Note: Descriptions are shown in the official language in which they were submitted.
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METHOD OF PRODUCTION OF NEW POLYMERIC MATERIAL
The present invention relates to a material as
interpenetrating polymer network (IPN) associating a gel with a
co-network of synthetic polymer and protein as well as a method
of the manufacture of such a material. In particular, the
invention relates to a material as interpenetrating polymer
network associating a fibrin gel with a co-network of
polyvinylalcohol and albumin thank to methacrylate bridges,
with improved biodegradability properties.
The invention also refers to use of this material in
biomedical applications, such as drug delivery vehicle,
encapsulation material, wound dressing, scaffold for tissue
engineering, hemostatic dressing, surgical dressing, as a
coating for medical devices such as stents, heart valves,
catheters, vascular prosthetic filters etc. Such materials may
also be used for eukaryotic cell culture, or as support for
active molecules such as healing agents, growth factors,
antibiotics, bactericidal, bacteriostatic or enzymes.
The definition of a gel corresponds to a nonfluid
colloidal network or polymer network that is expanded
throughout its whole volume by a fluid.
A gel can be made of:
(i) a covalent polymer network, e.g., a network formed by
crosslinking polymer chains or by nonlinear polymerization.
This type of gel is usually called "chemical gel". A chemical
gel is a supramolecular assembly whose molecules are associated
by high energy bonds (covalent bonds). The stability of this
assembly is very large. These gels have improved chemical
stability, the only way to degrade them is to destroy the
covalent network bonding.
(ii) a polymer network formed through the physical aggregation
or supramolecular assembly of polymer chains, caused by low
energy bonds such as hydrogen bonds or Van der Waals bonds,
crystallization, helix formation, complexation, that results in
regions of local order acting as the network junction points.
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This type of gel is usually called "physical gel". The
stability of this assembly is associated with a specific range
of physico-chemical conditions (pH, concentration of molecules,
temperature, solvent quality, ionic strength, etc). Outside
this range, the mixture is liquid. Thus, a change in the
parameters of the medium can lead to the destruction of the
scaffold and induce a gel/sol transition. Some "physical" gels
are well known such as ones obtained from macromolecules or
polymers of natural origin like proteins or polysaccharides.
Gelatin, alginate, pectin or starch may be mentioned as example
of macromolecules capable of forming physical gels.
Obtaining a material in the form of a hydrogel is a real
need especially in the medical field. Hydrogels are insoluble
three-dimensional (3-D) networks of hydrophilic homopolymers,
co-polymers or macromers with a swelling capacity in aqueous
environments. The hydrogels have the property of being very
elastic and rich in solvent. Their size and shape may vary
according to environmental stimuli such as pH, temperature,
solvent composition, ionic strength, electric and magnetic
fields. Hydrogels obtained from biological origin, usually
called biogels, are hydrophilic networks that hold large
amounts of water without dissolving and thus have similar
physical characteristics to those of biological soft tissue.
Furthermore, hydrogels synthesized from natural macromers
generally induce fewer immunogenic reactions as they are
produced from basic molecules already used by the body.
Synthetic derivative hydrogels have the advantage of
controlling the gel properties, which can be tailored through
the chemical nature of the synthetic precursors, their specific
molecular weight and cross-linking densities for example.
Biological molecules may be incorporated into synthetic
hydrogel networks to improve cell attachment. Hydrogels have a
high permeability to oxygen, nutrients and other water-soluble
metabolites and are well adapted for growth of tissue
structures. Hydrogels are attractive for drug delivery due to
their good compatibility with hydrophilic and macromolecular
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compounds (proteins, polysaccharides, oligonucleotides), their
biocompatibility and possibility to regulate the distribution
of drugs through the control of their swelling, their
crosslinking density or degradation for example. Hydrogels are
interesting for a localized diffusion since they can be formed
in situ and thus adhere to and comply with the target tissues.
This use for local delivery of active ingredients appears to
act synergistically with the barrier effect of the hydrogels.
Hydrogels have also been investigated as potential materials
for tissue regeneration or skin substitute which use constitute
one very promising application.
Fibrin is made from fibrinogen, a 340 kDa glycoprotein of
hepatic origin. Fibrinogen is an anionic polyelectrolyte under
physiological conditions and therefore cannot join or
aggregate. The hydrolysis of fibrinogen by thrombin, a serum
protease, leads to the release of small peptide fragments and
obtaining fibrin. Intermolecular interactions become possible,
through low energy forces, thereby creating a physical network
fibrin gel.
Fibrin gels physiologically form the basis of wound repair
and haemostasis. Fibrin gels are used as surgical glue and are
attractive as biocompatible and biodegradable materials.
However, at physiological concentrations (4.8 mg mL-1), fibrin
hydrogels are soft, cannot be easily handled and do not self-
stand.
Fibrin matrices with improved mechanical properties have
been designed by increasing fibrinogen Or thrombin
concentrations. However, while more rigid materials are needed,
the physiological-like organization of fibrin networks, which
is perfectly adapted to the various cells implied in
physiological processes such as wound healing, must be
maintained. Another approach to improving self-supported fibrin
gels may be used, consisting of the combination of hydrogel
with a synthetic polymer. A material showing the strength and
elasticity necessary to be easily handled would thus be
obtained, while maintaining all the properties of a fibrin gel.
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However, the macromolecule blend is generally
thermodynamically unfavorable, and macroscopic phase separation
generally occurs. Thus, an original pathway was developed to
obtain an easily handled fibrin-based material: fibrin gel is
associated with a synthetic polymer network inside
interpenetrating polymer network (IPN) architecture. IPN are
defined as a combination of two or more polymer networks, at
least partially interlaced, synthesized in juxtaposition. The
entanglement of two polymers, leads to forced miscibility,
compared with usual blends.
W02010/058132 discloses a material in the form of an
interpenetrating network involving fibrin gel with a network of
synthetic polymer (obtained by polymerization and crosslinking
monomer such as a polyethylene oxide-functionalized vinyl,
acrylate, methacrylate or allyl, optionally with a second
monomer derivative such as a functionalized polyethylene
glycol). While the mechanical properties of such a material
appear good, their biodegradability and resorption in situ is
limited and not adequate for all applications.
In Acta Biomaterialia 7 (2011) 2418-2427, are disclosed,
materials made of a fibrin gel interpenetrated with
polyethylene oxide dimethacrylate polymerized/crosslinked
network. It is stated that the mechanical properties of the
fibrin gel can thus be improved by the presence of the
polyethylene oxide network. The teaching of this document is
basically the same as W02010/058132.
Kundu et al (Biomaterialia Acta 8 (2012) 1720-1729)
disclose hydrogels comprising a network of PVA functionalized
(via function methacrylate) and associated with a silk protein
gel. In this document, the silk protein does not participate in
the crosslinking of the PVA polymer and the biodegradability is
also limited.
Fibrin has often been associated with type I collagen to
improve the mechanical properties of collagen. Alginate has
been also included in fibrin-based hydrogels. Fibrin-alginate
IPN shows dynamic mechanical properties that can be employed to
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enhance tissue development relative to a single hydrogel. This
association provides a dynamic mechanical environment, which
facilitates the growth of organized cell clusters. However, the
gels formed from organic molecules are soft materials and while
5 biodegradable. Their main drawback is a poor mechanical
resistance.
Furthermore, gels and materials of the prior art can
hardly keep their properties after being dehydrated and
rehydrated.
There is thus a real need for a material that can be easily
handled, which has the properties of a gel dimensionally
stable, improved degradation ability, that can be stored
dehydrated and can be fully rehydrated upon extemporaneous use.
There is also a need for a material that can contain living
cell embedded within it and can sustain the growth of those
embedded cell and that can also be frozen while preserving the
viability of the encapsulated cells.
As stated earlier, there is a need to have a material gel
which would combine the properties of physical and chemical
gels. Such a material would be stable and maintains its
structural integrity and thus its properties throughout the
duration of its handling and use, but would be degradable or
labile within biological environment, i.e. can be destabilized
and degraded when desired. To date, most of the easily handled
gels lack lability and biodegradability within the biological
environment, i.e. after implantation in the body. This is of
particular importance in the cases of patients victims of
wounds and burns in which cases it is important to ensure that
the material creates the least possible reaction from the host.
The present invention is specifically intended to meet this
need by providing a method for preparing a material as
interpenetrating polymer network (IPN) associating a gel with a
co-network of functionalized synthetic polymer crosslinked with
a functionalized protein, comprising the steps of:
i) preparing a first mixture by introducing into a buffer
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a. a gel forming solution or a gel forming precursor
solution,
b. a synthetic polymer selected in the group consisting of
polyvinyl alcohol (PVA), polyglycolic acid (PGA), polylactic
acid (PLA), polycaprolactone (PCL), poly(N-vinylpyrolidone),
poly(2-hydroxy ethyl methacrylate) (PHEMA), polyethyleneoxide
(PEO), and derivatives thereof, wherein the polymer is
functionalized with grafted chemical groups X, X being selected
in the group consisting of acrylate, methacrylate, vinyl, allyl
and styrene derivatives,
c. a protein functionalized with grafted chemical groups
X, wherein X is has the same meaning than in b)
d. a polymerization initiator,
ii) preparing a second reaction mixture by optionally adding a
gelification activator of the formation of the gel to the first
mixture prepared in i),
iii) incubating the reaction mixture obtained in i) or in ii)
at a temperature and during a time sufficient to allow
formation of the gel, and
iv) performing a polymerizing and crosslinking of the
functionalized synthetic polymer with the functionalized
protein.
The method of the invention allows preparing a material in
the form of interpenetrating polymer network (IPN) which is
homogeneous, self-supported, easy to handle, stable over time,
dimensionally stable and biodegradable.
Gelation on the one hand and polymerization and cross-linking
on the other hand can take place simultaneously or
sequentially.
In the remaining part of the description, the synthetic
polymer which is functionalized with grafted chemical groups X,
X being selected in the group consisting of acrylate,
methacrylate, vinyl, allyl and styrene is called functionalized
synthetic polymer.
Similarly, the protein which is functionalized with
grafted chemical groups X, wherein X is selected in the group
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consisting of acrylate, methacrylate, vinyl, ally' and styrene
is called functionalized protein.
In order to ensure the establishment of the co-network by
covalent bonds between the polymer and the protein, the
functionalized synthetic polymer and the functionalized protein
each comprise at least two grafted chemical groups X. The
functionalized protein thus comprises at least two grafted
chemical group X, preferably at least three, more preferably at
least 4, even more preferably at least five. The maximal number
of grafted chemical groups X on the functionalized protein is
not critical and is dependent on the number and accessibility
of amino acid side chain available for grafting. The highest
number of grafted chemical groups X, the highest the strength
of the link between the protein and the synthetic polymer.
