Hydrogels

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Introduction

Hydrogels are hydrophilic polymers that absorb water and are insoluble in water
at physiologic temperature, pH, and ionic strength because of the presence of
a three-dimensional network. The cross-links can be formed by covalent bonds,
or electrostatic, hydrophobic, or dipole–dipole interactions. The hydrophilicity is
due to the presence of hydrophilic groups, such as hydroxyl, carboxyl, amide, and
sulfonic groups along the polymer chain.

The area of hydrogel research has expanded dramatically in the last 10 years,

primarily because hydrogels perform well for biomedical applications. This is true
for both the synthetic and natural hydrogels. Hydrogels work well in the body
because they mimic the natural structure of the body’s cellular makeup. Recent
advances in the use of hydrogels for tissue engineering, drug delivery, and contact
lens application, to name but a few of the many biomedical applications of hydro-
gels, have led to (for the ﬁrst time) the potential to design artiﬁcial organs in a
controlled fashion, to deliver drugs to speciﬁc sites in the body, and to fabricate
the ﬁrst true extended wear contact lenses.

This article focuses on the biomedical applications of hydrogels. Several ar-

eas of nonbiomedical applications are also discussed. Hydrogels can be classiﬁed
as synthetic, natural, “smart,” and biodegradable hydrogels. There may be many
more classes of hydrogels, but this class division is based on the current focus
of hydrogel research. Hydrogels which are interpenetrating polymer networks
(IPNs) and block copolymers, for example, are discussed throughout the article.
The preparation of hydrogels focuses on the details of free-radical cross-linking
polymerization, chemical cross-linking, irradiation cross-linking, and physical in-
teraction cross-linking techniques.

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Properties

A hydrogel is a cross-linked polymer that swells in water to an equilibrium value.
The dry hydrogel is called a xerogel or dry gel. When the hydrogel is dried, wa-
ter evaporates from the gel and causes collapse of the gel structure. If water is
removed without disturbing the network, either by freeze-drying or by organic
solvent extraction techniques, the resultant hydrogel is extremely porous. These
materials are referred to as aerogels (1). The amount of water a material needs
to absorb to be classiﬁed as a hydrogel remains undeﬁned, but most researchers
generally agree that if a material absorbs at least 10% water and is insoluble in
water, it can be classiﬁed as a hydrogel. The swollen equilibrated state of a hydro-
gel results from a balance between the osmotic driving forces that cause the water
to enter the hydrophilic polymer and the cohesive forces exerted by the polymer
chains in resisting expansion (2,3). They attain an equilibrium swelling state that
depends on the osmotic driving forces and the cross-link density. An equilibrated
state is reached quickly following immersion of the dry (xero) gel in water. Most
hydrogels, in fact, reach an equilibrium concentration of water within 15 min of
hydration time. The degree of hydration (water content) can be expressed using
the following equation:

%Water(weight)

= [(hydrated weight − dry weight)hydrated weight]×100

The degree of water absorption related to the dry state of the polymer is

called percent hydration. This is calculated using the following equation:

%Hydration

= [(hydrated weight − dry weight)dry weight]×100

The more hydrophilic the polymer and/or monomers used to prepare the

hydrogel, the higher the degree of hydration. This hydration can also, to some
extent, be controlled through cross-link density. The higher the degree of cross-
linking for a given polymer system will result in a corresponding decrease in
water content. An expression for polymer swelling is found in the Flory–Huggins
equation where the volume fraction of the polymer in the swollen gel is expressed
as

− 1

+ m

− 1

where m

and m

are the weights of dry polymer and solvent, respectively, and

and d

are the densities of dry polymer and solvent at 25

◦

C (4,5).

The hydrophilicity of the hydrogel system can also be expressed by means

of an interaction parameter (

χ). This parameter deﬁnes the interaction energy

during the process of hydration. The parameter is determined experimentally
from stress–strain curves and the swelling characteristics of the hydrogel using
the Flory–Huggins equation,

χ = −

2
0

)

− v

+ ln(1 − v

)

+ v

− 2

where

is the concentration of elastically effective chains in a volume unit of

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unswollen polymer, v

is the molar volume of solvent, and

α the isotropic di-

lation factor. One can approximate

to equal (v

)

if the polymerization is

performed in a solvent (v

in this case is the volume fraction of the monomer in

the original mixture or the volume fraction of the polymer at network formation),
and

can be approximated from stress–strain curve.

The water that is contained in hydrogels is believed to consist of a “bound”

water and “free” water (6). When the dry hydrogel polymer is placed in water, the
hydrophilic groups along the polymer chain are hydrated ﬁrst. Water will form
a hydration sphere around these hydrophilic groups. This type of water is called
bound water. The bound water molecules are believed to be tightly held in the hy-
drogel matrix through a series of chemical interactions, such as hydrogen bonding.
As the hydrogel continues to hydrate, additional water that is absorbed by the hy-
drogel is referred to as unbound or free water. This water ﬁlls the voids and pores
of the hydrogel. The water in this hydration sphere has less structure and higher
mobility than does bound water. Higher water content hydrogels contain more
free water. It is the free water that is believed to be responsible for the “end of
day” dehydration characteristics of high water content hydrogel contact lenses.
These types of water are determined experimentally using dsc (differential scan-
ning calorimetry) and nmr (nuclear magnetic resonance) techniques. It is impor-
tant to note, however, that not all researchers believe that free and bound water
exists. From heat-capacity measurements in the hydrogel poly(HEMA) [poly(2-
hydroxyethyl methacrylate)] and a study of the thermodynamics from mixing
water with poly(HEMA), researchers have indicated that strong interactions of
water with the polar groups along the polymer chain does not occur. Evidence as
determined by pulsed gradient nmr indicates that all of the water in poly(HEMA)
hydrogels diffuses as a homogeneous water phase (7).

The amount and type of water that is contained in the hydrogels determines

the diffusion and transport characteristics of the hydrogel. This is important for
drug delivery and contact lens application. It has been shown that the low water
transport characteristics of some contact lens hydrogel system is the major factor
responsible for contact lens adhesion. The oxygen permeability and water trans-
port characteristics of conventional hydrogels increase with increasing levels of
hydration. Details of the diffusion and dynamic transport behavior or hydrogels
can be found in excellent reviews (8). The design of a high water content hydro-
gels for biomedical application where mechanical integrity is required can be a
serious problem because of the fact that high water content hydrogels typically
possess very poor tear strengths.

Classes of Hydrogels

Hydrogels can be classiﬁed as either synthetic or natural according to their ori-
gin, degradable or stable depending on their stability characteristics, and “intelli-
gent” or conventional depending on their ability to exhibit signiﬁcant dimensional
changes with variations in pH, temperature, or electric ﬁeld.

Conventional Synthetic Hydrogels.

The class of conventional synthetic

hydrogels is prepared by free-radical polymerization of vinyl-activated monomers.
These monomers can be classiﬁed according to type and charge. The water content

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of the resultant hydrogel can vary widely, depending on the hydrophilicity of the
monomer and degree of cross-linking. A difunctional monomer is added to cross-
link the polymer chains.

One of the most important classes of monomers used to prepare synthetic

hydrophilic polymers and hydrogels are the methacrylates and acrylates. The prin-
cipal monomer is shown in (1) where R H (acrylate) and R CH

(methacrylate).

A wide variety of commercially available hydrophilic acrylates and methacrylates
exist. A huge advantage of this class of monomers is their relative ease of poly-
merization and low cost. The monomer 2-hydroxyethyl methacrylate (HEMA) (2)
has been used extensively in the contact lens industry. Poly(HEMA) possesses a
water content of 38% and has excellent mechanical strength. The monomer glyc-
eryl methacrylate (3) when polymerized and cross-linked results in a hydrogel of
approximately 70% water, depending on its purity and degree of cross-linking.

Another very important class of monomers used to prepare synthetic hy-

drogels is the acrylamide/methacrylamide (4) (R CH

or H) based monomers.

These include the monomers acrylamide (AA), N-methacrylamide (MAA), N,N-
dimethylacrylamide (DMA), and diacetone acrylamide (DAA). The polymers pre-
pared from AA, MAA, and DMA are all “super water absorbent” polymers, each
capable of absorbing several times their weight in water. DMA is a particu-
larly useful monomer for biomedical application in that it possesses excellent
hydrolytic and thermal stability characteristics and moderately low levels of tox-
icity, unlike AA, which is extremely toxic. The polymerization of the monomer
2-hydroxyethylmethacrylamide (5) also results in a “super water absorbent” poly-
mer. Despite its similar structure to the methacrylate analogue HEMA, this poly-
mer is capable of water contents as high as 85%, simply because of substitution
of the ester for an amide linkage.

The class of hydrophilic monomers based on the cyclic lactams is also an

important class of monomers. The most widely used lactam monomer is N-vinyl
pyrrolidinone (NVP) (6). This is also a “super absorber.” Hydrogels prepared from

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this monomer absorb several times their weight in water, despite the hydrophobic
aliphatic ring structure. Cross-linking agents for NVP and NVP-based copoly-
merizations include N,N-methylenebisacrylamide and allyl methacrylate. The
monomer NVP is widely used for contact lens application, and in fact, most com-
mercial contact lens materials contain NVP.

Monomers that contain ionic functionality are also widely used for the prepa-

ration of hydrogels. These include methacrylate, methacrylamide, and styrene-
based monomers that contain acidic or basic functionality. In this class of
monomers is included methacrylic acid, acrylamidomethylpropylsulfonic acid (7),
and p-styrene sulfonate. These are typically used as comonomers at low concen-
tration. The ionic functionality in a buffered saline environment dramatically
increases the water content of the resultant hydrogel. For example, copolymer-
ization of 2 wt% methacrylic acid with HEMA results in a hydrogel having a
water content of 60% (compared with a 38% water content for HEMA alone). In a
similar fashion, the cationic monomer methacryloyloxyethyltrimethylammonium
chloride (MAC) is extremely hydrophilic. A hydrogel containing 60% NVP, 10%
MAC, and 30% HEMA results in a material having an equilibrium water content
of 87% (9).

