Note: Descriptions are shown in the official language in which they were submitted.
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BIO-SYNTHETIC MATRIX AND USES THEREOF
FIELD OF THE INVENTION
The present invention pertains to the field of tissue engineering and in
particular to a
bio-synthetic matrix comprising a hydrogel suitable for ifa viva use.
BACKGROUND
Tissue engineering is a rapidly growing field encompassing a number of
technologies
aimed at replacing or restoring tissue and organ function. The key objective
in tissue
engineering is the regeneration of a defective tissue through the use of
materials that
can integrate into the existing-tissue so as to restore normal tissue
function. Tissue
engineering, therefore, demands materials that can support cell over-growth,
in-
growth or encapsulation and, in many cases, nerve regeneration.
Polymer compositions are fording widespread application in tissue engineering.
Natural bio-polymers such as collagens, fibrin, alginates and agarose are
known to be
non-cytotoxic and to support over-growth, in-growth and encapsulation of
living cells.
1 S Matrices derived from natural polymers, however, are generally
insufficiently robust
for transplantation. W contrast, matrices prepared from synthetic polymers can
be
formulated to exhibit predetermined physical characteristics such as gel
strength, as
well as biological characteristics such as degradability. Reports that
synthetic
analogues of natural polymers, such as polylysine, polyethylene imine), and
the like,
can exhibit cytotoxic effects [Lynn & Langer, J. Arraef~. Chena. Soc.,
122:10761-10768
(2000)] have lead to the development of alternative synthetic polymers for
tissue
engineering applications.
Hydrogels are crosslinked, water-insoluble, water-containing polymers which
offer
good biocompatibility and have a decreased tendency to induce thrombosis,
encrustation, and inflammation and as such are ideal candidates for tissue
engineering
purposes. The use of hydrogels in cell biology is well known [see, for
example, A.
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Atala and R.P. Lanza, eds., "Methods in Tissue Engineering" Academic Press,
San
Diego, 2002]. A wide variety of hydrogels for in vivo applications have been
described [see, for example, the review by Jeong, et al., Adv. Drug Deliv.
Rev., 54:37-
51 (2002)]. Hydrogels based on N-isopropylacrylamide (NiPAAm) and certain co-
y polymers thereof, for example, are non-toxic and capable of supporting
growth of
encapsulated cells in vitro [Vernon, et al., Macromol. Synap., 109:155-167
(1996);
Stile, et al., Macromolecules, 32:7370-9 (1999); Stile, et
al.,Biomacf°omolecules 3:
591- 600. ( 2002); Stile, et al., Biomaeromolecules 2: 185 -194. ( 2001);
Webb, et
al., MUSC Orthopaedic J., 3:18-21 (2000); An et al., U.S. Patent No.
6,103,528].
Temperature-sensitive NiPAAm polymers have also been described for use in
immunoassays [U.S. Patent No. 4,780,409]. However, despite manipulations of
synthesis conditions and improvements to enhance biocompatibility, it is still
difficult
to obtain a seamless host-implant interface and complete integration of the
hydrogel
implant into the host [Hicks, et al. Surv. Ophthalmol. 42: 17f-189 (1997);
Trinkaus-
Randall and Nugent, J. Controlled Release 53:205-214 (1998)].
Modifications of synthetic polymer gels with a second naturally derived
polymer to
generate an interpenetrating polymer network ("IPN") structure have been
reported
[For example, see Gutowska et al., Macromolecules, 27:4167 (1994); Yoshida et
al.,
Nature, 374:240 (1995); Wu & Jiang, U.S. Patent No. 6,030,634; Park et al.,
U.S.
Patent No. 6,271,278]. However, these structures are frequently destabilised
by
extraction of the naturally derived component by culture media and by
physiological
fluids. Naturally derived polymers also tend to biodegrade rapidly within the
body
resulting in destabilisation of in vivo implants.
More robust hydrogels comprising cross-linked polymer compositions have also
been
described. For example, U.S. Patent No. 6,388,047 describes a composition
consisting
of a hydrophobic macromer and a hydrophilic polymer that are cross-linked to
form a
hydrogel by free-radical polymerisation. U.S. Patent No. 6,323,278 describes a
cross-
linked polymer composition which can form in situ and which comprises two
synthetic polymers, containing multiple electrophilic groups and the other
containing
multiple nucleophilic groups. Both U.S. Patent No. 6,388,047 and 6,384,105
describe
systems that must be cross-linked by free radical chemistry, which requires
the use of
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initiators that are well known to be cytotoxic (azo compounds, persulfates),
thus
leading to possible side effects if the hydrogel was to be used in the tissue
or with
encapsulated cells.
U.S. Patent No. 6,384,105 describes injectable, biodegradable polymer
composites
comprising polypropylene fumarate) and polyethylene glycol)-dimethacrylate
which
can be cross-linked ih situ. The hydrogels described in this patent are
largely based on
polymers with a polyethylene oxide backbone polymers. Although these polymers
are
known to be biocompatible, their ability to support cell growth is uncertain.
U.S. Patent No. 6,566,406 describes biocompatible cross-linked hydrogels that
are
formed from water soluble precursors having electrophilic and nucleophilic
groups
capable of reacting and cross-linking i~ situ. The precursors are described as
being a
polyalkylene oxide polymer and a cross-linker. As indicated above, the ability
of
polyalkylene oxide backbone polymers to support cell growth is uncertain.
There remains a need therefore, for an improved matrix that is biocompatible,
sufficiently robust to function as an implant and that can support cell growth
in vivo.
This background information is provided for the purpose of making known
information believed by the applicant to be of possible relevance to the
present
invention. No admission is necessarily intended, nor should be construed, that
any of
the preceding information constitutes prior art against the present invention.
SUMMARY OF THE INVENTION
An obj ect of the present invention is to provide a bio-synthetic matrix and
uses
thereof. In accordance with an aspect of the present invention, there is
provided a
synthetic co-polymer suitable for the preparation of a bio-synthetic matrix,
comprising
one or more N-alkyl or N,N-dialkyl substituted acrylamide co-monomer, one or
more
hydrophilic co-monomer and one or more acryl- or methacryl- carboxylic acid co-
monomer derivatised to contain a pendant reactive moiety capable of cross-
linking
bioactive molecules, said synthetic polymer having a number average molecular
mass
between about 2,000 and about 1,000,000.
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In accordance with another aspect of the invention, there is provided a bio-
synthetic
matrix comprising the synthetic co-polymer, a bio-polymer and an aqueous
solvent,
wherein the synthetic co-polymer and bio-polymer are cross-linked to form a
hydrogel.
In accordance with another aspect of the invention, there are provided uses of
the bio-
synthetic matrix as a scaffold for tissue regeneration, for replacement of
damaged or
removed tissue in an animal, or for coating surgical implants.
In accordance with another aspect of the invention, there are provided
compositions
comprising: one or more bioactive agent or a plurality of cells; a synthetic
co-polymer
of the invention; a bio-polymer; and an aqueous solvent.
In accordance with another aspect of the invention, there is provided an
implant for
use in tissue engineering comprising a pre-formed bio-synthetic matrix, said
matrix
comprising an aqueous solvent and a bio-polymer cross-linked with a synthetic
co-
polymer of the invention.
In accordance with another aspect of the invention, there is provided a use of
the
implant as an artificial cornea.
In accordance with another aspect of the invention, there is provided a
process for
preparing a synthetic co-polymer comprising: (a) dispersing one or more N-
alkyl or
N,N-dialkyl substituted acrylamide co-monomer, one or more hydrophilic co-
monomer and one or more acryl- or methacryl- carboxylic acid co-monomer
derivatised to contain a pendant cross-linkable moiety in a solvent in the
presence of
an initiator; (b) allowing the N-alkyl or N,N-dialkyl substituted acrylamide
co-
monomer, hydrophilic co-monomer and acryl- or methacryl- carboxylic acid co-
monomer to polymerise to form a synthetic co-polymer, and (c) optionally
purifying
the synthetic co-polymer; and a process for preparing a bio-synthetic matrix
comprising preparing a synthetic co-polymer, dispersing the synthetic co-
polymer and
a bio-polymer in an aqueous medium and allowing the synthetic co-polymer and
the
bio-polymer to cross-link to provide the bio-synthetic matrix.
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BRIEF DESCRIPTION OF THE FIGI1RES
Figure 1 depicts the general structure of a terpolymer according to one
embodiment
of the invention comprising N-isopropylacrylamide, (NiPAAm), acrylic acid
(AAc)
and N-acryloxysuccinimide (ASS.
Figure 2 presents the clinical results from the transplantation into pigs of
artificial
corneas prepared from a bio-sylthetic matrix according to one embodiment of
the
invention.
Figure 3 presents the results of in vivo confocal microscopy at 6 weeks post-
operative
of artificial corneas prepared from a bio-synthetic matrix according to one
embodiment of the invention and transplanted into pigs.
Figuf~e 4 depicts iya vivo testing for corneal sensitivity of artificial
corneas prepared
from a bio-synthetic matrix according to one embodiment of the invention and
transplanted into pigs.
Figures 5, 6 and 7 present the results of morphological and biochemical
assessment
of artificial corneas prepared from a bio-synthetic matrix according to one
embodiment of the invention and transplanted into pigs.
Figure 8 shows (A) the structure of a terpolymer containing a cross-linked
bioactive
according to one embodiment of the invention, (B) a corneal scaffold composed
of
cross-linked collagen and the terpolymer shown in (A) and (C) shows a corneal
scaffold composed of thermogelled collagen only.
Figuf~es 9 and 10 depict the results of delivery of a hydrogel containing
collagen and a
terpolymer-bioactive agent according to one embodiment of the invention into
mouse
and rat brains.
Figure 11 shows modulus (A) and stress at failure (B) from suture pull out
measurements as a function of the concentration ratios of N-acryloxysuccinimde
to
collagen amine groups for hydrogel matrices according to one embodiment of the
invention.
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Figure 12 depicts transmission (A) and back scattering (B) of light across the
visible
region as a function of the concentration ratios of N-acryloxysuccinimde to
collagen
amine groups for hydrogel matrices according to one embodiment of the
invention.
Figure 13 depicts transmission (A) and back scattering (B) of light across the
visible
region as a function of the concentration ratios of N-acryloxysuccinimde to
collagen
amine groups for a hydrogel matrix according to another embodiment of the
invention.
Figure 14 depicts restoration of touch sensitivity for a pig corneal implant
comprising
a hydrogel according to one embodiment of the invention.
Figure I S depicts corneal implantation procedure by lainellar keratoplasty in
pigs and
clinical in vivo confocal microscopic images of 6-week implants comprising a
hydrogel according to one embodiment of the invention. Bar = 25 pm for D-F, 15
~,m
for G-O.
Figure 16 depicts post-surgical corneal regeneration in pigs receiving corneal
implants comprising a hydrogel according to one embodiment of the invention.
Bar =
100 ~.m for A-F, 40 ~m for G-I, 200 nm for J-L, 20 ~.m for M-O, 30 pm for P-R
Figure 17 depicts implant-host integration post-surgery at 6 weeks post
surgery in
pigs receiving corneal implants comprising a hydrogel according to one
embodiment
of the invention. Bar =100 ~m in all cases.
Figure 18 depicts corneal touch sensitivity in implants in pigs receiving
corneal
implants comprising a hydrogel according to one embodiment of the invention.
Figure 19 depicts the results of innervation compatibility tests on various
hydrogel
matrices.
Figure 20 depicts epithelial cell growth and stratification on various
hydrogels. (A)
low magnification views of epithelial growth on the hydrogels (inset is higher
magnification) and (B) counts of the cell thickness of the epithelium grown
over the
hydrogels.
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DETAILED DESCRIPTION OF THE INVENTION
It should be understood that this invention is not limited to the particular
process steps
and materials disclosed herein, but is extended to equivalents thereof as
would be
recognised by those ordinarily skilled in the relevant arts. It should also be
understood
that terminology employed herein is for the purpose of describing particular
embodiments only and is not intended to be limiting.
DEFINITIONS
Unless defined otherwise, all technical and scientific terms used herein have
the same
meaning as commonly understood by one of ordinary skill in the art to which
this
invention pertains.
The term "hydrogel," as used herein, refers to a cross-linked polymeric
material which
exhibits the ability to swell in water or aqueous solution without dissolution
and to
retain a significant portion of water or aqueous solution within its
structure.
lThe term "polymer," as used herein, refers to a molecule consisting of
individual
monomers joined together. In the context of the present invention, a polymer
may
comprise monomers that are joined "end-to-end" to form a linear molecule, or
may
comprise monomers that are joined together to form a branched structure.