Similarly the functionalized synthetic polymer thus
comprises at least two grafted chemical group X, preferably at
least three chemical groups X, more preferably at least 4
chemical groups X, even more preferably at least five chemical
group X. Because of the chain length of the synthetic polymer,
the number of grafted chemical groups X can be up to 20,
particularly up to 30, particularly up to 40, even more
particularly up to 50 and even more particularly up to 60. The
number of grafted chemical groups X will be dependent upon the
molar weight of the polymer and the number of repeat units.
Typically in the case of PVA, particularly PVA 16000, used as
synthetic polymer, the number of grafted chemical groups X can
be such as 1 group X is grafted for 5 to 100 repeat units,
particularly 1 group X for 30 to 70 repeat units, more
particularly 1 group X for about 50 repeat units. In the case
of PVA 16000 the number of groups X, particularly methacrylate
groups, is about 5 to 20, particularly 5 to 10, even more
particularly around 7 groups per molecule.
In the context of the present invention, by "buffer" is
meant an aqueous solution whose water content is at least 30%
by weight, up to 99% by total weight of solvent and maintains
approximately the same pH despite the addition of small amounts
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of an acid, a base or a dilution. The water content of the
buffer is in particular 70 to 98% by total weight of the
solvent.
Groups X grafted on the synthetic polymer and Groups X
grafted on the protein may be the same or different provided
that they can react to create a covalent bond crosslinking the
protein and the synthetic polymer.
The buffer may further comprise other solvents selected
from the group comprising methanol, ethanol, pyridine, acetone,
acetic acid, DMSO, dichloromethane. The buffer may also
comprise cosolvants such as glycerol, sorbitol, mannitol or any
other polyol.
For example buffer includes phosphate buffers, Hepes,
Tris, sodium barbital, Tris-maleate.
Throughout the duration of the process, the pH of the
aqueous medium is maintained between 6 and 8 by means of a
buffer solution, for example a trihydroxyaminomethane solution
(Tris).
The gel can be a physical gel or a chemical gel. In a
preferred embodiment it is a physical gel.
The gel forming solution or gel forming precursor solution
can be a solution containing a gel forming polymer able to form
a gel such as polysaccharide polymer or polypeptide polymer
without or with the presence of a gelification activator. As a
matter of example, such a gelification activator maybe thrombin
and calcium when the gel forming precursor solution is
fibrinogen containing solution and in that case the obtained
gel is a physical gel, a fibrin gel.
In one embodiment of the present invention, the physical
gel forming precursor solution is a fibrinogen containing
solution. When a fibrinogen containing solution is used, the
gelification activator of the formation of the physical gel is
selected in the group consisting of thrombin and calcium, used
alone or in combination. Indeed, in the case the gel forming
precursor solution is fibrinogen, the activator is a
combination of thrombin and calcium. If the fibrinogen solution
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already contains calcium, then the activator will be thrombin
alone. In the case of the use of plasma as gel forming
precursor solution, since plasma already contains thrombin, the
activator will be calcium alone.
In an embodiment the physical gel forming precursor
solution containing fibrinogen is serum which naturally
contains fibrinogen.
In another embodiment, the physical gel forming precursor
solution containing fibrinogen is plasma, which contains
fibrinogen.
While plasma does contain thrombin and calcium, it may be
useful to add additional calcium and/or thrombin in the
reaction mixture in order to facilitate the formation of fibrin
gel.
Besides being an autologous fluid which can be collected
from the patient, plasma is also useful as a tank of healing
biomolecules, including proteins (albumins and globulins),
growth factors, and also fibrinogen which when placed in the
presence of thrombin and its cofactor, calcium, may be cleaved
to form a fibrin gel.
Plasma, as blood derivative, presents a real interest in
the field of tissue engineering because of its bioavailability,
because it contains growth factors, and molecules that are
useful for wound healing, such as fibronectin which is a cell
adhesion factor.
Plasma contains a number of proteins capable of forming a
gel network (especially fibrinogen forming fibrin gel). In the
context of the present invention, it is thus possible to
isolate the plasma of a patient in need of care, for
manufacturing a fully autologous material according to the
present invention, such material would participate in wound
healing without causing rejection or secondary effect.
The material manufactures with the use of blood products,
such as plasma, present the advantage of mimicking the tissue
microenvironment.
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Materials based on human proteins stemming from plasma
also have the enormous potential as matrix that can be
colonized by cells and that is biodegradable in situ, ie within
the body, as well as self-supported. A very promising property
5 of the present material according to the present invention is
that it can be of particular interest with patients suffering
from wounds and burns. Indeed, the use of blood plasma from the
patient himself ensures that the material causes the least
possible reaction from the host. That promising solution
10 according to the invention wherein the synthesis of the
material according to the invention is achieved with the use of
the blood plasma of the patient is particularly advantageous.
In the context of the present invention, the expression
"gelification activator" relates to a compound which is able to
and necessary for the formation of the gel. Such a gelification
activator may be an enzyme and/or an ion for example.
In the case of fibrinogen or fibrinogen containing solution,
the formation of fibrin as physical gel is triggered by the
action of thrombin, a serum protease, and calcium, which leads
to the release of small peptide fragments allowing creation of
intermolecular interactions through low energy forces, thereby
creating a physical network fibrin.
Fibrinogen may be for example human fibrinogen, pig
fibrinogen or beef fibrinogen. In a particular embodiment the
fibrinogen is human fibrinogen, obtained from human serum or
from human plasma.
In another embodiment, the physical gel forming precursor
solution may be a polysaccharide such as alginate or pectine
and the gelification activator is calcium.
In a further embodiment, the gel precursor solution may
also be a gelatin solution and the gelification activator is a
crosslinking agent such as glutaraldehyde or the enzyme
transglutaminase which allows the formation of gelatin gel,
which in this case is a chemical gel.
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In an embodiment of the invention, the co-network may
contain from 1 to 99% by weight of functionalized synthetic
polymer and between 99 and 1% by weight of functionalized
protein.
In a particular embodiment, the co-network may contain
from 70 to 30% by weight of functionalized synthetic polymer
and between 30 to 70 % by weight of functionalized protein.
In another particular embodiment, the co-network may
contain from about 50% by weight of functionalized synthetic
polymer and about 50 % by weight of functionalized protein.
In a particular embodiment, the co-network may contain
from 70 to 80 % by weight of functionalized synthetic polymer
and between 30 and 20 % by weight of functionalized protein.
In a particular embodiment, the material according to the
present invention may contain from 5 to 20% in weight of
functionalized synthetic polymer and from 5 to 20 % in weight
of functionalized protein.
In a preferred embodiment the material according to the
present invention contains 5 % in weight of functionalized
synthetic polymer and 5 % in weight of functionalized protein.
In a preferred embodiment the material according to the present
invention contains 10 % in weight of functionalized synthetic
polymer and 10 % in weight of functionalized protein.
In a preferred embodiment the material according to the
present invention contains 3 % in weight of functionalized
synthetic polymer and 7 % in weight of functionalized protein.
In a preferred embodiment the material according to the
present invention contains 7 % in weight of functionalized
synthetic polymer and 3 % in weight of functionalized protein.
The protein to be functionalized may be of any origin such
as human, animal, vegetal or bacterial origin.
In a particular embodiment, the said protein may be
selected from the group consisting of albumin, globulin,
lysozyme. The albumin may be human serum albumin or bovine
serum albumin, or pig serum albumin. In one embodiment the
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serum protein is bovine serum albumin. In one particular
embodiment the serum protein is human serum albumin.
In the context of the present invention, the grafting
chemical groups X on the protein can be achieved by the
covalent attachment of a functional chemical group to a side
chain of amino acid of the protein thus creating a
functionalized protein.
In a particular embodiment, the protein is grafted with
functional groups X that are able to create chemical covalent
bonds or bridges with other groups X when activated by the
polymerization initiator, thus creating a network.
In an embodiment of the invention, the grafted chemical
group X attached to the functionalized protein is the same as
the grafted chemical group X attached to the functionalized
synthetic polymer.
In one preferred embodiment, the grafted chemical group X
attached to the protein is methacrylate. In one further
preferred embodiment, the grafted chemical group attached to
the synthetic polymer is methacrylate.
In a preferred embodiment, the present invention thus
relates to a method of preparing a material as interpenetrating
polymer network (IPN) associating a gel and a co-network of
functionalized synthetic polymer cross-linked with a
functionalized protein. In a preferred embodiment the protein
is albumin (BSA or HSA) functionalized with methacrylate groups
and the synthetic polymer is PVA functionalized with
methacrylate groups. The material according to the invention is
thus formed by a co-network of functionalized protein cross-
linked with a functionalized synthetic polymer, that network is
associated with a gel, particularly a fibrin gel, thus forming
the IPN. There is no crosslinks between the gel, particularly
the fibrin gel, and the co-network of functionalized protein
and functionalized synthetic polymer.
The functionalization of the synthetic polymer by grafting
of groups X can be achieved according to techniques known to
the man skilled in the art. As a matter of example such a
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method for functionalization of PVA with X being methacrylate
is one described in Biomacromolecules, DOI: 10.1021/bm400991k
wherein the functionalization of PVA with methacrylate groups
is described in detail. Particularly, PVA can be solubilized in
DMSO with hydroquinone. 3 mol% 2-ICEMA (with respect to the
hydroxyl function of PVA) can be added in the PVA solution.
Reaction can therefore be mixed for about 4 h at about 60 C and
about 12 h at around 20 C under argon atmosphere. The obtained
solution is then purified by precipitation in acetone at room
temperature. The modified PVA (denoted PVAm), i.e.
functionalized synthetic polymer according the definition of
the present invention, is thus filtered and dried under vacuum
at 30 C for 48 h before dissolution in Tris buffer 250 mM at
pH 7.4.
The functionalization of the protein by grafting of group
X can be achieved according to techniques known to the man
skilled in the art. Protein can be solubilized in an adequate
buffer such acid boric buffer 250 mM at pH 7.4 for example.
Methacrylic acid N-hydroxysuccinimide ester (NHSm) solubilized
in acetone can be added by drop wise up to a 0.7/1 molar ratio
of NHSm/lysine of the protein. Reaction is carried out at room
temperature for about 12 h in dark. Protein modified with
methacrylate group, ie functionalized protein, can be purified
by dialysis (Mcut off 1 kDa) against Tris buffer 50 mM at pH
7.4 in order to eliminate unreacted NHSm. In order to increase
the functionalized protein concentration in solution, it is
possible to lyophilize and to solubilize it at 20% (w/v) in
Tris buffer 50 mM at pH 7.4.
