U.S. patent application number 11/529064 was filed with the patent office on 2007-03-29 for biodegradable nanocomposites with enhance mechanical properties for soft tissue.
This patent application is currently assigned to NORTHWESTERN UNIVERSITY. Invention is credited to Guillermo Ameer, Antonio R. Webb.
Application Number | 20070071790 11/529064 |
Document ID | / |
Family ID | 37900436 |
Filed Date | 2007-03-29 |
United States Patent
Application |
20070071790 |
Kind Code |
A1 |
Ameer; Guillermo ; et
al. |
March 29, 2007 |
Biodegradable nanocomposites with enhance mechanical properties for
soft tissue
Abstract
The present invention is directed to a novel poly(diol
citrates)-based nanocomposite materials created using completely
biodegradable and biocompatible polymers that may be used in tissue
engineering. More specifically, the specification describes methods
and compositions for making and using nanocomposites comprised of
citric acid copolymers and polymers including but not limited to
poly(L-lactic acid) (PLLA) and poly(lactic-co-glycolic acid)
(PLGA).
Inventors: |
Ameer; Guillermo; (Chicago,
IL) ; Webb; Antonio R.; (Evanston, IL) |
Correspondence
Address: |
MARSHALL, GERSTEIN & BORUN LLP
233 S. WACKER DRIVE, SUITE 6300
SEARS TOWER
CHICAGO
IL
60606
US
|
Assignee: |
NORTHWESTERN UNIVERSITY
Evanston
IL
|
Family ID: |
37900436 |
Appl. No.: |
11/529064 |
Filed: |
September 28, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60721687 |
Sep 28, 2005 |
|
|
|
Current U.S.
Class: |
424/423 |
Current CPC
Class: |
C08L 67/04 20130101;
A61L 27/18 20130101; A61K 47/30 20130101; A61K 31/74 20130101; A61K
31/765 20130101; A61K 47/34 20130101; A61L 2400/12 20130101; A61K
9/0024 20130101; A61L 27/18 20130101; A61L 27/58 20130101; A61L
27/50 20130101 |
Class at
Publication: |
424/423 |
International
Class: |
A61F 2/02 20060101
A61F002/02 |
Claims
1. A composition comprising a composite of: a) a citric acid
polyester having the generic formula (A-B-C)n, wherein A is a
linear aliphatic dihydroxy monomer; B is citric acid, C is a linear
aliphatic dihydroxy monomer, and n is an integer greater than 1;
and b) a biodegradable polymer used for implantable tissue
devices.
2. The composition of claim 1, wherein A is a linear diol
comprising between about 2 and about 20 carbons.
3. The composition of claim 1, wherein C is a linear diol
comprising between about 2 and about 20 carbons.
4. The composition of claim 1, wherein both A and C are the same
linear diol.
5. The composition of claim 4, wherein said linear diol is 1,8,
octanediol.
6. The composition of claim 1, wherein A and C are different linear
diols.
7. The composition of claim 5, wherein said linear aliphatic
dihydroxy poly 1,8-octanediol co-citric acid.
8. The composition of claim 5, wherein said linear aliphatic
dihydroxy poly 1,10-decanediol co-citric acid.
9. The composition of claim 1, wherein said polymer is selected
from the group consisting of poly(hydroxyvalerate),
poly(lactide-co-glycolide), poly(hydroxybutyrate),
poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,
poly(glycolic acid), poly(glycolide), poly(L-lactic acid),
poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),
poly(caprolactone), poly(trimethylene carbonate), and polyester
amide.
10. The composition of claim 1, wherein said polymer is
poly(L-lactic acid.
11. The composition of claim 1, wherein said polymer is
poly(lactic-co-glycolic acid) (PLGA).
12. The composition of claim 1, wherein said composition further
comprises a drug.
13. A substrate for use in tissue engineering comprising a citric
acid polymer of any of claims 1-11 formulated into a cell culture
substrate.
14. The substrate of claim 13, wherein said substrate further
comprises a surface modification.
15. The substrate of claim 13, wherein said polymer is
biodegradeable.
16. A method of producing engineered tissue, comprising: a.
preparing a scaffold of a composition of any of claims 1-11; and b.
culturing cells of said tissue on the scaffold of part (a).
17. The method of claim 16, wherein said cells are selected from
the group consisting of endothelial cells, ligament tissue, muscle
cells, bone cells, cartilage cells.
18. The method of claim 16, wherein said engineered tissue is
produced in a bioreactor.
19. A drug delivery device comprising a drug interspersed in the
polymer composition of any of claims a composition of any of claims
1-11.
20. The composition of claim 1 wherein said biodegradable polymer
is fabricated into one or more nanostructures.
21. The composition of claim 1, wherein said biodegradable polymer
is fabricated into nanofibers.
22. The composition of claim 1, wherein said biodegradable polymer
is formulated into nanoparticles.
23. A biocompatible, biodegradable composition comprising: a. an
elastomeric component comprised of a biodegradeable biocompatible
polymer and b. a strengthening component comprised of a second
biodegradable biocompatible polymer, wherein said second
biodegradable, biocompatible polymer is fabricated into one or more
nanostructures.
Description
[0001] The present application claims the benefit of priority of
U.S. Provisional Application No. 60/721,687 which was filed Sep.
28, 2005. The entire text of the aforementioned application is
incorporated herein by reference.
FIELD OF THE INVENTION
[0002] The present invention is generally directed to a substrate
used for tissue engineering. The substrate is a biodegradable
elastomeric polymer. Methods and compositions for testing and using
the same are disclosed.
BACKGROUND
[0003] The need for biodegradable polymers in emerging technologies
such as tissue engineering, drug delivery, and gene therapy has
been fueling a quest for novel biodegradable polymers [13-16]. In
particular, biodegradable polymers with elastomeric properties have
recently received attention for their potential use in the
engineering of soft tissues such as blood vessel, heart valves,
cartilage, tendon, and bladder, which exhibit elastic properties.
Due to their long history of use in clinical applications,
poly(hydroxyortho esters) such as polyglycolic acid (PGA), poly
lactic acid (PLA) and copolymers thereof are often used to
fabricate three-dimensional porous scaffolds to support cell
attachment, proliferation, migration, and extracellular matrix
synthesis. The development of a novel family of biodegradable
elastomeric polymers referred to as poly(diol citrates) has been
reported [1].
