Post on 20-Oct-2015
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
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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.
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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.
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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
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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
xii
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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