In a preferred embodiment, the method according to the
present invention comprises a method wherein the gel is a
physical fibrin gel and the gel precursor solution is
fibrinogen solution.
In another preferred embodiment, the method according to
the present invention comprises a method wherein the gel is a
physical fibrin gel and the gel precursor solution is plasma
solution containing fibrinogen.
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In a particularly preferred method according to the
present invention the synthetic polymer is PVA and the grafted
functional group X is methacrylate.
In a further preferred method according to the present
invention the protein functionalized with at least one type of
identical grafted chemical group X is a serum protein,
particularly serum albumin functionalized with methacrylate.
In another preferred method according to the present invention
the gelification activator of the formation of the gel is
thrombin and/or calcium.
In a further preferred method according to the present
invention the step of incubating the reaction mixture obtained
in i) or in ii) is achieved at a temperature comprised between
C and 40 C, particularly 37 C.
15 In another embodiment of the invention, in the method
according to the present invention the polymerization initiator
is a photopolymerization initiator selected from the group
consisting of Irgacures, particularly Irgacure 2959.
In a particularly preferred embodiment of the present
20 invention, herein is provided a method for preparing a material
as interpenetrating polymer network (IPN) associating a fibrin
gel with a co-network of functionalized PVA crosslinked with a
functionalized serum albumin, comprising the steps of:
i) preparing a first mixture by introducing into a buffer
a. a fibrinogen containing solution,
b. a functionalized synthetic polymer selected in the
group consisting of polyvinyl alcohol (PVA) grafted with
methacrylate,
c. a functionalized serum albumin protein grafted with
methacrylate
d. a photopolymerization initiator, preferably Irgacure
2959,
ii) preparing a second reaction mixture by adding thrombin
and/or calcium to the first mixture prepared in i),
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iii) incubating the reaction mixture obtained in i) or in ii)
at a temperature and during a time sufficient to allow
formation of the gel, and
iv) performing photopolymerization and crosslinking reaction of
5 the functionalized polyvinyl alcohol (PVA) grafted with
methacrylate with the functionalized serum albumin protein
grafted with methacrylate.
When the polymerization or reticulation is initiated,
those functional grafted chemical groups X are able to create
10 covalent bridges or bonds between functionalized synthetic
polymers and functionalized proteins.
The polymerization initiator is a compound that generates
free radicals. Any free radical generator known to the skilled
person, for the polymerization, can be used.
15 For example, the initiator can be a thermal initiator
selected from the group consisting of peroxides, persulfates,
or diazo compounds, such as potassium persulfate, ammonium
persulfate and hydrogen peroxide.
For example, the initiator may be a photopolymerization
initiator. It may be selected from the group consisting of
Irgacures (eg 2959, 184, 651, 819), 2-
hydroxy-4-(2-
hydroxyethoxy)-2-methylpropiophenone, benzophenone, 2-A
dimethylbenzophenone, benzoin, benzophenone ion selected from
the group comprising chloride trimethylmethylammonium
benzophenone-4, the sodium salt of 4-sulfomethylbenzyl,
aromatic ketones and aldehydes such as benzaldehyde,
acetophenone, biacetyl, the parachlorobenzophenone, the ferulic
acid, and all compounds producing radicals under visible or 0v
light. The Irgacure 2959 is preferably selected.
The concentrations of the different compounds in the
mixture (step i) may be:
- if the gel is fibrin gel, then fibrinogen or plasma is used
as gel precursor solution and the fibrinogen concentration may
range from 1 to 50 mg / mL, preferably between 5 and 25 mg / mL
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- the concentration of functionalized synthetic polymer,
particularly PVA functionalized with methacrylate, may range
from 1 to 400 mg / mL, preferably from 1 to 100 mg / mL
- the concentration of functionalized protein, particularly
albumin functionalized with methacrylate, may range from 1 to
200 mg / mL, preferably from 1 to 100 mg / mL.
When thrombin is used as gelification activator in order to
produce fibrin gel from fibrinogen solution or plasma in step
ii), thrombin can be derived for example from human plasma or
from pig or beef plasma.
The concentration of thrombin in the mixture of step (ii) can
be for example 0.1 to 2.5 U / mL, preferably 0.2 to 1 U / mL.
The concentration of calcium in the mixture of step (ii) can be
for example from 5 to 50 mM, preferably 20 mM.
- the concentration of polymerization initiator, particularly
Irgacure 2959, may range from 0.1 to 2 mg / mL, preferably of
0.6 mg / mL.
In step iii) the temperature of incubation can be a
temperature allowing the formation of the gel this temperature
may range for example between about 20 and 50 C, preferably,
C to 40 C, more preferably around 37 C.
It is during this step iii) that the gel is formed from
the gel forming solution or from gel forming precursor
solution.
25 If the gel forming precursor solution is fibrinogen
solution or plasma solution containing fibrinogen, it is during
this stage, step iii) that the fibrin physical gel is formed
from fibrinogen by enzymatic hydrolysis. The enzymatic
hydrolysis is catalyzed by thrombin and calcium. Unexpectedly,
30 the gel formation, the enzyme activity and the usual properties
of the fibrin gel are not affected by the presence of
functionalized synthetic polymer and of functionalized protein
nor by their polymerization and reticulation. Moreover most of
the properties, such as mechanical properties, of the fibrin
gel are improved.
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When the initiator is a thermal initiator, the
polymerization of the functionalized synthetic polymer and
functionalized protein can take place simultaneously with step
(iii) or after step (iii). When polymerization takes place
after step (iii), it can be implemented for example at a
temperature between 25 and 60 C. Preferably the temperature is
chosen such that the gel is not distorted. This is important
when the gel is fibrin gel and precaution should be taken to
avoid distortion of fibrin gel.
The thermal polymerization can be carried out between 0.5
and 10 hours, preferably between 1 and 3 hours in step iv).
When the initiator is a photopolymerization initiator, for
example those mentioned above, the polymerization can be
carried out in a UV / visible light at a wavelength between 190
and 800 nm, preferably under UV radiation of wavelength between
300 and 380 nm.
Photopolymerization in the step iv) may be performed for
30 minutes to 10 hours, preferably 30 minutes to 1 hour 30
minutes ,even more preferably 1 hour.
The different solutions were prepared in Tris buffer or
HEPES, preferably Tris-HC1 at a concentration of 25 to 100 mM,
preferably 50 mM, pH 7.2 to 7.5, preferably 7.4, containing
CaC12 10 to 50 mM, preferably 20 mM and NaCl 50 to 250 mM,
preferably 150 mM.
The final concentrations of the different compounds in the
mixture (step ii) may, for example:
- if the gel is fibrin gel, then fibrinogen solution or plasma
is used as gel precursor solution and the fibrinogen
concentration may range from 1 to 50 mg/mL, particularly
between 2 and 25 mg/mL, more particularly between 1 and
10 mg/mL,
- for the functionalized polymers, particularly PVA
functionalized with methacrylate, from 1 to 400 mg/mL,
preferably around 100 mg/mL ,
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- for the functionalized protein, particularly albumin
functionalized with methacrylate, from 1 to 200 mg/mL,
preferably from 1 to 100 mg/mL,
- for the initiator 0.1 to 1 mg/mL, preferably 0.42 mg/mL, and
- for thrombin, if fibrinogen is used as gel precursor
solution, of 0.1 to 2.5 U/mL, preferably 0.2 to 1 U/mL.
After step (iv), the method of the invention may further
comprise a step of enzymatic crosslinking of physical fibrin
gel. In this additional step can be used for example as
transglutaminase or lysyl oxidase enzyme.
Transglutaminase is an enzyme which bridges covalently (or
links) between them certain amino acids (primarily lysine and
glutamine) which contribute to the crosslinking of proteins
during the gelling process. This may well lead to a more
resistant system having improved mechanical properties.
According to this embodiment, the enzyme transglutaminase
for example, can be used. The IPN sample obtained after step
(iv) is then immediately or subsequently immersed in a solution
of transglutaminase in a concentration ranging from 0.1 to
10 U / mL, preferably between 0.5 and 3 U / mL for time
variable from 5 minutes to 6 hours, preferably 30 minutes to 3
hours.
The method of the invention can be implemented in any
container, beaker, crystallizer, pillbox, which can preferably
be closed. The container can also be custom made with, for
example, two glass plates or any UV transparent material and
waterproof material. These plates may be separated by a Teflon
gasket or polyethylene of various thicknesses of 100 mm to
3 mm, preferably 1 mm.
The invention also relates to the material in the form of
interpenetrating polymer network (IPN) associating a gel,
particularly a physical gel, more particularly a fibrin gel,
with a co-network of functionalized polyvinyl alcohol
covalently linked to a functionalized protein thank to
methacrylate bridges, obtainable by the method according to the
invention.
CA 02838126 2013-12-18
19
The material obtainable by the method according to the
invention has the advantage of being in the form of a gel,
homogeneous and self-supported while having dimensional
stability over time.
It is thus an object of the present invention to provide a
material obtainable by a method as herein described.
As mentioned above, in a preferred embodiment, the
material of the invention is in the form of interpenetrating
polymer network (IPN) combining a physical fibrin gel with a
co-network of functionalized polyvinyl alcohol (PVA)
crosslinked with functionalized protein, wherein there is no
chemical bond nor covalent bonding between the fibrin gel
component and the co-network components. In this
interpenetrating polymer network (IPN),
synthetic
functionalized polymer and functionalized protein form a
chemical network having a three-dimensional 3D existence. For
example, PVA network may be synthesized by crosslinking poly
(vinyl alcohol) modified with (meth) acrylate groups. PVA
network swells in water or in the buffer. According to an
advantageous embodiment of the invention, the material, the
interpenetrating polymer (IPN) network combines physical fibrin
gel and a network of polyvinyl alcohol prepared with poly
(vinyl alcohol) functionalized with methacrylate, wherein the
functionalized PVA is crosslinked to a functionalized protein.
In one preferred embodiment the invention concerns a material
as interpenetrating polymer network (IPN) associating a fibrin
gel and a co-network of functionalized synthetic polymer
crosslinked with a functionalized protein, wherein the
functionalized protein is albumin functionalized with
methacrylate groups and functionalized synthetic polymer is PVA
functionalized with methacrylate groups.
