and provided a

longer residence time for drug release. Shen et al. [2002] reported an intestinal patch

design for oral delivery. A longer residence time and uni-directional diffusion were

achieved for better drug diffusion through the intestinal barrier by using a mucoadhesive

layer of Carbopol/ pectin. Tao et al. [2003] combined microfabrication techniques with

the use of mucoadhesive plant lectins to design a microdevice with a long residence time.

111

The present work focuses on the design of an assembled DDS that can integrate

multiple functions in a single system. Specifically, the drug protection and self-regulated

oscillatory release were demonstrated by using a bilayered self-folding design of

hydrogel. PHEMA coated with a PEO-PPO-PEO surfactant was utilized to enhance

mucoadhesion on the device surface for targeted uni-directional release. The release

profiles of two model drugs, acid orange 8 (AO8) and bovine serum albumin (BSA) were

studied in this assembled system. The results were compared with the conventional

hydrogel entrapped with drugs and enteric-coating systems.

5.2 Experimental

5.2.1 Materials

The monomer, methyacrylic acid (MAA) (Aldrich), and a crosslinking agent,

tri(ethylene glycol) dimethacrylate (TEGDMA) (Aldrich), were used to prepare

pH-sensitive hydrogels [Zhang et al., 2000]. HEMA (Aldrich) was used to prepare neutral

hydrogels, while diethylene glycol dimethacrylate (DEGDMA, Aldrich, Milwaukee, WI)

was the crosslinking agent [Lu et al., 1999]. Both hydrogels contained 0.01~0.02 mol of

crosslinking agent/mol of monomer. A photoinitiator,

2,2-dimethoxy-2-phenylacetophenone (Irgacure 651, Aldrich), was used at around 1 wt.%

of the monomer mixture. The swelling tests were performed at pH=3.0 and 7.3 to

characterize the swelling behavior of hydrogels. The buffer solutions with different pH

values were prepared by mixing the citric acid solution and appropriate amounts of

sodium phosphate solution. Sodium chloride was used to adjust the ionic strength of all

solutions to I=0.1M. For the swelling test, the dried hydrogel samples were weighed and

112

placed in the buffer solution at room temperature. The hydrogels were taken out of the

solution at pre-selected time intervals. After the extra water on the surface was removed

by laboratory tissue, the weight of the wet hydrogels was measured. The weight swelling

ratio was calculated by the weight of the swollen sample to the weight of the dried

sample.

The agent used to enhance mucoadhesion was a surfactant, Pluronic F127 Prill

(BASF Corporation). The major component of this surfactant is a tri-block polymer

PEO-PPO-PEO. Mucin (type III) was obtained from Sigma-Aldrich. Enteric coating

materials were prepared from MAA monomer using a low level of crosslinking agent.

Two hydrophilic model drugs, acid orange 8 (AO8) and bovine serum albumin (BSA),

were purchased from Sigma-Aldrich. Their molecular weights and physical properties are

listed in Table 5.1.

Solute MW(Da) Stokes radius ( Α& ) Solubility in water

at 25 °C (mg/ml)

AO8 386.4 3.4a 1

BSA 65000 34.8 40

Table 5.1 Physical properties of model drugs.

Note: a The stokes radius of AO8 is approximately calculated based on the stokes radius

of three different model drugs[Zhang et al., 2000].

113

5.2.2 Device design and drug loading

For most conventional delivery systems, drugs are either entrapped in a polymeric

matrix or encapsulated by a protective coating. Besides these simple systems, more

complicated DDS can be developed to control the drug release. Decisions as to which

type of device is most appropriate for an intended application must consider the need for

response time, drug release pattern, cost, safety, and therapeutic uses.

A Entrapped devices

The hydrogel matrix with the entrapped drugs was prepared as follows. First, the

hydrogel matrix was prepared by free-radical photo-polymerization at room temperature.

5.0 grams of MAA, together with TEGDMA (crosslinking ratio 0.01) and 1.0 wt%

Irgacure 651, were mixed at the ambient temperature. The monomer mixture was diluted

with a solvent mixture of 50 wt% double deionized water and ethanol to make a 50 wt%

monomer solution. The monomer solution was then injected between two glass slides

separated by teflon spacers with 0.8 mm in thickness and exposed to a low intensity 365

nm UV light at a light intensity of 1.8 mw/cm2 for 20 minutes under nitrogen flow. The

cured hydrogels were then rinsed in double deionized water for 5 days to remove

unreacted monomer, initiator and sol fraction. Subsequently, the monomer-free disks

were cut into samples with a 5 mm diameter and 0.8 mm thickness. These hydrogel disks

were placed in a 10 ml buffer solution with pH of 7.3 and AO8 concentration of 0.3 wt%

for 24 hours to load the model drug, then dried to a constant weight in a vacuum oven at

37°C. In addition to AO8, bovine serum albumin was selected as a model protein drug

with a large molecular size. A dried and weighed hydrogel sample was placed in 10 ml of

2.0 wt% BSA solution and allowed to swell for 2 days at 2~4°C under gentle shaking.

114

The swollen hydrogel sample was wiped dry using laboratory tissue and weighed, then

dried to a constant weight at room temperature.

B Assembled devices

The assembled device consists of two parts: a drug reservoir with targeting

function and a bilayered hydrogel gate. The drug reservoir was made of PHEMA gels,

which were prepared by the same approach described in the previous section. The gate

(5.0 mm in diameter and 60 µm in thickness) was made of two partially cured layers

using different hydrogels, PHEMA and PMAA. The bilayered gate and the drug reservoir

loaded with drug were bonded together by photo-polymerization of the residual monomer

in the bilayered gate under UV light as shown in Figure 4.1. For BSA loading, a

photomask was used to cover the area loaded with drug to prevent protein denaturing by

UV light. In order to completely remove the residual monomers, the loaded area was

totally cured using a large dose of high intensity light before bonding with the reservoir,

while the circle area of bilayered gate was masked. After loading, the residual monomers

within the area of the bilayered gate and the reservoir were cured with a large dose to

ensure the complete conversion of hydrogels. By using photo-differential scanning

calorimetry, it was found that the conversion for HEMA monomer is higher than 98% at a

light intensity of 3.2 mw/cm2 for 10 minutes.

The concentration of the solvent during polymerization determines the

homogeneous or heterogeneous structure of the gel produced. In this study, the

pH-sensitive hydrogels, PMAA, were synthesized with 50% distilled water to achieve a

good balance between high mechanical strength and a high swelling response to pH

changes. PHEMA hydrogels were prepared with 40% distilled water to ensure the optical

115

transparency of homogeneous hydrogels [Lu et al., 1999].

For comparison, an enteric-coating was also used as the drug release gate in the

DDS. MAA was mixed with a very small amount of crosslinking agent TEGDMA at a

concentration of 0.3 mol%. Irgacure 651 was added around 1wt% of the monomer

mixture. The monomer mixture spread on a microscopic slide was exposed to UV light

for 20 minutes under nitrogen flow. The film was quickly washed by DI water several

times to remove the unreacted monomer and then dissolved in a pH=7.3 solution to form

a homogeneous solution. The solution was poured in a petri dish and dried in a vaccum

oven at 37 °C overnight to form an enteric-coating layer. This layer was bounded to the

DDS following the same procedure as that used for the bilayered gate.

5.2.3 In vitro drug release

AO8 release from the hydrogel systems was measured by monitoring its

absorbance at 495 nm using a UV-vis Spectrophotometer (Varian Cary UV-Visible

Spectrophotometer). Drug release tests were performed in a buffer solution with pH

values of 3.0 and 7.3. Hydrogel devices with 5 mm diameter were placed in 30 ml of

buffer solution at room temperature (25°C) and subjected to constant shaking. At

pre-selected time intervals, 2.5 ml buffer solution was taken out of the vials for the UV

test, then placed back into the vials. The concentration of AO8 in the buffer solution was

obtained from a calibration curve, and the amount of AO8 release at time t (Mt) was

calculated from accumulating the total AO8 release up to that time. The fractional drug

release, Mt/M0, could then be calculated. Here M0 is the amount of initially loaded AO8.

For the BSA release experiment, protein concentrations were measured by monitoring

116

their absorbance at 270 nm by the same UV-vis Spectrophotometer.

Due to the reversible swelling response of PMAA to pH changes in the aqueous

environment, the bilayered gate can offer self-regulated release. To demonstrate this

function, the assembled device was immersed in a buffer solution with pH=3.0 at 25 °C.

UV-vis Spectrophotometer monitored the absorbance changes of the buffer solution.

After 10 minutes, the device was transferred to a pH=7.3 buffer for 10 minutes. This

cycle was repeated three times.