The copolymerization of MAC with methacrylic acid results in ampholytic

hydrogels. These are hydrogels that contain both anionic and cationic character-
istics (10). These materials exhibit interesting pH-dependent behavior. Zwitteri-
onic hydrogels, based on 2-(methacryloyloxy)ethyl-2

-(trimethylammonium)ethyl

phosphate inner salt (8), have also been designed for improved biocompatibility.
These hydrogels mimic the zwitterionic structure of phospholipids such as phos-
phatidylcholine, which is a major component of the outer membrane of all living
cells (11). Zwitterionic monomers contain both a cationic and anionic charge on
the same molecule.

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It is important to note that almost all high water content hydrogels pos-

sess poor mechanical properties. This is a huge disadvantage of hydrogels, es-
pecially when one needs to design a material for an application that requires
some mechanical integrity. A homopolymer of poly(NVP), for example, resembles
a hydrated gelatin. In order to overcome the poor strength of these materials, it
is necessary to copolymerize these “super absorbent” monomers with monomers
that are capable of improving the overall tear strength of the material. Several
monomers have been successfully used. One such monomer is methyl methacry-
late (MMA). Low concentrations of MMA in a copolymer formulation typically
result in a reduction in water content, but a dramatic improvement in mechan-
ical integrity is also observed. Monomers such as cyclohexyl methacrylate and
t-butylcyclohexyl methacrylate also work well (12). These monomers improve the
mechanical properties of hydrogel formulations by imparting rigidity to the poly-
mer network. This approach, however, is seriously limited in that phase separation
often occurs before acceptable properties are achieved. One approach to overcome
this phase separation issue, while improving the overall hydrogel strength, is
through the use of “hydrophilic-bulky” strengthening agents, such as 4-t-butyl-
2-hydroxycyclohexyl methacrylate (TBE) (9) (13). Excellent mechanical charac-
teristics and optical clarity have been achieved through the use of TBE. A 80/20
(w/w) copolymer of NVP and TBE results in an 85% water-containing hydrogel
that possesses a tear strength equivalent to the 38% water-containing hydrogel
poly(HEMA).

Silicone Hydrogels.

A new class of hydrogels based on silicone has been

developed (14). These materials were developed in an attempt to combine the high
oxygen permeability of polydimethylsiloxane and the excellent water absorption
characteristics of conventional non-silicone hydrogels. This new class of hydrogels
was developed primarily for contact lens application. These silicone hydrogels are
also prepared by free-radical polymerization techniques.

The biggest limitation in the design of silicone hydrogels is that silicone-

based monomers are hydrophobic and insoluble in hydrophilic monomers.
The copolymerization of methacrylate-functionalized silicones with hydrophilic
monomers generally results in opaque, phase-separated materials. There have
been essentially three approaches in the design of transparent silicone hydro-
gels, ie, in minimizing the phase separation that occurs during polymerization of
methacrylate-functionalized silicones with hydrophilic monomers. One approach
has involved a protection–deprotection procedure (15). This involves the protec-
tion of a hydrophilic monomer with a hydrophobic protecting group. Copolymer-
ization of the protected monomer with a methacrylate-functionalized silicone,
followed by removal of the protecting group, results in a transparent silicone
hydrogel. For example, trimethylsilyl-protected 2-hydroxyethyl methacrylate is

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readily soluble in silicone-based monomers. Copolymerization of the protected
HEMA with the monomer methacryloyloxyethyl tris(trimethylsilyloxy)silane
forms a transparent, hydrophobic material that, following immersion in a mild
basic or acidic deprotecting solution and extraction to remove hexamethyldis-
iloxane, gives a transparent silicone-based hydrogel (10). This is a good ap-
proach in principle; however, it is extremely expensive because of the large
number of processing steps and length of time required for complete depro-
tection.

Another approach has been the use of a solubilizing co-solvent. For example,

the addition of a co-solvent that is capable of solubilizing an incompatible mix-
ture of silicone and a hydrophilic monomer, in many cases, results in transparent
materials following cure. Examples of co-solvents include isopropyl alcohol, hex-
anol, and methyl dodecanoate. This approach has the additional advantage of con-
trolling the glass-transition temperature of the material. The addition of higher
concentrations of co-solvent results in a lower glass-transition temperature as a
result of a higher polymer chain ﬂexibility (higher free volume). It is imperative
that the glass transition of the material is below the cure temperature in order to
effect a complete cure.

The third and most successful approach has been the preparation of

siloxanes containing hydrophilic groups to improve the solubility of silicone-
based materials with hydrophilic monomers. In this approach several syn-
thetic avenues have been pursued including the synthesis of hydrophilic-
TRIS

[tris(trimethylsiloxysilyl)propyl

methacrylate]

derivatives,

siloxanes

containing hydrophilic blocks, and siloxanes containing hydrophilic grafts, such
as polyethyleneglycol-grafted silicones (11) (16–19).

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Natural Hydrogels.

Naturally occurring water-soluble polymers include

polynucleotides, polypeptides, and polysaccharides. These polymers are derived
from a variety of naturally occurring sources such as plants, animals, and hu-
mans, or are synthesized. The polymer collagen, for example, is obtained from
cows, pigs, and humans, depending on the type of collagen required. Polypep-
tides can be synthesized by a protection/solid support scheme or through recom-
binant DNA techniques. Hydrogels of naturally occurring polymers are prepared
by the chemical or physical cross-linking of these polymers. The chemical cross-
linking reaction of polysaccharides (alginate, chitin, chitosan, cellulose, oligopep-
tides, and hyaluronic acid (12)) and proteins (albumin, gelatin) leads to a vari-
ety of well-deﬁned hydrogels (20–22). Hydrogels prepared from these polymers
exhibit excellent biocompatibility, primarily because they mimic the structural
components of the body. In humans, glycoaminoglycans are hydrogels that exist
in the connective tissue, such as skin, tendon, and bone (23,24). Additional in-
formation on the naturally occurring polymers can be found in excellent reviews
(25,26).

“Smart” Hydrogels.

A truly amazing class of hydrogels that has found po-

tential use for a wide variety of applications is the class of “smart” or “intelligent”
hydrogels. Smart hydrogels are, in most respects, very much similar to conven-
tional hydrogels. They are synthesized using similar methods and absorb water.
Smart hydrogels, however, deserve to be in their own class. The uniqueness of this
class is due to the unusual volume changes that these polymers exhibit under the
application of very speciﬁc stimuli. Smart hydrogels exhibit signiﬁcant volume
changes in response to stimuli such as changes in pH, temperature, electric ﬁeld,
and light (27–29). The major reason for the interest in these polymer systems

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is that the smart polymers, in most respects, behave like biopolymers in nature;
small changes of some speciﬁc stimuli dictate most of the biochemical response in
the body. The desire to form hydrogel systems that can mimic biological systems
drives a huge area of research. This volume change is abrupt and occurs with only
a small change in stimuli. In solution these changes are referred to as the lower
critical solution temperature (LCST), where at a speciﬁed temperature the poly-
mer precipitates from solution. For a hydrogel system, these changes are marked
by an order of magnitude change in the size, shape, and water content of the hydro-
gel. The hydrogel returns to its original state when the stimuli is removed. These
shifts are triggered by changes in the physical state of the hydrogel as a result of
changes in the hydrophilic/hydrophobic microstructure of the hydrogel. In design-
ing such systems, the goal is to control the balance of the hydrophobic/hydrophilic
nature of the hydrogel system. The driving force behind these transitions varies,
with common stimuli including (1) the neutralization of charged groups by either
pH shift, (2) the addition of an oppositely charged polymer, and (3) change in the
efﬁciency of hydrogen bonding with an increase in temperature or ionic strength
(30).

Biodegradable Hydrogels.

Biodegradable hydrogels, much similar to

that of the smart hydrogels category, has expanded at such a fast pace that it
now deserves to be in its own class. The uses of biodegradable hydrogels now
encompass a wide variety of applications, for both biomedical and nonbiomedi-
cal uses. In the design of biomedical products, the basic objective is to fabricate
materials that resorb or degrade in a physiological environment so that the de-
vice ultimately disappears with no adverse reaction (31,32). Degradable polymers
undergo chain scission to form low molecular weight oligomers or monomers. Ul-
timately, the oligomers and monomers are either fully degraded to biosubstances
or are eliminated by the body. Degradation is characterized by loss of molecu-
lar weight, loss of mass and mechanical strength. The deﬁnition of biodegrada-
tion is broad and includes a variety of degradative mechanisms, depending on
whether the degradation follows (33). Degradation hydrolytic, thermal, or en-
zymatic degradation pathways (see B

IODEGRADABLE

OLYMERS

, M

EDICAL

PPLI

CATIONS

Methods of Preparation

Free-Radical Polymerization.

Free-radical polymerization cross-linking

is the preferred route used to prepare hydrogels from the class of acrylates, amides,
and vinyl lactams. It can also be used to prepare hydrogels from the naturally oc-
curring polymers if the polymer backbone or chain end of the natural polymer has
been functionalized with a radically polymerizable group. It is also the preferred
route to prepare interpenetrating network hydrogels (IPNs) using either synthetic
monomers or natural polymers (again funtionalized with a radically polymeriz-
able group). The IPNs hydrogels are networks that contain two polymer systems,
each in its own cross-link network.

To form a hydrogel by free-radical polymerization, a difunctional cross-

linking agent must be added to the polymerization. The classical gelation

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theory can be applied to most of these polymerizations, especially for
methacrylate–dimethacrylate monomer combinations that have similar reactivity
(34). This theory allows for a fairly precise determination of the overall kinetics of
the system, which includes estimates of the gelation and vitriﬁcation point. There
are a variety of cross-linkers one can choose from. For methacrylate polymeriza-
tions, a common cross-linker is ethylene glycol dimethacrylate (13) that is added
at a concentration of 0.1–1.0%. For acrylamide systems, methylene bisacrylamide
is very common. For lactam-based systems, ethylene divinylurea works extremely
well.

Problems arise when one attempts to copolymerize monomers of unlike re-

activity, such as the copolymerization of HEMA with the vinyl lactam NVP. For
this type of polymerization, a block-type copolymer possessing high levels of un-
reacted NVP results. This is due to the slow kinetics of NVP polymerization
and unsuitable reactivity ratio of this comonomer system. The reactivity ratios
(r

and r

values) result in a “blocky” backbone. This high level of unreacted

NVP can be overcome by designing cross-linkers that are capable of polymer-
izing with both lactam and methacrylate functionality. For example, the cross-
linker methacryloxyethyl vinyl carbonate (14) (MEVC) is an excellent example
of a cross-linker that possesses reactivity for both a methacrylate and vinyl lac-
tam functionality. The vinyl carbonate group has the same reactivity as the vinyl
bond in NVP (35). The copolymerization of NVP and HEMA with MEVC results
in a signiﬁcant increase in water content (when compared with conventional
methacrylate cross-linkers) because of the higher incorporation of NVP in the
copolymer.