The material according to the present invention shows
improved properties compared to those of a sole fibrin gel. In
particular mechanical properties, biodegradability and
reversible dehydration and hydration properties are key
features of the material according to the present invention.
CA 02838126 2013-12-18
This is not the case for example for the fibrin gel alone which
after a dehydration cycle, cannot rehydrate more than 25% of
the initial value.
The material according to the present invention is self-
5 supported and can be dehydrated in order to be stored and
transported dried.
The material according to the present invention can thus
be reversibly dehydrated and rehydrated. The material according
to the invention can be dehydrated and then rehydrated between
10 25 and 100%, preferably between 45 and 95%.
Thus, in a particular embodiment of the present invention,
here is provided a material according to the present invention
which is dehydrated and has a moisture content of about 2 to
10% in weight, particularly about 2 to 5% in weight. The thus
15 obtained dry material can be stored at room temperature and can
be used upon necessity thank to extemporaneous rehydration
within aqueous medium.
The drying of the material according to the invention can
be achieved according to any technique known by the man skilled
20 in the art and typically, the material can be disposed into an
oven at a temperature comprised between 40 and 60 C for a
sufficient time to evaporate the contained water and to obtain
a material with a moisture content of about 2 to 10 % in
weight, particularly about 3 to 5 % in weight.
The material according to the invention has a glass
transition temperature (Tg) between -100 and +100 C and
preferably between -50 and 80 C. The glass transition
temperature Tg of a polymer material is defined as the
temperature below which the polymer chains have a low relative
mobility. Below Tg the polymer is in a glassy state, above Tg,
the polymer is in a rubbery state. The glass transition
temperature of a material can be measured, for example, by
differential scanning calorimetry, DSC.
The material of the invention may also have a storage
modulus (measured in shear mode, G ') of between 100 Pa and 10
MPa, for instance between 1000 Pa and 6 MPa, in the hydrated
CA 02838126 2013-12-18
21
state at 37 C and a storage modulus (measured in voltage
mode E ') of between 0.01 and 3000 MPa, for example between 1
and 3000 MPa, in the dry state.
The module G of a material subjected to a sinusoidal
strain is written in the complex form:
G* = G' +iG"
i represents the imaginary part of a complex number.
In shear mode, the storage modulus of elasticity (G') is
used to estimate the elasticity of the gel. The loss modulus or
viscous modulus (G") characterizes the liquid phase of the gel.
When G" G', the sample is regarded as liquid and when G'> G",
the sample is considered as a gel. The gel time is considered
as the time at which G' and G" moduli are equal.
The material obtainable by the method of the invention is
preferably biocompatible. When one wishes to perform a sterile
material synthesis (after the implementation of the method of
the invention), the material is advantageously obtained by
photopolymerization.
Another object of the invention is the use of a material
according to the invention or obtainable by the method of the
invention, such as:
- Wound dressing,
- Hemostatic dressing,
- Surgical dressing,
- A device for delivering therapeutic agents,
- Coating for medical devices selected from the group
consisting of stents, heart valves, catheters, vascular
prosthetic filters,
- Active carrier molecules selected from the group consisting
of growth factors, antibiotics, bactericides, bacteriostats and
enzymes,
- Support for eukaryotic cells culture
In a particular embodiment, the invention concerns the use
of the present material as skin substitute in the treatment of
burns and skin repair.
CA 02838126 2013-12-18
22
In another embodiment, the invention concerns the use of
the present material as skin model for research and screening
applications including penetration studies, pigmentation
studies and toxicity, corrosivity and irritation testing,
particularly in cosmetic.
Using functionalized protein such as serum albumin
functionalized with methacrylate in addition to functionalized
synthetic polymer such PVA functionalized with methacrylate,
various IPN materials including a fibrin gel were obtained. The
one pot-one shot method was applied in each case leading to a
rapid synthesis (from 2 to 15 minutes). The mechanical
properties of all IPN materials according to the present
invention are sufficient to obtain self-supported IPN
materials. The IPN architecture allows improving the storage
modulus of the fibrin gel by a factor ranging from 8 to 40
depending on the IPN material composition. All materials are
well synthesized and present interpenetrating polymer network
architecture as shown by low soluble fractions. The protein
based co-network (synthetic polymer, such as PVA, and protein
such as serum albumin, both functionalized with methacrylate)
is homogeneously spread.
The material according to the present invention is easily
biodegradable due to the presence of the functionalized
protein, particularly serum albumin, included and cross-linked
with the functionalized synthetic polymer inside the material.
This new type of material is the first co-network IPNs ever
described in literature to be potentially biodegradable through
fragmentation, then elimination within the biological
environment, i.e. after implantation in the body. In addition,
degradability helped to increase very significantly the
bioactivity of materials.
Indeed, previously existing biomaterials are hardly
biodegradable since they are made of synthetic polymers which
are not easily degraded in situ by enzymes.
When functionalized protein, particularly serum albumin
functionalized with methacrylate, is used as a constituent of
CA 02838126 2013-12-18
23
the material according to the invention, the surface of the
obtained material can be completely covered by fibroblasts
within two weeks. In addition, cells at the surface of the
material secrete the extracellular matrix biomacromolecules
(fibronectin, hyaluronic acid) playing a key role in the early
stages of wound healing and are able to structure them into
fibers (fibronectin). The affinity properties of the molecules
are identical to those of dermal fibroblasts (co-localization
of fibronectin and hyaluronic acid). The good spreading of
fibroblast cells on the surface of the materials allows them
producing type I collagen which is a protein characteristic of
late maturation phase. All these properties indicate that the
materials according to the present invention are excellent
supports for 2D- cell culture. Indeed, they could help to
increase significantly the speed of wound healing. These
innovative materials could therefore be used in their present
state as dressings or to fill large defect of substance as they
may be hydrolyzed and colonized progressively by surroundings
cells. Moreover, they are also able to sustain 3D culture, so
that they could be used to provide dermis in skin construct.
The material according to the present invention is
biodegradable by the action of endogeneous or exogeneous
proteolytic enzymes may be stored dry and rehydrated without
shrinkage upon desire.
In a preferred embodiment the use of serum albumin
functionalized with methacrylate (BSAm or HSAm) copolymerized
with PVA functionalized with methacrylate (PVAm) (co-network)
into IPN architecture containing a fibrin gel allows the
synthesis of a hybrid material with all the properties
described above and biodegradable by enzymes, within the
biological environment, i.e. after implantation in the body.
The material according to the present invention has been
tested in contact with human cells, such as fibroblasts. It has
been demonstrated that in addition to be non-cytotoxic, the
material according to the present invention could be completely
colonized by these cells. The development of a material
CA 02838126 2013-12-18
24
according to the present invention with a homogeneous cellular
distribution in the three dimensions of the material
constitutes an advantageous use of the present invention.
Advantageously, the even distribution of cells within the whole
matrix of the material according to the invention allows
obtaining cellularized dressings or skin substitutes. The
living cells herein contained in the material of the invention
will contribute to the reformation of the cellular matrix and
wound healing by sustaining neoangiogenesis and the rebuilding
of the burned, injured or damaged tissue.
A homogeneous distribution of cells, particularly
fibroblasts, in three dimensions of a piece of material
according to the present invention can be achieved through the
encapsulation of a viable population of cells, particularly
fibroblasts, which are thus embedded in the material. Cells
cultured on the surface or encapsulated within the material
according to the invention show a high viability as well as
proliferation and matrix remodeling abilities.
It is thus a further object of the present invention to
provide a material according to the herein described invention
characterized in that it contains living cells encapsulated.
The encapsulation of cells in the material according to
the present invention can support their viability for at least
5 weeks. Once encapsulated in the material of the invention,
the cells slowly proliferate in the early phase of the culture,
and their proliferative activity decreases after that early
phase. The cells synthesize extracellular matrix components and
are able to transform collagen type IV into collagen I. It is
possible to freeze the material according to the present
invention wherein living cells are encapsulated without losing
viability and cell proliferative activity.
By the expression "encapsulated" used herein, it is to be
understood that the cells are evenly distributed with the whole
matrix of the material according to the invention in the three
directions.
CA 02838126 2013-12-18
It is thus an object of the present invention to provide a
material according to the present invention characterized in
that it contains encapsulated living cells.
The encapsulation of cells within the material according
5 to the present invention may be achieved by the addition of
cells suspension to the first reaction mixture as herein
described.
In a particular embodiment the method of the present
invention for the preparation of a material according to the
10 present invention further comprises the step of adding a cell
suspension in the first mixture, prior to the step of
introduction of polymerization initiator.
The quantity of cells added in the first reaction mixture
is such that the final quantity of cells with the material is
15 comprised between 1*106 and 3*106 cells per unit of gel material
obtained. A unit of gel may have any form or size and would
typically have the size and the shape of the mold used for its
manufacture. A unit may thus be a cylinder, a cube or a
parallelepiped for example which dimension will be dependent
20 upon the intended use.
The encapsulation of cells such as fibroblasts can be
achieved by adding the so called first mixture containing the
gel forming solution or gel forming precursor solution with
functionalized synthetic polymer, functionalized protein,
25 polymerization initiator, particularly a photopolymerization
initiator, to a cell suspension. The resulting solution was
placed in a mold and then exposed to UV for the synthesis of
IPN as exposed above. The obtained material is then immersed in
a culture medium at 37 C for sufficient time to remove traces
of synthesis buffer. The density of cells obtained is dependent
upon the concentration of cells in the suspension which can be
chosen in order to obtain a final density of about 1000 to
10000 cells/mm3, particularly between 4000 and 8000 cells/mm3;
more particularly around 7000 cells/mm3.
It is particularly noticeable that the photopolymerization
under UV does not affect the viability of cells embedded with
CA 02838126 2013-12-18
26
the material of the invention. U.V exposure does not cause
mortality of cells on the surface nor within the material.
After 1 and 3 weeks of culture, cell population is at least
maintained and the distribution remains homogeneous. In
addition, the cells initially present on the surface are
strongly expanded and do cover the material. This distribution
on the one hand, and the lack of mortality on the other hand,
confirms that there is no impact of UV exposure on the cell
population comprised on and within the material encapsulating
cells.
Cell count shows an overall increase in cell population
with the cultivation time. Indeed, after three weeks, the cell
population is 150% of the initial value. Thus, even if the
cells are encapsulated in a matrix comprising the
functionalized PVA coreticulated with functionalized protein,
they are able to proliferate within the material.