5.2.4 Diffusion studies

A side-by-side diffusion cell made by CNC machining was used to measure the

permeability and the diffusion coefficient of AO8 and BSA through hydrogel layers. The

hydrogel layers were swollen in pH=7.3 buffer solutions until reaching an equilibrium

state, then cut into a disc shape 2.2 cm in diameter and placed between the two cells (the

effective diffusion area was 2.83 cm2). Subsequently, 8 ml of 0.3 mg/ml AO8 (or 5 mg/ml

BSA) solution was injected into the donor cell (Cell A), while 8 ml buffer solution

without any model drug was simultaneously injected into the receptor cell (Cell B). The

cells were subjected to constant shaking at room temperature (25°C). At predetermined

time intervals, 2.5 ml buffer solution was taken from Cell B for UV Spectrophotometry

test [Zhang et al., 2000].

5.2.5 Targeted unidirectional release

Besides drug protection, the carrier design for many gene-, vaccine-, and protein-

based drugs must offer local targeting. There are many bioadhesive agents for

117

site-specific targeting. As an example, a mucoadhesive agent was used to simulate the

targeting function for the assembled device. The same principle can be applied for other

bioadhesive agents. Conventional emulsification or enteric-coating techniques provided a

similar targeting function. However, drugs tend to release in all directions after targeting.

In contrast, this assembled device can provide targeted uni-directional release because

only the releasing surface of the device is modified. Enhanced mucoadhesion was

achieved by UV-curing of HEMA with 5wt% PEO-PPO-PEO surfactant as shown in

Figure 5.1.

The targeted unidirectional release of a food dye AO8 was measured by video in a

side-by-side diffusion cell. 25mg mucin was gently blended with 500mg distilled water to

form a homogeneous solution, which was then evenly spread over a 25mm diameter

millipore membrane and allowed to dry at room temperature to create a mucin-coated

membrane. Prior to testing, a digital camcorder was set to record the drug targeting and

release. Subsequently, the device was placed in the donor cell and the mucus-coated

membrane was then placed between the two cells. 8 ml of pH=7.3 buffer solution was

simultaneously injected into both cells and the set-up was subjected to constant shaking

at 25°C.

118

Figure 5.1 Schematic of the assembled device.

PEOPEO

P(MAA-g-EG)HEMA

Glass slide

MAAUV exposure

(partial cure)

UV exposure (partial cure)

Photomask

Bilayered gate

HEMA/5% PEO-PPO-PEO

PEOPEO

Poly(HEMA)

PEOPEO

Device assembling

Preparation of release gate Surface targeting

Loaded drug

PEOPEO

P(MAA-g-EG)HEMA

Glass slide

MAAUV exposure

(partial cure)

UV exposure (partial cure)

Photomask

Bilayered gate

HEMA/5% PEO-PPO-PEO

PEOPEO PEOPEO

Poly(HEMA)

PEOPEO PEOPEO

Device assembling

Preparation of release gate Surface targeting

Loaded drug

119

5.3 Results and Discussion

5.3.1. Swelling properties of hydrogels

The delivery device is based on the swelling properties of different hydrogels. The

dynamic swelling behaviors of the two hydrogels in different buffer solutions are shown

in Figure 5.2. As can be seen, the dried hydrogels swell at all pH conditions due to the

adsorption of water into the porous structure. However, compared with PHEMA

hydrogels, PMAA hydrogels have a much more sensitive response. In the high pH buffer

solution, the PMAA hydrogels swell rapidly and can achieve a much higher equilibrium

swelling ratio than PHEMA hydrogels. This is because ionization of the carboxyl groups

(the pendent group of MAA) occurs as the solution becomes less acidic, resulting in

dissociation of the hydrogen bonds between the carboxylic acid groups of MAA and the

oxygens of the ether groups of TEGDMA. The dissociation of hydrogen bonds,

combined with the electrostatic repulsion force, causes the hydrogel network to swell

quickly, thus more water is imbibed into the hydrogels and a higher swelling ratio is

obtained. On the other hand, PHEMA is a neutral hydrogel, which has no ionizable

groups on its side chain. With a change of pH values, this material exhibits very small

swelling in buffer solutions. In addition, since the solvent content in the HEMA monomer

solution (40 wt.%) was less than that in the MAA solution (50 wt.%), PHEMA hydrogels

should have a more compact structure than PMAA gels with the same crosslinking ratio.

Although DEGDMA has a shorter chain than TEGDMA, its contribution could be

neglected when considering the low amounts of crosslinker.

120

Figure 5.2 Dynamic swelling behavior of hydrogels. Samples were 5.0 mm in diameter and 0.8 mm in thickness. ( ) PMAA hydrogel in pH=7.3 buffer. ( ) PMAA

hydrogel in pH=3.0 buffer. ( ) PHEMA hydrogel in pH=7.3 buffer. ( ) PHEMA hydrogel in pH=3.0 buffer.

0

4

8

12

16

20

24

0 30 60 90 120 150 180Time (min)

Sw

ellin

g R

atio

(g

/g)

121

5.3.2 Model drug release from entrapped devices

There are two general loading methods for entrapped hydrogels as drug carriers.

In one method, the mixture of monomer, initiator, crosslinking agent, and model drug was

cured by free-radical photopolymerization to form a hydrogel matrix with uniform

entrapment of the model drug. However, two major drawbacks limit this method’s

application. One is the UV adsorption of model drug, which inhibits the hydrogel

polymerization, thus limiting the amount of loaded drug. The other drawback is drug

instability. The carried drugs, such as peptides and proteins, become unstable under the

UV light. Therefore, a different method is usually adopted, which overcomes the

disadvantages of the direct curing method. In this method, cured hydrogels are allowed to

swell to an equilibrium state in a drug solution, and then dried to obtain the drug-loaded

hydrogel matrix. However, the long drug loading time is the major drawback of this

method. In this experiments, the second approach was used to make the entrapped

samples.

In order to investigate the effect of gel structure on the drug release, such model

drugs as AO8 and BSA were entrapped into the gel matrix in pH 7.3 buffer solutions. For

small molecular AO8, 24-hour loading time was long enough to reach an equilibrium

state and homogeneous distribution. While for large molecule, even 48-hour loading time

is not long enough to get a homogeneous distribution. Based on the picture of confocal

Microscopy (not shown here), the closer to the surface the distance, the more entrapped

BSA. Figure 5.3 represents the AO8 release from 5mm entrapped samples with different

crosslinking agent. According to the data, it was concluded that the AO8 entrapped into

the lower crosslinking agent could fast release AO8, which corresponds to the swelling

122

behavior of PMAA. The experimental results in Figure 5.4 illustrate the effect of

crosslinking agent on the BSA release pattern. As expected, the gels with lower

crosslinking agent swell quickly and release the BSA with a fast release rate. Moreover,

all curves show a typical first-order release behavior: an initial high release rate followed

by a declining drug release rate.

To compare the release behavior of drugs with different sizes, AO8 and BSA

release from 5mm entrapped samples are presented in Figure 5.5. As can be seen, under

the acidic condition (pH=3.0), AO8 is released very slowly. The concentration gradient

drives AO8 release from the polymeric matrix to the buffer solution. At neutral condition,

the pH-sensitive hydrogel is capable of imbibing a large amount of water, enlarging the

mesh size and causing AO8 to be easily released. As shown in Fig. 5.5, about 40% AO8

can be released after 150 minutes at 25°C. In the experiment, the BSA loading

concentration was about 6 times higher than that of AO8 in order to be easily detected by

UV spectroscopy. This figure also shows the BSA release profile from the entrapped

hydrogels in pH=7.3 buffer solutions. The BSA release profile can be divided into two

stages: an initial fast release for 60 minutes, followed by a slow release. This release

profile may be explained as follows. In the first stage, due to the compact structure of

swelling hydrogels, the predominant transport is due to the movement of hydrogel chains.

Therefore, BSA and AO8 have a similar release profile in the first 60 minutes. After 60

minutes, the drug size becomes the major factor dominating the drug release rate. The

BSA release rate becomes about 43% that of AO8 because large molecules usually

diffuse slower than small molecules.

123

Figure 5.3 Acid Orange 8 release to pH 7.3 buffer solution from the entrapped 5.0 mm PMAA samples at 25 °C. The samples were 0.8 mm in thickness.

0

0.2

0.4

0.6

0.8

1

0 30 60 90 120Time(min)

AO

8 F

ract

ion

al R

elea

se

0.75%TEGDMA

1.00%TEGDMA

2.00%TEGDMA

124

Figure 5.4 BSA release to pH 7.3 buffer solution from the entrapped 5.0 mm PMAA samples at 25 °C. The samples were 0.8 mm in thickness.