[0004] Tissues such as blood vessels, cartilage, ligament, and
tendon have specific biomechanical requirements for successful
functional tissue engineering. These tissues are often subjected to
relatively large tensile or compressive forces, so it is important
that synthetic scaffolds or implants intended to model such tissues
have the necessary tensile strength, elasticity, and compressive
modulus to withstand such forces. Ideally, the mechanical
properties of the scaffold or implant would approximate those of
the natural tissue it is designed to mimic. The reported tensile
strength of human cartilage and ligament are 3.7-10.5 MPa and
24-112 MPa, respectively. The reported Young's modulus of cartilage
and ligament are 0.7-15.3 MPa and 65-541 MPa, respectively. The
reported tensile strength of human coronary arteries is 1.4-11.14
MPa.
[0005] It is well known in the art that the mechanical properties
of elastomers can be enhanced by the fabrication of composites in
which a second component or phase is added to the elastomeric
phase. One method by which elastomers can be strengthened and
stiffened is by incorporating nanoparticles into the elastomeric
matrix [Lavik, E. and R. Langer, Tissue engineering: current state
and perspectives. Appl Microbiol Biotechnol, 2004. 65: p. 1-8;
Okada, M., Chemical syntheses of biodegradable polymers. Prog Polym
Sci, 2002. 27: p. 87-133; Griffith., Polymeric Biomaterials. Acta
Mater, 2000. 48: p. 263-277; MacDonald, J. Biomed. Mater. Res. A,
2005, 74, 489-496]. This is the case in the rubber industry where
carbon black nanoparticles can be added to greatly increase the
mechanical properties [Lavik, E. and R. Langer, Tissue engineering:
current state and perspectives. Appl Microbiol Biotechnol, 2004.
65: p. 1-8]. The nanoparticles act as additional crosslink points
to reinforce the network chains and in general, the increase in
mechanical properties is inversely proportional to the nanoparticle
diameter [Lavik, E. and R. Langer, Tissue engineering: current
state and perspectives. Appl Microbiol Biotechnol, 2004. 65: p.
1-8]. Although this method has been used for industrial
applications, there have been no reports involving the use of
biocompatible, biodegradable nanoparticles to strengthen matrices
intended for in vivo use.
[0006] Poly(hydroxyortho esters) or other polymers have been mixed
with ceramics, glass microparticles, glass nanoparticles, glass
nanofibers, or carbon nanotubes to strengthen scaffolds for bone
tissue engineering applications and, to a lesser extent, for soft
tissue regeneration. However, most of these approaches introduce
inorganic and non-biodegradable components into the polymer
composite. A non-degradable second phase may interfere with the
body's natural remodeling mechanisms as the continuous presence of
a foreign material may induce long-term inflammatory responses.
Furthermore, the resulting composite does not exhibit the
elasticity and flexibility that is important for soft tissue
engineering.
[0007] Chun et al (U.S. patent application Ser. No. 10/383,369) and
Melican et al (U.S. patent application Ser. No. 09/747,489)
disclose tissue implants comprising a biodegradable mesh
reinforcement component and a biodegradable elastomeric foam
component. Ma et al (U.S. Pat. No. 6,146,892) disclose
three-dimensional biodegradable matrices comprised of nanofibers.
However, Chun et al, Melican et al, and Ma et al do not disclose
composites having mechanical properties approaching those of
natural soft tissue.
[0008] Analogous to rubber which is a three dimensional network of
crosslinked polymer chains, poly(diol citrates) are composed of
three-dimensional polyester networks formed by reacting citric acid
with various aliphatic diols. The mechanical properties could be
varied depending on the selection of diols and the applied
post-polymerization conditions. In general, longer chain diols have
a lower tensile strength and modulus, while increasing
polymerization time and/or temperature increase the tensile
strength and modulus. Preliminary in vitro cell culture evaluation
of poly(diol citrates) showed their great potential as
"cell-friendly" materials, as both smooth muscle and endothelial
cells attach and proliferate on the surface. Methods of preparation
of poly(diol) citrates are described in detail in U.S. patent
application Ser. No. 10/945,354 (incorporated herein by reference
and also shown in the Examples below). In vivo biocompatibility
results show a thin vascularized collagenous capsule after 4 months
of implantation with no inflammation. The thickness of this capsule
was smaller than that reported for poly(L-lactide-co-glycolide)
(PLGA) [17]. A thin vascularized capsule is considered to be
beneficial for mass transfer between a cell-based implant and
surrounding tissues.
[0009] Although the mechanical properties of synthetic polymers, in
particular poly(diol citrates), can be varied to meet specific
applications, it can be desirable to further increase the strength
and stiffness while maintaining the ability to be elongated to many
times their original length before rupture. The present invention
is directed to optimizing the strength and elasticity of
biocompatible scaffolds by preparing a composite comprising an
elastomeric polymer strengthened by the presence of a biodegradable
polymeric nanostructure.
SUMMARY OF THE INVENTION
[0010] The present invention describes a composition comprising a
biodegradable elastomeric polymeric component and a biodegradable
polymeric nano-structure. The present invention describes a
composition comprising a composite of a citric acid polyester
having the generic formula (A-B-C)n, wherein A is a linear
aliphatic dihydroxy monomer; B is citric acid, C is a linear
aliphatic dihydroxy monomer, and n is an integer greater than 1;
and a biodegradable polymer used for implantable tissue devices.
Preferably, the biodegradable polymer is fabricated into a
nanostructure such as a nanofiber, a nanoparticle, or the like.
[0011] In specific embodiments, A is a linear diol comprising
between about 2 and about 20 carbons. In other embodiments, C is
independently a linear diol comprising between about 2 and about 20
carbons. While in certain embodiments, both A and C may be the same
linear diol, other embodiments contemplate that A and C are
different linear diols. A particularly preferred linear diol is
1,8, octanediol. In other embodiments, one or both of A and C may
be 1,10-decanediol. The diol also may be an unsaturated diol, e.g.,
tetradeca-2,12-diene-1,14-diol, or other diols including
macromonomer diols such as polyethylene oxide, and
N-methyldiethanoamine (MDEA). This family of elastomers is named as
poly(diol citrate). In particularly preferred embodiments, the
composition of the invention is dihydroxy poly 1,8-octanediol
co-citric acid. Poly(diol citrate) can also form hybrids with other
materials like hydroxyapatite to form elastomeric composites.