As a matter of example for material containing cells
embedded within it, as it is observed with cell culture onto
the material of the invention, fibronectin is present at 24
hours of culture. Between 1 and 3 weeks of culture, a
significant increase in the presence of fibronectin is visible
on the depth of the material. After 3 weeks of culture,
different protein structures are observed depending on depth
studied. Indeed, on the surface, fibronectin adopts a typical
fibrillar form indicating that the protein has been remodeled
and integrated into the matrix. In depth, fibronectin is
located around the cells. All of these results suggest that it
is possible to produce cellularized dressings by encapsulation
of fibroblasts within a material according to the invention.
This type of material would be able to promote wound healing
and secretion of extracellular matrix macromolecules such as
fibronectin and collagen. The possibility of encapsulating
living cells within the material of the Invention and the
capacity of the cells to proliferate in it makes this material
capable of being used as a cellularized skin substitute for the
treatment of burns for example.
CA 02838126 2013-12-18
27
It is also noticeable that the material according to the
invention containing encapsulated cells can be frozen, kept
frozen and defrosted without alteration of the viability of the
cells embedded in it. The viability of the cells encapsulated
are not affected by freezing at -80 C and the material of the
invention containing encapsulated cells can be stored in the
frozen state at -80 C for several months without alteration of
the contained cells that can restart their growth once the
material is unfrozen. Typically the cryopreservation of
encapsulated cells is achieved by traditional means known by
the man skilled in the art. A material containing cells
encapsulated according to the present invention is incubated in
fetal bovine serum for few hours with or without cryoprotectant
such as glycerol prior to freezing at -80 C during few hours.
The material of the invention containing encapsulated cells can
then be stored at -80 C for months. The thawing is preferably
achieved by a rapid warming of the frozen sample, for example
by incubation 1 37 C. The defrosted material containing
encapsulated cells can be cultivated under classical conditions
and the cell viability, proliferative activity as well
extracellular protein production is not affected.
It is thus a further object of the present invention to
provide a material according to the invention that can contain
encapsulated cells and that is frozen.
The material according to the present invention, IPN
containing a fibrin gel scaffold at physiological concentration
and a co-network of functionalized synthetic polymer
crosslinked with a functionalized protein, has excellent
properties that make it suitable to be used as wound dressings
or cell culture support or matrix for skin regeneration. In
addition, thanks to the ability of the material according to
the invention to encapsulate cells, particularly fibroblasts,
such a material based on co-network of PVA functionalized with
methacrylate and albumin functionalized with methacrylate,
particularly serum albumin functionalized with methacrylate,
CA 02838126 2013-12-18
28
makes it suitable as a skin substitute in the treatment of
burns and skin repair.
It is thus a further object of the present invention to
provide the use of a material according to the present
invention as a skin substitute in the treatment of burns and
skin repair.
The material of the invention can be used as hemostatic
dressing. In case of bleeding, platelets aggregate on fibrin,
release growth factors of coagulation factors leading to blood
thrombus organization. The material according to the present
invention is biodegradable and can thus serve as support for
cell culture. Also, in the case of loss of ground substance as
can occur upon severe burning, the material according to the
present invention can stimulate angiogenesis and can thus
sustain the rebuilding of the affected tissues.
The material according to the present invention exhibits
very good rehydration properties after dry storage, as well as
enzymatic degradation by proteases once in place. The material
allows the maintenance of a humid atmosphere for the wound and
cell growth on the surface and the secretion an extracellular
matrix. The material according to the present invention is able
to support and sustain the reformation of normal tissue at the
wound sites.
Other advantages may also occur to those skilled in the
art upon reading the following examples, illustrated by the
accompanying drawings, given by way of illustration.
Brief description of the figures
Figure 1 shows :
- A- Hydrogel IPNs (1, 2, 3, 4, 5) just after synthesis with
respectively 0, 3, 5, 7 and 10 % (w/v) of BSAm and supplement
with PVAm to a final 10% (w/v) concentration.
- B - Hydrogel IPNs (1, 2, 3, 4, 5) after coloration with
Coomassie blue, respectively (1', 2', 3', 4', 5').
CA 02838126 2013-12-18
29
Figure 2 shows (1)PVAm(10%)/Fb; (2) PVAm(5%)coBSAm(5%)/Fb;
(3) BSAm(10%)/Fb IPNs (A) just after synthesis and (B )after 24
h incubation in a thermolysin solution (20 U. mL-1) at 37 C.
Figure 3 shows observation by confocal microscopy of the
repartition of BSAm and fibrin Fb (both light grey) in a
BSAm(10%)/Fb IPN (A) and in PVAm(5%)coBSAm(5%)/Fb IPN (B). -
Objectif x63
Figure 4 shows cell morphology with cytoskeleton and
nucleus, after 1 day, 1 and 3 weeks of culture on BSAm(10%)
single network (A) and PVAm(5%)coBSAm(5%)/Fb (B),
PVAm(3%)coBSAm(7%)/Fb (C) and BSAm(10%)/Fb (D) IPNs. Cells were
observed with a x20 objective by CLSM (Zeiss LSM710). Scale bar
: 20 pm.
Figure 5 shows (A) Cell density and (B) cell viability on
the surface of various materials after 24h (0), 72h (0), 1 week
(Li), 2 weeks (I) and 3 weeks (I) of culture. Materials are : I
: BSAm (10%) single network, II : PVAm(5%)-co BSAm(5%)/Fb, III
: PVAm(3%)-co BSAm(7%)/Fb and IV : BSAm(10%)/Fb IPNs.
Figure 6 shows cell secretion and remodeling of
fibronectin with cytoskeleton and nucleus after 1 day (first
row), 1 week (second row) and 3 weeks (third row) of culture on
PVAm(5%)-co BSAm(5%)/Fb (first column A), PVAm(3%)-co
BSAm(7%)/Fb (second column B) and BSAm(10%)/Fb (third column C)
IPNs. Cells were observed with a x 20 objective by CLSM (Zeiss
LSM710). Scale bar: 20 pm.
Figure 7 shows cell secretion and remodeling of hyaluronic
acid ( A (first row), B (second row)) and type I collagen ( C
(third row), D (forth row)) with nucleus after 1 day (A (first
row) and C (third row)) and 3 weeks (B (second row) and D
(froth row)) of culture on PVAm(5%)coBSAm(5%)/Fb (first
column), PVAm(3%)coBSAm(7%)/Fb (second column) and BSAm(10%)/Fb
(third column) IPNs. Cells were observed with a x20 objective
by CLSM (Zeiss LSM710). Scale bar: 20 pm.
Figure 8 shows the cell viability expressed as % of living
cells according to Live/Dead test for different cultivation
CA 02838126 2013-12-18
times (76 to 840 h) for fibroblast encapsulated in
PVAm(5%)coBSAm(5%)/Fb IPN.
Figure 9 shows the cell viability for different
cultivation times (96 h and 168 h) after thawing for
5 cryopreserved fibroblasts encapsulated in PVAm(5%)coBSAm(5%)/Fb
IPN with 3 % glycerol (dark grey) and without glycerol (middle
grey) in comparison with the unfrozen control sample (light
grey).
Figure 10 shows the proliferative activity cells for
10 different cultivation times (96 h and 168 h) after thawing for
cryopreserved fibroblasts encapsulated in PVAm(5%)coBSAm(5%)/Fb
IPN with 3 % glycerol (dark grey) and without glycerol (middle
grey) in comparison with the unfrozen control sample (light
grey).
EXAMPLES
Materials and methods
Chemicals
Thrombin (BP 25432), sodium chloride, sodium
dodecylsulfate, were purchased from Fisher Reagents and bovine
fibrinogen (Fg 341573) from
Calbiochem.
Tris(hydroxymethyl)aminomethane were purchased from VWR.
Calcium chloride and magnesium chloride were obtained from
Riedel-deHaen and Prolabo, respectively. Azide, Brillant Blue
R, sodium carbonate, dietholamine, 2-isocyanatoethyl
methacrylate (2-ICEMA), paraformaldehyde (PFA, P-6148), 4,6-
diamino-2-phenyl-indole (DAPI, D-9564), albumin from bovine
serum (purity 98%,
A7906), methacrylic acid N-
hydroxysuccinimide ester (purity = 98%, 730300), thermolysin
from Bacillus thermoproteolyticus rokko (P1512), anti-human
fibronectin IgG (produced in rabbit, F3648) and alkaline-
phosphatase conjugated anti-rabbit IgG (A3687) were obtained
from Sigma. Anti-fibrinogen from rabbit (A0080) was obtained
from Dako. P-nitrophenyl phosphate (pNPP, 71768) was purchased
from Fluka. Anti-human type I collagen IgG (produced in mouse,
MAB3391) and anti-human hyaluronic acid IgG (produced in sheep,
CA 02838126 2013-12-18
31
AB53842) were purchased respectively from Millipore and Abcam.
Polyvinyl alcohol (PVA 98%, M = 16000 g/mol) and hydroquinone
were purchased from Acros. 2-hydroxy-1-[4-(2-hydroxyethoxy)
phenyl]-2-methyl-1-propanone (Irgacure - 12959) was purchased
from Ciba. Acid boric and dimethylsulfoxide (DMSO) were
purchased from Merck. DMSO was dried before utilization and
stored in the dark on molecular sieve under argon. Phalloidin-
Alexa0 532 (P-5282), Alamar blue (DAL 11000) and LIVE/DEAD kit
(L3224) were obtained from Invitrogen. Immunoglobulin (IgG)
against mouse (produced in goat, A11029), sheep (produced in
donkey, A11057) and rabbit (produced in goat, A21071) labeled
respectively with Alexa 488, 568 and 633 were also obtained
from Invitrogen. Penicillin-streptomycin (1416), culture medium
DMEM high glucose (31966021), Trypsine EDTA, and PBS (foetal
bovine serum, 10270-106) were obtained from Gibco.
Synthesis of polyvinyl alcohol (PVAm) modified with
methacrylate groups
PVA was modified with methacrylate groups as described.
Briefly, 20% (w/v) PVA was solubilized in DMSO with
hydroquinone. 3 mol% 2-ICEMA (with respect to the hydroxyl
function of PVA) were added in the PVA solution. Reaction was
mixed for 4 h at 60 C and 12 h at 20 C under argon atmosphere.
This solution was purified by precipitation in acetone at room
temperature. The modified PVA (denoted PVAm) was filtered and
dried under vacuum at 30 C for 48 h before dissolution in Tris
buffer 250 mM at pH 7.4. Tg of PVAm is 75 C.