0

0.2

0.4

0.6

0.8

1

0 30 60 90 120Time(min)

BS

A F

ract

ion

al R

elea

se

0.75%TEGDMA1.00%TEGDMA

2.00%TEGDMA

125

Figure 5.5 AO8 and BSA release from the entrapped 5.0 mm PMAA samples at 25 °C. The samples were 0.8 mm in thickness. ( ) AO8 at pH=3.0. ( ) AO8 at pH=7.3. ( )

BSA at pH=7.3.

0

0.2

0.4

0.6

0.8

1

0 30 60 90 120 150 180

Time(min)

Fra

ctio

nal

Rel

ease

126

5.3.3 Diffusion studies

To further investigate the transport behavior of model drugs with different sizes in

the hydrogel matrix, permeation experiments of AO8 or BSA across the hydrogel layers

were carried out as described. Permeability can be calculated by the following equation

[Schwarte et al., 1998]:

Here, Ct is the solute concentration in Cell B at time t, C0 is the initial solute

concentration of Cell A, V is cell volume, A is the effective area of permeation, and P is

the membrane permeability coefficient. By plotting − (V/2A)*ln[1 - 2(Ct/C0)] versus time

t, the slope is the permeability coefficient.

The diffusion coefficient can be obtained from the permeability P, the solute

partition coefficient Kd, and the membrane thickness L in the swollen state. Their

relationship is shown in the following equation:

To determine the diffusion coefficient, the solute partition coefficient, Kd, needs to be

calculated from the experimental data by the following equation:

Here, Cm is the concentration in the membrane at equilibrium, Cs is the concentration in

the surrounding solution at equilibrium, C0 is the initial concentration in the surrounding

solution, V0 is the initial solution volume, Vm is the solution volume in the membrane at

1)1( 00 +×−==mss

md V

V

C

C

C

CK ( 3 )

dm K

PLD = ( 2 )

PtV

A

C

Ct 2)

21ln(

0

−=− ( 1 )

127

equilibrium, and Kd is a measure of the solubility of the solute in the membrane. A low

value of Kd means that a solute molecule is not easily soluble in the membrane. A high

value of Kd indicates that there may be binding between the solute and the polymer, thus

the solute molecule can be easily soluble in the membrane phase.

The permeability is defined as a particular solute through a particular membrane.

The solute size, membrane mesh size, pH, temperature, and the affinity of the solute with

the membrane may affect the permeation of the solute. In this experiment, the

temperature and pH were maintained constant. Two model drugs with significantly

different sizes were used in the experiment. The hydrodynamic radius of BSA is about 10

times larger than that of AO8. Figure 5.6 shows the solute permeation of AO8 and BSA

through swollen PMAA and PHEMA membranes at 25°C in pH=7.3 buffer solution. As

can be seen, − (V/2A)*ln[1 - 2(Ct/C0)] increases linearly with time. The slope of each

linear curve represents the permeability for a particular solute. As expected, for PMAA

membrane, the permeability of AO8 is higher than that of BSA. And, for a solute like

AO8, a PMAA membrane with a larger mesh size in the swollen state has a much higher

permeability than a PHEMA membrane.

128

Figure 5.6 Permeation of AO8 and BSA through different swollen hydrogel membranes at pH 7.3 and 25 °C. ( ) AO8 through PMAA. ( ) BSA through PMAA. ( ) AO8

through PHEMA.

0

0.03

0.06

0.09

0.12

0.15

0 50 100 150 200 250 300

Time(min)

-V/2

A*l

n(1-

2Ct/

C0)

0

0.03

0.06

0.09

0.12

0.15

0 50 100 150 200 250 300

Time(min)

-V/2

A*l

n(1-

2Ct/

C0)

129

Based on the permeability and partition coefficient, the diffusion coefficient can

be calculated as listed in Table4.2. The diffusion coefficient of AO8 (MW= 386.4g/mol)

within the PMAA film matrix is 2.03× 10-6 cm2/s. Zhang et al. [2000] investigated the

release kinetics of oxprenolol HCl (MW= 302g/mol) from a swollen

poly(MAA-g-NIPAA) hydrogel (weight swelling ratio=18.2) at 25°C in pH=7.3 buffer

solution. The reported diffusion coefficient was 4.68× 10-6 cm2/s. Therefore, this

measured AO8 diffusion coefficient in the PMAA film matrix is reasonable. The BSA

diffusion coefficient in the PMAA film matrix at 25°C was 8.00× 10-7 cm2/s, which is

close to the BSA diffusion coefficient estimated by Mariah et al. [2001]. The diffusion

coefficient of BSA within the PMAA hydrogel matrix is about 40% that of AO8. This

agrees with the measured results that the BSA release rate is 43% that of AO8 through the

same hydrogel matrix.

Membrane Model drug

Permeability P

)/(105 scm×

Partition

coefficient Kd

Diffusion coefficient

)/(10 27 scm×

PMAA AO8 2.83 0.99 20.03

PMAA BSA 0.33 0.35 8.00

PHEMA AO8 0.17 0.99 0.67

Table 5.2 Permeability and diffusion coefficient of model drugs through different membranes.

130

The mesh size of the polymer matrix also affects the diffusion coefficient. As

shown in Table 4.2, the diffusion coefficient of AO8 in the PMAA film matrix is about 30

times larger than that of AO8 in the PHEMA film matrix. This is due to significantly

different polymeric structures at equilibrium. At pH=7.3, the PMAA hydrogel has a much

looser structure than the PHEMA hydrogel and the model drug can diffuse quickly and

easily in the matrix. The large BSA molecule is not easily soluble in the membrane, while

AO8 can be entrapped in the hydrogel matrix easily.

5.3.4 Model drug release from assembled devices

The entrapped device based on pH-sensitive hydrogels can control the drug

release rate under different conditions. However, it has the disadvantages of low

drug-loading efficiency and a long loading time. Small molecules can be entrapped in a

hydrogel matrix easily and quickly due to their stable structure, high solubility, and large

diffusion coefficient. However, macromolecular drugs such as proteins cannot be easily

entrapped. In addition, these molecules are very sensitive to the environment. Therefore,

an assembled device was designed to solve these problems.

5.3.4.1 Drug protection

In this design, a controlled release was achieved by self-folding of the bilayered

hydrogel. The PHEMA layer is a major factor to control the drug release. First, its

swelling property influences the folding direction with increasing pH values. In pH=3.0

medium, PMAA hydrogels have a similar swelling response as PHEMA, thus the

bilayered gate would not open for drug delivery. With increasing pH, the swelling ratio of

131

the PMAA hydrogel layer increases significantly, while the PHEMA layer has a relatively

constant swelling ratio independent of pH values. Therefore, the bilayered gate folds

outward until the bonding between the gate and the reservior breaks. As a result, the

model drug can be released quickly. Figure 5.7 describes the AO8 release from the

assembled device at pH=7.3. In this system (Fig. 5.7A), AO8 particles were loaded in the

reservoir. After the device was placed in the buffer solution, the bilayered gate started to

fold outward due to water imbibing into the device and a small amount of AO8 was

released from a small interstice (Fig. 5.7B). With increasing time, the interstice became

larger and larger. After 80 minutes, the swelling properties of hydrogels caused the gate

to fold like a roll (Fig. 5.7C). Figure 5.7D shows the schematic of AO8 release from the

side view.

Since the PHEMA layer has a much lower permeability to the model drug than the

PMAA layer, it also serves as a barrier to protect the proteins through the stomach. Figure

5.8 presents the AO8 release from the 5.0 mm assembled device with different gates at

pH=3.0 and 25°C. As can been seen, in pH=3.0 medium, AO8 fractional release was

nearly zero after 4 hours for the bilayered gate. Actually, after 24 hours, AO8 could not

be released in pH 3.0 buffers based on the experiment. However, for PMAA gate, the gate

did not have sufficient mechanical strength to resist the enlarged volume of drug solution,

such that the PMAA gate was broken in the center and AO8 started to release quickly

from the device at 160 minutes. This protection function can protect biomolecules from

gastric acids and enzymes in the stomach.

132

Figure 5.7 AO8 release from the assembled device at pH=7.3 and 25°C. The diameter of the device is 5.0 mm. The thickness of bilayered gate is 60 µm and the thickness of the

drug reservoir is 1.0 mm. (A) Dry assembled device. (B) Releasing at t= 40 minutes. (C) Released at t= 80 minutes. (D) Schematic of AO8 release from assembled device.

(A)

5.0 mm

(A)

5.0 mm5.0 mm

(C)

5.0 mm

(C)

5.0 mm5.0 mm

(B)

5.0 mm

(B)

5.0 mm5.0 mm

(D)

133

Figure 5.8 AO8 release from the 5.0 mm assembled devices with different gates at pH=3.0 and 25°C. The gate thickness is 60 µm and the reservoir thickness is 1.0mm. ( )

PMAA hydrogel gate. ( ) PHEMA and PMAA bilayered gate.