[0012] In some embodiments the nano-structure of the invention is
fabricated from a polymer. Preferably the polymer is a
biodegradable polymer selected from the group consisting of
poly(hydroxyvalerate), poly(lactide-co-glycolide),
poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate),
polyorthoester, polyanhydride, poly(glycolic acid),
poly(glycolide), poly(L-lactic acid), poly(L-lactide),
poly(D,L-lactic acid), poly(D,L-lactide), poly(caprolactone),
poly(trimethylene carbonate), and polyester amide. In more
preferred embodiments, the polymer is poly(L)-lactic acid,
poly(lactic-co-glycolic acid) (PLGA) or a combination of the
two.
[0013] In specific embodiments, the composition further comprises a
drug.
[0014] Another aspect of the invention contemplates a substrate
that may be formulated for tissue culture and/or tissue engineering
wherein the substrate is made of a composition as described herein
above. In preferred embodiments, the substrate may further comprise
a surface modification that allows cellular attachment. Preferably,
the polymer of the invention employed as cell/tissue culture
substrate is biodegradable. Preferably, the polymer also is
biocompatible. The term "biocompatible" is intended to encompass a
polymer that may be implanted in vivo or alternatively may be used
for the growth of cells that may be implanted in vivo without
producing an adverse reaction, such as an immunological response or
otherwise altering the morphology of the cells grown thereon to
render the cells incompatible with being implanted in vivo or used
to model an in vivo organ.
[0015] Also contemplated herein is a method of producing engineered
tissue, comprising providing a biodegradable composition of the
present invention as a scaffold for the growth of cells and
culturing cells of said tissue on the scaffold. In preferred
methods, the polymer is composite of a poly
1,8-octanediol-co-citric acid, or a derivative thereof, or a poly
1,10-decanediol-co-citric acid or derivative thereof in combination
with a biodegradable polymer such as PGLA or PLLA. In specific
embodiments, the cells are selected from the group consisting of
connective tissue cells, organ cells, muscle cells, nerve cells,
and any combination thereof. In more specific embodiments, the
cells are selected from the group consisting of tenocytes,
fibroblasts, ligament cells, endothelial cells, lung cells,
epithelial cells, smooth muscle cells, cardiac muscle cells,
skeletal muscle cells, islet cells, nerve cells, hepatocytes,
kidney cells, bladder cells, urothelial cells, chondrocytes, and
bone-forming cells. In other preferred embodiments, the tissue
engineering method comprises growing the cells on the scaffold in a
bioreactor.
[0016] In further embodiments, the compositions of the invention
may be used as bandages, patches or sutures for implantation during
surgery. In still other embodiments, the compositions of the
present invention may be used to form a drug delivery device
comprising a drug interspersed in the polymer composition of the
invention.
[0017] Other features and advantages of the invention will become
apparent from the following detailed description. It should be
understood, however, that the detailed description and the specific
examples, while indicating preferred embodiments of the invention,
are given by way of illustration only, because various changes and
modifications within the spirit and scope of the invention will
become apparent to those skilled in the art from this detailed
description.
BRIEF DESCRIPTION OF THE DRAWINGS
[0018] The following drawings form part of the present
specification and are included to further illustrate aspects of the
present invention. The invention may be better understood by
reference to the drawings in combination with the detailed
description of the specific embodiments presented herein.
[0019] FIG. 1: Poly(diol citrate) reaction scheme.
[0020] FIG. 2: PLGA Nanoparticle Formation.
[0021] FIG. 3: SEM micrographs of PLLA nanoscaffolds before and
after penetrating the pores with PDC. (Scale bars=1 .mu.m).
[0022] FIG. 4A-C: (A) Tensile strength, (B) Young's modulus, (C)
Elongation at break of PLLA-PDC nanoscaffold composites and PDC
control that were polymerized at 120.degree. C. for 1 day with
vacuum (n=4). (* p<0.01 from control).
[0023] FIG. 5A-C: SEM micrographs of (A) PLGA nanoparticles (Scale
Bar=1 .mu.m), (B) 5% PLGA-PDC nanocomposite (Scale Bar=2 .mu.m),
(C) 10% PLGA-PDC nanocomposite (Scale Bar=2 .mu.m).
[0024] FIG. 6A-C: (A) Tensile strength, (B) Young's modulus, (C)
Elongation at break of PLLA-PDC nanoscaffold composites and PDC
control that were polymerized at 80.degree. C. for 3 days without
vacuum (n=4). (* p<0.01 from control).
[0025] FIG. 7A-D--(A) Tensile strength, (B) Young's modulus, (C)
Elongation at break of PLLA-PDC nanocomposites and controls that
were polymerized at 80.degree. C. for 3 day without vacuum (n=4).
(d) Representative stress-strain curves. PLLA controls were not
shown for clarity. (* p<0.05 from PDC control, # p<0.05 from
PLLA control, $ p<0.05 comparing 5% PLLA-PDC to 10%
PLLA-PDC).
[0026] FIG. 8A-C--(A) Tensile strength, (B) Young's modulus, (C)
Elongation at break of PLLA-PDC composites and controls that were
polymerized at 80.degree. C. for 3 day without vacuum or
120.degree. C. for 1 day without vacuum then 120.degree. C. for 1
day with vacuum (n=4). (* p<0.05 from PDC controls, # p<0.05
comparing between polymerization conditions for identical
materials).
[0027] FIG. 9A-C--(A) Tensile strength, (B) Young's modulus, (C)
Elongation at break of PLLA-PDC and PLLA-POC composites and
controls that were polymerized at 80.degree. C. for 3 day without
vacuum (n=4) (* p<0.05 from poly(diol citrate) controls, #
p<0.05 comparing between composite and PLLA control).
[0028] FIG. 10--Compressive modulus of PDC-PLLA nanofibrous
scaffold composites. (* p<0.05 from PDC control, # p<0.05
from PLLA control, $ p<0.05 comparing between 5% PLLA-PDC and
10% PLLA-PDC)
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0029] Poly(diol citrates) are a family of biodegradable and
biocompatible elastomers that have shown significant potential for
soft tissue engineering. As noted above, it is desirable to
increase the strength and stiffness while maintaining the ability
to be elongated to many times their original length before rupture.