The chemistry of typical free-radical polymerizations involves an initia-

tion, propagation, chain transfer, and termination step leading to the forma-
tion of a cross-linked polymer system (36). The initiation step (radical formation
step) utilizes chemistries that when subjected to thermal or ultraviolet radia-
tion form radicals that react with activated monomers, such as a methacrylate.
A wide variety of thermal, ultraviolet, visible, and redox initiators are commer-
cially available. Typical thermal initiators include the class of azo compounds,
such as azobisisobutylonitrile (AIBN), and peroxide initiators, such as the per-
oxydicarbonates and the hindered peroctoates. Polymerization conditions vary

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widely depending on the type of initiator and its half-life. For example, the
respective 10-h half-life data of AIBN is 64

◦

C, and for cumylperoxydecanoate

the 10-h half-life is 38

◦

C. This is extremely helpful in controlling polymer-

ization exotherms, reactivity ratios, and monomer solubility. Typical commer-
cially available ultraviolet and visible light initiators include the benzoin methyl
ethers, acetophenones, and benzoyl phenyl phosphine oxides, to name but only a
few.

The polymerizations can be performed either in solution or neat. There are

advantages to both methods depending on the product end use. Solution polymer-
izations can be helpful when the preparation of large quantities of hydrogel are
required. The solvent for most reactions is water; however, a wide variety of polar
solvents can be used with the only requirement that they can be exchanged for
water in the hydration step. The polymerization exotherm can be controlled by
choice of solvent. For many copolymerizations, the addition of a suitable “solubiliz-
ing” solvent is necessary to solubilize monomers of widely varying hydrophilicity.
Also, a solvent can aid in the molecular weight control. By use of a chain-transfer
solvent such as an alkyl mercaptan, molecular weight and end-group functionality
can be controlled (37).

Neat polymerizations, also sometimes referred to as bulk polymerizations,

are typically performed between metal or glass plates utilizing a ﬂexible spacer
to accommodate shrinkage. Clamps are used to assure a complete seal. The metal
plates are many times treated with a ﬂuoro polymer, and the glass plates are
treated with a chloromethylsilane to facilitate plate separation and to elimi-
nate adhesion of the polymer ﬁlm. The ﬁlled plates are cured in an oven or
under ultraviolet or visible lights, depending on the desired initiation mode.
Following the cure, the ﬁlms are removed and extracted with an appropriate
solvent. Neat polymerizations are very fast, usually requiring only minutes for to-
tal monomer conversion. This, however, limits the reaction size in that exotherms
are difﬁcult to control. A big advantage of neat polymerizations is that there
is no need for solvent removal. Solvent removal can be very time-consuming,
requiring either extensive thermal devolatilization or solution extraction
steps.

Emulsion and suspension polymerizations are also an important route to

obtain hydrogels by free-radical polymerization (38). For some applications, this
is the preferred route, particularly when droplets or spheres of the hydrogel are
desired. This is an important route to prepare hydrogels for drug delivery appli-
cation. For these polymerizations, the initiator, solvent, and monomer are added
together with a suspending agent and/or emulsiﬁer. The cross-linker can be added
but is not necessary, depending on whether one desires a soluble polymer. A ma-
jor disadvantage of this route is that the emulsiﬁer and/or suspending agent is
sometimes difﬁcult to remove.

Chemical Cross-Linking of Polymers.

Another important method used

to prepare hydrogels is by the cross-linking of hydrophilic polymers (39). This route
can be employed for both the synthetic and naturally occurring polymers. In this
reaction, a bifunctional cross-linking agent is added to a dilute solution of the
hydrophilic polymer. The hydrophilic polymer contains functionality that is capa-
ble of reacting with the cross-linking agent. The reaction is typically in solution,

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but may also be performed through a suspension reaction where microparticles
or spheres are desired.

Most of the naturally occurring polymers can be cross-linked in this fash-

ion. For example, the polymers albumin and gelatin can be cross-linked with
formaldehyde or a difunctional dialdehyde (40,41). The aldehyde reacts with the
amino group along the albumin polymer backbone (15). Also, using a similar
approach, chrondroitin sulfate can be cross-linked with diaminododecane cat-
alyzed by dicyclohexycarbodiimide (42). In this example, increasing the concen-
tration of diaminododecane increased the degree of cross-linking. Cystein-bearing
polypeptides can be cross-linked through cystein bonds (43,44). Another exam-
ple is the cross-linking of poly(vinyl alcohol) using a diisocyanate-terminated
poly(ethylene oxide). Hydrogels have also been prepared from functionalized
poly(ethylene glycol) (PEG) through enzymatic cross-linking. PEG functional-
ized with a glutaminamide and a lysine-containing polypeptide were cross-linked
by the reaction of a natural tissue enzyme, transglutaminase (16). Transpar-
ent gels of water contents as high as 90% have been prepared by this route
(45).

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Irradiation Cross-Linking of Polymers.

Hydrogels can also be obtained

by ionizing-radiation techniques. This route can be employed for both the syn-
thetic and naturally occurring polymers. Ionizing radiation is a radiation that
possesses enough energy to ionize simple molecules either in air or water (46).
The radiation can be in the form of an electron beam irradiation or

γ –radiation.

This cross-linking reaction can be accomplished by irradiation of a hydrophilic
polymer in bulk or in solution. These reactions are usually performed in wa-
ter. The preferred method, however, is irradiation of a polymer solution. The
solution method is preferred because it requires less energy for formation of
a macroradical, and radical efﬁciency is increased because of the reduced vis-
cosity of the reaction mixture. When a polymer solution is irradiated, reactive
sites along the polymer backbone are formed. The main reactive species (when
water is used as the reactive solvent) are hydrated electrons, hydroxyl radi-
cals, and hydrogen atoms. It is the hydroxy radicals, however, that lead to ab-
straction of hydrogen atoms along the polymer chain with formation of a “mul-
tisite” radical functionalized macromolecule. When these radicals combine, a
cross-link is formed. This is not true for all polymer systems; some polymers,
in fact, will degrade under ionizing radiation (47). Many polymers in solution
will undergo simultaneous cross-linking and degradation reactions. Each poly-
mer system is unique and the optimum irradiation conditions need to be deter-
mined experimentally to minimize chain degradation and maximize cross-linking
reactions.

Gels through Physical Interactions.

Hydrogels can also be formed

through a series of physical interactions. It is this type of reaction that, in fact,
provides most of the cellular network in the body. These physical interactions
include polyelectrolyte complexation, hydrogen bonding, hydrophobic association,
and crystalline entanglements. Typical methods to prepare ﬁlms utilizing physical
interactions involves solvent casting or precipitation techniques.

Polyelectrolyte Complexation.

Hydrogels can be easily formed through the

formation of polyelectrolyte complexes (48). The bonds formed through polyelec-
trolytic complexation occur between pairs of charged sites along the polymer back-
bone. The hydrogels formed through electrolytic complexation are insoluble in
water, and the electrolytic bonds can be very stable depending on the pH of the
system (7–10 kcal).

Polyelectrolyte complexes are divided into four subclasses depending on the

basicity and acidity of the polyelectrolytes. These include strong acid–strong base,
strong acid–weak base, weak acid–strong base, and weak acid–weak base sub-
classes. The composition of these complexes is dictated by the degree of dissocia-
tion of the electrolytic components.

The

complexation

poly(sodium

styrene

sulfonate)

and

poly(4-

vinylbenzyltrimethylammonium chloride) is easy to form (17) (49). By com-
bining two water-based solutions of each polymer at equimolar concentration,
the complex forms immediately as evident by the formation of white pre-
cipitate. Another example is the polyelectrolyte complexation of an amino
group containing chitosan with the polymer sodium alginate (50). Sodium
alginate also forms a strong complex through the association of calcium ions
(51).

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Hydrogen Bonding.

Hydrogels can also be formed through the hydrogen

bonding of macromolecular chains. A hydrogen bond is formed through the associ-
ation of electron-deﬁcient hydrogen and a functional group of high electron density.
Similar to the polyelectrolytic complexes described above, hydrogen-bonded com-
plexes occur in many biological systems. The hydrogen-bonded complex (18) of
poly(acrylic acid) and poly(NVP) is one of the more common (52,53). This complex
is affected by a variety of factors, such as the molar ratio of each polymer, the solu-
tion temperature, polymer concentration, type of solvent, and polymer structure
(degree of association between complexing functionalities). This complex will not
form a gel at neutral pH.

Hydrophobic Association.

Hydrogels can also be formed through hy-

drophobic associations (54–58). Polymer systems such as block copolymers, graft
copolymers, and polymer blends form microphase/microdomain separated struc-
tures. The hydrophobic domains in these structures behave as associated cross-
link sites. These polymers combine hydrophobic segments that form unique
hydrophobic phases dispersed (or surrounded) by hydrophilic water absorbing
regions. These hydrophobic-associated polymer blends typically possess poor me-
chanical properties because of poor interfacial adhesion, and ﬁlms are usually
opaque because of macrophase polymer separation. Control of the size of the hy-
drophobic phase tends to improve the optical transparency and mechanical in-
tegrity. A precise balance needs to be established.

One of the major advantages of this approach is the resultant low cost eco-

nomics of the system. Commercially available polymers can be used to generate
a wide variety of high strength, low cost hydrogels. This is particularly true of
polymer blends. The hydrophobic phases of blend hydrogels form multicross-link
sites along the polymer backbone and, so loss of the water-soluble polymer is

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minimized. This is extremely important for biomaterials where polymer leaching
cannot be tolerated. Another advantage of this approach is that these polymer
blends are soluble in organic solvents and ﬂow at elevated temperatures. This
allows for the processing of these hydrophobic blends by injection molding. Co-
valently cross-linked systems are insoluble in organic solvents and do not ﬂow
even at elevated temperatures. An excellent example of a hydrophobically associ-
ated hydrogel is the copolymer of poly(butyl methacrylate-co-methacrylamide-co-
acrylic acid) with a blend of poly(N,N-dimethylacrylamide-co-N-vinylpyrrolidone)
(59,60). Hydrophobic association is the predominant associated force holding this
polymer blend together. It is reported that this polymer is stable in water at a pH
range of 1–11.