0

0.2

0.4

0.6

0.8

1

0 30 60 90 120 150 180 210 240

time(min)

Frac

tiona

l Rel

ease

Time(min)

134

Figure 5.9 shows AO8 and BSA release from the 5 mm assembled device at room

temperatures. In pH=3.0 medium, there was no drug release for 2 hours. In pH= 7.3

buffer solution, it took about 40 minutes to open the device at 25°C and then reached

90% drug release quickly. Compare with the AO8 release, BSA had a similar release

pattern, which confirms that the release mechanism of this device is based on hydrogel

folding and is independent of the model drug size.

The thickness of the bilayered gate and the thickness of the drug reservoir also

influence the release time as shown in Figure 5.10. When the thickness of the gate was

reduced from 90 µm to 60 µm, the open time would be reduced to about 20 minutes.

Decreasing the thickness of the drug reservoir would reduce the bonded area between the

bilayered gate and the drug reservoir. Thus, the model drugs would be released more

quickly. Because the controlling mechanism is based on the hydrogel swelling behavior,

not the height of the drug reservoir, the height can be varied to adjust the amount of drug

loading.

135

Figure 5.9 AO8 and BSA release from the 5.0 mm assembled device at 25°C. The thickness of the bilayered gate is 60 µm and the thickness of the drug reservoir is 1.0 mm.

( ) AO8 at pH=3.0. ( ) AO8 at pH=7.3. ( ) BSA at pH=7.3.

0

0.2

0.4

0.6

0.8

1

0 30 60 90 120

Time(min)

Fra

ctio

nal

Rel

ease

136

Figure 5.10 Thickness effects of the bilayered gate and reservoir on AO8 release behavior at pH=7.3 and 25 °C. ( ) The gate thickness is 60 µm and the reservoir

thickness is 0.5 mm. ( ) The gate thickness is 60 µm and the reservoir thickness is 1.0 mm. ( ) The gate thickness is 90 µm and the reservoir thickness is 0.5 mm.

0

0.2

0.4

0.6

0.8

1

0 30 60 90 120

time(min)

Fra

ctio

nal

Rel

ease

Time(min)

137

5.3.4.2 Self-regulated oscillatory release

Since PMAA gels possess unique swelling properties, the device can sense the

environmental pH change and provide oscillatory release behavior. To demonstrate the

oscillatory regulation, the device was tested with a varying pH field. Figure 5.11 presents

the oscillatory release behavior of this device. It is evident from the graph that a pulsatile

release rate was obtained when the pH was increased from 3.0 to 7.3 due to self-folding

of the bilayered gate. When the pH decreased from 7.3 to 3.0, the small value of the

release rate indicates that the bilayered gate has reversed to its flat shape and blocked the

drug release. In contrast, the assembled device with an enteric gate can only provide a

one-time irreversibly pulsatile release profile.

This gate design has the limitation that the response time is in minutes. By

controlling the chemical structure of hydrogels, gate thickness, and the bilayer ratio, the

response time can be reduced to seconds. By using various stimuli-sensitive hydrogels,

assembled devices can be activated by pH, temperature, pressure, ionic strength,

electromagnetic radiation, buffer composition or the concentration of glucose [Peppas

1991].

5.3.4.3 Targeted unidirectional release

Spatial localization of the therapeutic payload in the target regions is very

important for high bioavailability of the administrated drug for therapeutic uses. Different

targeted molecules can be attached to the surface of delivery devices by covalent or non-

138

Figure 5.11 The oscillatory release behavior of the assembled device. The gate thickness is 50 µm and the thickness ratio for PHEMA to PMAA layer is 4.

0

0.005

0.01

0.015

0.02

0 10 20 30 40 50 60Tim e (m in)

Rel

ease

Rat

e (m

g/m

in)

0

3

6

9

12

Bu

ffer

pH

0

0.005

0.01

0.015

0.02

0 10 20 30 40 50 60Tim e (m in)

Rel

ease

Rat

e (m

g/m

in)

0

3

6

9

12

Bu

ffer

pH

139

covalent binding to improve the device bioadhesion. Typical examples of bioadhesion

include mucoadhesive hydrogels for mucosal route of delivery, plant lectins for mucosal

route, and carbohydrate antibody for cell surface receptors. The mucoadhesion was

demonstrated in this study by modifying the targeted area with photo-curing of

HEMA/5% PEO-PPO-PEO surfactant. Pappes has proposed to the enhancement of

mucoadhesion by tethered chains of poly(ethylene glycol) (PEG) grafted on a polymer

backbone. Along the polymeric structure of this surfactant, each domain plays a specific

role in the resulting surface function: the hydrophobic PPO backbone prefers to

interpenetrate in the PHEMA hydrogels, while the hydrophilic tethered chains of PEO act

as adhesion promoters to enhance mucoadhesion due to tether diffusion. The addition of

bioadhesive polymer chains increases the entanglement between the polymer and mucus

network, resulting in strong interaction binding [Ascentiis et al., 1995]. This surface

modification prolongs the residence time at delivery sites and improves drug absorption.

Figure 5.12 compares the targeted unidirectional release with an untargeted

release in the side-by-side diffusion cell. As shown in Figure 5.12A, the device can attach

on the mucin-coated membrane due to the mucoadhesive modification on the device

surface. When the bilayered gate self-folded, the imbibed water in the reservoir pushed

the dissolved drugs from the donor cell to the receptor cell through the membrane. On the

other hand, the unmodified device could not attach to the membrane surface and the

released AO8 was in the donor cell as shown in Figure 5.12B.

140

Figure 5.12 The comparison of the targeted uni-directional release with untargeted release: (A) Targeted release. (B) Untargeted release.

(A)

(B)

141

5.4 Conclusions

There is considerable interest in the development of “smart therapeutics” DDS for

bioactive drugs. Such a desirable carrier needs to offer multiple functions in a single

device. An assembled DDS was demonstrated in this study to achieve multifunctions

such as drug protection, self-regulated oscillatory release, and targeted uni-directional

delivery by a bilayered hydrogel design and simple surface mucoadhesion. A PHEMA

layer not only affects the folding direction but also serves as a barrier to protect the model

drug. A cylindrical drug reservoir design provides easy loading of large amount of drugs.

The release time can be controlled by the thickness of the bilayered gate and the

thickness of the drug reservoir. Due to the reversible swelling behavior of PMAA gels,

the bilayered gate can sense the environmental pH change and achieve an oscillatory

release pattern. Surface modification with PEO chains can act as adhesion promoters to

enhance the device mucoadhesion. The local targeting and uni-directional release have

been successfully demonstrated in vitro. Based on the self-folding mechanism,

optimization of gate and device design, as well as the proper choice of hydrogel materials,

the DDS described in this study has the potential to provide the desired release pattern for

a broad range of therapeutic uses.

For biomolecular delivery used in inter- and intra-vascular applications, the

device need be reduced to micron-sized or smaller. Current work focuses on the design of

miniaturized DDS by using polymer micro-fabrication and integration techniques.

142

CHAPTER 6

AN ORAL DELIVERY DEVICE BASED ON SELF-FOLDING HYDROGELS

SYNOPSIS

A self-folding miniature device has been developed to provide enhanced

mucoadhesion, drug protection, and targeted unidirectional delivery. The main part of the

device is a finger like bilayered structure composed of two bonded layers. One is a

pH-sensitive hydrogel based on crosslinked poly(methyacrylic acid) (PMAA) that swells

significantly when in contact with body fluids, while the other is a non-swelling layer

based on poly(hydroxyethyl methacrylate) (PHEMA). A mucoadhesive drug layer is

attached on the bilayer. Thus, the self-folding device first attaches to the mucus and then

curls into the mucus due to the different swelling of the bilayered structure, leading to

enhanced mucoadhesion. The non-swelling PHEMA layer can also serve as a diffusion

barrier, minimizing any drug leakage in the intestine. The resulting unidirectional release

provides improved drug transport through the mucosal epithelium. The functionality of

this device is successfully demonstrated in vitro using a porcine small intestine.

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6.1 Introduction

Many protein- and DNA-based drugs exhibit high sensitivity to the surrounding

physiological conditions as a result of their delicate physicochemical characteristics and

the susceptibility to degradation by proteolytic enzymes in biological fluids. They need to

be properly protected during administration and their release needs to be precisely targeted

and controlled. Typically, the intramuscular or intravenous injection is used for the

administration of peptides and proteins. However, due to the undesirable nature of this

method, such as pain, inconvenience and inconsistent pharmacokinetics, other routes have

been considered. They include pulmonary, oral, nasal, buccal, rectal, ocular, vaginal, and

transdermal delivery [Kopecek et al., 1998], among which oral administration is the most

convenient and ideal route.