The methods of the present invention are directed to strengthening
such elastomers. One method by which elastomers can be strengthened
and stiffened is by incorporating nanoparticles into the
elastomeric matrix [2-4, 18]. This is case in the rubber industry
where carbon black nanoparticles can be added to greatly increase
the mechanical properties [2]. The nanoparticles act as additional
crosslink points to reinforce the network chains and in general,
the increase in mechanical properties is inversely proportional to
the nanoparticle diameter [2].
[0030] Since poly(diol citrates) are targeted for soft tissue
engineering, the nanophase used to reinforce the material must be
made from a biodegradable and biocompatible polymer. Two polymers
were chosen, poly(L-lactic acid) (PLLA) and poly(lactic-co-glycolic
acid) (PLGA). These polymers were chosen as they are rigid and
strong and have been used in many tissue engineering applications
[19]. Furthermore, the rate of degradation could be tailored to
match that of the surrounding elastomeric matrix. Poly(L-lactic
acid) has a degradation time of greater than two years while
poly(glycolic acid) has a degradation time of 1-2 months [20]. By
changing the ratio of lactic to glycolic acid, the degradation rate
could be varied from fast (1-2 months) to slow (>2 years). For
tissue engineering, the rate of degradation of the polymer scaffold
should match that of tissue regrowth.
[0031] Two methods were used to create biodegradable
nanocomposites: (A) Nanoscaffolds of PLLA, (B) Nanoparticles of
PLGA. The nanoscaffold method was chosen because it was believed
that a rigid nano-fiber could scaffold could provide reinforcement
to the material. Furthermore, a method to produce synthetic
nanofibrous matrices was already developed elsewhere [21]. The
nanoparticle method was chosen as it analogous to what is commonly
done with elastomers to increase the mechanical properties of
rubber [2-4, 18]. In addition, a simple method was found to produce
large quantities of nanoparticles [22, 23]. Furthermore, it has
been demonstrated that nanoparticles are a viable method by which
drugs can be incorporated and released, thus allowing the
possibility for drug release from nanocomposite materials [9-12].
These methods are exemplified in Example 2 below.
[0032] Thus, in order to enhance the mechanical properties of
poly(diol citrates) nanocomposite materials were created using
completely biodegradable and biocompatible polymers. This is the
first report of nanocomposites for tissue engineering where the
macro and nano phases are completely made from biodegradable and
biocompatible synthetic co-polymers. Composites were created using
nanoscaffolds and nanoparticles. The nanoparticle method is
analogous to what is commonly done with rubber materials which are
often doped with silica or carbon black nanoparticles to greatly
increase their mechanical properties. However, because poly(diol
citrates) are being targeted for tissue engineering, nondegradable
carbon black or silica nanoparticles are not appropriate for the
reinforcing nanophase. Poly(L-lactic acid) and
poly(lactic-co-glycolic acid), two commonly used polymers for
tissue engineering were used. These polymers were selected for
their biocompatibility, rigid mechanical properties, and ease of
fabrication into nanoscaffolds and nanoparticles. Creation of
nanocomposites significantly increased the tensile strength and
modulus compared to control materials. In particular, it should be
noted that the tensile strength and elongation at break of the most
preferred compositions of the present invention are greater than
the strength and elongation at break of either of the polymer
components individually. Since the materials are fabricated from
two biodegradable and biocompatible materials, it is believed that
the resulting nanocomposites could be used for tissue engineering
applications. Furthermore, drug encapsulation and release has from
nanoparticles has been demonstrated, thus raising the possibility
of drug release from nanocomposites.
[0033] In preferred embodiments, the compositions of the present
invention can be reinforced with a biocompatible, biodegradable
mesh reinforcement component and/or with a three dimensional matrix
comprised of nanofibers. The compositions of the present invention
are such that they have a tensile strength and elasticity that is
similar to soft tissue. For example, the compositions of the
invention are prepared from a nanostructure that could be mesh or
nanofiber matrix that has a tensile strength similar to that of
human cartilage, human coronary arteries, or human ligament. As
used herein, the term "nano-structure" is intended to encompass an
object with at least one characteristic length between about 10 nm
and about 500 nm. In specific embodiments of the invention, the
biodegradable polymeric nano-structure is a nanoparticle. In a
preferred embodiment the nano-structure is a nanofiber. In the most
preferred embodiment, the nano-structure is a nanofibrillar
network. Uses for such networks are described in for example, U.S.
Pat. No. 6,599,323 and U.S. Pat. No. 6,146,892 (each incorporated
herein by reference).
[0034] The compositions described herein in some aspects comprise a
biodegradable elastomeric polymeric component and a biodegradable
polymeric nano-structure. In certain embodiments, the
nanocomposites of the invention are comprised of poly(diol citrate)
molecules and a polymer that strengthens the poly(diol citrate)
matrix. The polymer may be a biodegradable polymer or a
non-biodegradable polymer, but preferably is a biodegradable
polymer. Biodegradable polymers include, but are not limited to
collagen, elastin, hyaluronic acid and derivatives, sodium alginate
and derivatives, chitosan and derivatives gelatin, starch,
cellulose polymers (for example methylcellulose,
hydroxypropylcellulose, hydroxypropylmethylcellulose,
carboxymethylcellulose, cellulose acetate phthalate, cellulose
acetate succinate, hydroxypropylmethylcellulose phthalate), casein,
dextran and derivatives, polysaccharides, poly(caprolactone),
fibrinogen, poly(hydroxyl acids), poly(L-lactide) poly(D,L
lactide), poly(D,L-lactide-co-glycolide),
poly(L-lactide-co-glycolide), copolymers of lactic acid and
glycolic acid, copolymers of .epsilon.-caprolactone and lactide,
copolymers of glycolide and epsilon.-caprolactone, copolymers of
lactide and 1,4-dioxane-2-one, polymers and copolymers that include
one or more of the residue units of the monomers D-lactide,
L-lactide, D,L-lactide, glycolide, .epsilon.-caprolactone,
trimethylene carbonate, 1,4-dioxane-2-one or 1,5-dioxepan-2-one,
poly(glycolide), poly(hydroxybutyrate), poly(alkylcarbonate) and
poly(orthoesters), polyesters, poly(hydroxyvaleric acid),
polydioxanone, poly(ethylene terephthalate), poly(malic acid),
poly(tartronic acid), polyanhydrides, polyphosphazenes, poly(amino
acids). The biodegradeable polymers used herein may be copolymers
of the above polymers as well as blends and combinations of the
above polymers. (see generally, Illum, L., Davids, S. S. (eds.)