The term "monomer," as used herein, refers to a simple organic molecule which
is
capable of forming a long chain either alone or in combination with other
similar
organic molecules to yield a polymer.
The term "co-polymer," as used herein, refers to a polymer that comprises two
or
more different monomers. Co-polymers can be regular, random, block or grafted.
A
regular co-polymer refers to a co-polymer in which the monomers repeat in a
regular
pattern (e.g. for monomers A and B, a regular co-polymer may have a sequence:
ABABABAB). A random co-polymer is a co-polymer in which the different
monomers are arranged randomly or statistically in each individual polymer
molecule
(e.g. for monomers A and B, a random co-polymer may have a sequence:
AABABBABBBAAB). In contrast, a block co-polymer is a co-polymer in which the
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different monomers are separated into discrete regions within each individual
polymer
molecule (e.g. for monomers A and B, a block co-polymer may have a sequence:
AAABBBAAABBB). A grafted co-polymer refers to a co-polymer which is made by
linking a polymer or polymers of one type to another polymer molecule of a
different
composition.
The term "terpolymer," as used herein, refers to a co-polymer comprising three
different. monomers.
The term "bio-polymer," as used herein, refers to a naturally occurring
polymer.
Naturally occurnng polymers include, but are not limited to, proteins and
carbohydrates. The term "bio-polymer" also includes derivatised forms of the
naturally occurring polymers that have been modified to facilitate cross-
linking to a
synthetic polymer of the invention.
The term "synthetic polymer," as used herein, refers to a polymer that is not
naturally
occurring and that is produced by chemical or recombinant synthesis.
The terms "alkyl" and "lower alkyl" are used interchangeably herein to refer
to a
straight chain or branched alkyl group of one to eight carbon atoms or a
cycloalkyl
group of three to eight carbon atoms. These terms are further exemplified by
such
groups as methyl, ethyl, r~-propyl, i-propyl, n-butyl, t-butyl, l-butyl (or 2-
methylpropyl), i-amyl, yz-amyl, hexyl, cyclopropyl, cyclobutyl, cyclopentyl,
cyclohexyl
and the like.
The term "bioactive agent," as used herein, refers to a molecule or compound
which
exerts a physiological, therapeutic or diagnostic effect in vivo. Bioactive
agents may
be organic or inorganc. Representative examples include proteins, peptides,
carbohydrates, nucleic acids and fragments thereof, anti-tumour and anti-
neoplastic
compounds, anti-viral compounds, anti-inflammatory compounds, antibiotic
compounds such as antifungal and antibacterial compounds, cholesterol lowering
drugs, analgesics, contrast agents for medical diagnostic imaging, enzymes,
cytokines,
local anaesthetics, hormones, anti-angiogenic agents, neurotransmitters,
therapeutic
oligonucleotides, viral particles, vectors, growth factors, retinoids, cell
adhesion
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factors, extracellular matrix glycoproteins (such as laminin), hormones,
osteogemc
factors, antibodies and antigens.
The term "biocompatible," as used herein, refers to an ability to be
incorporated into a
biological system, such as into an organ or tissue of an animal, without
stimulating an
immune and / or inflammatory response, fibrosis or other adverse tissue
response.
As used herein, the term "about" refers to a +/-10% variation from the nominal
value.
It is to be understood that such a variation is always included in any given
value
provided herein, whether or not it is specifically referred to.
1. BIO-SYNTHETIC MATRL~
The present invention provides a bio-synthetic matrix comprising a hydrogel
which is
formed by cross-linking a synthetic polymer and a bio-polymer. The bio-polymer
may
be in its naturally-occurring form, or it may be derivatised to facilitate
cross-linking to
the synthetic polymer. The matrix is robust, biocompatible and non-cytotoxic.
The
matrix can be formed in aqueous solution at neutral pH and can be tailored to
further
comprise one or more bioactive agents such as growth factors, retinoids, cell
adhesion
factors, enzymes, peptides, proteins, nucleotides, drugs, and the like. The
bioactive
agent can be covalently attached to the synthetic polymer, or it may be
encapsulated
and dispersed within the final matrix depending on the end use demands for the
matrix. The matrix may also comprise cells encapsulated and dispersed therein,
which
are capable of proliferation and/or diversification upon deposition of the
matrix in
vivo.
In one embodiment of the present invention, the bio-synthetic matrix supports
cell
growth. Such cell growth may be epithelial and/or endothelial surface coverage
(i. e.
two dimensional, 2D, growth) andlor three-dimensional (3D) cell growth
involving
growth into the matrix itself.
In another embodiment of the invention, the bio-synthetic matrix supports
nerve in-
growth. As is known in the art, nerve growth into transplanted tissue takes
place over
an extended period of time, typically in the order of months or years. Growth
of
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nerves into the matrix can occur more rapidly than growth of nerves into
transplanted
tissue thus leading to more rapid regeneration of functional tissue, for
example, nerve
in-growth may occur within weeks.
The bio-synthetic matrix can be tailored for specific applications. For
example, the
matrix can be used in tissue engineering applications and may be pre-formed
into a
specific shape for this purpose. The matrix can also be used as a drug
delivery vehicle
to provide sustained release of a therapeutic or diagnostic compound at a
particular
site within the body of an animal.
In order to be suitable for i~ vivo implantation for tissue engineering
purposes, the
bio-synthetic matrix must maintain its form at physiological temperatures, be
substantially insoluble in water, be adequately robust, and support the growth
of cells.
It may also be desirable for the matrix to support the growth of nerves.
1.1 Synthetic Polymer
In accordance with the present invention, the synthetic polymer that is
incorporated
into the bio-synthetic matrix is a co-polymer comprising one or more
acrylamide
derivatives, one or more hydrophilic co-monomers and one or more derivatised
carboxylic acid co-monomers which comprise pendant cross-linkable moieties.
As used herein, an "acrylamide derivative" refers to a N-alkyl or N,N-dialkyl
substituted acrylamide or methacrylamide. Examples of acrylamide derivatives
suitable for use in the synthetic polymer of the present invention include,
but are not
limited to, N-methylacrylamide, N-ethylacrylamide, N-isopropylacrylamide
(NiPAAm), N-octylacrylamide, N-cyclohexylacrylamide, N-methyl-N-
ethylacrylamide, N-methylmethacrylamide, N-ethylmethacrylamide, N-
isopropylmethacrylamide, N,N-dimethylacrylamide, N,N-diethylacrylamide, N,N-
dimethylmethacrylamide, N,N-diethylmethacrylamide, N,N-dicyclohexylacrylamide,
N-methyl-N-cyclohexylacrylamide, N-acryloylpyrrolidine, N-vinyl-2-
pyrrollidinone,
N-methacryloylpyrrolidine, and combinations thereof.
A "hydrophilic co-monomer" in the context of the present invention is a
hydrophilic
monomer that is capable of co-polymerisation with the acrylamide derivative
and the
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derivatised carboxylic acid components of the synthetic polymer. The
hydrophilic co-
monomer is selected to provide adequate solubility for polymerisation, aqueous
solubility of the co-polymer and freedom from phase transition of the final co-
polymer
and hydrogel. Examples of suitable hydrophilic co-monomers are hydrophilic
acryl- or
methacryl- compounds such as carboxylic acids including acrylic acid,
methacrylic
acid and derivatives thereof, acrylamide, methacrylamide, hydrophilic
acrylamide
derivatives, hydrophilic methacrylamide derivatives, hydrophilic acrylic acid
esters,
hydrophilic methacrylic acid esters, vinyl ethanol and its derivatives and
ethylene
glycols. The carboxylic acids and derivatives may be, for example, acrylic
acid,
methacrylic acid, 2-hydroxyethyl methacrylate (HEMA), or a combination
thereof.
Examples of hydrophilic acrylamide derivatives include, but are not limited
to, N,N-
dimethylacrylamide, N,N-diethylacrylamide, 2-[N,N-
dimethylamino]ethylacrylamide,
2-[N,N-diethylamino]ethylacrylamide, N,N-diethylmethacrylamide, 2-[N,N-
dimethylamino] ethylmethacrylamide, 2-[N,N-diethylamino]ethylmethacrylamide, N-
vinyl-2-pyrrollidinone, or combinations thereof. Examples of hydrophilic
acrylic
esters include, but are not limited to, 2-[N,N-diethylamino] ethylacrylate, 2-
[N,N-
dimethylamino]ethylacrylate, 2-[N,N-diethylamino]ethylmethacrylate, 2-[N,N-
dimethylamino]ethylmethacrylate, or combinations thereof.
As used herein, a "derivatised carboxylic acid co-monomer" refers to a
hydrophilic
acryl- or methacryl- carboxylic acid, for example, acrylic acid, methacrylic
acid, or a
substituted version thereof, which has been chemically derivatised to contain
one or
more cross-linking moieties, such as succinimidyl groups, imidazoles,
benzotriazoles
and p-nitrophenols. The term "succinimidyl group" is intended to encompass
variations of the generic succinimidyl group, such as sulphosuccinimidyl
groups.
Other similar structures such as 2-(N-morpholino)ethanesulphonic acid will
also be
apparent to those skilled in the art. In the context of the present invention
the group
selected as a cross-linking moiety acts to increase the reactivity of the
carboxylic acid
group to which it is attached towards primary amines (i.e. NHS groups) and
thiols
(i.e. -SH groups). Examples of suitable groups for derivatisation of the
carboxylic
acid co-monomers for use in the synthetic polymer include, but are not limited
to, N-
succinimide, N-succinimide-3-sulphonic acid, N-benzotriazole, N-imidazole andp-
nitrophenol.
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In one embodiment of the present invention, the synthetic polymer comprises:
(a) one or more acrylamide derivative of general formula I:
1
C C
O C R3
N
R ~ R5
wherein:
Rl, R2, R3, R4 and RS are independently selected from the group of: H and
lower alkyl;
(b) one or more hydrophilic co-monomer (which may be the same or different to
(a))
having the general formula II:
16 I7
Y C C II
R9 R8
wherein:
Y is O or is absent;
R6, and R~ are independently selected from the group of H and lower alkyl;
R$ is H, lower alkyl, or -OR', where R' is H or lower alkyl; and
R9 is H, lower alkyl, or -C(O)Rlo, and
Rlo is NR4R5 or-OR", where R" is H or CH2CH20H;
and (c) one or more derivatised carboxylic acid having the general formula
III:
11 ~ 12
C C III
R~3 C O
O
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wherein:
Rm Riz and R13 are independently selected from the group of: H and lower
alkyl and
Q is N-succinimido, 3-sulpho-succinimdo (sodium salt), N-benzotriazolyl, N-
imidazolyl orp-nitrophenyl.
In one embodiment, the synthetic polymer comprises one or more acrylamide
derivative of general formula I, one or more hydrophilic co-monomer of general
formula II and one or more derivatised carboxylic acid of general formula III,
as
described above, wherein the term "lower alkyl" refers to a branched or
straight chain
alkyl group having 1 to 8 carbon atoms.
In another embodiment, the synthetic polymer comprises one or more acrylamide
derivative of general formula I, one or more hydrophilic co-monomer of general
formula II and one or more derivatised carboxylic acid of general formula III,
as
described above, wherein the term "lower alkyl" refers to to a cycloalkyl
group having
3 to 8 carbon atoms, such as cyclopropyl, cyclobutyl, cyclopentyl and
cyclohexyl.
The co-polymer may be linear or branched, regular, random or block. In
accordance
with the present invention, the final synthetic polymer comprises a plurality
of
pendant reactive moieties available for cross-linking, or grafting, of
appropriate
biomolecules.
A worker skilled in the art will appreciate that the group of compounds
encompassed
by the term "acrylamide derivative" and the group of compounds encompassed by
the
term "hydrophilic co-monomer" overlap substantially and that a single monomer
could be selected that fulfils the functions of both these components in the
co-
polymer. Thus, for example, when an acrylamide derivative is selected that is
sufficiently hydrophilic to confer on the synthetic polymer the desired
properties, then
a hydrophilic co-monomer component may be chosen that is identical to the
selected
acrylamide derivative resulting in a co-polymer that comprises two different
monomers only (i.e. the acrylamide derivative/hydrophilic co-monomer and the
derivatised carboxylic acid co-monomer). On the other hand, when enhanced
hydrophilicity beyond that provided by the selected acrylamide derivative is
desired,
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then one or more different hydrophilic co-monomers may be chosen resulting in
a co-
polymer comprising at least three different monomers.