Biomedical Applications of Hydrogels

One of the ﬁrst areas of commercial application for hydrogels was contact lenses
(61). In the 1950s and 1960s Otto Wichterle in Prague discovered the hydrogel
poly(HEMA), and a very simple process to prepare contact lenses (62). The pro-
cess consisted of spinning a monomer solution in a preformed optical quality mold.
This ingenious work was completed on an old-fashioned “erector set” in the early
morning hours of Christmas. Since that time the area of hydrogel research has
expanded rapidly. Hydrogels have been used for a wide variety of nonbiomedical
applications, but the primary area of hydrogel research has focused on biomedical
applications. This is the result of the generally excellent biocompatibility exhib-
ited by hydrogels. Hydrogels have been successfully used for a wide variety of
biomedical applications including contact lenses, intraocular lenses, drug deliv-
ery devices, implants, and scaffolds for living cell encapsulations.

Contact Lenses.

Conventional Hydrogels.

With the discovery of poly(HEMA), the contact

lens industry began to ﬂourish. This is because HEMA is a transparent, soft ma-
terial that, when hydrated, absorbs 38% water, and as a result is very comfortable
to wear. The HEMA lens also has excellent wetting characteristics and biocompat-
ibility. Within a few years following this discovery, a number of companies began
to market their own version of the HEMA lens. A multitude of new hydrogel con-
tact lenses emerged (63–66). Most were based on copolymers of HEMA, NVP, and
glyceryl methacrylate, together with cross-linkers and initiators. These materials
were marketed as daily wear lenses, ie, wear the lens for one day, remove it at
night for cleaning, and again wear the lens the next day.

With the huge success of the daily wear contact lens market, researchers

began to look for ways to increase the contact lens wearing time. This started
a huge research effort to design contact lens materials for extended wear ap-
plication, with the primary goal to simplify the patients cleaning and wearing
schedule. Such design was not a trivial task. The design of a new contact lens
material for extended wear application requires a material to satisfy a number
of very strict design requirements. The material must be optically transparent,
possess chemical and thermal stability, biocompatibility, and be wettable to tears.
It must also have suitable mechanical properties. This requires a material to
have a low modulus for patient comfort and high tear strength for lens handling

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durability. In addition, it is important that the material can be processed through
free-radical polymerization techniques. Finally, the material must be permeable
to oxygen. There is a lack of blood vessels in the corneal network, and so the
cornea obtains almost all of its oxygen requirements from the atmosphere. Plac-
ing a contact lens over the eye that does not “breathe” may result in a number of
physiological problems (microcysts, inﬂammation, infections, and corneal ulcers)
(67).

There have been several approaches in the design of high oxygen permeable

hydrogels for extended wear contact lens application. The ﬁrst and simplest ap-
proach reduces the lens thickness to increase the oxygen transmissibility. The key
intrinsic material properties that are a measure of oxygen diffusion are oxygen
permeability (Dk), where D is the diffusion coefﬁcient and k is a proportionality
coefﬁcient called the Henry’s law coefﬁcient, and oxygen transmissibility (DkL

− 1

where the actual amount of oxygen reaching the cornea is inversely proportional
to the lens thickness L. Several thin-lens designed contact lenses were intro-
duced to the market with, however, limited success. These lenses were extremely
difﬁcult to handle, and most thin-lens designs did not provide enough oxygen
permeability for extended lens wear application.

The second approach consists of designing hydrogels with a high water con-

tent. A direct correlation exists between water content and oxygen permeability,
and so the higher the water content, the higher the hydrogel’s oxygen permeabil-
ity (Fig. 1,

). There are two methods of designing high water hydrogels. The

ﬁrst involves the polymerization of highly hydrophilic monomers, such as NVP.
The second utilizes the use of ionic monomers. The polymerization of HEMA with
small concentrations of methacrylic acid, in a buffered saline environment results
in a signiﬁcant increase in water content. For example, with the addition of 2 wt%
of methacrylic acid to HEMA, a 60% water content is realized (as compared with
38% for pure HEMA). Both methods require the use of cross-linkers and initiators
and may need strengthening agents.

120

140

100

120

% Water

Fig. 1.

Relationship between Dk and percent water for conventional hydrogels (

) and

TPVC-based silicone hydrogels (

♦) (Dk in units of Barrers).

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707

At the time of the discovery of high water content hydrogels, these materi-

als were considered a huge breakthrough with the potential, many practitioners
believed, to be the “ultimate contact lens material.” This promise, however, was
short-lived. There are several limitations of high water content hydrogels. The
ﬁrst is that high water content hydrogels typically possess poor tear strengths
and often exhibit a high afﬁnity for protein, particularly for hydrogels possessing
an ionic functionality (68). Deposits can affect material wetting, patient comfort,
visual acuity, and may even cause inﬂammatory responses. In addition, in dry
environments, high water content hydrogels can induce a clinical response known
as epithelial dehydration (a dehydration of the corneal epithelial cells) (69). This
results in damage to the corneal epithelial cells. This phenomenon is a result
of the high rate of water evaporation that occurs with the high water content
hydrogels.

Silicone Hydrogels.

Another approach in the design of an extended wear

contact lens consisted of the development of materials based on polydimethyl-
siloxane elastomers (PDMS). PDMS appeared to be an ideal polymer for use in
an extended wear lens because it possessed a low modulus of elasticity, excel-
lent transparency, and high oxygen permeability (70). The major drawback is that
PDMS is completely nonwettable to tears, thus requiring surface treatment to
impart wettability. These surface treatments were typically ineffective, resulting
in surfaces that possessed marginal wetting characteristics and a high afﬁnity for
lipids. Another signiﬁcant drawback was that silicone lenses under normal wear
conditions adhered to the cornea. This was attributed to the low water transport
and high recovery characteristics of silicone (See S

ILICONES

In an attempt to combine the high oxygen permeability of PDMS and the

excellent comfort, wetting, and deposit resistance of conventional hydrogels, the
design of silicone-based hydrogels for contact lens application has been studied.
This approach was also an attempt to design materials that did not adhere to the
cornea. But most of all, it would provide lenses possessing levels of oxygen per-
meability high enough for extended wear application with minimal physiological
impact.

One simple, yet elegant approach in the design of a transparent sil-

icone hydrogel has been reported (71,72). This material has been intro-
duced commercially under the trade name Balaﬁlcon. This new silicone
hydrogel system is based on a vinyl carbamate substituted TRIS deriva-
tive: (Tris(trimethylsiloxy)silylpropylvinylcarbamate) (TPVC, (19)). The TPVC
molecule contains the hydrophobic silicone portion, and attached to this silicone
is a hydrophilic vinyl carbamate group. This “direct hydrophilic” attachment now
gives the silicone signiﬁcant hydrophilic character. The TPVC molecule is soluble
in all proportions with hydrophilic monomers, such as HEMA and NVP. In addi-
tion, the vinyl carbamate group provides a “polymerization link” for copolymeriza-
tion with hydrophilic monomers. The polymerization reaction of TPVC with NVP
(and a suitable cross-linker) cast using uv initiation methods, results in transpar-
ent, high Dk, low modulus gels that are insoluble in water. Figure 1 (

♦) shows

the relationship between oxygen permeability (Dk) versus percent water for the
TPVC-based silicone hydrogels. This is an important relationship in that it clearly
shows the Dk advantage in designing hydrogels based on silicone. In contrast
to conventional non-silicone-based hydrogels, the Dk decreases with increasing

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water content because of the lower concentration of TPVC in the higher water
content copolymers.

Intraocular Lenses.

An intraocular lens (IOL) is an implant that is used

to replace the diseased or damaged natural lens of the eye (73–77). The lens is
supported in the eye through the use of haptics or loops that are attached directly
to the IOL optics. There have been a wide variety of materials that have been
successfully used as IOLs, ranging from the rigid poly(methyl methacrylate) to the
soft elastomer silicone. The current trend in IOL research is to design materials
that can be folded and inserted through a small corneal incision. The natural
lens is removed using emulsiﬁcation techniques, and the IOL lens is placed in
the original capsule bag that held the lens. Once the IOL is inserted, the lens
recovers to its original shape. The small incision is desirable to reduce the degree
of induced astigmatism (loss of sphericity) in the operated eye and to minimize
corneal trauma.

Hydrogels have also been successfully used for small incision IOLs. There

are a number of commercially available hydrogel IOLs based predominately on
HEMA, that range in water content from 17 to 28%. Aromatic-based methacry-
lates, such as phenylethyl methacrylate, are added to increase the refractive in-
dex of the lens. A higher refractive index will allow for a thinner IOL design (for
power considerations) and result in an IOL that can be inserted through a smaller
opening and recover its shape faster. A recent research effort has focused on an
expandable IOL hydrogel where the lens is inserted in a dehydrated state and
allowed to hydrate in the eye. This allows for the insertion of IOLs with reduced
dimensions that make insertion through a small opening possible (78). Once the
lens is implanted in the eye, it quickly hydrates, expands, and reaches its ﬁnal
dimensions within minutes.

Drug Delivery.

The goal of drug delivery is to maintain the drug concen-

tration in the body (plasma) within therapeutic limits for long periods of time.
A constant drug release rate (zero-order) is desired. Conventional drug admin-
istration (oral delivery, injection) usually results in poor control of the plasma
drug concentration. The controlled release of drugs from polymeric matrices has,
however, been very successful. Many polymeric devices that deliver drugs at a

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sustained release rate are now commercially available. Controlled release devices
offer the desired therapeutic range of drug dose. Controlled drug delivery applica-
tions include sustained delivery (time) and target delivery systems (insertion at
the diseased site). The delivery of drugs such as protein-based drugs, eg, insulin
and growth factor, and also conventional drugs, eg, steroids and antibiotics, can
be achieved.