Although oral administration is a non-invasive route of drug delivery, peptides

and proteins delivery through the gastrointestinal (GI) tract remains a highly challenging

task because of their low bioavailability resulting from the pH fluctuation, proteolytic

degradation, low transport efficiency, and short residence time. Enteric-coated systems

have been commercially used for releasing drugs through oral administration [Brogmann

et al., 2001]. The encapsulation of drugs within lipid vesicles also has the potential

advantage of drug protection and high drug loading [Gregoiraidis, 1995]. The inclusion

of enhancers/promoters, protease inhibitors, and/or specific adhesion may help the

diffusion of large molecules across the epithelial membrane. However, a major limitation

is that these systems cannot fully protect the drugs and release them in a targeted area

with a precisely controllable rate over a long period of time.

Mucoadhesive drug delivery systems (MDDSs) have attracted considerable

144

interest because of their sustained drug release profile at the absorption site and increased

drug bioavailability due to the intimate contact with the absorbing tissue. MDDSs

typically present in the form of symmetric micro- and nano-spheres or asymmetric

patches. Mucoadhesion occurs through surface-to-surface contact. Micro-/nano-particles

prepared by phase separation, microemulsion and spray drying have been successfully

used as drug delivery carriers. [Jain, 2000; Langer, 2000; Li, 2000]. These particles

usually have polydisperse sizes and relatively simple structures. Additionally, the

symmetric shape leads to drug release to all directions. Recently, several research groups

have made efforts to design patch-like asymmetric delivery devices with functionalities

such as drug protection and targeted unidirectional release [Dorkoosh et al., 2002; Shen et

al., 2002; Whitehead et al., 2003; Tao et al., 2004; He et al., 2004]. However, the

surface-to-surface adhesion for all these systems leads to the limited residence time due

to the continuous shedding of surface mucus.

In this study, a novel particulate-like miniature device is developed based on the

integration of a number of micro-manufacturing modules such as soft-lithography,

micro-imprinting, and polymer self-folding. Approaches that are able to improve oral

bioavailability, such as protective coating, mucoadhesive binding and mechanical

grabbing are also applied in the device design.

6.2 Experimental

6.2.1 Materials

The pH-sensitive hydrogel was prepared from the monomer, methyacrylic acid

(MAA, Sigma-Aldrich), and a crosslinking agent, tri(ethylene glycol) dimethacrylate

145

(TEGDMA, Sigma-Aldrich). Hydroxyethyl methacrylate (HEMA, Sigma-Aldrich)

crosslinked with diethylene glycol dimethacrylate (DEGDMA, Sigma-Aldrich) was used

to prepare the non-swelling hydrogel. Both hydrogels contained 0.01mol of crosslinking

agent/mol of monomer. A photoinitiator, 2,2-dimethoxy-2-phenylacetophenone (Irgacure

651, Aldrich), was used at 1 wt% of the monomer mixture. The free-radical

photopolymerization of MAA/TEGDMA system was carried out in a water/ethanol

mixture ( 1vs.1 ratio). The ratio of monomer to solvent during synthesis was 50:50 (w/w).

The HEMA/DEGDMA system was polymerized in a water solution with a 40 wt.%

solvent ratio. Poly(dimethylsiloxane) (PDMS) resin was purchased from Dow-Corning. A

degradable poly(ε -caprolactone) (PCL) and a water-soluble poly(vinyl alcohol) (PVA)

were purchased from Sigma-Aldrich. Carbopol 934 was purchased from BF Goodrich

(Cleveland, OH). All reagents, unless specified, are of analytical grade and were used

without further purification. Two hydrophilic model drugs, acid orange 8 (AO8) and

bovine serum albumin (BSA) were also purchased from Sigma-Aldrich. Fresh porcine

small intestines were collected from The Ohio State University Lab Animal Resource.

6.2.2 Device design and fabrication

The device mainly consists of three functional layers: a backing layer, a foldable

bilayer (a swelling layer/a non-swelling layer), and a mucoadhesive layer entrapped with

drugs (shown in Figure 6.1A). The swelling bilayer was made of MAA crosslinked by

TEGDMA and the non-swelling layer was HEMA crosslinked by DEGDMA.

Soft-lithographic techniques were used to produce hydrogel bilayered microstructures.

The devices were fabricated following the procedures shown in Figure 6.2. A PDMS

146

mold with a desirable surface pattern was made by casting a prepolymer and a curing

agent at 10:1 weight ratio onto a complementary relief structure from the standard

photolithographic process [Guan et al., 2005; Xia et al., 1998]. The HEMA monomer

solution was brushed onto the PDMS mold with an applicator. The solution was trapped

in the discrete wells due to discontinuous dewetting. After being subjected to UV

radiation for 10 minutes, the MAA monomer solution was brushed onto the cured

PHEMA layer to prepare a bilayered structure under another 15-minute UV radiation. A

high light intensity and large dosage were applied to ensure high monomer conversion

(around 99%). Our experimental observations showed no loosening or separation

between these bilayers. This is because the MAA solution diffused into the PHEMA layer

before the PMAA layer was solidified. To remove the residue monomer and unreacted

initiator, distilled water was used to continuously wash the cured structures covered by a

10�m-thick isopore membrane in the wells for 2 hours. To take out the bilayered

structures, the PDMS mold was placed on a PHEMA film covered on a glass slide by

briefly exposing the film to water vapor generated from a hot water bath. A solid weight

(50g/cm2) was placed on the PDMS mold for 10 minutes. The mold was then removed

with the bilayered structures stuck to the PHEMA/glass slide. In this study, the model

drug was mixed with Carbopol 934 and PVA to form a drug/mucoadhesive layer. A

homogeneous solution of these materials in distilled water (1:1:1, 10wt.%) was brushed

onto the PDMS mold. Water was allowed to evaporate and the drug/mucoadhesive layers

were formed in the wells. The PDMS mold was then aligned and placed onto the

bilayered structures. A solid weight (around 500g/cm2) was placed on the PDMS mold so

the sticky drug/mucoadhesive layer would adhere to the bilayered structures due to the

147

compression force. After the drug/mucoadhesive layers were totally dried out in 10

minutes, the PDMS mold was removed. By using this simple approach, we can make

both micro- and millimeter sized devices (240 �m − 4 mm). The typical dimensions of

device used in this study are shown in Figures 6.1(A) and (B). When the device is

conveyed into the small intestine, it may directly target onto the small intestine surface

due to the Carbopol mucoadhesion. Then the bilayered structures may fold into the

mucosa in a ‘grabbing’ manner, resulting in better drug protection and enhanced

mucoadhesion (Figure 6.1C).

148

Figure 6.1 Schematic of the 3-layer device from (A) side view and (B) top view, (C) folding on the small intestine surface, and (D) a capsule containing devices.

(A) (B)

2.0mm

4.0mm0.2mm

Swelling layer(PMAA, 50µm)

Drug/Mucoadhesive layer(PVA, Carbopol, Drug,

300µm)

Non-swelling layer(PHEMA, 50µm)

Thin non-adhesive layer(PHEMA, 1~10µm)

2.0mm

4.0mm0.2mm

Swelling layer(PMAA, 50µm)

Drug/Mucoadhesive layer(PVA, Carbopol, Drug,

300µm)

Non-swelling layer(PHEMA, 50µm)

Thin non-adhesive layer(PHEMA, 1~10µm)

(C)

Small intestineSmall intestine

4mm4mm

(D)

149

1. Non-swelling monomer solution(HEMA/DEGDMA, 60 wt.%)

2. UV partial curing and drying

3. Swelling monomer solution(MAA/TEGDMA, 50 wt.%)

4. UV curing

5. Stamping of bilayered structures

6. Stamping of prepared drug layer

PDMS Mold

Folding bilayer

Drug/mucoadhesive layer

Thin PHEMA layer

Thin PHEMA layer

1. Non-swelling monomer solution(HEMA/DEGDMA, 60 wt.%)

2. UV partial curing and drying

3. Swelling monomer solution(MAA/TEGDMA, 50 wt.%)

4. UV curing

5. Stamping of bilayered structures

6. Stamping of prepared drug layer

PDMS Mold

Folding bilayer

Drug/mucoadhesive layer

Thin PHEMA layer

Thin PHEMA layer

Figure 6.2 Fabrication procedure of the miniature devices.