In one embodiment of the present invention, the acrylamide derivative and the
hydrophilic co-monomer used in the preparation of the synthetic polymer are
the
same. In another embodiment, the acrylamide derivative and the hydrophilic co-
monomer used in the preparation of the synthetic polymer are different.
The overall hydrophilicity of the co-polymer is controlled to confer water
solubility at
0°C to physiological temperatures without precipitation or phase
transition. In one
embodiment of the present invention, the co-polymer is water soluble between
about
0°C and about 37°C.
The co-polymer should be sufficiently soluble in aqueous solution to
facilitate
hydrogel formation. In accordance with one embodiment of the present
invention,
therefore, the term "water soluble" is intended to refer to an aqueous
solubility of the
co-polymer of at least about 0.5 weight/volume (w/v) %. In another embodiment,
the
co-polymer has an aqueous solubility of between about 1.0 w/v % and about 50
w/v
%. In a further embodiments, the co-polymer has an aqueous solubility of about
5 w/v
and about 45 w/v % and between about 10% w/v and about 35% w/v.
As is known in the art, most synthetic polymers have a distribution of
molecular mass
and various different averages of the molecular mass are often distinguished,
for
example, the number average molecular mass (M") and the weight average
molecular
mass (MW). The molecular weight of a synthetic polymer is usually defined in
terms of
its number average molecular mass (M"), which in turn is defined as the sum of
n;M;
divided by the sum of n; ,where n; is the number of molecules in the
distribution with
mass M;. The synthetic polymer for use in the matrix of the present invention
typically
has a number average molecular mass (Mn) between 2,000 and 1,000,000. In one
embodiment of the present invention, the M" of the polymer is between about
5,000
and about 90,000. In another embodiment of the present invention, the M" of
the
polymer is between about 25,000 and about 80,000. In a further embodiment, the
M"
of the polymer is between about 30,000 and about 50,000. In another
embodiment, the
Mn of the polymer is between about 50,000 and about 60,000.
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It is also well-known in the art that certain water-soluble polymers exhibit a
lower
critical solution temperature (LCST) or "cloud point." The LCST of a polymer
is the
temperature at which phase separation occurs (i. e. the polymer begins to
separate from
the surrounding aqueous medium). Typically, for those polymers or hydrogels
that are
clear, the LCST also corresponds to the point at which clarity begins to be
lost. It will
be readily apparent that for certain tissue engineering applications, the
presence or
absence of phase separation in the final hydrogel may not be relevant provided
that the
hydrogel still supports cell growth. For other applications, however, a lack
of phase
separation in the final hydrogel may be critical. For example, for optical
applications,
clarity (and, therefore, the absence of any phase transition) will be
important.
Thus, in accordance with one embodiment of the present invention, co-polymers
with
a LCST between about 35°C and about 60°C are selected for use in
the hydrogels. In
another embodiment, co-polymers with a LOST between about 42°C and
about 60°C
are selected for use in the hydrogels It is also known in the art that the
LCST of a
polymer may be affected by the presence of various solutes, such as ions or
proteins,
and by the nature of compounds cross-linked or attached to the polymer. Such
effects
can be determined empirically using standard techniques and selection of a
synthetic
polymer with an appropriate LCST for a particular application is considered to
be
within the ordinary skills of a worker in the art.
In order for the synthetic polymer to be suitably robust and thermostable, it
is
important that the ratio of acrylamide derivatives) to hydrophilic co-
monomer(s) is
optimised when different monomers are used for these components. Accordingly,
the
acrylamide derivatives) are present in the synthetic polymer in the highest
molar
ratio. W addition, the number of derivatised carboxylic acid co-monomer(s)
present in
the final polymer will determine the ability of the synthetic gel to form
cross-links
with the bio-polymer in the bio-synthetic matrix. Selection of suitable molar
ratios of
each component to provide a final synthetic polymer with the desired
properties is
within the ordinary skills of a worker in the art.
In accordance with one embodiment of the present invention, when different
monomers are being used as the acrylamide derivative. and hydrophilic co-
monomer
components, the amount of acrylamide derivative in the polymer is between 50%
and
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90%, the amount of hydrophilic co-monomer is between 5% and 50%, and the
amount
of derivatised carboxylic acid co-monomer is between 0.1% and 15%, wherein the
sum of the amounts of acrylamide derivative, hydrophilic co-monomer and
derivatised
carboxylic acid co-monomer is 100%, wherein the % value represents the molar
ratio.
In accordance with another embodiment of the invention, the synthetic polymer
is
prepared using the same monomer as both the acrylamide derivative and the
hydrophilic co-monomer and the molar ratio of the acrylamide
derivative/hydrophilic
co-monomer is between about 50% and about 99.5% and the molar ratio of the
derivatised carboxylic acid co-monomer is between about 0.5% and about 50%.
In accordance with a further embodiment, the combined molar ratio of the
acrylamide
derivative and the hydrophilic co-monomer is between about 80% and about 99%
and
the molar ratio of the derivatised carboxylic acid co-monomer is between about
1
and about 20%.
One skilled in the art will appreciate that the selection and ratio of the
components in
the synthetic polymer will be dependent to varying degrees on the final
application for
the bio-synthetic matrix. For example, for ophthalmic applications, it is
important that
the final matrix be clear, whereas for other tissue engineering applications,
the clarity
of the matrix may not be an important factor.
In one embodiment of the present invention, the synthetic polymer is a random
or
block co-polymer comprising one acrylamide derivative, one hydrophilic co-
monomer
and one derivatised carboxylic acid co-monomer (a "terpolymer"). In another
embodiment, the synthetic polymer is a terpolymer comprising NiPAAm monomer,
acrylic acid (A.Ac) monomer and a derivatised acrylic acid monomer. In a
further
embodiment, the synthetic polymer is a terpolymer comprising NiPAAm monomer,
acrylamide (A.Am) monomer and derivatised acrylic acid monomer. In another
embodiment, the derivatised acrylic acid monomer is N-acryloxysuccinimide. In
another embodiment, a terpolymer is prepared with a feed ratio that comprises
NiPA_Am monomer, AAc monomer and N-acryloxysuccinimide in a ratio of about
85:10:5 molar %.
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In an alternate embodiment of the invention, the synthetic polymer is a random
or
block co-polymer comprising one acrylamide derivative/hydrophilic co-monomer
and
one carboxylic acid co-monomer. In another embodiment, the synthetic polymer
comprises DMAA monomer and a derivatised acrylic acid monomer. In a further
embodiment, the derivatised acrylic acid monomer is N-acryloxysuccinimide. In
another embodiment, a synthetic polymer is prepared with a feed ratio that
comprises
DMAA monomer and N-acryloxysuccinimide in a ratio of about 95:5 molar %.
1.2 Bio pol~nzzer
Bio-polymers are naturally-occurnng polymers, such as proteins and
carbohydrates. In
accordance with the present invention, the bio-synthetic matrix comprises a
bio-
polymer or a derivatised version thereof cross-linked to the synthetic polymer
by
means of the pendant cross-linking moieties in the latter. Thus, for the
purposes of the
present invention the bio-polymer contains one or more groups which are
capable of
reacting with the cross-linking moiety (e.g. a primary amine or a thiol), or
can be
derivatised to contain such a group. Examples of suitable bio-polymers for use
in the
present invention include, but are not limited to, collagens (including Types
I, II, III,
IV and V), denatured collagens (or gelatins), recombinant collagens, fibrin-
fibrinogen,
elastin, glycoproteins, alginate, chitosan, hyaluronic acid, chondroitin
sulphates and
glycosaminoglycans (or proteoglycans). One skilled in the art will appreciate
that
some of these bio-polymers may need to be derivatised in order to contain a
suitable
reactive group as indicated above, for example, glucosaminoglycans need to be
deacetylated or desulphated in order to possess a primary amine group. Such
derivatisation can be achieved by standard techniques and is considered to be
within
the ordinary skills of a worker in the art.
Suitable bio-polymers for use in the invention can be purchased from various
commercial sources or can be prepared from natural sources by standard
techniques.
1.3 Bioactive Agefzts
As indicated above, the synthetic polymer to be included in the bio-synthetic
matrix
the present invention contains a plurality of pendant cross-linking moieties.
It will be
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apparent that sufficient cross-linking of the synthetic and bio-polymers to
achieve a
suitably robust matrix can be achieved without reaction of all free cross-
linking
groups. Excess groups may, therefore, optionally be used to covalently attach
desirable bioactive agents to the matrix. Non-limiting examples of bioactive
agents
that may be incorporated into the matrix by cross-linking include, for
example, growth
factors, retinoids, enzymes, cell adhesion factors, extracellular matrix
glycoproteins
(such as laminin, fibronectin, tenascin and the like), hormones, osteogenic
factors,
cytokines, antibodies, antigens, and other biologically active proteins,
certain
pharmaceutical compounds, as well as peptides, fragments or motifs derived
from
biologically active proteins.
In one embodiment of the present invention, the cross-linking groups are
succinimidyl
groups and suitable bioactive agents for grafting to the polymer are those
which
contain either primary amino or thiol groups, or which can be readily
derivatised so as
to contain these groups.
2. METHOD OF PREPARING THE BIO-SYNTHETIC MATRIX
2.1 Preparatio~z of the Syfithetie Polyyraer
Co-polymerization of the components for the synthetic polymer can be achieved
using
standard methods known in the art [for example, see A. Rawe "Principles of
Polymer
Chemistry", Chapter 3. Plenum Press, New York 1995]. Typically appropriate
quantities of each of the monomers are dispersed in a suitable solvent in the
presence
of an initiator. The mixture is maintained at an appropriate temperature and
the co-
polymerisation reaction is allowed to proceed for a pre-determined period of
time. The
resulting polymer can then be purified from the mixture by conventional
methods, for
example, by precipitation.
The solvent for the co-polymerisation reaction may be a non-aqueous solvent if
one or
more monomer is sensitive to hydrolysis or it may be an aqueous solvent.
Suitable
aqueous solvents include, but are not limited to, water, buffers and salt
solutions.
Suitable non-aqueous solvents are typically cyclic ethers (such as dioxane),
chlorinated hydrocarbons (for example, chloroform) or aromatic hydrocarbons
(for
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example, benzene). The solvent may be nitrogen purged prior to use, if
desired. In one
embodiment of the present invention, the solvent is a non-aqueous solvent. In
another
embodiment, the solvent is dioxane.
Suitable polymerisation initiators are known in the art and are usually free-
radical
initiators. Examples of suitable initiators include, but are not limited to,
2,2'-
azobisisobutyronitrile (AIBN), other azo compounds, such as 2,2'-azobis-2-
ethylpropionitrile; 2,2'-azobis-2-cyclopropylpropionitrile; 2,2'-
azobiscyclohexanenitrile; 2,2'-azobiscyclooctanenitrile, and peroxide
compounds,
such as dibenzoyl peroxide and its substituted analogues, and persulfates,
such as
sodium, potassium, and the like.
Once the polymer has been prepared, and purified if necessary, it can be
characterised
by various standard techniques. For example, the molar ratio composition of
the
polymer can be determined by nuclear magnetic resonance spectroscopy (proton
and/or carbon-13) and bond structure can be determined by infrared
spectroscopy.
Molecular mass can be determined by gel permeation chromatography andlor high
pressure liquid chromatography. Thermal characterisation of the polymer can
also be
conducted, if desired, for example by determination of the melting point and
glass
transition temperatures using differential scanning calorimetric analysis.
Aqueous
solution properties such as micelle and gel formation, and LCST can be
determined
using visual observation, fluorescence spectroscopy, UV-visible spectroscopy
and
laser light scattering instruments.
In one embodiment of the present invention, the synthetic polymer is prepared
by
dispersing the monomers in nitrogen-purged dioxane in the presence of the
initiator
AlBN and allowing polymerisation to proceed at a temperature of about
60°C to 70°C.
The resulting polymer is purified by repeated precipitation.