"Polymers in Controlled Drug Delivery" Wright, Bristol, 1987;
Arshady, J. Controlled Release 17:1-22, 1991; Pitt, Int. J. Phar.
59:173-196, 1990; Holland et al., J. Controlled Release 4:155-0180,
1986).
[0035] In particular preferred embodiments, the biodegradable or
resorbable polymer is one that is formed from one or more monomers
selected from the group consisting of lactide, glycolide,
e-caprolactone, trimethylene carbonate, 1,4-dioxan-2-one,
1,5-dioxepan-2-one, 1,4-dioxepan-2-one, hydroxyvalerate, and
hydroxybutyrate. In one aspect, the polymer may include, for
example, a copolymer of a lactide and a glycolide. In another
aspect, the polymer includes a poly(caprolactone). In yet another
aspect, the polymer includes a poly(lactic acid),
poly(L-lactide)/poly(D,-L-Lactide) blends or copolymers of
L-lactide and D,L-lactide. In yet another aspect, the polymer
includes a copolymer of lactide and e-caprolactone. In yet another
aspect, the polymer includes a polyester (e.g., a
poly(lactide-co-glycolide). The poly(lactide-co-glycolide) may have
a lactide:glycolide ratio ranges from about 20:80 to about 2:98, a
lactide:glycolide ratio of about 10:90, or a lactide:glycolide
ratio of about 5:95. In one aspect, the poly(lactide-co-glycolide)
is poly(L-lactide-co-glycolide; see e.g., U.S. Pat. No. 6,531,146
and U.S. application No. 2004/0137033.). Other examples of
biodegradable materials include polyglactin, and polyglycolic
acid.
[0036] Representative examples of non-biodegradable compositions
include ethylene-co-vinyl acetate copolymers, acrylic-based and
methacrylic-based polymers (e.g., poly(acrylic acid),
poly(methylacrylic acid), poly(methylmethacrylate),
poly(hydroxyethyl methacrylate), poly(alkylcynoacrylate),
poly(alkyl acrylates), poly(alkyl methacrylates)), polyolefins such
as poly(ethylene) or poly(propylene), polyamides (e.g., nylon 6,6),
poly(urethanes) (e.g., poly(ester urethanes), poly(ether
urethanes), poly(carbonate urethanes), poly(ester-urea)),
polyesters (e.g., PET, polybutyleneterephthalate, and
polyhexyleneterephthalate), olyethers(poly(ethylene oxide),
poly(propylene oxide), poly(ethylene oxide)-poly(propylene oxide)
copolymers, diblock and triblock copolymers, poly(tetramethylene
glycol)), silicone containing polymers and vinyl-based polymers
(polyvinylpyrrolidone, poly(vinyl alcohol), poly(vinyl acetate
phthalate), poly(styrene-co-isobutylene-co-styrene), fluorine
containing polymers (fluoropolymers) such as fluorinated ethylene
propylene (FEP) or polytetrafluoroethylene (e.g., expanded
PTFE).
[0037] The polymers may be combinations of biodegradable polymers,
and also may be combinations of biodegradable and non-biodegradable
polymers. Further examples of polymers that may be used are either
anionic (e.g., alginate, carrageenin, hyaluronic acid, dextran
sulfate, chondroitin sulfate, carboxymethyl dextran, caboxymethyl
cellulose and poly(acrylic acid)), or cationic (e.g., chitosan,
poly-I-lysine, polyethylenimine, and poly(allyl amine)) (see
generally, Dunn et al., J. Applied Polymer Sci. 50:353, 1993;
Cascone et al., J. Materials Sci.: Materials in Medicine 5:770,
1994; Shiraishi et al., Biol. Pharm. Bull. 16:1164, 1993;
Thacharodi and Rao, Int'l J. Pharm. 120:115, 1995; Miyazaki et al.,
Int'l J. Pharm. 118:257, 1995). Preferred polymers (including
copolymers and blends of these polymers) include
poly(ethylene-co-vinyl acetate), poly(carbonate urethanes),
poly(hydroxyl acids) (e.g., poly(D,L-lactic acid) oligomers and
polymers, poly(L-lactic acid) oligomers and polymers, poly(D-lactic
acid) oligomers and polymers, poly(glycolic acid), copolymers of
lactic acid and glycolic acid, copolymers of lactide and glycolide,
poly(caprolactone), copolymers of lactide or glycolide and
.epsilon.-caprolactone), poly(valerolactone), poly(anhydrides),
copolymers prepared from caprolactone and/or lactide and/or
glycolide and/or polyethylene glycol.
[0038] In specific embodiments, the compositions of the invention
(i.e., the compositions that are made up of a poly(diol citrate)
polymer and a polymer that stiffens the poly diol citrate polymer)
are molded or otherwise formed into flexible compositions that can
be used as "patches." Such "patches" may be comprises of just the
poly(diol citrate) polymer and a polymer, or they may be
impregnated or otherwise loaded with a drug or other biologically
active agent to be delivered (e.g., in a controlled-release manner)
or they may be seeded with cells so that they can act as cellular
tissue patches or tissue grafts.