Controlled release systems can generally be divided into three sections de-

pending on their mode of release: diffusion controlled, chemical erosion, and sol-
vent activated (79). In a diffusion-controlled device, the drug is surrounded by an
inert barrier and diffuses from a reservoir, or the drug is dispersed throughout
a polymer and diffuses from the polymer matrix. In a chemical erosion device,
the drug is dispersed in a bioerodible polymer system or is covalently linked to a
polymer backbone via a hydrolyzable linkage. As the polymer or hydrolyzable link
degrades, the drug is released. In a solvent-activated device, the drug is dispersed
within polymeric matrix and the device is swelled with a suitable solvent (usually
water). As the device swells, the drug is released.

Conventional Hydrogels.

Hydrogels have been used extensively in the ﬁeld

of controlled drug delivery (80–85). The advantages of using hydrogels in drug
delivery systems are that they can be used at a local level, ie, insertion at the
diseased site. This has become important because many of the new protein-based
drugs require delivery in this fashion. The release of a growth factor, for exam-
ple, to a speciﬁc site is highly desirable. This is because many biologically active
polypeptides have very short half-lives and can not be administered orally (86).
The hydrogel can be cross-linked at the diseased site by photopolymerization tech-
niques, complexation, and enzymatic cross-linking. The hydrogel will also conform
to the local anatomy (organs, vessels).

A wide variety of conventional cross-linked homo- and copolymeric hydrogels

have been used for drug delivery application. In most hydrogels, the rate of diffu-
sion through the bulk depends on two primary factors: the extent of cross-linking
and water content. The extent of cross-linking determines the extent of swelling
and the distance between chains within the hydrogel network. When entrapped
drugs are diffusing within the network, the rate of diffusion can depend on the
interchain separation and the size of the diffusing drug. The rate of diffusion is
also controlled by water content. The general approach is to design hydrogels with
very speciﬁc levels of hydrophilicity/hydrophobicity. This hydrophilic/hydrophobic
balance controls water content and drug diffusion, where each drug will diffuse
according to its dissolution proﬁle (87). For water-soluble drugs dispersed in a
high water content hydrogel, drug release will be rapid.

Hydrogel drug delivery systems can be used in the hydrated or dry states.

The drug is incorporated during the polymerization process or through diffusion
techniques using preformed ﬁlms, tablets, etc (equilibrium absorption from con-
centrated drug solutions). A major limitation of incorporating the drug during the
polymerizing process is that the drug will release during the extraction process
(the process used to remove unreacted monomers) and, as a result, this technique
is rarely used for biomedical applications. The diffusion method is a viable al-
ternate to disperse and deliver drugs. In this method, the hydrogel is soaked in
a drug solution until an equilibrium dispersion is reached. The hydrogel is then
dried, packaged, and sterilized before use. Dried hydrogels release the drug as the

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material hydrates (the drug will diffuse very slowly in an unhydrated hydrogel).
The release of water-soluble drugs from a dry hydrogel involves the simultaneous
absorption of water and desorption of drug through a swelling-controlled diffusion
mechanism (88). As water penetrates a glassy hydrogel matrix that contains a dis-
solved or dispersed drug, the polymer swells and its glass-transition temperature
is lowered. As the water enters the material, a water front forms that separates the
glassy polymer from the swelled hydrated polymer. In regard to drug distribution
and release, the solvent front separates the drug in the unhydrated core from the
swollen water phase. As the water phase continues to grow and expand, the drug
diffuses and is released. The drug release kinetics during this process can range
from a square root of time dependence (Fickian) to a linear time (Case II transport)
dependence. The Case II transport is governed by the rate of polymer relaxation.
For most cases, the intermediate situation, which is termed non-Fickian (rate of
Fickian diffusion and polymer relaxation are comparable), is observed (89). For
example, thiamine HCl release from an initial dry poly(HEMA) hydrogel bead, as
plotted versus the square root of time, shows an initial non-Fickian behavior with
linearity established only after long periods of time.

Many conventional hydrogel designs that achieve zero-order drug-release are

available. The manufacture of reservoir-based hydrogels (reservoir of drug encap-
sulated and surrounded by a polymer membrane) and matrix diffusion systems
(drug is homogeneously dispersed in the hydrogel matrix) using common hydro-
gels, such as HEMA and glyceryl methacrylate, are relatively easy. The zero-order
release rate can be achieved in many cases using either approach; however, the
release rate of drugs from a matrix-type system typically declines continuously
proportional to the square root of time. Soaking a poly(HEMA)/progesterone ma-
trix, followed by uv irradiation in the presence of a cross-linker, can result in a
rate-controlled membrane that exhibits a fairly steady release rate (90). Another
use of a controlled release device makes use of a semicrystalline hydrogel (91). The
hydrogels in these examples are based on poly(vinyl alcohol) and poly(ethylene
oxide). In these systems, ﬁlms are prepared through solution casting techniques.
The ﬁlms are subjected to an annealing process that creates varying degrees of
crystallinity depending on the annealing time and temperature. The degree of
crystallinity can be controlled and the degree of crystallinity is the controlling
factor of drug release. The rate of crystal dissolution in water controls the rate of
drug release. A commercial drug delivery application utilizing hydrogels has been
introduced by Geltex pharmaceuticals (92). In this work a nonabsorbed hydrogel
based on poly(allylamine) substituted with quaternary amine was designed for
hypercholesterolemia (high cholesterol). The quaternary amine binds selectively
with cholesterol. This polymer acts in the intestine as a bile acid scavenger to lower
serum cholesterol levels in patients with increased risk of vascular disease. Bile
acid sequestrants act by binding bile salts in the intestine and prevent them from
passing through the intestinal wall and back into the blood stream. A U.S. NDA
(new drug application) was ﬁled in July 1999 with approval received in May 2000.

Degradable Hydrogels.

Many research groups have studied the use of

degradable hydrogels for drug delivery (93). The encapsulated drug will release
at a controlled rate depending on such factors as pH, temperature, and cross-link
density. Many examples of degradable drug delivery hydrogels can be found in the
literature (39).

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One of the ﬁrst successful degradable release systems was based on the

polyester hydrogels. These hydrogels are prepared by the copolymerization of
malonate-type polyesters (polyesters that contain pendent backbone unsatura-
tion) with monomers, such as NVP and acrylamide. Release of serum albumin
occurs from these systems by a bulk erosion process. Other drugs can also be
used. As the ester linkage of the polyester backbone hydrolyzes, the entrapped
drug is gradually released to the surrounding environment. The release rate de-
pends on the cross-link density. The degradation components of these erodible
cross-linked polyesters result in nontoxic, water-soluble by-products. In a similar
system, polyester macromers of a poly(ethylene glycol)-co-(lactic acid or glycolic
acid) were prepared and evaluated for use as bioerodible hydrogels (94). These
gels degrade upon hydrolysis of the oligo(

α-hydroxy acid) regions and release the

by-products poly(ethylene glycol), the hydroxy acid, and oligo(

α-acrylic acid) (20).

The degradation rate can be controlled by tailoring the concentration of

α-hydroxy

acid. If polymerized in contact with tissue, the gels adhere, and if polymerized prior
to tissue placement, the gels are nonadhesive. The protein albumin was entrapped
within these hydrogels and shown to release at a consistent rate. The release rate
was dependent on the cross-link density and the molecular weight of the protein.

Drug-delivery devices combining cross-linked conventional hydrogels, where

the drug is covalently linked via a degradable oligopeptide side chain, have
been designed for targeted anticancer drugs. In this example, hydrogels of N-(2-
hydroxypropyl)methacrylamide were prepared. The drug was attached syntheti-
cally by a degradable oligopeptide linkage. In vitro experiments have shown that
these devices target ovarian carcinoma cells with promising results (95). Another
interesting example of hydrogels for drug delivery utilizes the concept of degrad-
able natural polymers (96). In this work, the natural oligosaccharide hyaluronic
acid (HA) is chemically modiﬁed through the use of a pendent hydrazido func-
tional group that is covalently linked to a wide variety of drugs, such as steroids.
The covalent link is hydrolytically unstable and under mild conditions it releases
the drug.

“Smart” Hydrogels.

Research efforts on the design of smart hydrogels for

drug delivery application has increased signiﬁcantly over the past few years
(30,31,97). Such systems show promise as drug-delivery devices because of
the rapid release of conventional and protein-based drugs during the expan-
sion/contraction of these hydrogel systems. Efforts have focused primarily on de-
signing systems that make use of change in pH and temperature. The idea behind
this approach is that smart hydrogels will both expand and contract, forming a
hydrogel “switch” that releases drug or protein in a controlled fashion. When a
drug is incorporated into a smart hydrogel, the diffusion of the polymer is depen-
dent on the gel-state. Smart hydrogels can be designed to release drugs either
above or below the lower critical solution temperature (LCST). Smart hydrogels

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have been designed to release drugs at low pH (gastric release application) and
at basic pH (intestinal release application).

Some brilliant work in this area has been described. Most of this has been

based on the polymer poly(isopropyl acrylamide) (IPA) (97–99). IPA has a critical
solution temperature close to that of body temperature (32

◦

C). A large variety of

IPA-based polymers (grafts, block) containing varying concentration of hydropho-
bic monomers have been synthesized. The concentration of hydrophobic monomer
controls the critical solution temperature. Recent examples include the design of
thermosensitive hydrogels as heparin-releasing polymers for the prevention of
thrombosis. In this system, copolymers such as of N-isopropylacrylamide, butyl
methacrylate, and acrylic acid were cross-linked together with heparin. At tem-
peratures below the LCST, the polymer swelled signiﬁcantly and solution load-
ing of the drug was performed. Above the LCST, the gel de-swelled and released
heparin. Another example makes use of a hydrogel based on an ether–urethane–
isopropylacrylamide IPN for use as a heparin-release system. In this example,
the “smart” polymer was applied as a coating on polyurethane catheters. This
heparin-release coating system resulted in a signiﬁcant reduction in thrombus
formation on test surfaces in contact with venous blood (100).