150

6.2.3 Swelling and self-folding studies

To prepare hydrogel samples for the swelling test, a monomer solution was

transferred to a glove box under a nitrogen atmosphere. Nitrogen was bubbled through

the solution for 20 minutes, then the mixture was pipetted between two glass slides

separated by a Teflon spacer. The thickness of the spacer was 0.3mm. The setup was then

placed under a UV light for photopolymerization at 2.0 mw/cm2. The cured hydrogels

were then rinsed in double deionized water overnight to remove unreacted monomer,

initiator and sol fraction. Subsequently, the monomer-free disks were cut into disk

samples with 5.0 mm in diameter.

Swelling tests were performed at various pH values ranging from 3.0 to 7.0 to

characterize the hydrogel behavior in the GI tract. The buffer solutions with different pH

values were prepared by mixing the citric acid with appropriate amounts of sodium

phosphate solution. Sodium chloride was used to adjust the ionic strength of all solutions

to I=0.1M, which is the near-physiological condition. For the swelling test, the dried

hydrogel samples were weighed and placed in the buffer solution at room temperature

(25°C). The hydrogels were taken out of the solution at pre-selected time intervals. After

the extra water on the surface was removed by laboratory tissue, the weight of the wet

hydrogels was measured. The weight-swelling ratio was calculated by the weight of the

swollen sample to the weight of the dried sample. Self-folding of the hydrogel bilayers

was observed and recorded in a buffer solution and on the porcine small intestine. All

animal procedures were performed based on the institutional protocols.

151

6.2.4 Mucoadhesion measurement

The detachment between the device and a segment of porcine small intestine was

measured in a flow trough and a microbalance (shown in Figure 6.3). First, a sacrificed

small intestine was longitudinally cut into small pieces (2cm × 3cm), sliced lengthwise to

spread flat, exposing the lumen side, bonded on the trough bottom by super glue, then

washed with 50 ml phosphate buffer saline (PBS) solution. Before the pump drove the

buffer solution through the trough, the sample was gently dropped on the intestinal

surface. The buffer solution with a high viscosity was prepared by mixing 0.2wt%

Xanthan Gum (CP Kelco, Wilmington, DE) in a pH=6.5 buffer for a solution viscosity of

87.9 cp. By controlling the flow rate, the residence time of samples on the intestinal

surface was determined through the microscope observation.

To prevent the acidic degradation in the stomach, the devices can be loaded in an

enteric capsule (shown in Figure 6.1D), so they can maintain the shape until the enteric

capsule is dissolved in the small intestine. The flow experiments were carried out to

evaluate the device adhesion in the small intestine. Briefly, a 15cm long porcine intestine

segment was placed horizontally on a bench top to form a flow channel and one end was

connected to a tube so that the lumen could be filled with a pH=6.5 buffer solution at a

volumetric flow rate of 1 ml/min. A capsule containing three devices shown in the

following figure was placed near the entrance of the tube and pushed into the intestine

channel. After 20 minutes, the flow test was stopped and a longitudinal incision was

carried out in the intestine to observe the device attachment. The experimental

temperature was maintained near 37°C.

152

Figure 6.3 Experimental setup for (A) flowing testing and (B) the detachment force measurement.

Sample

Buffer CollectorPump

Microscope

Sample

Buffer CollectorPump

Microscope

(A)

DCA

Buffer solution aroundthe tested sample

Small intestine

DCADCA

Buffer solution aroundthe tested sample

Small intestineSmall intestineSmall intestine

(B)

153

The detachment forces were also quantitatively measured by a microbalance

attached to a dynamic contact angles analyzer (Cahn DCA-322). A 3.0 cm section of

intestine was cut and bonded on a beaker bottom as in the flow test, and covered with

pH=6.5 PBS solution at room temperature. The beaker was then placed in the

microbalance enclosure and fixed on the stage. A cylindrical sample (the bottom area: 2

mm × 2 mm) or a miniature device, mounted on a clamp and hung from the sample loop

of the microbalance, was brought in contact with the tissue by moving up the stage. The

polymeric sample was left in contact with the tissue for three minutes with an applied

force of approximately 100 mN and then pulled vertically away from the tissue sample by

moving down the stage while recording the required force for detachment. The

mucoadhesion force was normalized by the contact area.

6.2.5 Delivery performance

To evaluate whether the self-folded device has any improved effect on drug

protection and transport, targeted unidirectional release was conducted for trans-

epithelium delivery of two model drugs in a side-by-side diffusion chamber. Having

rinsed with PBS buffers, the jejunum part of the intestine was cut into a disc shape of 2.2

cm in diameter and placed on a support between the two chambers (the effective

diffusion area was 2.83 cm2). Before the experiment, the prepared device (the dimension

4mm×4mm, shown in Figures 1A and B) was placed onto the jejunum surface in the

donor chamber. Subsequently, 8 ml of pH=6.5 buffer solution was simultaneously

injected into both the donor chamber and the receptor chamber at room temperature

(25°C). The setup was subjected to constant shaking at 180 rpm. At predetermined time

154

intervals, 0.15 ml buffer solution was taken from the receptor chamber for concentration

test. To maintain a constant volume, 0.15 ml fresh PBS buffer was added after each

sample was withdrawn.

AO8 release was measured by monitoring its absorbance at 490 nm using a

microplate reader (GS Spectra MAX250). The concentration of AO8 in the buffer

solution was obtained from a calibration curve, and the amount of AO8 release at time t

(Mt) was calculated from accumulating the total AO8 release up to that time. The

fractional drug release, Mt/M0, could then be calculated. Here M0 is the amount of

initially loaded AO8. For the BSA release experiment, 0.1 ml samples were taken and

replaced by fresh buffer. After accounting for dilution caused by previous measurements,

protein concentrations were measured with a Bio-Rad protein assay using the microplate

assay protocol. The color change of the dye in response to the concentration change was

monitored by measuring the absorbance at 595 nm on the same microplate.

6.3 Results and Discussion

6.3.1 Swelling and self-folding studies

The pH-sensitive hydrogel, PMAA has been studied extensively as a promising

candidate for oral delivery of peptide and protein drugs through the gastrointestinal tract

because of its unique swelling property. Figure 6.4 exhibits the dynamic swelling

behavior of the hydrogels in different buffer solutions. As can be seen, the dried

hydrogels swelled at all pH conditions due to the adsorption of water into the porous

structure. In the high pH buffers, PMAA hydrogels swelled rapidly and achieved a much

higher weight-swelling ratio. This was because ionization of the carboxyl groups (the

155

pendent group of MAA) occurred as the solution become less acidic, resulting in

dissociation of the hydrogen bonds between the carboxylic acid groups of MAA and the

oxygen of the ether groups of TEGDMA. The dissociation of hydrogen bonds, combined

with the electrostatic repulsion force, caused the hydrogel network to swell quickly and

greatly under an osmotic pressure. Below a pH of 6.5, the swelling ratio drastically

decreased to a small value. This implied that the hydrogel was in a relatively collapsed

state. On the other hand, PHEMA is a neutral hydrogel, which has no ionizable groups on

its side chain. With a change of pH values, this material exhibited very little swelling in

buffer solutions. In addition, since the solvent content in the HEMA monomer solution

(40 wt.%) was less than that in the MAA solution (50 wt.%), PHEMA hydrogels should

have a more compact structure than PMAA gels with the same crosslinking ratio.

Although DEGDMA has a shorter chain than TEGDMA, its contribution could be

neglected when considering the low amounts of crosslinker.

Due to different swelling of the two layers, the bilayered structures would curl in

the buffer solutions. To demonstrate the self-folding function, a dried bilayer is shown in

Figures 6.5(A) and (B), respectively. The dried bilayer consisted of a PHEMA layer at the

top and a PMAA layer at the bottom. The bilayers represent a convex curvature after

becoming completely dried out. Figure 6.5(C) shows the folded bilayer in a buffer

solution (pH=6.5). It was observed that this structure folded like a fist.

156

Figure 6.4 Dynamic swelling behavior PMAA and PHEMA hydrogels.

0

4

8

12

16

20

24

0 30 60 90 120 150 180

Time (min)

Wei

gh

t S

wel

ling

Rat

io (

g/g

)

PMAA in pH=7.3

PHEMA in pH=7.3

PMAA in pH=6.5

PMAA in pH=3.0

PMAA in pH=7.3

PHEMA in pH=7.3

PMAA in pH=6.5

PMAA in pH=3.0

157

Figure 6.5 Optical graphs of a bilayered structure at dried state (A) top view, (B) side

view, (C) a curled bilayered structure in a buffer solution. Scale bars=2.0 mm.

A B CA B C

158

6.3.2 Mucoadhesion measurement

The layered shape of the device maximizes its contact area with the intestinal wall,

while the thin side areas minimize its exposure to the liquid flow through the intestine.