[0039] In specific embodiments, the elastomeric composites made
according to the methods of the present invention will be useful
both as substrata for the growth and propagation of tissues cells
that may be seeded on the substrata and also as implantable
devices. In those embodiments where the elastomeric composites are
used as bioimplantable devices, the substrate may be formulated
into a shape suitable for implantation. For example, as described
in U.S. Pat. No. 6,423,092 (incorporated herein by reference), it
may be desirable to fabricate a biodegradable stent for
implantation into a lumen. The composites of the invention may also
be used in craniofacial applications, for example, as an
alternative to spring-mediated cranial reshaping for the treatment
of craniosynostosis [27]. In certain embodiments of the invention,
it may be desirable to produce prosthetic organ tissue for
implantation into an animal, such as e.g., testicular tissue
described in U.S. Pat. No. 6,620,203(incorporated herein by
reference). Other organs for which tissue implantation patches may
be generated include, but are not limited to skin tissue for skin
grafts, myocardial tissue, bone tissue for bone regeneration,
testicular tissue, endothelial cells, blood vessels, and any other
cells from which a tissue patch may be generated. Thus, those of
skill in the art would understand that the aforementioned
organs/cells are merely exemplary organs/cell types and it should
be understood that cells from any organ may be seeded onto the
biocompatible elastomeric composites of the invention to produce
useful tissue for implantation and/or study.
[0040] The cells that may be seeded onto the elastomeric composites
of the present invention may be derived from commercially available
cell lines, or alternatively may be primary cells, which can be
isolated from a given tissue by disaggregating an appropriate organ
or tissue which is to serve as the source of the cells being grown.
This may be readily accomplished using techniques known to those
skilled in the art. Such techniques include disaggregation through
the use of mechanically forces either alone or in combination with
digestive enzymes and/or chelating agents that weaken cell-cell
connections between neighboring cells to make it possible to
disperse the tissue into a suspension of individual cells without
appreciable cell breakage. Enzymatic dissociation can be
accomplished by mincing the tissue and treating the minced tissue
with any of a number of digestive enzymes either alone or in
combination. Digestive enzymes include but are not limited to
trypsin, chymotrypsin, collagenase, elastase, and/or hyaluronidase,
Dnase, pronase, etc. Mechanical disruption can also be accomplished
by a number of methods including, but not limited to the use of
grinders, blenders, sieves, homogenizers, pressure cells, or
sonicators to name but a few. For a review of tissue disaggregation
techniques, see Freshney, Culture of Animal Cells. A Manual of
Basic Technique, 2d Ed., A. R. Liss, Inc., New York, 1987, Ch. 9,
pp. 107-126.
[0041] Once the primary cells are disaggregated, the cells are
separated into individual cell types using techniques known to
those of skill in the art. For a review of clonal selection and
cell separation techniques, see Freshney, Culture of Animal Cells.
A Manual of Basic Techniques, 2d Ed., A. R. Liss, Inc., New York,
1987, Ch. 11 and 12, pp. 137-168. Media and buffer conditions for
growth of the cells will depend on the type of cell and such
conditions are known to those of skill in the art.
[0042] In certain embodiments, it is contemplated that the cells
attached to the biocompatible elastomeric composite substrates of
the invention are grown in bioreactors. A bioreactor may be of any
class, size or have any one or number of desired features,
depending on the product to be achieved. Different types of
bioreactors include tank bioreactors, immobilized cell bioreactors,
hollow fiber and membrane bioreactors as well as digesters. There
are three classes of immobilized bioreactors, which allow cells to
be grown: membrane bioreactors, filter or mesh bioreactors, and
carrier particle systems. Membrane bioreactors grow the cells on or
behind a permeable membrane, allowing the nutrients to leave the
cell, while preventing the cells from escaping. Filter or mesh
bioreactors grow the cells on an open mesh of an inert material,
allowing the culture medium to flow past, while preventing the
cells from escaping. Carrier particle systems grow the cells on
something very small, such as small nylon or gelatin beads. The
bioreactor can be a fluidized bed or a solid bed. Other types of
bioreactors include pond reactors and tower fermentors. Any of
these bioreactors may be used in the present application for
regenerating/engineering tissues on the citric acid based
elastomeric composites of the present invention.
[0043] Certain tissues that are regenerated by use of the citric
acid based elastomeric composites of the invention may be
encapsulated so as to allow the release of desired biological
materials produced by the cells at the site of implantation, while
sequestering the implanted cells from the surrounding site. Cell
encapsulation can be applied to all cell types secreting a
bioactive substance either naturally or through genetic engineering
means. In practice, the main work has been performed with insulin
secreting tissue.
[0044] Encapsulation procedures are most commonly distinguished by
their geometrical appearance, i.e. micro- or macro-capsules.
Typically, in microencapsulation, the cells are sequestered in a
small permselective spherical container, whereas in
macroencapsulation the cells are entrapped in a larger
non-spherical membrane, Lim et al. (U.S. Pat. Nos. 4,409,331 and
4,352,883) discloses the use of microencapsulation methods to
produce biological materials generated by cells in vitro, wherein
the capsules have varying permeabilities depending upon the
biological materials of interest being produced, Wu et al, Int. J.
Pancreatology, 3:91-100 (1988), disclose the transplantation of
insulin-producing, microencapsulated pancreatic islets into
diabetic rats.
[0045] As indicated above, the cells that are seeded on the
elastomeric composites of the present invention may be cell lines
or primary cells. In certain preferred embodiments, the cells are
genetically engineered cells that have been modified to express a
biologically active or therapeutically effective protein product.
Techniques for modifying cells to produce the recombinant
expression of such protein products are well known to those of
skill in the art. In particular preferred embodiments, the
compositions of the invention may be used to form of a tissue graft
or tissue patch. Such a tissue graft may be an autograft,
allograft, biograft, biogenic graft or xenograft. Tissue grafts may
be derived from various tissue types. Representative examples of
tissues that may be used to prepare biografts include, but are not
limited to, rectus sheaths, peritoneum, bladder, pericardium,
veins, arteries, diaphragm and pleura. For such grafts the cells
may be endothelial cells, ligament tissue, muscle cells, bone
cells, cartilage cells. Such cells may be grafted into the
compositions of the invention alone or in combination with a drug
or biologically active agent to be delivered to an in vivo site.
For example, such cells for the biograft may be harvested from a
host, loaded with the agent of interest and then applied in a
perivascular manner at the site where lesions and intimal
hyperplasia can develop. Once implanted, the agent of interest
(e.g., paclitaxel) is released from the graft and can penetrate the
vessel wall to prevent the formation of intimal hyperplasia at the
treatment site. In certain embodiments, the biograft may be used as
a backing layer to enclose a composition (e.g., a gel or paste
loaded with anti-scarring agent).