2.2 Preparation of the Hydrogel
Cross-linking of the synthetic polymer and bio-polymer can be readily achieved
by
mixing appropriate amounts of each polymer at room temperature in a suitable
solvent. Typically the solvent is an aqueous solvent, such as a salt solution,
buffer
solution, cell culture medium, or a diluted or modified version thereof. One
skilled in
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the art will appreciate that in order to preserve triple helix structure of
polymers such
as collagen and to prevent fibrillogenisis and/or opacification of the
hydrogel, the
cross-linking reaction should be conducted in aqueous media with close control
of the
pH and temperature. The significant levels of amino acids in nutrient media
normally
used for cell culture can cause side reactions with the cross-linking moieties
of the
synthetic polymer, which can result in diversion of these groups from the
cross-
linking reaction. Use of a medium free of amino acids and other proteinaceous
materials can help to prevent these side reactions and, therefore, increase
the number
of cross-links that form between the synthetic and bio-polymers. Conducting
the
cross-linl~ing reaction in aqueous solution at room or physiological
temperatures
allows both cross-linking and the much slower hydrolysis of any residual cross-
linking
groups to take place.
Alternatively, a termination step can be included to react any residual cross-
liking
groups in the matrix. For example, one or more wash steps in a suitable buffer
containing glycine will terminate any residual cross-linking groups as well as
removing any side products generated during the cross-linking reaction.
Unreacted
cross-linking groups may also be terminated with a polyfunctional amine such
as
lysine or triethylenetetraamine leading to formation of additional short,
inter-chain
cross-links. Wash steps using buffer alone can also be conducted if desired in
order to
remove any side products from the cross-linking reaction. If necessary, after
the cross-
linking step, the temperature of the cross-linked polymer suspension can be
raised to
allow the hydrogel to form fully.
In accordance with the present invention, the components of the hydrogel are
chemically cross-linked so as to be substantially non-extractable, i. e. the
bio-polymer
and synthetic polymer do not exude extensively from the gel under
physiological
conditions. In accordance with one embodiment of the present invention, the
amount
of bio-polymer or synthetic polymer that can be extracted from the matrix into
aqueous media under physiological conditions over a period of 24 hours is less
than
5% by weight of either component. In another embodiment, the amount of bio-
polymer or synthetic polymer that can be extracted from the matrix into
aqueous
media under physiological conditions over a period of 24 hours is less than 2%
by
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weight of either component. In further embodiments, the amount that can be
extracted
over a period of 24 hours is less than 1% by weight, less than 0.5% by weight
and less
than 0.2% by weight.
The amount of bio-polymer and/or synthetic polymer that can be extracted from
the
matrix into aqueous media can be determined ira vitYO using standard
techniques (for
example, the USP Basket Method). Typically, the matrix is placed in an aqueous
solution of a predetermined pH, for example axound pH 7.4 to simulate
physiological
conditions. The suspension may or may not be stirred. Samples of the aqueous
solution are removed at predetermined time intervals and are assayed for
polymer
content by standard analytical techniques.
One skilled in the art will understand that the amount of each polymer to be
included
in the hydrogel will be dependent on the choice of polymers and the intended
application for the hydrogel. In general, using higher initial amounts of each
polymer
will result in the formation of a more robust gel due to the lower water
content and the
presence of a greater amount of cross-linked polymer. Higher quantities of
water or
aqueous solvent will produce a soft hydrogel. In accordance with the present
invention, the final hydrogel comprises between about 20 and 99.6 % by weight
of
water or aqueous solvent, between about 0.1 and 30 % by weight of synthetic
polymer
and between about 0.3 and 50 % by weight of bio-polymer.
In one embodiment of the present invention, the final hydrogel comprises
between
about 40 and 99.6 % by weight of water or aqueous solvent, between about 0.1
and 30
by weight of synthetic polymer and between about 0.3 and 30 % by weight of bio-
polymer. In another embodiment, the final hydrogel comprises between about 60
and
99.6 % by weight of water or aqueous solvent, between about 0.1 and 10 % by
weight
of synthetic polymer and between about 0.3 and 30 % by weight of bio-polymer.
In a
further embodiment, the final hydrogel comprises between about 80 and 98.5 %
by
weight of water or aqueous solvent, between about 0.5 and 5 % by weight of
synthetic
polymer and between about 1 and 15 % by weight of bio-polymer. In other
embodiments, the final hydrogel contains about 95 to 97 % by weight of water
or
aqueous solvent and between about 1- 2 % by weight of synthetic polymer and
about
2 - 3 % by weight of bio-polymer; and about 94 to 98 % by weight of water or
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aqueous solvent and between about 1 - 3 % by weight of synthetic polymer and
about
1- 3 % by weight of bio-polymer.
Similarly, the relative amounts of each polymer to be included in the hydrogel
will be
dependent on the type of synthetic polymer and bio-polymer being used and upon
the
intended application for the hydrogel. One skilled in the art will appreciate
that the
relative amounts bio-polymer and synthetic polymer will influence the final
gel
properties in various ways, for example, high quantities of bio-polymer will
produce a
very stiff hydrogel. One skilled in the art will appreciate that the relative
amounts of
each polymer in the final matrix can be described in terms of the weight :
weight ratio
of the bio-polymer : synthetic polymer or in terms of equivalents of reactive
groups. In
accordance with the present invention, the weight per weight ratio of bio-
polymer
synthetic polymer is between about 1 : 0.07 and about 1 : 14. In one
embodiment, the
w/w ratio of bio-polymer : synthetic polymer is between 1 : 1.3 and 1 : 7. In
another
embodiment, the w/w ratio of bio-polymer : synthetic polymer is between 1 : 1
and 1
3. In a further embodiment, the w/w ratio of bio-polymer : synthetic polymer
is
between 1 : 0.7 and 1 : 2.
In an alternative embodiment of the present invention, the matrix comprises a
proteinaceous bio-polymer and a synthetic polymer comprising pendant N-
acryloxysuccinimide groups. In this embodiment of the invention, the ratio of
bio-
polymer : synthetic polymer is described in terms of molar equivalents of free
amine
groups in the bio-polymer to N-acryloxysuccinimide groups and is between 1 :
0.5 and
1: 20. In another embodiment, this ratio is between 1 : 1.8 and 1 : 10. In a
further
embodiments, the ratio is between 1 : 1 and 1 : 5, and between 1 : 1 and 1 :
3.
2.3 Izzcorporatiozz of Bioactive Age~zts i~zto the Bio-synthetic Matrix
Bioactive agents can be incorporated into the matrix if desired either by
covalent
attachment (or "grafting") to the synthetic polymer through the pendant cross-
linking
moieties, or by encapsulation within the matrix. Examples of bioactive agents
that
may be covalently attached to the synthetic polymer component of the matrix
are
provided above. If necessary, the bioactive agent may be first derivatised by
standard
procedures to provide appropriate reactive groups for reaction with the cross-
linking
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groups. For covalent attachment of a bioactive agent, the synthetic polymer is
first
reacted with the bioactive agent and then this modified synthetic polymer
subsequently cross-linked to the bio-polymer as described above. Reaction of
the
bioactive agent with the synthetic polymer can be conducted under standard
conditions, for example by mixing the bioactive agent and the synthetic
polymer
together in a non-aqueous solvent, such as N,N-dimethyl formamide, dioxane,
dimethyl sulphoxide and N,N-dimethylacrylamide. The use of a non-aqueous
solvent
avoids hydrolysis of the reactive groups during incorporation of the bioactive
agent.
Alternatively, the reaction may be conducted in an aqueous solvent as
described above
for the cross-linking reaction.
Bioactive agents which are not suitable for grafting to the polymer, for
example, those
that do not contain primary amino or free thiol groups for reaction with the
cross-
linking groups in the synthetic polymer, or which cannot be derivatised to
provide
such groups, can be entrapped in the final matrix. Examples of bioactive
agents which
may be entrapped in the matrix include, but are not limited to, pharmaceutical
drugs,
diagnostic agents, viral vectors, nucleic acids and the like. For entrapment,
the
bioactive agent is added to a solution of the synthetic polymer in an
appropriate
solvent prior to mixture of the synthetic polymer and the bio-polymer to form
a cross-
linked hydrogel. Alternatively, the bioactive agent can be added to a solution
containing both the synthetic and bio-polymers prior to the cross-linking
step. The
bioactive agent is mixed into the polymer solution such that it is
substantially
uniformly dispersed therein, and the hydrogel is subsequently formed as
described
above. Appropriate solvents for use with the bioactive agent will be dependent
on the
properties of the agent and can be readily determined by one skilled in the
art.
2.4 Ezztrapnzezzt of Cells irz the Bio-syzztlzetic Matzix
The bio-synthetic matrix according to the present invention may also comprise
cells
entrapped therein and thus permit delivery of the cells to a tissue or organ
irz vivo. A
variety of different cell types may be delivered using the bio-synthetic
matrix, for
example, myocytes, ocular cells (e.g. from the different corneal layers),
adipocytes,
fibromyoblasts, ectodermal cells, muscle cells, osteoblasts (i.e. bone cells),
chondrocytes (i.e. cartilage cells), endothelial cells, fibroblasts,
pancreatic cells,
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hepatocytes, bile duct cells, bone marrow cells, neural cells, genitourinary
cells
(including nephritic cells), or combinations thereof. The matrix may also be
used to
deliver totipotent stem cells, pluripotent or committed progenitor cells or re-
programmed (dedifferentiated) cells to an in vivo site such that cells of the
same type
as the tissue can be produced. For example, mesenchymal stem cells, which are
undifferentiated, can be delivered in the matrix. Examples of mesenchymal stem
cells
include those which can diversify to produce osteoblasts (to generate new bone
tissue), chondrocytes (to generate new cartilaginous tissue), and fibroblasts
(to
produce new connective tissue). Alternatively, committed progenitor cells
capable of
proliferating to provide cells of the same type as those present at the.irz
vivo site can be
used, for example, myoblasts, osteoblasts, fibroblasts and the like.
Cells can be readily entrapped in the final matrix by addition of the cells to
a solution
of the synthetic polymer prior to admixture with the bio-polymer to form a
cross-
linked hydrogel. Alternatively, the cells can be added to a solution
containing both the
synthetic and bio-polymers prior to the cross-linking step. The synthetic
polymer may
be reacted with a bioactive agent prior to admixture with the cells if
desired.
Typically, for the encapsulation of cells in the matrix, the various
components (cells,
synthetic polymer and bio-polymer) are dispersed in an aqueous medium, such as
a
cell culture medium or a diluted or modified version thereof. The cell
suspension is
mixed gently into the polymer solution until the cells are substantially
uniformly
dispersed in the solution, then the hydrogel is formed as described above.
2.5 Other Elenzefzts
The present invention also contemplates the optional inclusion of one or more
reinforcing material in the bio-synthetic matrix to improve the mechanical
properties
of the matrix such as the strength, resilience, flexibility and/or tear
resistance. Thus,
the matrix may be reinforced with flexible or rigid fibres, fibre mesh, fibre
cloth and
the like. The use of such reinforcing materials is known in the art, for
example, the
use of fibres, cloth, or sheets made from collagen fibrils, oxidised cellulose
or
polymers such as polylactic acid, polyglycolic acid or polytetrafluoroethylene
in
implantable medical applications is known.
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The reinforcing material can be incorporated into the matrix using standard
protocols.
For example, an aqueous solution of synthetic and bio-polymers in an
appropriate
buffer can be added to a fibre cloth or mesh, such as Interceed (Ethicon Inc.,
New
Brunswick, N.J.). The aqueous solution will flow into the interstices of the
cloth or
mesh prior to undergoing cross-linking and will thus form a hydrogel with the
cloth or
mesh embedded therein. Appropriate moulds can be used to ensure that the
fibres or
fibre mesh are contained entirely within the hydrogel if desired. The
composite
structure can subsequently be washed to remove any side products generated
during
the cross-linking reaction. Typically, the fibres used are hydrophilic in
nature to
ensure complete wetting by the aqueous solution of polymers.
One skilled in the art will appreciate that, for applications requiring high
optical
clarity, the structure of the reinforcement should be selected to prevent
light scattering
from the final composite matrix, for example, by the use of nano-fibers and/or
careful
refractive index matching of reinforcement and hydrogel.
3. TESTING THE BIO-SYNTHETIC MATRIX
In accordance with the present invention, the bio-synthetic matrix comprises a
hydrogel with or without added bioactive agents and/or encapsulated cells. In
order to
be suitable for ih vivo implantation for tissue engineering purposes, the bio-
synthetic
matrix must maintain its form at physiological temperatures, be adequately
robust, be
substantially insoluble in water, and support the growth of cells. It may also
be
desirable for the matrix to support the growth of nerves. It will be readily
appreciated
that for certain specialised applications, the matrix may require other
characteristics.