Additionally, since the bilayers curl into the mucus in the mode of “grabbing”, it is

expected to provide more resistance to mucus shedding than conventional mucoadhesion.

Thus, the residence time can be significantly increased due to the combination of the

“grabbing” adhesion of the folding bilayers and the conventional adhesion of the

mucoadhesive layer. This enhanced performance was demonstrated in the flow test. At

5cm height, samples with similar dimensions were randomly dropped on the intestinal

surface using tweezers without external force and the flow rate was gradually adjusted

from 4.0 to 5.5 ml/s. Figure 6.6(A) summarizes the number of bound samples remaining

on the mucus surface as a function of the flow time. For each case, the initially bounded

samples were the same. Within three minutes, all samples with a PHEMA surface were

washed away at a flow rate of 4.0 ml/s. For the PCL patches (i.e. the drug layer adhered

onto a PCL layer) and the folded devices, all samples still stayed on the mucosal surface

after 60 minutes. A higher flow rate (5.5 ml/s) was then used in the measurement.

According to the Figure 6.6(B), the average residence time for the PCL patch was around

72 minutes. The folded devices showed the longest average residence time, around 103

minutes.

To visually demonstrate the folding behavior and enhanced mucoadhesion, a

folding device tinted with blue dyes was placed on the mucus surface and a digital

camcorder recorded its folding process from the side view. Figure 6.7(A) shows the

folding behavior of a bilayered device with each layer having a thickness of 10 µm.

159

Figure 6.6 (A) Number of bound samples and (B) residence time for different samples

attached to intestinal mucus in the flow test.

0

1

2

3

4

0 30 60 90 120Time(min)

Nu

mb

er o

f B

ou

nd

Sam

ple

s

(A)

(B)

0

20

40

60

80

100

120

PHEMA PCL Patch Folded Device

Res

iden

ce T

ime

(min

)

160

In the beginning, the device adhered on the mucus surface. Around 2 minutes later, the

bilayered structure started to fold into the mucus and at 4 minutes the structure completed

the folding. Temperature is a very important factor, which may influence the swelling

ratio of gels, response time of folding bilayer, and residence time of the folded device. At

the typical body temperature 37°C, the swelling ratio of PMAA in pH=6.5 buffer was

increased from 10.39 to 11.01 and the response time was improved to 2 minutes as a

result of temperature increase. The residence time of the folded device also increased due

to the increased extent of folding.

Snapshots shown in Figure 6.7(B) describe the device attachment in the flow test.

As a control, a PCL patch was also placed on the mucosal surface. At the beginning, both

devices attached onto the surface tightly in the flow field. After 65 minutes, the PCL

patch started to detach from the surface. Around 70 minutes, the patch was completely

washed away from the mucosal surface. Due to the combined effect of mucoadhesion and

self-folding, the folded device could stay on the mucus for a longer time. It started to

detach at 82 minutes and was finally washed away at approximately 108 minutes. The

detachment was due to the mucus shedding, not the unfolding of the bilayered structure.

PMAA is a typical mucoadhesive material with a strong detachment force. To

ensure that only the drug/Carbopol layer would stick on the mucosal surface, a thin

PHEMA layer was added onto the PMAA side. Since the compression pressure from a

solid weight (50g/cm2) was weak, this layer could be delaminated from the folded bilayer

after the device was immersed in the buffers. The presence of the thin PHEMA layer also

offered a delay time for device folding. Figure 6.8 compares the attachment of two

devices with different contact sides on the porcine intestine. For the left one (S1), the

161

Figure 6.7 Dynamic processes for (A) folding behavior and (B) enhanced mucoadhesion. Buffer pH=6.5 and 25°C.

A

B

Folded device

Time (min)0 2 4

4mm

Folded device

Time (min)0 2 4

4mm

Time (min)0 2 4 Time (min)0 2 4

4mm

Folded device

Folded device Patch

Time (min)0 65 82 108

3mm

Folded device Patch

Time (min)0 65 82 108

Folded device Patch

Time (min)0 65 82 108 Time (min)0 65 82 108

3mm

162

PHEMA-side contacted with the mucus, while the right one (S2) showed the

drug/Carbopol layer in contact with the mucosal surface. S1 was washed away

immediately at a flow rate of 4.0 ml/s, while S2 stayed on the surface. This thin PHEMA

layer was completely peeled off in several minutes (about 10 minutes for this case) when

the bilayered arms curled into mucus. When the enteric capsule dissolved in the flow

experiment, the devices were able to adhere to the mucos and fold. This experiment was

repeated three times. Eight out of nine devices were found adhered to the lumenal wall by

the drug/Carbopol-side.

The enhanced mucoadhesion of the self-folding device was also revealed in the

detachment force measurement. As shown in Figure 6.9, the one-layer PCL and PHEMA

samples exhibited very weak adhesion. The major component of Carbopol is acrylic acid,

which is a mucoadhesive material. To prepare the sample for the detachment

measurement, Carbopol 934, PVA and the model drug were mixed to form a

homogeneous solution in distilled water (1:1:1, 10.0wt.%), which was then poured into a

petri dish. Water was allowed to evaporate and a drug/mucoadhesive layer was formed.

Samples of 2 mm×2 mm dimensions were cut for the detachment measurement. The

Carbopol/PVA/Drug sample showed a much stronger detachment force, which could be

explained by the formation of hydrogen bond due to the carboxylic acid groups [Peppas

et al., 1996]. The strongest force was observed for the folded device. These results agree

with what was observed in the flow test.

163

Figure 6.8 Compared attachments for the devices with different contact sides in the flow

test. Buffer pH=6.5 and 25°C.

Time (min) 0 0.5 10

S1: PHEMA-side

S2: Carbopol-side

2mm

S1S2

Time (min) 0 0.5 10

S1: PHEMA-side

S2: Carbopol-side

2mm

S1S2 S1S2

164

Figure 6.9 The detachment force of different samples on the small intestinal surface.

Buffer pH=6.5 and 25 °C. Error bar = SD, n = 3.

0

2

4

6

8

PCL PHEMA Carbopol /PVA /Drug

FoldedDevice

Det

achm

ent F

orce

(m

N/c

m2)

165

6.3.3 Delivery performance

A side-by-side diffusion chamber was used for drug release studies. When the

device was attached to the intestinal surface, the drug concentration change in the donor

chamber indicated the leakage in the small intestine. Figure 6.10 compares AO8 leakage

of delivery systems with different protection layers. Due to good mucoadhesion, a simple

PMAA layer could adhere to the mucus surface tightly and the leakage was very low in

the beginning. After 60 minutes, the high swelling of PMAA hydrogel, however, led to a

very large permeability resulting in severe drug leakage through the protection layer. For

the PCL layer, the drug could gradually leak into the donor chamber from the edge of the

patch. For the bilayered structure, since the PHEMA protection layer has a lower

permeability than the PMAA layer, it served as a barrier to provide protection from drug

leakage. Furthermore, the folded structure prevented the leakage from the edges.

Consequently, the total leakage from the folded device was very low, less than 30% of

loaded drugs after 2 hours.

For in vitro drug transport across the mucosal epithelium, we separated the

mucosal membrane from the serosal compartment of the small intestine. The isolated

mucosal membrane was loaded in the side-by-side diffusion chamber for the diffusion

measurement. The drug concentration in the receptor chamber indicates the transferred

drugs. Figure 6.11 compares the AO8 transport from different systems across the mucosal

epithelium. The squares indicate the homogeneous solution loaded into the donor

chamber. The triangles and the circles are for the PCL patch system and the folded device,

respectively. All three systems had an equal amount of loaded drug. The figure shows

166

Figure 6.10 The fractional leakage of AO8 from the drug reservoir with different protection layers (thickness=20 µm) at pH=6.5 and 25°C. Error bar = SD, n = 3.

0

0.2

0.4

0.6

0.8

1

0 20 40 60 80 100 120

Time(min)

Fra

ctio

nal

Lea

kag

e (M

t/M

o)

PMAA PCL PMAA and PHEMAPMAA PCL PMAA and PHEMA

167

Figure 6.11 AO8 transport from different systems across the mucosal epithelium at

pH=6.5 and 25°C. Error bar = SD, n = 3.

0

0.2

0.4

0.6

0 30 60 90 120

Time(min)

Fra

ctio

nal

Rel

ease

(M

t/M

o)

● Folded device ▲ PCL Patch ■ Solution

168

Figure 6.12 BSA transport from different systems across the mucosal epithelium at pH=6.5 and 25°C. Error bar = SD, n = 3.