The design of smart polymers as glucose-sensitive systems that release

insulin has recently been described (101,102). The ultimate goal in this work is
to design an artiﬁcial pancreas. This is accomplished by incorporating glucose
oxidase into a pH- or temperature-sensitive smart polymer. An insulin-releasing
drug-delivery system was produced by the incorporation of glucose oxidase,
bovine serum albumin, and insulin into a gel of poly(N,N-dimethylamino)ethyl
methacrylate-co-ethylacrylamide)(103,104). Exposure of this system to glucose
in the desired physiological concentration range resulted in a reduction in the
system pH and swelling of the polymer (as brought about by protonation of
the polymer). This swelling causes a release of insulin. Another system based
on a polymer complex of poly(N-vinyl-2-pyrrolidinone-co-phenylboronic acid
acrylamide-co-dimethylaminopropylacrylamide) and poly(vinyl alcohol) also
shows glucose sensitivity. In this system, a stable gel forms at pH 7.4 because of
the “covalent” linkage of the boronate–hydroxy groups. The gel contains insulin.
Through the addition of glucose, the complex dissociates and transforms the gel
to a sol-state (the gel begins to solubilize) resulting in swelling and subsequent
insulin release (105).

Tissue Engineering.

The use of biomedical materials over the last 20

years has, in most respects, been extremely successful. A wide variety of materi-
als have been developed that have been successfully used to replace hips, heart
valves, and the natural lens of the eye, to name but a few successful applications.
The major drawback in this area of research, however, is that these materials,
for almost all cases, can only perform for short periods of time. No long-term re-
placements have to date been developed. The clinical needs are even greater in
the case of organs, where replacement by organ transplantation is the only option.
The area of tissue engineering, however, offers some intriguing possibilities.

The goal of tissue engineering is to create living, three-dimensional tis-

sue/organs using cells obtained from readily available sources, such as cells ob-
tained directly from the patient (31). The ultimate goal in this work is to grow
cells speciﬁc to a particular organ and then to direct the cell growth to form the

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actual organ. This theoretically can be accomplished through the attachment of
speciﬁc cells to a scaffolding matrix that directs cell attachment, differentiation,
and growth. In order to do this amazing feat, a physical scaffold is required to al-
low for the organization of cells to form the speciﬁc organ (to simulate cell growth,
migration, and differentiation). These scaffolds need to interact with the cells
through highly speciﬁc bioreactions that control cell adhesion and growth factor
responses. The scaffolds are ideally biodegradable. Some fabulous work in this
area has recently been reported.

The use of biodegradable hydrogels as a temporary-support template for

cartilage has been reported (106). Cartilage is a biphasic material made up of
collagen as the solid support suspended in a gel of proteoglycans. Within this
gel are the chondrocytes that are responsible for maintaining the extracellular
matrix of cartilage (glycosaminoglycans and type II collagen). For this work the
hydrogel calcium alginate was used as the scaffold. The alginate polymer was
prepared by dissolving in water and adding calcium ions to form a “complexed”
cross-link polymer network. The immobilization of chondrocytes in alginate was
accomplished simply by soaking the alginate in a solution of chondrocytes. Re-
searchers have demonstrated that the chondrocytes maintained their structure,
were capable of proliferating at rates signiﬁcantly higher than that of monolayer
cultures, maintained their production of glycosaminoglycans and collagen, and
formed a mechanically functional matrix in a hydrogel network. This has all led
to the ﬁrst successful experiments that have shown that new cartilage can be
created in vivo using hydrogel scaffolding.

In a similar concept, the design of biodegradable hydrogels for bone regenera-

tion through growth factor release has been reported (107). In this research effort,
gelatin was cross-linked with either glutaraldehyde or carbodiimide. The growth
factor was added by solution adsorption into the preformed gelatin hydrogels.
The cationic growth factor was held in the gelatin matrix through complexation
with the anionic sites along the gelatin backbone. The gelatin enzymatically de-
grades in the body to release the growth factor. When implanted into a bone defect,
the growth factor resulted in accelerated bone regeneration and closed the bone
defect.

Wound Dressings.

Hydrogels have also been used as wound dressings

because most hydrogels are soft, ﬂexible, conform to the wound, are biocompati-
ble, and are permeable to water vapor and metabolites. As wound dressings, they
absorb the exudate, do not stick to the wound, allow for access of oxygen to the
wound site, and accelerate healing. This is actually a very large market with to-
tal sales of synthetic wound dressings, as reported in 1999, of $350 million (108).
The dressing is usually applied as a thin preformed/prehydrated ﬁlm. The hydro-
gel wound dressings are based on hydrocolloid/hydrogels and alginate dressings.
Hydrophilic monomers and polymers used to prepare the hydrogel bandages are
based on NVP, poly(ethylene oxide), and poly(vinyl alcohol).

Biosensors.

A biosensor is a compact device or probe that detects, records,

and transmits information regarding a physiological change or the presence of var-
ious chemical or biological materials in the environment. A biosensor is a probe
that integrates a biological component, such as a biological product (enzyme or an-
tibody), with an electronic component to yield a measurable response. Biosensors
are used to monitor changes in the physiological environment (109). The usual

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aim of a biosensor is to produce either discrete or continuous electronic signals
that are proportional to a single analyte or a related group of analyte (110). The
biosensors typically comprise a biological sensing element, a transducer, a sig-
nal conditioner, a data processor, a signal generator, and one or more organic or
inorganic membranes. Hydrogels have been used as reactive matrix membranes
in biosensors. Hydrogels possess several advantages over other materials in that
they exhibit rapid and selective diffusion characteristics of the analyte, as well as
provide support. For example, there has been a signiﬁcant amount of success in the
area of lipid bilayer based biosensor membranes. One of the early disadvantages
of these bilayer probes was their poor ion diffusion characteristic. Hydrogels were
successfully used to provide both the required ion diffusion and also the support
for the thin lipid bilayer membrane (111). Among the various types of biosensors,
those that measure glucose have received the most attention, due primarily to
the fact that nearly 6% of the population suffers from diabetes (112). In these
biosensors, the consumption of oxygen or the formation of hydrogen peroxide is
monitored (enzyme glucose oxidase catalyzes the reaction of glucose and oxygen
to form gluconic acid and hydrogen peroxide). Hydrogels are used as enzyme im-
mobilization matrices in these type of biosensors. Enzyme entrapped in a hydro-
gel of poly(HEMA), polyacrylamide, and poly(vinyl alcohol) have been reported
(113).

Surface Coatings Applications of Hydrogels.

The surface treatment

of polymeric materials is one of the most active research areas in the ﬁeld of
biomedical materials. It is generally agreed that the surface of any material is
what dictates its cellular response in the body. Attempts to design biocompatible
surfaces have been explored with signiﬁcant success for many years. This ap-
proach, ie, minimization of chemical and physical interactions between the sub-
strate polymer and blood, is the most promising avenue for short-term clinical
success. It has been shown that materials functionalized with surfaces consist-
ing of groups such as carboxylate, sulfate, or sulfonate groups act as antithrom-
bogenic agents in that they repel plasma proteins and platelets. The grafting
of poly(ethylene oxide) functionalized surfaces onto a variety of material sub-
strate results in an increase in hydrophilicity and provides a reduction of comple-
ment activation and platelet adhesion (114). In a similar example, the grafting
of poly(HEMA) onto a variety of surfaces has also provided an improvement in
antithrombogenic properties. Novel techniques of surface modiﬁcation, based on
molecular imprinting, have led to hydrogel systems with the ability to recognize
biological and pharmaceutical compounds. Gel surfaces can be molecularly de-
signed for speciﬁc applications (115,116). For example, polyacrylamide gels mod-
iﬁed with an adhesion receptor, asioaloglycoprotein, were prepared to study cell
adhesion. This work, as well as other similar research efforts, has helped gain
an understanding of the required surface functionality for improved receptor re-
sponse (117,118). Also, in the area of contact lens research, the formation of a
hydrogel surface on nonwetting substrates, such as silicone, provides an avenue
for a wettable, biocompatible material possessing excellent comfort characteris-
tics. These have all been made possible utilizing techniques such as plasma ox-
idation, the surface-initiated polymerization of hydrophilic monomers, and the
graft functionalization of high molecular hydrophilic polymers through covalent
coupling.

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HYDROGELS

715

Plasma Surface Modiﬁcation.

One of the earliest approaches to design

wettable, hydrogel-like surfaces was by glow discharge plasma and/or corona dis-
charge techniques. Corona treatment of polymer surfaces consists of the reaction
of oxygen with the polymer surface under an electric discharge to create polar
functionality such as carboxyl, ether, carbonyl, or hydroxyl groups. The forma-
tion of polar groups raises the surface free energy and allows for surface wetting.
Low pressure glow discharge oxygen plasma has also been used successfully on
silicone-based materials. This results in a wettable surface that consists primar-
ily of hydrated cross-linked silicate/silanol functionality. In fact, the Silsight

silicone elastomer lens, commercialized primarily for pediatric aphakic (loss of
natural crystalline lens) patients, is made wettable through oxygen plasma tech-
niques.

Since these early efforts to generate wettable surfaces through oxidative

treatments, a variety of novel plasma processes have been developed (119). These
efforts have focused primarily on plasma polymerization (deposition). This effort
is in an attempt to deposit a continuous thin hydrophilic polymer ﬁlm layer. In
this process, polymerizable gases (monomers) are introduced to the plasma reac-
tor during or after the glow discharge treatment. The plasma conditions initiate
polymerization that results in deposition of a thin polymer surface, where the sur-
face chemistry can be manipulated by choice of monomer. This technique results
in primarily a covalent polymer attachment.

Direct Covalent Coupling.

Similar approaches to covalently attached poly-

mers via glow discharge plasma have also been pursued. In these efforts, the
polymers were solution-adsorbed and surface-grafted under inert gas plasma. A
free-radical mechanism for the grafting has been proposed. This approach has
been used to coat a variety of substrates with hydrogels. Upon plasma treatment
radicals are formed on the surface that react with oxygen to form hydroperoxy
compounds. The homolytic decomposition of the hydroperoxides form radicals that
initiate the polymerization (thermal or ultraviolet) of hydrophilic monomers, such
as HEMA (120). The adsorption can best be accomplished by using surface-active
compounds that preferentially immobilize on the surface before introduction of
the plasma.