For example, for surgical purposes, the matrix may need to be relatively
flexible as
well as strong enough to support surgical manipulation with suture thread and
needle,
and for ophthalinic applications, such as cornea repair or replacement, the
optical
clarity of the matrix will be important.
3.1 Physical/ Clze~rzical testing
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When used for tissue engineering applications, the bio-synthetic matrix needs
to meet
the mechanical parameters necessary to prevent the matrix tearing or rupturing
when
submitted to surgical procedures and to provide adequate support for cell
growth once
in place. The ability of matrix to resist tearing is related to its intrinsic
mechanical
strength, the form and thickness of the matrix and the tension being applied.
The ability of the bio-synthetic matrix to withstand shearing forces, or
tearing can be
roughly determined by applying forces in opposite directions to the specimen
using
two pairs of forceps. Alternatively, a suitable apparatus can be used to
measure
quantitatively the ability of the matrix to withstand shearing forces.
Tensiometers for
this purpose are available commercially, for example, from MTS, Instron, and
Cole
Parmer.
For testing, the matrix can be formed into sheets and then cut into
appropriately sized
strips. Alternatively, the matrix can be moulded into the desired shape for
tissue
engineering purposes and the entire moulded matrix can be tested. To calculate
tensile
strength, the force at rupture, or "failure," of the matrix is divided by the
cross-
sectional area of the test sample, resulting in a value that can be expressed
in force (I~
per unit area. The stiffiiess (modulus) of the matrix is calculated from the
slope of the
linear portion of the stress/strain curve. Strain is the real-time change in
length during
the test divided by the initial length of the test sample before the test
begins. The
strength at rupture is the final length of the test sample when it ruptures
minus the
length of the initial test sample, divided by this initial length.
One skilled in the art will appreciate that because of the softness of
hydrogels and
exudation of the aqueous component when clamped, meaningful tensile data can
be
difficult to obtain from hydrogels. Quantitative characterisation of tensile
strength in
hydrogels can be achieved, for example, through the use of 'suture pull-out
measurements on moulded matrix samples. Typically, a suture is placed about 2
mm
from the edge of a test sample and the peak force that needs to be applied in
order to
rip the suture through the sample is measured. For example, for a test sample
of
matrix intended for ophthalmic applications that has been moulded in the shape
and
thickness of a human cornea, two diametrically opposed sutures can be inserted
into
the matrix, as would be required for the first step in ocular implantation.
The two
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sutures can then be pulled apart at about 10 mm/min on a suitable instrument,
such as
an Instron Tensile Tester. Strength at rupture of the matrix is calculated,
together with
elongation at break and elastic modulus [see, for example, Zeng et al., J.
Biomech.,
34:533-537 (2001)]. It will be appreciated by those skilled in the art that,
for those
bio-synthetic matrices intended for surgical applications, the matrices need
not be as
strong (i.e. have the same ability to resist tearing) as mammalian tissue. The
determining factor for the strength of the matrix in such applications is
whether or not
it can be sutured in place by a careful and experienced surgeon.
If desired, the LCST of the bio-synthetic hydrogel matrix can be measured
using
standard techniques. For example, LCST can be calculated by heating samples of
the
matrix at about 0.2°C per minute and visually observing the cloud point
(see, for
example, H. Uludag, et al., J. Appl. Polym. Sci. 75:583 - 592 (2000)).
Permeability of the bio-synthetic matrix can be determined by assessing the
glucose
permability coefficient andlor the average pore sizes for the matrix using
standard
techniques such as PBS permeability assessment using a permeability cell
and/or
atomic force microscopy. In accordance with one embodiment of the present
invention, the bio-synthetic matrix has an average pore size between about 90
nm and
about 500 nm. hz another embodiment, the matrix has an average pore size
between
about 100 nm and about 300 nm.
Optical transmission and light scatter can also be measured for matrices
intended for
ophthalmic applications using a custom-built instrument that measures both
transmission and scatter [see, for example, Priest and Monger, havest.
Ophthalmol.
Vis. Sci. 39: 5352 (1998)].
3.2 In vitro Testisig
It will be readily appreciated that the bio-synthetic matrix must be non-
cytotoxic and
biocompatible in order to be suitable for in vivo use. The cytotoxicity of the
bio-
synthetic matrix can be assessed using standard techniques such as the Ames
assay to
screen for mutagenic activity, the mouse lymphoma assay to screen for the
ability of
the matrix to induce gene mutation in a mammalian cell line, ifZ vitro
chromosomal
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aberration assays using, for example, Chinese hamster ovary cells (CHO) to
screen for
any DNA rearrangements or damage induced by the matrix. Other assays include
the
sister chromatid assay, which determines any exchange between the arms of a
chromosome induced by the matrix and ira vitro mouse micronucleus assays to
determine any damage to chromosomes or to the mitotic spindle. Protocols for
these
and other standard assays are known in the art, for example, see OECD
Guidelines fog
tlae Testing of Chemicals and protocols developed by the ISO.
The ability of the matrix to support cell growth can also be assessed ifZ
vits°o using
standard techniques. For example, cells from an appropriate cell line, such as
human
epithelial cells, can be seeded either directly onto the matrix or onto an
appropriate
material surrounding the matrix. After growth in the presence of a suitable
culture
medium for an appropriate length of time, confocal microscopy and histological
examination of the matrix can be conducted to determine whether the cells have
grown over the surface of and/or into the matrix.
The ability of the matrix to support in-growth of nerve cells can also be
assessed iya
vitro. For example, a nerve source, such as dorsal root ganglia, can be
embedded into
an appropriate material surrounding the matrix or directly inserted into the
matrix. An
example of a suitable material would be a soft collagen based gel. Cells from
an
appropriate cell line can then be seeded either directly onto the matrix or
onto an
appropriate material surrounding the matrix and the matrix can be incubated in
the
presence of a suitable culture medium for a pre-determined length of time.
Examination of the matrix, directly and l or in the presence of a nerve-
specific marker,
for example by immunofluorescence using a nerve-specific fluorescent marker
and
confocal microscopy, for nerve growth will indicate the ability of the matrix
to
support neural in-growth.
Growth supplements can be added to the culture medium, to the matrix or to
both in
experiments to assess the ability of the matrix to support cell growth. The
particular
growth supplements employed will be dependent in the type of cells being
assessed
and can be readily determined by one skilled in the art. Suitable supplements
for nerve
cells, for example, include laminin, retinyl acetate, retinoic acid and nerve
growth
factors for nerve cells.
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3.3 In vivo Testing
In order to assess the biocompatibility of the bio-synthetic matrix and its
ability to
support cell growth ira vivo, the matrix can be implanted into an appropriate
animal
model for immunogenicity, inflammation, release and degradation studies, as
well as
determination of cell growth. Suitable control animals may be included in the
assessment. Examples of suitable controls include, for example, unoperated
animals,
animals that have received allografts of similar dimensions from a donor
animal
and/or animals that have received implants of similar dimensions of a
standard,
accepted implant material.
At various stages post-implantation, biopsies can be taken to assess cell
growth over
the surface of and/or into the implant and histological examination and
immunohistochemistry techniques can be used to determine whether nerve in-
growth
has occurred and whether inflammatory or immune cells are present at the site
of the
implant. For example, various cell-specific stains known in the art can be
used to
assess the types of cells present as well as various cell-specific antibodies,
such a anti-
neurofilament antibodies that can be used to indicate the presence or absence
of nerve
cells. In addition, measurement of the nerve action potentials using standard
techniques will provide an indication of whether the nerves are functional.
Ifa vivo
confocal microscopic examination can be used to monitor cell and nerve growth
in the
animal at selected post-operative times. Where appropriate, touch sensitivity
can be
measured by techniques known in the art, for example, using an esthesiometer.
Restoration of touch sensitivity indicates the regeneration of functional
nerves.
4. APPLICATIDNS
The present invention provides a bio-synthetic matrix which is robust,
biocompatible
and non-cytotoxic and, therefore, suitable for use as a scaffold to allow
tissue
regeneration in vivo. For example, the bio-synthetic matrix can be used for
implantation into a patient to replace tissue that has been damaged or
removed, for
wound coverage, as a tissue sealant or adhesive, as a skin substitute or
cornea
substitute, or as a corneal veneer. The matrix can be moulded into an
appropriate
shape prior to implantation, for example it can be pre-formed to fill the
space left by
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damaged or removed tissue. Alternatively, when used as an implant, the matrix
may
be allowed to form ifa situ by injecting the components into the damaged
tissue and
allowing the polymers to cross-link and gel at physiological temperature.
In one embodiment of the present invention, the matrix is pre-formed into an
appropriate shape for tissue engineering purposes. hi another embodiment the
matrix
is pre-formed as a full thickness artificial cornea or as a partial thickness
matrix
suitable for a cornea veneer.
The bio-synthetic matrix can be used alone and as such will support the in-
growth of
new cells in situ. Alternatively, the matrix can be seeded with cells prior to
implantation and will support the outgrowth of these cells in vivo to repair
andlor
replace the surrounding tissue. It is contemplated that the cells may be
derived from
the patient, or they may be allogeneic or xenogenic in origin. For example,
cells can
be harvested from a patient (prior to, or during, surgery to repair the
tissue) and
processed under sterile conditions to provide a specific cell type such as
pluripotent
cells, stem cells or precursor cells. These cells can then be seeded into the
matrix, as
described above and the matrix can be subsequently implanted into the patient.
The matrix can also be used to coat surgical implants to help seal tissues or
to help
adhere implants to tissue surfaces, for example, through the formation of
cross-links
between unreacted cross-linking groups on the synthetic polymer component of
the
hydrogel and primary amino or thiol groups present in the tissue. For example,
a layer
of the matrix may be used to patch perforations in corneas, or be applied to
catheters
or breast implants to reduce fibrosis. The matrix may also be applied to
vascular grafts
or stems to minimise blood or serosal fluid leakage, to artificial patches or
meshes to
minimise fibrosis and to help adhesion of the implants to tissue surfaces.
The matrix may also be used as a delivery system to deliver a bioactive agent
to a
particular region in a patient. The bioactive agent can be delivered as a
solution
together with the synthetic and bio-polymers such that the matrix comprising
the
bioactive agent can form ih situ, or the matrix comprising the bioactive agent
can be
pre-formed and implanted. Once within the body, the bioactive agent may be
released
from the matrix, for example, through diffusion-controlled processes or, if
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bioactive agent is covalently bound to the matrix, by enzymatic cleavage from
the
matrix and subsequent release by diffusion-controlled processes.
Alternatively, the
bioactive agent may exert its effects from within the matrix.
In one embodiment of the present invention, the bio-synthetic matrix is used
as an
artificial cornea. For this application, the matrix is pre-formed as a full
thickness
artificial cornea or as a partial thickness matrix suitable for a cornea
veneer. In
accordance with this embodiment, the hydrogel is designed to have a high
optical
transmission and low light scattering. For example, hydrogels comprising a
synthetic
p(NiPAAm-co-AAc-co-ASI) terpolymer or p(DMAA-co-ASI) co-polymer cross-
linked to collagen have high optical transmission, very low light scattering
and are
capable of remaining clear up to 55°C.
S. KITS
The present invention also contemplates kits comprising the bio-synthetic
matrix. The
kits may comprise a "ready-made" form of the matrix or they may comprise the
individual components required to make the matrix in appropriate proportions
(i.e. the
synthetic polymer and the bio-polymer. The kit may optionally further comprise
one
or more bioactive agent either pre-attached to the synthetic polymer, or as
individual
components that can be attached to the synthetic polymer during preparation of
the
matrix. The kits may further comprise instructions for use, one or more
suitable
solvents, one or more instruments for assisting with the injection or
placement of the
final matrix composition within the body of an animal (such as a syringe,
pipette,
forceps, eye dropper or similar medically approved delivery vehicle), or a
combination
thereof. Individual components of the kit may be packaged in separate
containers. The
kit may further comprise a notice in the form prescribed by. a governmental
agency
regulating the manufacture, use or sale of biological products, which notice
reflects
approval by the agency of the manufacture, use or sale for human or animal
applications.
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To gain a better understanding of the invention described herein, the
following
examples are set forth. It should be understood that these examples are for
illustrative
purposes only. Therefore, they should not limit the scope of this invention in
any way.