[0046] The patches made of the compositions of the present
invention may be combined with drugs for delivery or therapeutic
agents that can form part of a tissue patch prepared from the
polymers of the invention. For example, the compositions of the
invention may be used to form a mesh or a patch made of the
biodegradable polymeric matrix that conforms to the tissue and
releases the agent (e.g., a therapeutic agent such as a drug or a
diagnostic agent such as a marker, dye or other marker of that will
allow visualization of a diseased state). In preferred examples,
the compositions are fashioned into a mesh or patch that
incorporates a drug that can be released in a controlled release
manner. See, e.g., U.S. Pat. No. 6,461,640. The mesh or patch made
of the compositions of the invention may be impregnated with an
antioxidant and/or antimicrobial. See, e.g., U.S. Pat. No.
6,572,878. The mesh or patch may be a tissue patch that is adapted
to cover a bony dissection in the spine. See, e.g., U.S. Pat. No.
5,868,745 and U.S. patent application No. 2003/0078588. The tissue
patch made of the compositions of the invention may be prepared to
be wrapped around a nerve in a canal to reduce fibroplasia. See,
e.g., U.S. Pat. No. 6,106,558. The tissue patch may be a resorbable
collagen membrane that is wrapped around the spinal chord to
inhibit cell adhesions. See, e.g., U.S. Pat. No. 6,221,109. The
tissue patch may be used as a dressing to cover a wound and promote
wound healing. See, e.g., U.S. Pat. No. 6,548,728. The compositions
of the present invention may be prepared as a bandage that contains
a scar treatment pad with a layer of silicone elastomer or silicone
gel. See, e.g., U.S. Pat. Nos. 6,284,941 and 5,891,076. The
compositions may be used to incorporate a biologically active
compound. See, e.g., U.S. Pat. Nos. 6,323,278; 6,166,130; 6,051,648
and 5,874,500.
[0047] Methods for incorporating the a biologically active material
onto or into the mesh or patch of the invention include: (a)
affixing (directly or indirectly) to the patch such a biologically
active material (e.g., by either a spraying process or dipping
process as described above, with or without a carrier), (b)
incorporating or impregnating into the patch made with the
composition with a biologically actice material (e.g., by either a
spraying process or dipping process as described above, with or
without a carrier (c) by coating the patch made with the
composition with a substance such as a hydrogel which will in turn
absorb the biologically active material, (d) constructing the patch
made with the composition itself with the biologically active
material in either the biodegradable polymer, the poly(diol
citrate) polymer, or in the mixture of the two, or (e) by
covalently binding the biologically active material directly to the
surface of the composition of the invention.
[0048] The compositions of the present invention also may be used
to coat devices such as medical stents and the like. For devices
that are coated, the coating process can be performed in such a
manner as to (a) coat only one surface of device with the
compositions of the invention or (b) coating all or parts of the
device with the compositions of the invention.
[0049] The nanocomposites of the present invention also will be
useful in soft tissue engineering as described in the examples
herein below.
[0050] The patches prepared from the compositions of the invention
or a device coated with the same may be made sterile either by
preparing them under aseptic environment and/or they may be
terminally sterilized using methods known in the art, such as gamma
radiation or electron beam sterilization methods or a combination
of both of these methods.
[0051] Patches made with the compositions of the invention may be
applied to any bodily conduit or any tissue. Prior to implantation,
the patch may be trimmed or cut from a sheet of bulk material to
match the configuration of the widened foramen, canal, or
dissection region, or at a minimum, to overlay the exposed tissue
area. Such a patch may be shaped to match the particular
configuration of the placement region. For example, the patch may
be rolled in a cuff shape or cylindrical shape and placed around
the exterior periphery of the desired tissue. It may be provided in
a relatively large bulk sheet and then cut into shapes to mold the
particular structure and surface topography of the tissue or device
to be wrapped. Alternatively, it may be pre-shaped into one or more
patterns for subsequent use. The patches made from the compositions
of the invention may be typically rectangular in shape and be
placed at the desired location within the surgical site by direct
surgical placement or by endoscopic techniques. The patches may be
secured into place by wrapping it onto itself (i.e.,
self-adhesive), or by securing it with sutures, staples, sealant,
and the like. Alternatively, the film or mesh may adhere readily to
tissue and therefore, additional securing mechanisms may not be
required.
[0052] The patches made from the compositions of the invention may
be used for a variety of indications, including, without
limitation: (a) prevention of surgical adhesions between tissues
following surgery (e.g., gynecologic surgery, vasovasostomy, hernia
repair, nerve root decompression surgery and laminectomy); (b)
prevention of hypertrophic scars or keloids (e.g., resulting from
tissue burns or other wounds); (c) prevention of intimal
hyperplasia and/or restenosis (e.g., resulting from insertion of
vascular grafts or hemodialysis access devices); (d) may be used in
affiliation with devices and implants that lead to scarring as
described herein (e.g., as a sleeve or mesh around a breast implant
to reduce or inhibit scarring); (e) prevention of infection (e.g.,
resulting from tissue bums, surgery or other wounds); or (f) may be
used in affiliate with devices and implants that lead to infection
as described herein.
[0053] The patches made of the compositions of the invention may be
used for hernia repair surgery or in other types of surgical
procedures as discussed in U.S. Pat. Nos. 6,638,284; 5,292,328;
4,769,038 and 2,671,444. Thus, the patches made from the
compositions of the invention also may be used to prepare hernia
meshes, which support the repaired hernia or other body structures
during the healing process. Hernias are often repaired surgically
to prevent complications. Conditions in which a hernia mesh may
need to be used include, without limitation, the repair of inguinal
(i.e., groin), umbilical, ventral, femoral, abdominal,
diaphragmatic, epigastric, gastroesophageal, hiatal, intermuscular,
mesenteric, paraperitoneal, rectovaginal, rectocecal, uterine, and
vesical hernias. Hernia repair typically involves returning the
viscera to its normal location and the defect in the wall is
primarily closed with sutures, but for bigger gaps, a mesh is
placed over the defect to close the hernia opening. Infiltration of
the subject polymer composition comprising an anti-scarring agent
into tissue adjacent to a hernia repair mesh may reduce or prevent
fibrosis proximate to the implanted hernia mesh, thereby minimizing
the possibility of adhesions between the abdominal wall or other
tissues and the mesh itself, and reducing further complications and
abdominal pain.