Synthesis of bovine serum albumin (BSAm) modified with
methacrylate groups
4% (w/v) BSA were solubilized in acid boric buffer 250 mM
at pH 7.4. Methacrylic acid N-hydroxysuccinimide ester (NHSm)
solubilized in acetone was added by drop wise up to a 0.7/1
molar ratio of NHSm/lysine of BSA. Reaction was carried out at
room temperature for 12 h in dark. BSA modified with
methacrylate groups (denoted BSAm) was purified by dialysis
(Mcut off 1 kDa) against Tris buffer 50 mM at pH 7.4 in order
to eliminate unreacted NHSm. To increase the BSAm concentration
CA 02838126 2013-12-18
32
in solution, it was lyophilized and solubilized at 20% (w/v) in
Tris buffer 50 mM at pH 7.4. Solution of BSAm was filtered
(0.22pm, Millipore) before storage (-20 C).
Gelation procedure
All reactants were solubilized in a 0.05 mol/L Tris-HC1
buffer pH 7.4 and incubated at 37 C for 15 min prior to
mixing. Same protocol was used for any materials regardless of
PVAm and BSAm ratio.
Co-network/Fibrin IPN was synthesized as follows: 1 mL of
Tris-HC1 buffer containing 10% (w/v) of polymers (PVAm 10%,
PVAm(5%)coBSAm(5%), PVAm(3%)coBSAm(7%) and BSAm 10%), 5 mg
fibrinogen, 0.20 units thrombin, 0.15 mol/L NaCl, 0.02 mol/L
CaC12, 0.042% (w/v) Irgacure 2959 was prepared in a microvial.
The mixture was placed into mold made by Teflon and glasses.
The mold was placed at 5 cm from a UV lamp (VL-6, Bioblock, 2 x
5W, 365 nm) for 1 h at 37 C. Self-supported materials were
obtained with areas varying from 1.5 to 78.5 cm2, according to
the mold size.
Soluble fraction
Soluble fraction of synthetic polymer networks was
extracted for 48 h in Soxhlet with deionized water. The samples
were weighted before (wd and after (wd and soluble fractions
were determined as follow:
SF (%) = [ (wl-wf+wsalt)/(141)]*100
Unreacted fibrinogen in IPNs was extracted by immersion of
materials in a 10-fold volume of buffer during one night.
Concentration of protein extracted from hydrogels was assessed
by Elisa test. Absorbance was detected with a plate reader (Bio
Tek) at 405 nm. Each sample was synthesized in triplicate, and
each measurement repeated three times.
Staining of materials by Coomassie blue
PVAm(10%) and BSAm(10%) single networks and co-networks
containing PVAm and BSAm at different ratios were incubated for
1.5 h in aqueous coloration solution containing 2.5% (w/v)
Coomassie blue, 40% (v/v) ethanol, and 7% (v/v) acetic acid.
*Then, the materials were rinsed in aqueous solution containing
CA 02838126 2013-12-18
33
20% (v/v) ethanol and 10 % (v/v) acetic acid until removing of
excessive coloration was done.
Rheology
Gelation study was performed by Rheological measurements
with an Anton Paar Physica NCR 301 rheometer equipped with CTD
450 temperature control device with cone-plate geometry (cone:
diameter 25 mm, angle 2 ; plate: polymerization system made
from glass coupled with U.V Source Omnicure). The solution of
precursors of materials was put between the two geometry and
measurements begin immediately. Polymerization was initiated by
U.V exposure (4.46 mW/cm2) at 37 C. Storage modulus (G') and
loss modulus (G") at 1% deformation imposed at 1 Hz were
recorded as a function of time. The final storage modulus G' is
determined after 1 h UV exposition (system equilibrium reached)
and gel time was determined at the intersection between storage
and loss modulus curves. The average values of the shear
modulus were measured 3 times minimum.
Enzymatic degradation of materials
Just after synthesis, materials were immersed in Tris
Buffer 50 mM pH 7.4, 0.02% (w/v) NaN3 containing 20 U/mL
thermolysin. The hydrolysis volume solution is 20 fold the
volume of the materials. Enzymatic degradation of the materials
was performed at 37 C for 24h.
BSAm and Fb repartition.
Proteins in BSAm based IPN are specifically labeled as
follows. After synthesis, materials are immersed in PBS-casein
buffer for 1 h at 37 C before incubation first in PBS buffer
with 1%(w/v) casein containing BSA-antibodies (1/50) for 1 h at
37 C. After 3 successive rinses in PBS buffer, they are then
incubated in PBS-casein buffer containing secondary rabbit-
antibodies 350 (1/100) for 1 h at 37 C. Finally, fibrin
specific staining is performed in PBS-casein buffer containing
FITC fibrin-antibodies (1/100) for 1 h at 37 C, followed by 3
successive rinses before confocal microscope observation with a
x63 (Plan-Apochromat 63x/1.4) objective.
CA 02838126 2013-12-18
34
Biocompatibility test
Cell culture
Cell line used for biocompatibility test was Human
fibroblast from foreskin (FB-BJ, ATCC CRL 2522). All cells were
cultured in DMEM high glucose medium supplemented with 10%
(v/v) fetal bovine serum (FBS) and penicillin-streptomycin in a
humidified 5% CO2 incubator. All experiments were performed
with cell between 9 to 11 scans after de-freezing of ATCC vial.
Cell seeding
For cell seeding on surface of materials, they were placed
in a sterile culture plate (P24) and immersed in 1 mL of
culture medium without FBS for at least 24 h. They were then
seeded with 1 mL of a cell suspension containing 1x105 cell/mL
(cell density 5x104 cell/cm2). The materials tested in
triplicate at different culture times (from one day to 3 weeks)
were the BSAm(10%) single network, PVAm(5%)coBSAm(5%)/Fb IPN,
PVAm(3%)coBSAm(7%)/ Fb IPN and BSAm(10%)/Fb IPN.
Live /dead() assay
Viability tests were performed by staining cells with a
solution containing 0.2 pM calcein AM and 0.2 pM ethidium
bromide dimer in phosphate buffer for 30 min at 37 C. Picture
were taken with Confocal Laser Scanning Microscope (CLSM, Zeiss
LSM 710, Germany) in sequential line mode (averaging 2) with a
x10 (EC Plan-Neofluar 10x/0.30 M27) objective. Living and dead
cells were stained respectively with calcein AM (Aex = 488nm,
Xem = 490-573 nm) and ethidiumbromide dimer (Xex = 561m, 2\em =
580-730 nm). Evaluation of cell population was performed using
ImageJO software (Cell counter plug in) and viability of cell
were calculated by the formula:
% Viability = [(full population of cell - quantity of dead
cell) / (full population of cell)] *100
Cell staining
A particular attention has been paid to the origin of
antibodies and to the emission wavelength of conjugated dyes,
so that 5 different stainings have been performed on every
CA 02838126 2013-12-18
sample to allow direct observation of cell morphology and ECM
remodeling at the same time.
Cell fixation was carried out with 3% (w/v)
paraformaldehyde solution for 12 h at 37 C, permeabilized with
5 a PBS-0.1% (v/v) Triton X-100 solution at room temperature (RT)
for 30 min and rinsed with PBS. Then non-specific binding sites
were blocked by immersion in saturation buffer composed of PBS
with 10% (v/v) PBS for 30 min at room temperature.
For cell density studies, nuclei were stained with DAPI (1
10 pg/mL) for 1 h at room temperature and then rinsed. Cells were
observed with CLSM with a x20 (Plan-Apochromat 20x/0.8 M27)
objective. Cell numbers were quantified by counting nuclei with
ImageM, software (Cell counter plug in).
For cell morphology and ECM remodeling studies, primary
15 antibody solutions against human collagen, hyaluronic acid and
fibronectin were deposited on the surface of materials. After
min incubation at room temperature, samples were rinsed and
labeled for 45 min with a staining solution containing DAPI (1
pg/mL), phalloidin AlexaC) 532 (dilution 1/10000) and secondary
20 antibodies. Materials surfaces were observed with CLSM with a
x20 (Plan-Apochromat 20x/0.8 M27) objective..
Results and discussion
The aim of this work is to introduce a biodegradable part
inside a synthetic polymer network incorporated into a fibrin-
25 based Interpenetrating Polymer Networks (IPN) architecture to
insure the progressive vanishing of the material.
Serum albumin was chosen as the biodegradable part as it
is the major proteic component of serum, the same biological
source as fibrinogen. Moreover, due to its polypeptidic nature,
30 serum albumin is naturally responsive to various proteases.
Bovine serum albumin (BSA) was grafted with methacrylate groups
to be able to generate a network - or to copolymerize with
another polymer precursor - through photopolymerization. An
average of 66% free amine groups of BSA, mainly located on the
35 lateral chains of lysine residues, was modified with a
CA 02838126 2013-12-18
36
methacrylate group after the procedure. The modified protein,
named BSAm, was synthesized with a 91% yield.
BSAm single networks and PVAm-BSAm co-networks
First the possibility to generate BSAm and PVAm-coBSAm
homogeneous networks was explored. BSAm is used to confer
biodegradability to the network and PVAm to provide good
mechanical and hydration properties to the biomaterial. In
order to maintain favorable conditions for the further
formation of a fibrin network, the synthesis was carried out in
buffer pH 7.4 with appropriate salt concentration. Irgacure
2959 was selected as photoinitiator because it is well known to
be weakly toxic and noxious towards different types of cells at
concentrations lower than 0.1% (w/v); the concentration here
used is 0.04%.
Under these conditions, different homogeneous materials
were obtained with BSAm concentration ranging from 0 to 10 wt%,
the final polymer concentration (BSAm + PVAm) being lOwt % in
all cases. The higher the BSAm ratio, the more transparent they
are (figure 1A).
It may also been observed that with a low concentration
(3%, A-4) or in the absence (A-5) of PVAm, the network deforms
and the initial regular round shape is not maintained after
removing the mould.
After their synthesis, the proteins in the different co-
networks were stained with Commassie blue. As shown in Figure 1
B, the coloration of the different materials containing BSAm is
regular, indicating a homogeneous repartition of proteins
(BSAm) inside the materials. The absence of coloration with
PVAm alone ascertains the method validity.
Having shown that a single BSAm network and various
BSAmcoPVAm co-networks may be obtained, materials including a
fibrin gel in IPN architecture were synthesized.