EXAMPLES
Abbreviations
RTT: rat-tail tendon
ddH20: distilled, de-ionised water
PBS: phosphate buffered saline
D-PBS: Dulbecco's phosphate buffered saline
AlBN: azobis-isobutyronitrile
NiPAAm: N-isopropylacrylamide
pNiPAAm: poly(N-iso-propylacrylamide)
Apc: acrylic acid
DMAA N,N-dimethylacrylamide
ASI: N-acryloxysuccinimide
pNIPAAm-co-AAc: copolymer of NiPAAm and AAc
poly(NiPAAm-co-AAc-co-ASI): terpolymer of NIPAAM, AAc and ASI
poly(DMAA-co-ASI) co-polymer of DMAA and ASI
GPC: gel permeation chromatography
NMR nuclear magnetic resonance
YIGSR: amide-terminated pentapeptide (tyrosine-
isoleucine-glycine-serine-arginine)
All gel matrices described in the Examples set out below used sterile collagen
I, such
as telocollagen (rat-tail tendon, RTT) or atelocollagen (bovine or porcine),
which can
be prepared in the laboratory or more conveniently is available commercially
(for
example, from Becton Dickinson at a concentration of 3.0-3.5 mg/ml in 0.02N
acetic
acid and in 0.012N hydrochloric acid for bovine and porcine collagen). Such
collagens
can be stored for many months at 4°C. In addition, such collagen
solutions may be
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carefully concentrated to give optically clear, very viscous solutions of 3 -
30 wt/vol
collagen, suitable for preparing more robust matrices.
Collagen solutions are adjusted to physiological conditions, i.e. saline ionic
strength
and pH 7.2 - 7.4, through the use of aqueous sodium hydroxide in the presence
of
phosphate buffered saline (PBS). PBS, which is free of amino acids and other
nutrients, was used to avoid depletion of cross-linking reactivity by side
reactions with
NH2 containing molecules, so allowing the use of the minimum concentration of
cross-linking groups and minimising any risk of cell toxicity.
pNiPAAm homopolymer powder is available commercially (for example, from
Polyscience). All other polymers were synthesized as outlined below.
EXAMPLE 1: PREPARATION OFA pNiPAAnz-COLLAGENHYDROGEL
A pNiPAAm-collagen hydrogel was prepared to provide an alternative hydrogel
against which the properties of the hydrogels of the present invention could
be
compared.
A 1 wt/vol% solution of pNiPAAm homopolymer in ddHzO was sterilised by
autoclaving. This solution was mixed with sterile RTT collagen solution [3.0-
3.5
mg/ml (w/v) in acetic acid (0.02N in water] (1:1 vol/vol) in a sterile test
tube at 4°C
by syringe pumping to give complete mixing without bubble formation. Cold
mixing
avoids any premature gelification or fibrilogenesis of the collagen. The
collagen-
pNiPAAm was then poured over a plastic dish (untreated culture dish) or a
mould
(e.g. contact lens mould) and left to air-dry under sterile conditions in a
laminar flow
hood for at least 2-3 days at room temperature. After drying to constant
weight (~7
water residue), the formed matrix was removed from the mould. Removal of the
matrix from the mould is facilitated by soaking the mould in a sterile PBS at
room
temperature. Continued soaking of the free sample in this solution gives a gel
at
physiological temperature, pH and ionic strength, which was subsequently
submitted
to testing for cell growth and ih vivo animal testing (see Examples 6 and 7).
EXAMPLE 2: PREPARATION OF A SYNTHETIC TERPOLYMER
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A collagen-reactive terpolymer, poly(NiPAAm-co-AAc-co-ASI) (Figure 1), was
synthesised by co-polymerising the three monomers: N-isopropylacrylamide,
(NiPAAm, 0.85 mole), acrylic acid (AAc, 0.10 mole) and N-acryloxysuccinimide
(ASI, 0.05 mole). The feed molar ratio was 85:10:5 (NiPAAm: AAc: ASI), the
free-
radical initiator AIBN (0.007 mole/mole of total monomers) and the solvent,
dioxane
(100 ml), nitrogen purged before adding AIBN. The reaction proceeded for 24 h
at
65°C.
After purification by repeated precipitation to remove traces of homopolymer,
the
composition of the synthesised terpolymer (82% yield) was found to be
84.2:9.8:6.0
(molar ratio) by proton NMR in THF-D8. The M" and MW of the terpolymer were
5.6
X 104 Da and 9.0 x 104 Da, respectively, by aqueous GPC.
A solution of 2 mg/ml of the terpolymer in D-PBS remained clear even up to
55°C,
consistent with a high LCST. A solution of 10 mg/ml in D-PBS became only
slightly
cloudy at 43°C. Failure to remove homopolymer formed in the batch
polymerisation
reaction (due to the NiPAAm reactivity coefficient being greater than that of
AAc or
ASI) from the terpolymer gave aqueous solutions and hydrogels which cloud at
~32°C
and above.
EXAMPLE 3: PREPARATION OF A SYNTHETIC POLYMER COMPRISING
A BIOACTIVE AGENT
A terpolymer, containing the pentapeptide YIGSR (a nerve cell attachment
motif),
was synthesised by mixing the terpolymer prepared in Example 2 (1.0 g) with
2.8~g
of laminin pentapeptide (YIGSR, from Novabiochem) in N,N-dimethyl formamide.
After reaction for 48 h at room temperature (21°C), the polymer
product was
precipitated out from diethyl ether and then vacuum dried. ASI groups
remaining after
reaction with the pentapeptide are available for subsequent reaction with
collagen.
The structure of this polymer is shown in Figure 8A.
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EXAMPLE 4: PREPARATION OF A COLLAGEN TERPOLYMER
HYDROGEL
A cross-linked, terpolymer-collagen hydrogel was made by mixing neutralised 4%
bovine atelocollagen (1.2 ml) with the terpolyrner prepared in Example 2
[0.34m1
(100 mg/ml in D-PBS)] by syringe mixing at 4°C (collagen : terpolymer
1.4 : 1 w/w).
After careful syringe pumping to produce a homogeneous, bubble-free solution,
aliquots were injected into plastic, contact lens moulds and incubated at room
temperature (21°C) for 24 hours to allow reaction of the collagen NHZ
groups with
ASI groups as well as the slower hydrolysis of residual ASI groups to AAc
groups.
The moulded samples were then incubated at 37°C for 24 hours in 100%
humidity
environment, to give a final hydrogel. The hydrogel contained 95.4 ~ 0.1 %
water,
2.3% collagen and 1.6% terpolymer. Matrices were moulded to have a final
thickness
between either 150 - 200 p.m or 500 - 600 Vim. Each resulting hydrogel matrix
was
removed from its mould under PBS solution and subsequently immersed in PBS
containing 1% chloroform and 0.5% glycine. This wash step removed N-
hydroxysuccinimide produced in the cross-linking reaction, terminated any
unreacted
ASI groups in the matrix, by conversion to acrylic acid groups and sterilised
the
hydrogel matrix. As an alternative, moulded gels rnay be treated with aqueous
glycine
to ensure that all ASI are terminated prior to cell contact.
Succinimide residues left in the gels prepared from collagen and terpolymer
were
below the 1R detection limit after washing.
EXAMPLE 5: PREPARATION OF A HYDROGEL COMPRISING A
BIOACTI1~E AGENT
Cross-linked hydrogels of collagen-terpolymer comprising YIGSR cell adhesion
factor were prepared by thoroughly mixing viscous, neutralised 4% bovine
collagen
(1.2 ml) with terpolymer to which laminin pentapeptide (YIGSR) was covalently
attached (from Example 3; 0.34m1, 100 mglml) at 4°C, following the
procedure
described in Example 4.
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The YIGSR content of extensively washed gels was 4.3 x 10-11 mole/ml (2.6 x 10-
8
g/ml) of hydrated gel quantified by labelling the tyrosine (primary amine-
bearing)
groups with lasl using the Iodogen method and measuring the radioactivity with
a
standardised gamma counter (Beckman, Gamma 5500). The final, total polymer
concentration in each hydrated, PBS-equilibrated hydrogel was 3.4 w/v %
(comprising
collagen and YIGSR terpolymer at 2.0 and 1.4 w/v %, respectively).
EXAMPLE 6: COMPARISON OF THE PHYSICAL PROPERTIES OF
HYDROGEL MATRICES
Collagen thermogels are frail and readily collapse and break and are obviously
opaque
(see Figure 8C). Collagen thermogels were prepared by neutralization of
collagen and
casting in the same moulds as described above in Examples 4 and 5. The moulded
collagen was then incubated, first for 24 h at 21°C then at
37°C, to spontaneously form
translucent thermogels (produced by self association of collagen triple
helices into
micro-fibrils).
The permeability coefficient of glucose in PBS (pH 7.4) through hydrogels
prepared
as described in Examples 5 was calculated from measurements in a permeation
cell by
periodically removing aliquots of permeate, adding adenosine triphosphate and
converting glucose to glucose-6-phosphate with the enzyme hexokinase. The
latter
was reacted with nicotinamide adenine dinucleotide in the presence of
dehydrogenase
and the resultant reduced dinucleotide quantified by its UV absorption at 340
run in
solution (Bondar, R. J. & Mead, D. C. (1974) Clih Claem 20, 586-90).
Topographies
of hydrogel surfaces, fully immersed in PBS solution, were examined by atomic
force
microscopy (Molecular Image Co., USA) in the "contact" mode. Pore sizes from
this
technique were compared with average pore diameters calculated from the PBS
permeability of the hydrogels as previously described (Bellamkonda, R.,
Ranieri, J. P.
& Aebischer, P. (1995) JNeu~osci Res 41, 501-9). The hydrogels had refractive
indices (1.343 ~ 0.003) comparable to the tear film (1.336-1.357) in the human
eye
(Patel, S., Marshall, J. & Fitzke, F. W., 3rd (1995) JRefYact Surg 11, 100-5).
They
showed high optical clarity compared to matrices that contain only collagen
(Fig. 8B
and C). The hydrogels had pore diameters of 140 -190 nm (from both atomic
force
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microscopy and PBS permeability) and a glucose diffusion permeability
coefficient of
2. 7x 10-6 cm2/s, which is higher than the value for the natural stroma (~0.7x
10-6 cmz/s,
calculated from published diffusion (2.4x10-6 cm2/s) and solubility (0.3)
coefficients
(McCarey, B. E. & Schmidt, F. H. (1990) Curf~ Eye Res 9, 1025-39)).
The following properties of the hydrogels prepared as described in Examples 4
and 5
indicate that they are cross-linked:
~ water insoluble,
~ strong enough to support surgical manipulation with suture thread and
needle, and attachment to a human corneal ring
~ relatively flexible in handling
~ demonstrate an increase in stress at failure and apparent modulus
during tensile testing by over x2 on going from ASI/-NH2 equivalent
ratio of 0.5 to 1.5.
Matrices prepared as described above, but with varying ratios of collagen
amine
groups to ASI groups in the synthetic polymer had high optical transmission
and low
scatter in the visible region (Fig. 8B, 12 and 13). In contrast, the collagen
thermogel,
prepared from collagen as described above, had low transmission and high
scatter,
consistent with its opaque appearance (Fig. 8C, 12). Such thermogel matrices
with up
to 3 wt/vol collagen were too weak to allow mounting for suture pull-out
testing.
Quantitative characterisation of the hydrogels came from the use of suture
pull-out
measurements on samples moulded into the shape and thickness of a human
cornea.
This involved insertion of two diametrically opposed sutures, as required for
the first
step in ocular implantation, and pulling these two sutures apart at 10 mm/min
on an
Instron Tensile Tester, a procedure that is well established for the
evaluation of heart
valve components. The sutures employed were 10-0 nylon sutures. Strength at
rupture
of the gel is calculated, together with elongation at break and elastic
modulus.
Modulus and stress at failure from suture pull-out measurments showed that
maxima
were reached at specific collagen amine to ASI group ratios (Figure 11).
The hydrogels prepared as described in Examples 4 and 5 have high optical
transmission and very low light scattering, comparable to the human cornea, as
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measured with a custom-built instrument that measures transmission and scatter
[Priest and Munger, Invest. Ophthalmol. Vis. Sci. 39: 5352 - 5361 (1998)].
Back
scattering and transmission of light across the visible region for hydrogels
prepared as
in Example 4 showed excellent performance except at high terpolymer
concentrations
(high collagen amine to terpolymer ASI ratios, Fig. 12). Similarly, a
thermogel (free
of cross-linking synthetic polymer) had a very low transmission and high back
scattering (Fig. 12). The hydrogels described in Example 5 also showed
excellent
performance in this analysis as shown in Fig. 12 (1:1 ratio of collagen to
terpolymer-
pentapeptide is represented by the solid squares).