0

0.1

0.2

0.3

0.4

0 30 60 90 120

Time(min)

Fra

ctio

nal

Rel

ease

● Folded device ■ Solution

169

that only about 12% AO8 in the solution was delivered through the mucosal epithelium in

120 minutes, while 20% AO8 loaded in the patch system could transfer across the

intestinal membrane. The self-folded device showed the highest drug transport fraction

(33%) due to its localized high drug concentration.

To compare the release behavior of drugs with different sizes, BSA was also used

as a model drug. In the experiment, the BSA loading concentration was about 3 times

higher than that of AO8 in order to provide easy detection by UV spectroscopy. Figure

6.12 shows the BSA transport profile from a folded device and the homogeneous solution

at room temperature. As can be seen, the self-folded device exhibited an improved BSA

transport fraction. Compared with Figure 6.11, the transport of large molecules across the

mucosal epithelium was much more difficult than small molecules.

6.4 Conclusions

A self-folding miniature hydrogel device has been developed based on the

integration of a number of micro-manufacturing modules. They demonstrated

multi-functionalities such as enhanced mucoadhesion, lower drug leakage, and improved

unidirectional delivery. The enhanced mucoadhesion due to self-folding increased the

residence time at the target site, and led to improved drug transport. The PHEMA layer

served as a diffusion barrier to provide good drug protection and prevented the drug

leakage.

170

CHAPTER 7

CONCLUSIONS AND RECOMMENDATION

7.1 Conclusions

This work determined the roles of the solvent composition and light intensity in

the photopolymerization of the MAA/TEGDMA resin system. It was found that the rate

of polymerization increased and more compact gels would form with a higher water

fraction in the 50wt% solvent/reactant mixture. This is because the weaker interactions

between MAA and solvent molecules give a higher opportunity for propagation and a

higher reaction rate. The hydrophobic TEGDMA and initiator tend to form aggregates in

the higher water solution, contributing to the inhomogeneous microgel formation. It was

also conlcuded that the rate of polymerization was enhanced as the light intensity

increased, especially at the low light intensity range and low conversion. At too high a

light intensity, a reduced MAA conversion was obtained. Additionally, the high light

intensity significantly shortened the reaction time to reach the macro-gelation and

increased the swelling ratio of formed hydrogels, which can be explained by the

mechanism of intra- vs. intermolecular reaction. With a high UV intensity, more free

171

radicals and more possibility for intramolecular reaction lead to a higher reaction rate and

faster gel formation. Since the intramolecular reaction contributes to less crosslinked

microgels, the resulting hydrogels have a higher swelling ratio.

By using these desired functional hydrogels cured under the optimal

polymerization conditions, an assembled and a self-folding DDS were developed based

on the selected integration of a number of micro-manufacturing modules to achieve

multi-functionalities such as drug protection, self-regulated oscillatory release, enhanced

mucoadhesion and targeted unidirectional release. The self-folding device first attached

to the mucosal surface and then curled into the mucus, leading to enhanced

mucoadhesion in the mode of “grabbing”. Furthermore, the folded layer served as a

diffusion barrier, minimizing the drug leakage in the small intestine. The resulting

unidirectional release provides improved drug transport through the mucosal epithelium

due to the localized high drug concentration. The functionalities of the devices have been

successfully demonstrated in vitro using a porcine small intestine.

The novel delivery devices will be of great benefit to the advancement of oral

administration of proteins and DNAs. Since the mucus layer covers many tissues at other

specific sites, the devices may be applied for ocular, buccal, vaginal and rectal

administrations. The polymer self-folding phenomena at the microscale can also be

applied as probe arrays for bio/chemical sensing, carriers in cell-based bioreactors, and

tissue clamping.

172

7.2 Recommendation

The developed self-folding device has significantly enhanced the mucoadhesion

and extended the residence time for drug transport. To further improve the

mucoadhesion, it would be desirable if a device can penetrate the loose-adherent layer

and adhere to the firmly adherent mucus layer such that longer retention than a few hours

may be achieved. This objective can be realized by reducing the device scale and adding

the nanotips. The device scale should be reduced to 5 µm or less such that they can move

into the microvilli for a longer residence time. Traditional fabrication protocols, such as

phase separation, microemulsion and spray drying, have been successfully used for the

production of micro-/nano-particles for drug delivery [Jain, 2000; Langer, 2000].

However, the resulting particles are usually polydisperse and relatively simple

structurally due to the surface-driven manufacturing process of these methods. To obtain

an ideal delivery vehicle, a series of methods for making micron-sized polymeric layered

structures has been developed in our laboratory using a soft lithography micro-transfer

molding technique [Guan et al., 2005]. PDMS molds with an array of micron-sized wells

can be made by the standard soft lithography technique. Figure 7.1 presents the

self-foldable microdevices for drug delivery. These soft lithographic techniques can

produce microparticles with similar structures but are simpler and of lower cost.

Compared to the conventional microspheres for drug delivery, the microfabricated

capsules are more uniform in size and shape, have a higher drug loading capacity, and

may be absent of the burst effect that is typically associated with microspheres prepared

by conventional methods. This basic fabrication operation has been successfully

demonstrated in our laboratory and by other researchers. The remaining challenges are to

173

extend the technique to smaller particle sizes (i.e. 1-10 µm and nanoscale) with different

shapes, to extend the imprinting area, and to be adopted to the high precision

manufacturing platform for mass production.

Figure 7.1 Schematic of fabrication of self-foldable microdevices. Optical micrographs

of (a-c) bilayered microdevices with different curvatures controlled by the composition of the primary swelling layer; (d) a self-folded microdevice in water, and (e) several

microdevices folded into the mucus of porcine intestine [Guan et al., 2005].

174

Controlling particle size and size distribution is most important for drug delivery

applications. However, this alone is insufficient for the increase of delivery efficiency.

Although the folding structure of the microvillis spatially restricts the mobility of small

delivery devices, it has a high possibility for the delivery devices moving out these fine

structures due to the peristalsis of the intestinal wall. The miniature device with flat and

layered shape maximizes the contact area with the intestinal wall. The thin side areas

minimize the exposure to the flow of liquids in the intestine. To extend a long duration

time, the self-foldable finger-like arms and enhanced nanotips are considered in the

device design. Figure 7.2 shows the schematic of a proposed device from the side view

and the top view. By using a novel low-cost sacrificial template imprinting (STI) process

developed by our group [Wang et al., 2004], the nanotips can be introduced on the drug

layer. The enhanced nanotips not only help the device adhere to the firmly adherent

mucus layer, but also may mechanically open the local tight junctions for improved

permeability.

Figure 7.2 Schematic of the self-foldable microdevice with enhanced nanotips.

175

Furthermore, the choice of other functional copolymers can be used to better

control the response performance and delivery behavior of DDS for various applications.

The designed devices are mainly used for the delivery of proteins and DNAs through the

GI tract. However, the physiological characteristics of each segment in small intestine

changes a lot. For example, the pH value in the duodenum is around 5.5, while this value

increases to 7.0 in the ileum. Other factors such as food compositions also influence the

pH values. To deliver the device at a more specific site, the transition range of hydrogels

between the swollen and the collapsed state with a pH change needs to be very narrow.

Although it is possible to localize a device within each part of small intestine, the

attainment of site-specific delivery in the rectum (pH=7.0) is even easier than in the small

intestine [Kim et al., 2002]. The monomer composition is adjusted to match the

requirement for pH-sensitivity of functional hydrogels at different specific sites. It is

known that the transition range becomes sharper for more hydrophobic hydrogels and

shifts to a higher pH for gels with the longer alkyl group. The transition range of

PMAA is around the pH of 6.1. With the addition of a single methylene unit,

poly(ethylacrylic acid) exhibited a sharper transition at the pH of 6.3. The addition of

another methylene unit with poly(propyl acrylic acid) (PPAA) shifted the pH profile even

further and PPAA displayed a much sharper transition close to the physiologic pH

[Stayton et al., 2005]. Since the mucus layer covers many tissues at various specific sites,

the device may be applied for ocular, buccal, vaginal and rectal administrations. Variation

of the transition range for acrylic polymers with similar molecular weight provides a

series of potential candidates for these applications.

Delivery systems developed in this study are likely to enhance the oral

176

bioavailability of proteins and DNAs. The major market could be for improving the

delivery of existing therapeutic agents with established markets such as protein drugs

insulin, human growth hormone, and interfereon-alpha, and nucleic acid drugs such as

antisense oligonucleotides (e.g., anti-bcl-2 oligo Genesence). Recent advances in

biomedical research have yielded many novel therapeutic candidates that are based on

proteins or nucleic acid, which have tremendous clinical potential but minimal oral

bioavailability. The development of this technology can lead to significant benefits to

improve patient compliance and cost savings, in addition to the reduction in pain and

inconvenience associated with parenteral administration.

177

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