BSAm/Fibrin and PVAm - BSAm/Fibrin IPNs
The free radical photopolymerization of the BSAm and PVAm,
initiated by U.V exposure, was performed in aqueous buffered
medium at 37 C all along with the fibrin gel formation that
CA 02838126 2013-12-18
37
was synthesized by the enzymatic hydrolysis of fibrinogen by
thrombin. Using the same conditions, the in situ syntheses of
BSAm(10%)/Fb IPN and of various PVAmcoBSAm/Fb IPNs have been
developed. The precursor solutions contain fibrinogen and 10%
methacrylate polymers (BSAm+PVAm) with BSAm concentrations
being 0, 5 or 10% (w/v).
In all cases, a material was obtained after 15 min. The
presence of fibrinogen has no significant impact on the
formation rate of the synthetic networks as gelation occurs at
similar times for single co-networks and IPNs.
After 60 min irradiation, the storage moduli were measured
and those for IPNs compared to that obtained for the fibrin
network. All materials containing either BSAm or/and PVAm are
self-supported (G' > 100 Pa) both as single co-networks and in
IPN architectures. The addition of any synthetic network to a
fibrin gel improves its mechanical properties as shown by the
large increase in storage modulus (G') from 8 to 40 fold. The
synthetic networks well fulfill their role of rendering a
fibrin gel easily handled. As suggested above through the
visual observation, the single BSAm network shows lower
mechanical properties (G' = 220 Pa) than co-network with 5%
PVAm (G' = 830 Pa) or than the PVAm single network (G'= 3160
Pa). However, the addition of the fibrin gel improves this
characteristic a lot (from 220 to 868 Pa).
To check the effective formation of the co-networks in IPN
architecture, extractions of both the synthetic and the fibrin
fractions of the materials were performed.
To check the effective formation of the networks in IPN
architecture, extractions of both the synthetic and the fibrin
fractions of the materials were performed. The results are
presented in Table 1.
CA 02838126 2013-12-18
38
Table 1: Soluble synthetic and fibrin fraction for various
single networks, IPNs and fibrin gel.
Sample SN PVAm IPN PVAm/Fb IPN IPN BSAm/Fb SN
PVAmcoBSAm/Fb Fb
% soluble synthetic fraction 8% 0,3 12% 2 16% 0,9 16% 5
% soluble fibrin fraction 0,189/. 022 0,47% 0,03 0,32% 0,03
0,03% 0,03
From the Soxhlet extraction, it appears that the part of
methacrylate polymers which is not included in the synthetic
network varies from 12 to 16 wt%. These values are high but
correct for polymer networks synthesized from a diluted
precursor solution.
The proper formation of the fibrin network was also
verified. The Elisa assay on extraction solution of IPNs shows
less than 1 wt% of extractable fibrinogen whatever the
composition of its partner network. These results confirm that
the synthetic network does not inhibit the enzymatic formation
of fibrin network by thrombin. Using a Western blot technique
after disruption of the IPNs, the presence of an intense y-y
band was also evidenced, confirming the transformation of
fibrinogen into fibrin upon thrombin cleavage and the
association of fibrin chains into a gel phase inside the
material.
Biodegradability
Next, PVAm(10%)/Fb, PVAm(5%)coBSAm(5%)/Fb and BSAm(10%)/Fb
IPNs were synthesized and observed (Figure 1, A). As for the
single networks, the BSAm/Fb IPN (A-3) is more transparent than
IPNs containing PVAm (A-2 and A-1). Accordingly to the rheology
results, the three materials keep their round shape just after
synthesis. Their degradation was thus tested. The different
IPNs were incubated with a concentrated metalloprotease
solution. The results are illustrated in figure 2-B.
As suspected, the PVAm(10%)/Fb IPN is not solubilized by
the protease under those conditions. Thermolysin may degrade
the fibrin network inside the IPN, but this protein represents
only 5 wt% of the solid fraction of the IPN, so its proteolysis
does not lead to the dislocation of the material. This
CA 02838126 2013-12-18
39
experiment also proves the good repartition of PVAm inside the
IPN architecture as the round shape of the material is not
affected by the hydrolysis of the fibrin network. This point
illustrates one of the best properties of the material here
generated; its shape stays constant, even in presence of
degrading enzymes.
The PVAm(5%)coBSAm(5%)/Fb IPN is partly degraded upon
enzyme action; many fragments are obtained, indicating that
BSAm was also well distributed in the co-network.
The BSAm(10%)/Fb IPN is totally degraded by the protease.
This material is thus well biodegradable. A similar result may
be obtained with a lower enzyme concentration (10 U/mL).
These experiments show that BSAm in the material is well
accessible to the protease and may be hydrolyzed even in the
IPN architecture. The homogeneous repartition of BSAm in the
BSAm(10%)/Fb IPN and in the PVAm(5%)coBSAm(5%)/Fb IPN has been
imaged in confocal laser scanning microscopy using specific
fluorescent probes for both proteins (figure 3).
Staining of the BSAm(10%)/Fb IPN with antibodies reveals
an homogeneous repartition of BSAm in the whole material.
Moreover, the fibrin network observed in the BSAm(10%)/Fb IPN
presents the characteristic honeycomb structure of a single and
biologically active fibrin network. This indicates that the
increase in gel time due to the presence of BSAm (from 3 to 15
minutes) is not related to a perturbation in the fibrin network
formation.
For the PVAm(5%)coBSAm(5%)/Fb IPN, BSAm is also present in
the whole volume while fibrin forms larger connected fibrils
forming an irregular network.
However, both proteins are present throughout the
material. Combination of all the results: similar extract
concentrations with a Soxhlet for a BSAm or a PVAmcoBSAm
network in the IPNs, homogeneous coloration with coomassie
blue, homogeneous repartition of BSAm in all the IPNs, increase
of mechanical properties and homogeneous biodegradability
CA 02838126 2013-12-18
suggest that photopolymerization successfully formed a
copolymer between PVAm and BSAm.
Material made with plasma
Blood plasma from the Etablissement Francais du Sang (EFS)
5 filtered through 0.22 microns, or not filtered, was used as a
source of fibrinogen.
The protocol includes solubilizing PVAm in Tris buffer,
followed by addition of plasma Instead of fibrinogen and the
formation of IPN is similar to the protocol used in the
10 previous examples. In all cases, calcium (20 mM) and thrombin
(0.2 U/mL) were added to enable the rapid formation of the
fibrin network.
IPN containing co-network of PVAm and HSAm as well as a
fibrin gel from plasma were synthesized according to the same
15 principle as in the previous examples. In this case, the
lyophilized HSAm is dissolved in the plasma, in order to have
the concentration as high as possible (2.8 mg/mL) of
fibrinogen. These materials are homogeneous and can be
manipulated. Obtained synthetic materials are homogeneous and
20 translucent. Following incubation in the Coomassie Blue and
successive rinses, the IPN made with plasma has a uniform blue
color. This indicates that the fibrin network is homogeneously
distributed in the materials. On extractable protein, the
single network PVAm (10 %) and PVAm (5 %) coHSAm (5%) co-
25 network have zero values, thus constituting robust negative
control and in the case of the PVAm (10%) / plasma IPN and co-
network PVAm (5%) coHSAm (5%) / plasma IPN, soluble fractions
are less than 0.1 %: the fibrin network is well formed, and the
presence of PVAm and HSAm does not alter its formation.
30 As regards to rheological characteristics as previously
measured, the module of co-network PVAm (5 %) coHSAm (5 %) /
plasma IPN is 61 Pa (38 Pa for the co-network PVAm (5 %) co
HSAm(5 %) / Fb IPN.
The use of plasma instead of fibrinogen seems to increase
35 the storage modulus.
CA 02838126 2013-12-18
41
Regarding biodegradability : PVAm (10 %) / plasma IPN
retains its integrity after being incubated in the enzyme
solution of thermolysin. The co-network PVAm (5 %) coHSAm (5 %)
/ plasma IPN was completely degraded in the enzyme solution.
The feasibility and biodegradability of the plasma based IPN
co-network materials according to the invention have been
verified.
Material cytotoxicity
The impact of these different materials on fibroblast
viability, proliferation and their extracellular matrix
synthesis was explored. A cell suspension (50 000 cells / cm2
confluent density) is put in contact with the
PVAm(5%)coBSAm(5%)/Fb, PVAm(3%)coBSAm(7%)/Fb and BSAm(10%)/Fb
IPNs and with BSAm(10%) single network as controls. After
incubation between 24 h and 3 weeks, a metabolic assay (Alamar
Blue , supplementary data) shows no alteration of metabolic
activity in comparison with the control. Then cells were imaged
by confocal microscopy to determine their morphology and their
density on the materials. Finally, the proportion of dead cells
was assessed (Live/ Dead test) to conclude on the
biocompatibility of these different materials.
First the cell morphology was observed. Nuclei and actin
cytoskeleton were stained with DAPI and phalloidin,
respectively.
Whatever the material, cell morphology evolves over time.
After 24 h culture on the single BSAm(10%) network or on
BSAm(10%)/Fb IPN, cells are round or star shaped. After 1 to 3
weeks of culture on the BSAm(10%) single network, all cells are
elongated but show little spreading and many pseudopodia. The
cytoskeleton shows many stress fibers. Thus, cells do not
easily adhere on this network surface on which they do not
acquire their normal fibroblast morphology and do not
proliferate much. During the same time, cells are much more
spread on the BSAm(10%)/Fb IPN where they tend to adopt a
stellate morphology. Their cytoskeleton is visible with coarse
stress fibers and a substantial increase of the cell population
CA 02838126 2013-12-18
42
is observed during the 3 weeks of culture. The introduction of
fibrin into the BSAm network (BSAm(10%)/Fb IPN) allows
fibroblasts to spread, to increase their cell surface and to
proliferate up to total recovery of the material surface.
Cells do not spread and grow on PVAm single network. Adding
fibrin (PVAm(10%)/Fb IPN) improves the cell viability but does
not allow an important cell growth. On the contrary, when PVAm
is copolymerized with BSAm in an IPN architecture containing
fibrin (PVAm(5%)coBSAm(5%)/Fb and PVAm(3%)coBSAm(7%)/Fb IPNs),
cells are plated after 24 h of culture. After 1 and 3 weeks,
the cells are already polarized and adopt the typical fusiform
fibroblast morphology. The cytoskeleton fibers are also clearly
visible. Cell populations are almost confluent after the first
week of culture and the whole material surface is covered
uniformly by cells after 2 weeks of culture.