In contrast, collagen-pNiPAAm homopolymer gels (as described in Example 1; 1.0
0.7 to 1.0 : 2.0 wt/wt) were opaque at 37°C. In addition, both collagen
and pNiPAAm
extracted out from this hydrogel into PBS (over 50 wt % loss in 48 h).
EXAMPLE 7: IN VIVO TESTING OF VARIOUS BIO-SYNTHETIC
MATRICES
Hydrogels formed as described in Examples 1, 4 and 5 were moulded to form
artificial
corneas and implanted into the eyes of pigs (Figure 2).
As ira vivo corneal implants, the gels from Example 1 exude white residue
after 5 to 6
days implanted in pigs' eyes.
The hydrogel prepared from 4% collagen and pentapeptide-terpolyrner as
described in
Example 5 demonstrated good biocompatibility as did the collagen-terpolymer
hydrogel prepared as described in Example 4. More rapid, complete epithelial
cell
overgrowth and formation of multiple layers occurred when the former hydrogel
was
used, as compared to collagen- terpolymer hydrogel which showed slower, less
contiguous, epithelial cell growth, without formation of multiple layers.
In vivo, confocal microscope images of full thickness hydrogel prepared from
collagen
and the pentapeptide-terpolymer (from Example 5; final concentration: collagen
2.3
wt %; terpolymer + pentapeptide 1.6 wt %) and implanted into a pig's eye
showed that
epithelium cells grew over this matrix and stratified. A basement membrane was
regenerated and hemidesmosomes, indicating a stably anchored epithelium, were
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present. Stromal cells were found to spread inside the matrix after only three
weeks.
The implants became touch sensitive within 3 weeks of implantation (Cochet-
Bonnet
Aesthesiometer) indicating functional nerve in-growth (Figures 4 and 14).
Nerve in-
growth was also observed directly by confocal microscopy and histology. No
clinical
signs of adverse inflammation or immune reaction were observed over an 8 week
period following implantation. See Figures 2, 3 and 5 - 7.
In more detail:
Figure 5 shows morphological and biochemical assessment of a section through
the
pig cornea at 3 weeks post-implantation (A) picro-sinus stain for collagen and
(B)
H&E stain for cells. Figure 7 shows (A) a section through the pig cornea at 3
weeks
1 S post-implantation, stained with picro-sirius red, which demonstrates the
stromal-
implant interface (arrowheads). The implant surface has been re-covered by a .
stratified epithelium. (B) a similar section at 8 weeks post-implantation.
Stromal cells
have moved into the implant and the implant appears to have been replaced by
tissue
sub-epithelially (arrows). (C) a higher magnification of the epithelium (H & E
stained)
showing the regenerated basement membrane (arrow). (D) a corresponding section
stained with anti-type VII collagen antibody that recognizes hemidesmosomes
attached to the basement membrane (arrow). (E) the hernidesmosomes (arrows)
attached to the underlying basement membranes are clearly visualized by
transmission
electron microscopy (TEM). (F) a flat mount of the pig cornea showing nerves
(arrowheads) within the implant, stained with an anti-neurofilament antibody.
Figure 3 shows whole mount confocal microscopic images of pigs corneas at 6
weeks
post-operation showing a regenerated corneal epithelium and basement membrane
on
the surface of the implant. Irt vitro nerve growth patterns within the
collagen-
terpolymer composite and within the underlying host stroma are shown, as are
in-
growing stromal cells.
Restoration of touch sensitivity was rapid (< 14 days post-operative) in
comparison
with minimal restoration in the transplanted allograft over the same time
period for an
additional six animals that received allografts of donor pig corneas of
similar
dimensions (Figure 14).
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EXAMPLE 8: DEPOSITION OF COLLAGEN TERPOLYMER MATRICES
INTO RODENT BRAIN
Following euthanasia, the whole brain of each mouse or rat used was excised
and
placed within a sterotaxic frame. Either two microlitres (2 ml) or three
microlitres (3
ml) of hydrogel containing collagen, terpolymer-pentapeptide at either 0.33%
collagen
- 0.23% terpolyner or 0.63% collagen-0.44% terpolymer was injected over a
period
of 6 to 10 min, respectively, into each individual mouse brain, at the
following
coordinates: 0.3 mm from bregma, 3.0 mm deep and 2.0 mm from the midline. For
rats, four to six microlitres of hydrogel was inj ected over 10 min. into each
brain, at
0.7-0.8 mm from bregma, 6 mm deep and 4 mm from the midline. The hydrogel
samples were mixed with Coomassie blue dye for visualization.
Results indicate successful direct, precise delivery of small amount of the
hydrogel
into the stratum of the brain, in these samples (Figures 9 and 10). This
suggests that it
is possible to use the hydrogel as a delivery system for cells or drugs into
specific
locations at very small volumes.
EXAMPLE 9: IN VIVO TESTING OFA HYDROGEL COMPRISINGA
BIOACTIT~E AGENT
Sterile hydrogels prepared as described in Example 5 were thoroughly rinsed in
PBS
before implantation. Following the Association for Research in Vision and
Ophthalmology guidelines for animal use, each tissue engineered (TE) corneal
matrix
(5.5 mm in diameter and 200 ~ 50 ~m thick) was implanted into the right cornea
of a
Yucatan micro-pig (Charles River Wiga, Sulzbach, Germany) (see Fig. 15A-C).
Contralateral unoperated corneas served as controls. Under general
anaesthesia, a
partial-thickness 5.0 mm diameter circular incision was made using a Barraquer
trephine (Geuder, Heidelberg, Germany). Host corneal tissue was removed and
replaced with an implant 0.50 mm larger in diameter to allow adequate wound
apposition between the graft and host tissue. After surgery, an amniotic
membrane
was sutured over the entire corneal surface for one week to keep implants in
place. In
sutured samples, implants were sutured into the host tissue using 8
interrupted 10-0
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nylon sutures. Post-operative medication consisted of dexamethasone (qid) and
gentamycin (qid) for 21 days. n= 3 pigs with sutures and 3 without sutures.
Follow-ups were performed daily on each pig up to 7 days post-operative, and
then
weekly. Examinations included slit-lamp examination to ensure corneas were
optically clear, sodium fluorescein staining to assess epithelial integrity
and barner
function (Josephson, J. E. & Caffery, B. E. (1988) Invest Oplathalmol Vis Sci
29,
1096-9), measurements of intraoculax pressure to ensure that corneas were not
blocking aqueous humour flow, and in vivo confocal microscopic examination
(ConfoScan3, Nidek, Erlangen, Germany) to assess cell and nerve in-growth.
Corneal
touch sensitivity was measured using a Cochet-Bonnet esthesiometer (Handaya
Co.,
Tokyo, Japan) at five points within the implant area of each cornea (four
peripheral,
one central) as previously described (Millodot, M. (1984) Ophthalmic Physiol
Opt 4,
305-18). Animals that received allografts of pig donor corneas were also
similarly
evaluated.
For immunohistochemistry and histopathological examination, tissues and
constructs
were fixed in 4% PFA in O.1M PBS. For nerve immunolocalization, flat mounts
were
permeabilized with a detergent cocktail (Brugge, J. S. & Erikson, R. L. (1977)
Nature
269, 346-8) (150 mM NaCI, 1 mM ethylenediamine tetraacetic acid, 50 mM Tris,
1%
Nonidet P-40, 0.5% sodium deoxycholate, 0.1% sodium dodecyl sulphate), blocked
for non-specific staining with 4% foetal calf serum in PBS and incubated in
anti-
neurofilament 200 antibody (Sigma, Oakville, Canada). They were then incubated
with FITC or Cy3-conjugated secondary antibodies (Sigma; Amersham, Baie
D'Urfe,
Canada, respectively) and visualization by confocal microscopy.
For histology and further immunohistochemistry, samples were processed,
paraffin
embedded and sectioned. Sections were stained with haematoxylin and eosin
(H&E)
for histopathological examination (Jumblatt, M. M. & Neufeld, A. H. (1983)
Invest
Ophthalnaol Vis Sci 24, 1139-43). Immunofluorescence was performed as
described
above on deparaffinized sections for expression of type VII collagen (Sigma,
Munich,
Germany), a hemidesmosome marker (Gipson, I. I~., Spurr-Michaud, S. J. &
Tisdale,
A. S. (1988) Dev Biol 126, 253-62). hnmunohistochemical staining using
peroxidase-
diaminobenzidine (DAB) visualization was performed with the following: with
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AEl/AE3 antibody (Chemicon, Temecula, CA, USA) for epithelial markers, anti-
vimentin antibody (Ruche, Laval, Canada) for stromal fibroblasts, anti-smooth
muscle
actin antibody, lA4 (Cell Marque, Austin, TX) for activated stromal
fibroblasts
(myofibroblasts) and SP1-D8 antibody (DHSS, Iowa, USA) for procollagen 1
synthesis (to localize sites of de yaovo collagen synthesis). CD15 and CD45
staining
for immune cells (Becton-Dickinson, Oakville, Canada) was performed using the
ARK peroxidase kit (DAKO, Mississauga, Canada) to pre-conjugate the primary
antibodies to their respective secondary antibodies and peroxidase for
visualization.
For anti-vimentin, anti-smooth muscle actin and SP8-D1 antibodies, antigen
retrieval
was preformed by pre-treating with Proteinase-K (2 mg/ml) for 30 min at
37°C prior
to incubation in primary antibody. Illex eu~opaeus aggultinin (IDEA) lectin
staining
was used to visualize tear film mucin deposition (Shatos, M. A., Rios, J. D.,
Tepavcevic, V., Kano, H., Hodges, R. & Dartt, D. A. (2001) Invest Ophthalmol
Vis
Sci 42, 1455-64). Samples were incubated with biotinylated UEA (Sigma), then
reacted with avidin-horseradish peroxidase and visualized with DAB. For
transmission electron microscopy (TEM), all samples were treated in
conventional
fixative, stain and potting resin (Karnovsky's, Os04, uranyl acetate, epoxy).
No adverse inflammatory or immune reaction was observed by clinical
examination
after implantation of either bio-synthetic matrices or pig corneas. Epithelial
cell in-
growth over the implant was complete by 4 days post-operative. By one week,
the
regenerated epithelium showed exclusion of sodium fluorescein dye, indicating
that
the epithelium was intact and had re-established barner properties.
Intraocular
pressures were between 10 and 14 mm mercury (Hg) pre-operatively, and 10-16 mm
Hg post-operatively throughout the study period of up to 6 weeks, showing that
the
implants did not block flow of aqueous humour within the eye. Implants
remained
optically clear (slit-lamp biomicroscopy) and epithelial re-stratification was
observed
in all animals at 3 weeks post-surgery. Clinical in vivo confocal microscopy
of the
implanted stromal matrices at 3 weeks post-surgery showed a regenerated
epithelium
(Fig. 15D), newly in-grown nerves (Fig. 15G), and stromal (Fig. 15J) and
endothelial
cells (Fig. 15M) with cellular morphology mimicking that of un-operated
controls
(Fig. 15F,I,L,O). Epithelial and endothelial cell morphology in the allografts
(Fig.
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15E,I~ was similar to that of controls. Sub-epithelial and stromal nerves were
not
observed in the allografts at 3 weeks post-surgery (Fig. 15H,K).
In more detail, Figure 15 shows:
(A-C) lamellar keratoplasty (LKP) procedure on a Yucatan micro-pig. (A): A
trephine
is used to cut a circular incision of pre-set depth (250 ~.m) into the cornea.
The
existing corneal layers are removed and (B) are replaced with a bio-synthetic
matrix
implant (arrow, 250 ~,m in thickness), which is sutured in place (C). Sutures
are
indicated by arrowheads.
(D-O) In vivo confocal microscopy of implanted bio-synthetic matrix. (D):
confocal
image showing regenerated corneal epithelium on the surface of the implant.
The
corresponding allograft control (E) contains donor epithelium, while the un-
operated
control (F) has an intact epithelium. (G): Regenerated nerves (arrowheads) axe
present
at the interface between implant and overlying regenerated epithelium. These
correspond to the sub-epithelial nerves in the un-operated control (I). In the
allograft
(H), however, sub-epithelial nerves are absent. (J-L): Stromal cells and
branching
nerve bundle (arrowhead) deeper within the underlying stroma of corneas with
implant (J), allograft (K) and in a corresponding region in the control (L).