In addition to the conventional plasma polymerization approaches, a sig-

niﬁcant number of research programs have focused on the design of wettable
surfaces utilizing a combination of glow discharge plasma and polymer graft-
ing through classical chemical reaction techniques (121). One approach demon-
strates the successful covalent attachment of a polysaccharide (21). In this
process, a silicone surface is aminated using glow discharge ammonia plasma.
The aminated surface is then reacted with an oxidized dextran to form a cova-
lently attached (via a Schiff-based linkage) dextran. Following reduction to a
stable secondary amine linkage, a highly wettable surface is formed (122). An-
other approach makes use of covalent grafting utilizing solution techniques only.
In this example, the vinyl acetate groups of a medical grade poly(ethylene-co-
vinyl acetate-co-vinyl chloride) was saponiﬁed using a solution of methanol and
sodium hydroxide and then reacted with hexamethylene diisocyanate to form a
highly reactive isocyanate-modiﬁed surface. The resultant surface is then reacted,
again via solution techniques, with an amine functionalized hydrophilic polymer
(123).

716

HYDROGELS

Vol. 2

Hydrogel Surfaces through In Situ Polymerization Techniques.

Several

research groups have reported on efforts to design wettable silicone-based ma-
terials through in situ polymerization techniques. These are attempts to design
wettable, biocompatible surfaces without the need for surface treatment. The pri-
mary driver for this research is cost. Without the need for a secondary surface
treatment, a large reduction in lens cost would be achieved.

Signiﬁcant progress in this area has been made. It has been shown recently

that the polarity of the casting substrate may provide an avenue for surface wet-
ting without surface treatment. The casting of silicone-based formulations from
polar mold materials results in a surface rich character. For example, a copoly-
mer based on a ﬂuorinated silicone and the hydrophilic monomer, DMA was cast
against the hydrophilic mold resin Barex (copolymer of acrylonitrile and styrene).
X-ray photoelectron spectroscopy analysis of the surface showed a threefold in-
crease in surface nitrogen concentration as compared with the same formulation
cast from a hydrophobic polypropylene resin. The Barex cast material resulted in

Vol. 2

HYDROGELS

717

a hydrogel-like surface as a result of the preferential polymerization–migration
of DMA at the Barex surface.

The use of polymerizable surfactants (surface-active macromer, SAM) in

silicone formulations has been explored to form hydrogel-like surfaces without
a surface treatment (124). Several surface-active water-soluble macromers were
evaluated. The SAMs were prepared using a two-step synthetic procedure. In this
procedure, hydrophobic alkylmethacrylates or ﬂuoromethacrylates monomers are
polymerized with polyoxyethylene methacrylates using a functional mercaptan as
a chain-transfer agent. The resultant hydroxyl or carboxylate capped macromer
(22) is further functionalized with a polymerizable methacrylate.

The SAMs are surface-active, yet possess signiﬁcant hydrophilic character.

They are simply added to a silicone formulation, at concentrations of 0.1–0.5%,
and polymerized. It has been shown that the polymerization of SAMs in silicone-
based formulations results in a signiﬁcant increase in surface polarity. This is
due to migration of the SAM to the mold–lens interface before polymerization.
Clinical performance has also been improved. The incorporation of SAMs has had
a positive effect on surface wetting and deposition characteristics.

Nonbiomedical Applications

This article has detailed a variety of biomedical applications of hydrogels.
It is important to note, however, that many applications of hydrogels for

718

HYDROGELS

Vol. 2

nonbiomedical use have been successfully designed. These include the use of hy-
drogels for chemical valves, bioseparation devices, biomimetic actuators, ther-
moresponsive surfaces, afﬁnity precipitation, anodes for bridge-building applica-
tion, water retention for soil application, and disposable diapers the details of
which can be found in other articles (10,31,125–128).

BIBLIOGRAPHY

“Contact Lenses” in EPST 1st ed., Suppl. Vol. 1, pp. 158–219, by M. F. Refojo, Eye Research
Institute of Retina Foundation; “Contact Lenses” in EPSE 2nd ed., Vol. 4, pp. 164–173, by
H. K. Singer and E. C. Bellantoni, American Optical Corp., and A. R. Leboeuf, Barnes-Hind
Pharmaceuticals, Inc.

1. K. Park, W. S. W. Shalaby, and H. Park, Biodegradable Hydrogels For Drug Delivery,

Technomic Publishing Co, Inc., Lancaster, Pa., 1993, pp. 2–3.

2. J. M. Guenet, Thermoreversible Gelation of Polymers and Biopolymers, Academic

Press, New York, 1992.

3. N. Peppas and Y. Yang, Contact Lens 7, 300 (1981).
4. P. J. Flory, J. Am. Chem. Soc. 78, 5222 (1956).
5. P. J. Flory, Principles of Polymer Chemistry, Cornell University Press, Ithaca, N.Y.,

1953, Chapt. “XII”, pp. 495–540.

6. J. M. Rosiak and F. Yoshii, Nucl. Instrum. Methods Phys. Res. 151, 56–64

(1999).

7. L. J. C. Peschier and co-workers, Biomaterials 14, 945–952 (1993).
8. N. A. Peppas and S. R. Lustig, in N. Peppas, ed., Hydrogels in Medicine and Pharmacy,

Vol. 1, CRC Press, Boca Raton, Fla., 1987, pp. 57–83.

9. V. Kudella, in H. Mark, N. Bikales, C. Overberger and G. Menges, eds., Encyclope-

dia of Polymer Science and Engineering, Vol. 7, Wiley-Interscience, New York, 1987,
p. 787.

10. S. D. Bruck, J. Biomed. Mater. Res. 7, 387 (1973).
11. U.S. Pat. 5,739,236 (1998), R. J. Bowers, S. Jones, and P. Stratford (to Biocompatibles

Ltd.).

12. U.S. Pat. 4,436,887 (1984), R. C. Chromecek and co-workers (to Bausch and

Lomb).

13. U.S. Pat. 5,270,418 (1993), G. D. Friends and J. F. K ¨

unzler (to Bausch and

Lomb).

14. J. K ¨

unzler, Trends Polym. Sci. 4, 52 (1995).

15. U.S. Pat. 4,649,184 (1987), T. Yoshikawa (to Toyo Contact Lens Co., Ltd.).
16. U.S. Pat. 5,321,108 (1994), J. K ¨

unzler and R. Ozark (to Bausch and Lomb).

17. J. K ¨

unzler and R. Ozark, J. Appl. Polym. Sci. 65, 1081 (1997).

18. U.S. Pat. 4,139,513 (1979), K. Tanaka and co-workers.
19. U.S. Pat. 4,260,725 (1981), P. L. Keogh, J. K ¨

unzler, and C. C. Niu.

20. L. Junqueira, J. Carneiro, and R. Kelley, Basic Histology, Appleton and Lange, Nor-

walk, conn., 1992.

21. W. S. W. Shalaby and K. Park, in J. C. Salamone, ed., Polymeric Materials Encyclope-

dia, CRC Press, Boca Raton, Fla., 1996, p. 2125.

22. A. Hoffman, in R. Kronenthal, Z. Oser and E. Martin, eds., Polymer in Medicine and

Surgery, Plenum, New York, 1975.

23. S. D. Brook, Properties of Biomaterials in Physiological Environments, CRC Press,

Boca Raton, Fla., 1980.

Vol. 2

HYDROGELS

719

24. J. Andrade, ed., Hydrogels for Medical and Related Applications, American Chemical

Society, Washington, D.C., 1976.

25. C. L. McCormick and J. Bock, in H. Mark, N. Bikales, C. Overberger, and G. Menges,

eds., Encyclopedia of Polymer Science and Engineering, Vol. 17, Wiley-Interscience,
New York, 1989, pp. 737–784.

26. A. L. Lehninger, Principles of Biochemistry, Worth Publishers, New York, 1982, pp.

127–137.

27. T. Takagi, K. Takahashi, M. Aizawa, and S. Miyata, in Proceedings of the First Interna-

tional Conference on Intelligent Materials, Technomic Publishing Co., Inc., Lancaster,
Pa., 1993.

28. H. G. Schield, Prog. Polym. Sci. 17, 163–249 (1992).
29. A. S. Hoffman, Artiﬁcial Organs 19, 458–467 (1995).
30. I. Y. Galaev and B. Mattiasson TIBTECH 17, 335–340 (1999).
31. L. G. Grifﬁth, Acta Mater. 48, 263–277 (2000).
32. R. K. Kulkarni and co-workers, J. Biomed. Mater. Res. 5, 169 (1971).
33. H. Park and K. Park, in R. Ottenbrite, S. Huang, and K. Park, eds., Hydrogels

and Biodegradable Polymers for Bioapplications, ACS Symposium Series, Vol. 627,
American Chemical Society, Washington, D.C., 1996, pp. 1–10.

34. Ref. 8, pp. 11–12.
35. U.S. Pat. 5,310,779 (1994), Y. C. Lai (to Bausch and Lomb).
36. D. E. Gregonis, C. Chen, and J. Andrade, in Ref. 24, ACS Symposium Series, Vol. 31,

p. 88.

37. Y. Tsukahara and co-workers, J. Polym. Sci. Part A: Polym. Chem. 27, 1099–1114

(1989).

38. G. Odian, Principles of Polymerization, 3rd ed., John Wiley and Sons, Inc., New York,

1991.

39. Ref. 1, pp. 35–66.
40. H. Akin and N. Hasirci, Proc. Am. Chem. Soc. Polym. Div. Polym. Prepr. 36, 384–385

(1995).

41. K. Yamada and co-workers, J. Neurosurg 86, 871, (1997).
42. A. Rubinstein, D. Nakar, and A. Sintov, Pharm. Res. 9, 276–278 (1992).
43. Y. M. Lizuka and co-workers, J. Appl. Polym. Sci. 56, 73–77 (1995).
44. N. Peppas, ed., Hydrogels in Medicine and Pharmacy, Vol. 2, CRC Press, Boca Raton,

Fla., 1987, pp. 1–48.

45. J. J. Sperinde and L. G. Grifﬁth, Macromolecules 30, 5255–5264 (1997).
46. A. S. Hoffman, Radiat. Phys. Chem. 9, 207 (1977).
47. J. M. Rosiak and F. Yoshii, Nucl. Instrum. Methods Phys. Res. 151, 56–64

(1999).

48. E. Tsuchida and K. Abe, in H. J. Cantow, ed., Advances in Polymer Science, Vol. 45,

Springer-Verlag, New York, 1982.