Cell density was quantified to ascertain these
observations. As shown in figure 5-A, this experiment
reinforces the imaging observations. For comparison, previous
experiments on non-biodegradable samples free of BSAm have
shown that cells remained round shaped and their population
decreased to 20 % after one week on PVAm single network. Cells
better adhered on PVAm(10%)/fibrin IPN, the cell population was
maintained on the material surface for 1 week and spread, but
no cell growth was observed. Here, on BSAm single network,
cells seem to survive over 3 weeks but they do not proliferate
while their number largely increases when cultivated on either
BSAm(10%)/Fb IPN or PVAmcoBSAm/Fb IPNs. In polymer
biomaterials, cell growth has been linked to biodegradability
and the present results confirm this assumption.
The measure of cell viability (figure 5-B) comforts these
observations. Cells rapidly die in contact to PVAm single
network surface; the presence of fibrin in the PVAm(10%)/Fb IPN
slow down their death but this effect is not sufficient to
allow their growth. When BSAm is present either into a single
network or IPNs, the cell viability stays close to 100% after 3
CA 02838126 2013-12-18
43
weeks, indicating that the protein environment favors cell
viability.
The IPNs synthesized with fibrin and BSAm have shown many
interesting properties. Hence, on these biodegradable and
fibrin based IPNs, cells truly survive and proliferate. However
to consider a potential application in tissue engineering the
maintaining of cell viability is not sufficient. Indeed, the
cells on the materials must also be able to generate their own
extracellular matrix and proceed to its remodeling to well
settle on the material. The appearance of fibronectin, the
first protein, and that of hyaluronic acid, the first
polysaccharide neo-synthesized after fibrin gel formation in
wound healing process (inflammatory phase) have been imaged by
confocal laser scanning microscopy. The synthesis of
fibronectin is important as this protein usually cross-link
with fibrin to form the basis matrix for healing. Moreover, in
healing, type III collagen is synthesized by fibroblasts during
the proliferative phase and then later replaced by type I
collagen during the remodeling phase allowing the further
epithelialization. This step is a marker of ECM remodeling and
maturation which indicates the good formation of ECM
(composition and dynamics) by fibroblasts.
Fibronectin, hyaluronic acid and collagen I syntheses were
followed on cell cultures also showing nuclei and the
cytoskeleton as a function of time using 5 concomitant
stainings. Fibronectin obviously appears upon time and this
increase in concentration is similar for all materials for 24 h
of culture. After a week, fibronectin is structured into
fibers, which number increases after 3 weeks. Due to the low
concentration of fibronectin in the culture medium, and the
increasing amount of fibronectin visualized along the 3 weeks
of cell culture it may be assumed that the observed fibronectin
has been secreted by fibroblasts during their growth and
development.
At 24 h, hyaluronic acid is detected at the surface and in
the cytoplasm of the cells. After 3 weeks of culture, its
CA 02838126 2013-12-18
44
presence has increased significantly. Comparing the stainings
of fibronectin and hyaluronic acid, a large colocation is
observable, which is consistent with the well-known affinity
between these two macromolecules. Thus, in addition of
fibronectin, fibroblasts secrete also hyaluronic acid during
their growth on the biomaterials. This macromolecule is known
to maintain good hydration of tissues.
From these analyses, it is shown that fibroblasts
developing at the surface of these IPNs are able to neo-
synthesize and secrete the first macromolecules that are
usually produced by this cell line after their adhesion on an
adapted support.
Monitoring of type I collagen is also particularly
interesting. Indeed, this protein is the main component of the
extracellular matrix of the dermis and its synthesis and
deposition are essential for the formation of healthy tissue.
However, in the healing process, fibroblasts first produce type
III collagen at the early stages of healing (from 10 h to 3
days) while type I collagen replaces type III collagen only
during the maturation phase. The appearance of collagen I thus
come later in the healing process than that of fibronectin and
hyaluronic acid and it is associated with a maturation of ECM.
In order to verify its presence, collagen type I present on the
surface of materials was immunostained. After 24 h of culture,
type I collagen is already detected on the edges of the cells
and then its presence increases significantly during the 3
weeks of culture. This staining intensification demonstrates
that the cell population present on the material surface is
capable of secreting it. It is known that induction of type I
collagen synthesis requires a good cell spreading; the
observation of type I collagen is an additional marker of the
good state of fibroblasts on the surface of the IPNs.
This synthesis of fibronectin, hyaluronic acid and type I
collagen was observed for all IPNs containing BSAm but was
maximal for the PVAm(5%)coBSAm(5%)/Fb IPN.
CA 02838126 2013-12-18
The various IPNs formed with BSAm and fibrin or with
different proportions of PVAm in addition to BSAm and fibrin
are all capable of supporting cell growth on their surface.
Fibroblasts cover the entirety of the material after 2 weeks of
5 culture and show a high rate of survival; this allows
qualifying these various materials as non-cytotoxic. Finally,
cells grown on the surface of these biomaterials and secrete
the key macromolecules of the human dermis ECM and are even
able to structure them as fibers as it could be observed for
10 fibronectin.
Conclusion
At this point, using BSAm alone or in addition to PVAm,
various IPN materials including a fibrin gel were obtained. To
our knowledge, the synthesis of a IPN constituted of a protein
15 gel obtained through an enzymatic reaction and of the co-
network of a functionalized synthetic polymer with a
functionalized protein had never been described in the
literature. Here several IPNs were obtained adding fibrin gel
to a co-network made of various proportions of PVAm and BSAm. A
20 one pot-one shot method was applied in each case leading to a
rapid synthesis (from 2 to 15 minutes). The mechanical
properties of all IPNs are good, they are all self-supported;
the IPN architecture allowed improving the storage modulus of
the fibrin gel by a factor ranging from 8 to 40 depending on
25 the polymer composition. All materials are well synthesized and
present an interpenetrated polymer networks architecture as
shown by low soluble fraction. The protein networks (fibrin and
BSAm) are homogeneously spread.
These biomaterials are biodegradable due to BSAm
30 introduction inside the material. This new type of materials is
the first co-network IPNs ever described in literature to be
potentially biodegradable through fragmentation, then
elimination. In addition, degradability helped to increase very
significantly the bioactivity of materials.
35 A pertinent choice of the dyes and antibodies used has
allowed us to perform concomitantly 5 different stainings on
CA 02838126 2013-12-18
46
the same sample and to observe at the same time both the cell
morphology and ECM remodeling. The utilization of such a large
number of different fluorescent probes is totally original.
When BSAm was used as a constituent of the Fb IPNs, the surface
of the obtained materials can be completely covered by
fibroblasts within two weeks, which was not the case of the
PVAm/Fb IPN. In addition, cells at the surface of the material
secrete the extracellular matrix
biomacromolecules
(fibronectin, hyaluronic acid) playing a key role in the early
stages of wound healing and are able to structure them into
fibers (fibronectin). The affinity properties of the molecules
are identical to those of dermal fibroblasts (co-localization
of fibronectin and hyaluronic acid). The good spreading of
fibroblast cells on the surface of the materials allows them to
produce type I collagen which is a protein characteristic of
late maturation phase.
Encapsulation of fibroblasts
To encapsulate fibroblasts in a PVAm (5 %) coBSAm (5 %) /
Fb, IPN the precursors of fibrin gel (fibrinogen and thrombin)
and those of the synthetic network (PVAm, BSAm and Irgacure
2959) were mixed in Tris buffer as detailed previously, and
then the solution was then added to a concentrated fibroblast
cells suspension. The resulting solution was placed in a mold
and then exposed to UV for 1 hour at 37 C for the synthesis of
IPN. The material is then immersed in a culture medium to 37 h
for 3 hours before repeating the medium to remove traces of
synthesis buffer. The measurements were made on materials
having 140 mm3 and 1000000 cells each (ie 7000 cells/mm3). Given
the high cell concentration culture medium is changed every 3
days.
The material containing cells were frozen at -80 C with
and without 3% glycerol. After a rapid thawing at 37 C, the
cells viability and cells proliferative activities were
measured.
CA 02838126 2013-12-18
47
Results
The cellular distribution and viability of the cells
population are homogeneous in thickness throughout the
material. The U.V exposure does not cause mortality of the
cells on the surface of the material. After 1 and 3 weeks of
culture, cell population is at least maintained and the
distribution remains homogeneous. On figure 8 it can be seen
that the cell viability remains identical during cultivation
time from 72 h to 840 h. In addition, the cells initially
present on the surface of the material are strongly expanded
and cover the material. This distribution and the absence of
mortality confirm that there is no impact of UV exposure on the
cell population. About 3 weeks of culture, cell death is low
and homogeneous within the material. The fact that the location
of the cells in the depth of the material does not cause
mortality suggests that the flow of nutrients and oxygen runs
smoothly and there is no phenomenon of cellular hypoxia. After
3 weeks of culture, the average viability remains at or above
95 %. These results confirm that cell encapsulation is possible
in this matrix without toxicity.
From the point of view of cell morphology, after a short
time of culture (24 hours) the cells are round regardless of
the depth where they are located. After 1 week of culture, all
the cells were elongated. Some cells show many very elongated
pseudopodia. After 3 weeks, cells with numerous pseudopodia are
preserved thoroughly. Similarly, the cells on the surface of
the material are elongated and have a profile of fibroblasts.
It has been noticed that fibronectin is present after 24
hours of culture. Between 1 and 3 weeks of culture, a
significant increase in the presence of fibronectin is visible
within the depth of the material. In depth, fibronectin is
located either around the cells or in the form of flakes.
Similarly, the presence of collagen is detected as from 24
hours of culture in the IPN and increases significantly after
one and three weeks of culture. Finally, it appears that its
presence is greater within depth in the material than at its
CA 02838126 2013-12-18
48
surface. These data confirm that fibroblasts retain their
ability to secrete collagen when encapsulated in a matrix PVAm
(5 %) coBSAm (5 %) / Fb IPN.
All of these results suggest that it is possible to
produce cellularized dressings by encapsulation of fibroblasts
within a material according to the invention such as PVAm (5 %)
coBSAm (5 %) /Fb IPN. This type of material would be able to
promote wound healing and secretion of extracellular matrix
macromolecules such as fibronectin and collagen. The
possibility of encapsulating living cells within the material
of the invention and the capacity of the cells to proliferate
in it makes this material capable of being used as a
cellularized skin substitute for the treatment of burns for
example.
As to frozen and unfrozen samples, the results shown in
figures 9 and 10 show that the viability of encapsulated
cryopreserved fibroblasts remain at a level of 95% after 96h
and 168h of cultivation after thawing (figure 9) with or
without glycerol. This level is almost identical to the control
level corresponding to non-frozen samples. As to the
proliferative activity of the cryopreserved encapsulated cells,
as seen from figure 10, the presence of 3 % glycerol allows a
quicker response in the restart of proliferative activity which
is however not inhibited by the freezing.