(M-O): The
endothelium in corneas with implant (IVI), allograft (I~ and un-operated
controls (O)
are intact and show similar morphology.
Histological sections through corneas with implants showed a distinct but
smooth,
implant-host tissue interface (Fig. 16A) that resembled that of control
corneas that
received allografts (Fig. 16B). In both corneas with implants or allografts,
the
regenerated epithelium was stratified. Detailed examination showed a fully
differentiated epithelium that was positively stained by AE1/AE3 antibody
markers
(Fig. 16D,E), overlying a regenerated basement membrane that was positive for
Type
VII collagen, a marker for hemidesmosomes at the basement membrane-epithelium
interface (Fig. 16G,H). TEM observations indicated morphology consistent with
the
presence of hemidesmosomes (Fig. 16J,K). In the implants, neurofilament-
positive in-
growing nerves had begun to re-establish a sub-epithelial network and showed
extension into the epithelial cells (Fig. 16M). However, no sub-epithelial
nerves were
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located in the allografted corneas (Fig. 16I~. The tear film was restored in
corneas
with implants (Fig. 16P) as in the allograft (Fig. 16Q).
In more detail, Figure 16 shows: post-surgical corneal regeneration.
(A-F) H&E stained sections are shown. Stromal cells are present in the implant
(A)
and the allograft control (B), and both appear to be seamlessly integrated
into the host.
(symbols are as follows: e, epithelium; i, implant; g, allograft; s, stroma).
(C):
Unoperated control. The regenerated epithelium of the implant (D) and donor
epithelium of the allograft control (E) expressed cytokeratin differentiation
markers,
similar to the un-operated control (F).
(G-I): Immunolocalization of type VII collagen, a marker for hemidesmosomes,
at the
epithelium-implant interface (arrows) in the implant (G), allograft (H) and
control (I).
(J-L): TEM of epithelium-implant interface. Hemidesmosome plaques (arrowheads)
and anchoring fibrils (arrows) have formed within the bio-synthetic matrix
between
the epithelial cells and underlying implant (J), emulating the structures
normally found
at the epithelial-stromal interface as demonstrated in the allograft (I~) and
control (L).
Flat mount of cornea showing nerve fibres (arrows) within the implant (M), and
un-
operated control (O) but absent in the allograft (I~, stained with an anti-
neurofilament
antibody. UEA binding (arrowheads) to the epithelial surface on the implant
(P), and
allograft (Q) indicate restoration of the tear film in all cases. Un-operated
control (R).
Immunohistochemistry results indicated that cells within both implant and
allograft
were synthesizing procollagen I. However, more procollagen synthesis occurred
in the
allografts as indicated by the more intense staining in allografts compared to
implants
(Fig. 17G,H). Both allografts and implants had stromal cells that were
vimentin
positive (Fig. 17A,B), indicating a fibroblastic phenotype. Both also showed
smooth
muscle actin staining and therefore the presence of activated stromal
fibroblasts,
although the implants showed fewer positive cells than the allografts (Fig.
17D,E).
In more detail, Figure 17 shows: implant-host integration at 6 weeks post
surgery.
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(A-C): Staining for procollagen type I. Positive staining is observed in
matrix of both
the implanted biosynthetic matrix (A) and the allograft control (B) indicating
sites of
new collagen deposition. Unoperated control (C) has no new collagen synthesis.
(D-F): Staining for vimentin throughout stroma identifies stromal fibroblasts.
Staining throughout the implanted biosynthetic matrix (D) demonstrates cell
invasion.
Cells may also been seen within the implanted allograft (E) and throughout the
un-
operated control (F)
(G-I): Smooth muscle actin staining indicates activated myofibroblasts and the
potential for scarnng. In the biosynthetic matrix implant (G), staining is
occasionally
present in the biosynthetic matrix, but is not found in the host stroma, nor
in the
transition zone between host and implant. Positive staining in the allograft
implanted
cornea (H) is identified both in the allograft, and the transition zone, but
not in the
intact host stroma. (I): Un-operated control.
Corneal touch sensitivity measured at S points on the corneal implant in 3
pigs pre-
and post-operatively using a Cochet-Bonnet esthesiometer, showed a dramatic
drop in
touch sensitivity after surgery. However, recovery occurred between 7 and 14
days
and by 21 days post-operative, sensitivity had returned to pre-operative
levels (Fig. 18;
All groups, n = 3. ~P < 0.01 by repeated measures ANOVA with Tukey 2-way
comparisons). Touch sensitivity returned at the same rate and to the same
plateau
level at all peripheral and central points tested on the implant. In control
animals that
had received donor corneal allografts, however, the conieas remained
anaesthetic over
the six-week period (Fig. 18).
Implants recovered after 6 weeks in vivo were examined by infrared
spectroscopy
(Midac M, FTIR spectrometer, ZnSe beam condenser and diamond cell) and clearly
indicated the presence of the terpolyrner.
EXAMPLE 1 D: PREPARATION OF A SYNTHETIC CO-POLYMER
A poly(DMAA-co-ASI) co-polymer was synthesised by co-polymerization of the
monomers: N,N-dimethyl acrylamide, (DMAA) and N-acryloxysuccinimide (ASI).
The feed molar ratio was 95:5 (DMAA: ASI). The free-radical initiator AIBN and
the
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solvent, dioxane, were nitrogen purged prior to use and polymerisation
reaction
proceeded at 70°C for 24h.
After purification by repeated precipitation to remove traces of homopolymer,
the
composition of the synthesized copolymer (70% yield) was found to be 94.8:5.2
(molar ratio) by proton NMR. Molecular mass (M") was determined at 4.3 x 104,
by
aqueous GPC. Polydispersity (PD; MW/Mn) =1.70 was also determined by GPC.
EXAMPLE 11: PREPARATION OF A COLLAGEN CO POLYMER HYDROGEL
A cross-linked collagen-co-polymer hydrogel was prepared by mixing neutralized
S%
bovine collagen (1.0 ml) with the synthetic co-polymer prepared in Example 9
[0.2m1
(200 mglml in D-PBS)] by syringe mixing. After careful syringe pumping to
produce
a homogeneous, bubble-free solution, aliquots were injected into plastic,
contact lens
moulds and incubated at room temperature for 24 hours to allow reaction of the
collagen NH2 groups with ASI groups in the co-polymer as well as the slower
hydrolysis of residual ASI groups to AAc groups.
The moulded samples were then incubated at 37°C for 24 hours in a 100%
humidity
environment to provide the final hydrogel. At gelation, the hydrogel contained
94.8%
water, 2.9% collagen and 2.3% synthetic co-polymer. Matrices were moulded to
have
a final thickness between either 150 - 200 p,m or 500 - 600 Vim. Each
resulting
hydrogel matrix was removed from its mould under PBS solution and subsequently
immersed in PBS containing 1% chloroform and 0.5% glycine. This wash step
removed N-hydroxysuccinimide produced in the cross-linking reaction and
terminated any residual ASI groups in the matrix, by conversion to acrylic
acid
groups.
Succinimide residues left in the gels prepared from collagen and copolymer
were
below the IR detection limit after washing.
EXAMPLE 12: PHYSICAL PROPERTIES OF COLLAGEN CO-POLYMER
HYl~ROGEL
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Light back scattering and light transmission across the visible region and
with white
light for hydrogels prepared as in Example 10 as a function of collagen amine
to
copolymer ASI ratios is shown in Fig 13A and B.
The copolymer from Example 10 and its hydrogels had no detectable cloud point
(LCST) at up to 60°C.
EXAMPLE 13: BIOLOGICAL PROPERTIES OF VARIOUS HYDROGELS
11.1 Biocompatibility
Three 12 mm diameter and 650 pm thick discs each of collagen-poly(DMAA-co-
ASI), collagen-poly(NiPAAm-co-AAc-co-ASI)-pentapeptide and 3% collagen
hydrogels were soaked for 30 minutes in PBS. They were each laid onto a 12 mm
membrane insert commercially available for a culture dish and adhered to the
membrane with a thin coating of gelatin. After drying for 10 minutes, 1 X 104
human
corneal epithelial cells (HCEC) cells suspended a serum-free medium containing
epidermal growth factor (Keratinocyte Serum-Free Medium (KSFM; Life
Technologies, Burlington, Canada)) were added to the top of the gels, and KSFM
without cells was added to the underlying well. Cultures were incubated at
37°C with
5% COa.
Within 12 hours the cells had adhered to the surface of the matrix in all
samples.
Medium was changed every second day with KSFM added to the inserts and to the
outside wells. HCEC were grown to confluence on the gels and reached
confluence
on the same day (5 days). The medium in the inserts and surrounding wells was
replaced by a serum-containing medium (modified SHEM medium (Jumblatt, M. M.
& Neufeld, A. H. (1983) Invest Ophthalnaol Vis Sci 24, 1139-43)). After 2 more
days,
the medium was removed from the inserts, and the volume of SHEM in the
underlying
wells reduced to 0.5 ml. The epithelium was allowed to stratify for a further
7 days
and the layer of cells visualized.
After 7 days, the membranes were fixed in 4% paraformaldehyde in PBS for 30
minutes at 4°C. Samples were prepared for cryosectioning by
equilibration in 30%
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sucrose in PBS followed by flash freezing in a 1:1 mixture of 30% sucrose in
PBS and
OCT. These were cryosectioned to 13 ~,m and the structure visualized by HandE
staining. The number of cell layers in the stratified epithelium was
determined by
counting nuclei and identifying cell borders. The collagen thermogel attained
an
epithelial thickness of approximately 2 cells, which contrasts poorly with the
human
corneal epithelium that contains between about 5 and 7 cell layers. HCEC
cultured
and induced to stratify on collagen-p(DMAA-co-ASI) and collagen-p(NiPAAm-co-
AAc-co-ASI)-pentapeptide, however, resulted in an epithelium about 4.5 cell
layers
thick that included apparently keratinized outer layers suggesting appropriate
differentiation of the epithelium (Fig. 20).
11.2. Inneyvation of HydYOgel
Twelve millimeter diameter and 650 ~m thick discs each of collagen p(DMAA-co-
ASI), collagen p(NiPAAm-co-AAc-co-ASI)-pentapeptide and 3% collagen thermogel
were soaked for 30 minutes in PBS. Discs were laid in a 6 cm culture dish, and
four 1
mm holes bored through each. The holes were filled a third of the way up with
a plug
of 0.3% collagen cross-linked with glutaraldehyde and quenched with glycine.
After
10 minutes, dorsal root ganglions from E8 chicks were dipped in the same
collagen
mixture and placed in the holes. The holes were filled the rest of the way
with cross-
linked collagen, and allowed to set for 30 minutes at 37°C. Cultures
were grown for 4
days in KSFM supplemented with B27, N2, and 1 nM retinoic acid for 4 days and
neurite extension monitored by brightfield microscopy. The innervated discs
were
fixed in 4% paraformaldehyde in PBS for 30 minutes room temperature, stained
for
NF200 immunoreactivity, and visualized by immunofluorescence. Localization was
visualized on the surface and in the centre of the polymer disc. While there
was some
neurite extension over the surface of the collagen thermogel, none could be
seen
extending into the thermogel itself. In the hydrogels, neurites could be seen
extending
into the matrix. As well, in both the hydrogels extensive innervation could be
seen
over the surface of the matrix suggesting a better surface innervation than
occurred
with the collagen thermogel (Fig. 19; A depicts the collagen thermogel, B
depicts the
collagen-p(NiPAAm-co-AAc-co-ASI)-pentapeptide hydrogel and C depicts the
collagen-p(DMAA-co-ASI) hydrogel). The left column represents
immunofluorescent
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visualizations of the middle of the polymers stained for the nerve
neurofilament
marker - NF200. The middle column depicts a brightfield view of the surface of
the
polymer with the neurites extending from the ganglion source. The right column
represents an immunofluorescent visualization of the same surface view of the
polymer stained for NF200 immuno-reactivity. The arrows indicate neurites
extending
in the middle of the polymer. The intact human cornea demonstrates both sub
epithelial surface and deep nerves suggesting that these matrices are both
biocompatible to nerves and can emulate the corneal stroma.
The invention being thus described, it will be obvious that the same may be
varied in
many ways. Such variations are not to be regarded as a departure from the
spirit and
scope of the invention, and all such modifications as would be obvious to one
skilled
in the art are intended to be included within the scope of the following
claims.
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