49. A. Rembaum, Appl. Polym. Symp. 22, 299 (1973).
50. Y. Hagino and S. J. Huang, Proc. ACS Div. Polym. Mater. Sci. Eng. 72, 249–250

(1995).

51. T. L. Bowersock and co-workers, Proc. Am. Chem. Soc., Div. Polym. Mater. 35, 405–406

(1994).

52. H. Ohno, K. Abe, and E. Tsuchida, Makromol. Chem. 179, 755 (1978).
53. F. E. Bailey and H. G. France, J. Polym. Sci. 49, 397 (1961).
54. U.S. Pat. 4,085,168 (1978), R. Milkovich and M. T. Chiang (to CPC International Inc.).
55. U.S. Pat. 4,095,877 (1978), O. Wichterle, V. Stoy, and A. Stoy (to Ceskoslovenska

akademic ved, (University in Prague).).

56. T. Nakashima and K. Takakura, J. Biomed. Mater. Res. 11, 787 (1977).
57. T. Okano, M. Kitayama, and I. Shinohara, J. Appl. Polym. Sci. 22, 369 (1978).

720

HYDROGELS

Vol. 2

58. T. Okano and co-workers, J. Biomed. Mater. Res. 15, 393 (1981).
59. K. R. Shah, in B. M. Culbertson, ed., Contemporary Topics in Polymer Science, Vol. 6,

Plenum, New York, 1989, p. 486.

60. K. R. Shah, in Ref. 21, pp. 3092–3097.
61. J. K ¨

unzler and J. McGee, Chem. Ind. 16, 651, 1995.

62. U.S. Pat. 3,220,960 (1965), O. Wichterle and D. Lim (to Ceskoslovenska akademic ved,

(University in Prague)).

63. S. Hosaka, J. Biomed. Mater. Res. 14, 557 (1980).
64. O. Wichterle and R. Chromecek, J. Polym. Sci. 16, 4677 (1969).
65. U.S. Pat. 3,957,362 (1976), W. Mancini, D. Korb, and M. Refojo (to Corneal Sciences,

Inc.).

66. U.S. Pat. 4,625,009 (1986), T. Izumitani and co-workers.
67. B. Holden, G. Mertz, and J. McNally, Invest. Ophthalmol. Vis. Sci. 24, 218

(1983).

68. E. J. Castillo and co-workers, Biomaterials 7, 9–16 (1986).
69. L. Minarik and J. Rapp, Contact Lenses (CLOA Journal) 15, 185–188 (1989);

Orsborn and S. Zantos, Contact Lenses (CLOA Journal) 14, 81–85 (1988).

70. J. Brandrup and J. Immergut, eds., Polymer Handbook, 3rd ed., Vol. VI, Wiley-

Interscience, New York, 1989, p. 435.

71. Eur. Pat. 396,364 (1990), R. Bambury and D. Seelye (to Bausch and Lomb).
72. G. L. Grobe and co-workers, Polym. Prepr. (Am. Chem. Soc., Div. Mater. Sci.) 80, 108

(1999).

73. U.S. Pat. 6,030,416 (2000), P. Huo, S. Zhou, and C. Liau (to Pharmacia & Upjohn AB).
74. U.S. Pat. 5,331,073 (1994), J. Weinschank and R. Christ.
75. R. Larsson and co-workers, Biomaterials 10, 511–516 (1989).
76. M. F. Refojo, in P. Christel, A. Meunier, and A. J. C. Lee, eds., Biological and Biome-

chanical Performance of Biomaterials, Elsevier Science, Amsterdam, 1986.

77. T. Kohnen, R. J. Lambert, and D. D. Koch, Ophthalmology 104, 1277–1286

(1997).

78. D. Schonfeld, Ocular Surgery News, 10–11 (2001).
79. S. J. Holland, B. J. Tighe, and P. L. Gould, J. Controlled Release 4, 155–180

(1986).

80. N. Peppas, in Ref. 8, pp. 102–113.
81. D. Shino and co-workers, in Proceedings of the First International Conference on Intel-

ligent Materials, Technomic Publishing Co., Inc., Lancaster, Pa., 1993, pp. 301–304.

82. S. Kitano and co-workers, J. Controlled Release 19, 162–170 (1992).
83. R. W. Baker, M. E. Tuttle, and R. Helwing, Pharm. Technol. 8(2), 26–30

(1984).

84. P. Hocking, J.M.S.-Rev. Macromol. Chem. Phys., Part C 32, 35–54 (1992).
85. J. A. Hubbell, MRS Bull. 21, 33 (1996).
86. J. Heller and co-workers, Biomaterials 4, 262 (1983).
87. W. Mark Saltzman, Drug Delivery, Oxford University Press, Oxford 2001, pp. 259–

261.

88. D. S. T. Hsieh, ed., Controlled Release Systems: Fabrication Technology, Vol. 2, CRC

Press, Boca Raton, Fla., 1988, pp. 61–73.

89. Ref. 88, pp. 74–75.
90. H. A. J. Struyker-Boudier, ed., Rate-Controlled Drug Administration and Action, CRC

Press, Boca Raton, Fla., 1985, pp. 24–27.

91. S. K. Mallapragada and N. A. Peppas, J. Controlled Release 45, 87–94 (1997).
92. R. Jones, Curr. Opin. Cardiovasc., Pulm. Renal Invest. Drugs 2, 284–291

(2000).

Vol. 2

HYDROGELS

721

93. S. J. Holland, B. J. Tighe, and P. L. Gould, J. Controlled Release 4, 155–180

(1986).

94. A. S. Sawhney, C. P. Pathak, and J. A. Hubbell, Macromolecules 26, 581–587

(1993).

95. J. Kopecek and P. Kopeckova, Chem. Listy 91, 600–603 (1997).
96. T. Pouyani and G. D. Prestwich, Bioconjugate Chem. 5, 339–347 (1994).
97. N. A. Peppas, Curr. Opinm. Colloid Interface Sci. 2, 531–537 (1997).
98. N. A. Peppas, in R. Gurny, H. E. Junginger, and N. A. Peppas, eds., Pulsatile Drug

Delivery, Wissenschaftlishe Verlagsgesellschaft, Stuttgart, 1993, pp. 41–56.

99. S. W. Kim, in N. Ogata, S. W. Kim, J. Feijen, and T. Okano, eds., Advanced Biomate-

rials in Biomedical Engineering and Drug Delivery Systems, Springer, Tokyo, 1996,
pp. 125–133.

100. A. Gutowska and S. W. Kim, Proc. Am. Chem. Soc. Div. Polym. Chem. 212, 115

(1996).

101. S. Kitano and co-workers, in Proceedings of the First International Conference on

Intelligent Materials, Technomic Publishing Co., Inc., Lancaster, Pa., 1993, pp. 383–
388.

102. T. Aoki and co-workers, Polym. J. 28, 371–374 (1996).
103. J. Kost, Pulsed and Self-Regulated Drug Delivery, CRC Press, Boca Raton, Fla.,

1990.

104. M. Goldraich and J. Kost, Clin. Mater. 13, 135–142 (1993).
105. T. Aoki and co-workers, Polym. J. 28, 371–374 (1996).
106. S. K. Ashiku, M. A. Randolph, and C. A. Vacanti, Mater. Sci. Forum 250, 129–150

(1997).

107. Y. Tabata, M. Yamamoto, and Y. Ikada, Pure Appl. Chem. 70, 1277–1282

(1998).

108. Synthetic Wound Dressings Market, Theta Report No. 610, New York, 1996.
109. K. Bruce Jacobson Oak Ridge Natl. Lab Rev. 29(3), 1–18, 1996.
110. A. P. F. Turner, I. Karube and G. S. Wilson, Biosensors: Fundamentals and Applica-

tions, Oxford University Press, Oxford, 1987.

111. H. T. Tien and co-workers, Anal. Sci. 14, 3–18, 1998.
112. W. J. Wizeman and P. Koﬁnas, Biomaterials 22, 1485–1491, 2001.
113. M. Koudelke-Hep, P. van der Wal, D. J. Strike, and N. F. de Rooij, in Ref. 21,

pp. 3368–3371.

114. D. Klee and H. Hoecker, in Advances in Polymer Science, Vol. 149, Springer, Verlag,

Berlin, 2000, pp. 1–57.

115. W. J. Wizeman and P. Koﬁnas, Biomaterials 22, 1485–1491 (2001).
116. P. Bures and co-workers, J. Controlled Release 72, 25–33 (2001).
117. P. H. Weigel and co-workers, J. Biol. Chem. 253, 330 (1978).
118. L. G. Grifﬁth and S. T. Lopina, Biomaterials 19, 979 (1998).
119. E. M. Liston, L. Martinu, and M. R. Wertheimer, in M. Strobel, C. S. Lyons, and K. L.

Mittal, eds., Plasma Surface Modiﬁcation of Polymers, VSP, Utrecht, the Netherlands,
1994, pp. 3–39.

120. P. W. Rose and E. M. Liston, Plast. Eng. 41, 41 (1985).
121. M. S. Sheu, A. S. Hoffman, B. D. Ratner, J. Feijen, and J. M. Harris, in Ref. 119,

pp. 197–208.

122. L. Dai and co-workers, in B. D. Ratner and D. G. Castner, eds., Surface Modiﬁcation

of Polymeric Materials, Plenum Press, New York, 1996, pp. 147–156.

123. J. Wirshing, D. Klee, H. Hoecker, A. Dekker, and C. Mittermayer, in Proc. 12th Euro-

pean Conference on Biomaterials, Porto, 1995, p. 104.

124. P. V. Valint, D. Ammon, G. L. Grobe, and J. McGee, in Ref. 122, pp. 21–26.

722

HYDROGELS

Vol. 2

125. M. Gupta and B. Mattiasson, in G. Street, ed., Highly Selective Separations in Biotech-

nology, Blackie Academic & Professional, London, 1994, pp. 7–33.

126. A. Hoffman, J. Controlled Release 6, 297–305 (1987).
127. W. Heitz and co-workers, Angew. Makromol. Chem. 123/124, 147 (1984).
128. G. Burillo and T. Ogawa, Radiat. Phys. Chem. 18, 1143 (1981).

RIEDRICH

K ¨

UNZLER

Bausch and Lomb Inc.

β-HYDROXYALKANOATES.

See P

OLY

3-(

HYDROXYALKANOATES

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