[0054] In yet another aspect, the patches made from the
compositions of the invention may be used for delivering a specific
therapeutic or other agent to an external portion (surface) of a
body passageway or cavity. Examples of body passageways include
arteries, veins, the heart, the esophagus, the stomach, the
duodenum, the small intestine, the large intestine, biliary tracts,
the ureter, the bladder, the urethra, lacrimal ducts, the trachea,
bronchi, bronchiole, nasal airways, Eustachian tubes, the external
auditory mayal, vas deferens and fallopian tubes. Examples of
cavities include the abdominal cavity, the buccal cavity, the
peritoneal cavity, the pericardial cavity, the pelvic cavity,
perivisceral cavity, pleural cavity and uterine cavity.
[0055] In the following Examples, Example 1 is directed to
teachings of methods of producing various poly(diol) citrate
molecules. Example 2 provides teachings of how to strengthen and
stiffen biocompatible membranes made from such poly(diol) citrate
molecules. These nanocomposite materials described herein can be
used in tissue engineering, drug delivery, or any application where
strong biodegradable and flexible elastomeric materials may be
necessary. The increasing and fast development of tissue
engineering applications will require this type of technology to
maximize flexibility for design requirements of scaffolds for
tissue engineering or drug delivery devices.
EXAMPLE 1
Biodegradable Elastomeric Polymers
[0056] The nanoparticle compositions of the invention are based on
biodegradable elastomeric polymers of poly(diol) citrate molecules.
Such molecules typically comprising a polyester network of citric
acid copolymerized with a linear aliphatic di-OH monomer in which
the number of carbon atoms ranges from 2 to 20. Polymer synthesis
conditions for the preparation of these molecules vary from mild
conditions, even at low temperature (less than 100.degree. C.) and
no vacuum, to tough conditions (high temperature and high vacuum)
according the requirements for the materials properties. By
changing the synthesis conditions (including, but not limited to,
post-polymerization temperature, time, vacuum, the initial monomer
molar ratio, and the di-OH monomer chain length) the mechanical
properties of the polymer can be modulated over a wide range. This
series of polymers exhibit a soft, tough, biodegradable,
hydrophilic properties and excellent biocompatibility in vitro.
[0057] The poly(diol)citrate polymers used herein have a general
structure of: (A-B-C).sub.n
[0058] Where A is a linear, aliphatic diol and C also is a linear
aliphatic diol. B is citric acid. The citric acid co-polymers of
the present invention are made up of multiples of the above
formula, as defined by the integer n, which may be any integer
greater than 1. It is contemplated that n may range from 1 to about
1000 or more. It is particularly contemplated that n may be 1, 2,
3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20,
21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 31, 32, 33, 34, 35, 36,
37, 38, 39, 40, 41, 42, 43, 44, 45, 46, 47, 48, 49, 50, or
more.
[0059] In preferred embodiments, the identity of "A" above is poly
1,10-decanediol and in another preferred embodiment the identity of
A is 1,8-octanediol. However, it should be understood that this is
merely an exemplary linear, aliphatic diol. Those of skill are
aware of other aliphatic alcohols that will be useful in
polycondensation reactions to produce poly citric acid polymers.
Exemplary such aliphatic diols include any diols of between about 2
carbons and about 20 carbons. While the diols are preferably
aliphatic, linear, unsaturated diols, with the hydroxyl moiety
being present at the C.sub.1 and C.sub.x position (where x is the
terminal carbon of the diol), it is contemplated that the diol may
be an unsaturated diol in which the aliphatic chain contains one or
more double bonds. The preferred identity for "C" in one embodiment
is 1,8, octanediol, however as with moiety "A," "C" may be any
other aliphatic alcohols. While in specific embodiments, both A and
C are both the same diol, e.g., 1,8-octanediol, it should be
understood that A and C may have different carbon lengths. For
example, A maybe 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15,
16, 17, 18, 19, 20 or more carbons in length, and C may
independently be 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15,
16, 17, 18, 19, 20 or more carbons in length. Exemplary methods for
the polycondensation of the citric acid with the linear diols are
provided in this Example. These materials are then used as starting
materials for the composites described in Example 2.
[0060] Synthesis of Poly(1,10-decanediol-co-citric acid) (PDC) In a
typical experiment, 19.212 g citric acid and 17.428 g
1,10-decanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,10-decanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
[0061] Preparation of Poly(1,8-Octanediol-co-citric acid) (POC) In
a typical experiment, 19.212 g citric acid and 14.623 g Octanediol
were added to a 250 mL three-neck round-bottom flask, fitted with
an inlet adapter and an outlet adapter. The mixture was melted
within 15 min by stirring at 160-165.degree. C. in silicon oil
bath, and then the temperature of the system was lowered to
140.degree. C. The mixture was stirred for another 1 hr at
140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time (from one day to 3
weeks depending on the temperature, with the lower temperatures
requiring longer times) to achieve the
Poly(1,8-octanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
[0062] Porous scaffolds of POC (tubular and flat sheets) were
prepared via a salt leaching technique as follows: POC pre-polymer
was dissolved into dioxane to form 25 wt % solution, and then the
sieved salt (90-120 microns) was added into pre-polymer solution to
serve as a porogen. The resulting slurry was cast into a
poly(tetrafluoroethylene) (PTFE) mold (square and tubular shape).
After solvent evaporation for 72 h, the mold was transferred into a
vacuum oven for post-polymerization. The salt in the resulting
composite was leached out by successive incubations in water
(produced by Milli-Q water purification system every 12 h for a
total 96 h. The resulting porous scaffold was air-dried for 24 hr
and then vacuum dried for another 24 hrs. The resulting scaffold
was stored in a dessicator under vacuum before use. Porous
scaffolds are typically preferred when cells are expected to
migrate through a 3-dimensional space in order to create a tissue
slice. Solid films would be used when a homogenous surface or
substrate for cell growth is required such as an endothelial cell
monolayer within the lumen of a vascular graft.
[0063] Using similar techniques porous scaffold of PDC or other
poly(diol)citrates can be prepared.
[0064] Synthesis of Poly(1,6-hexanediol-co-citric acid) (PHC). In a
typical experiment, 19.212 g citric acid and 11.817 g
1,6-hexanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in a silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for a predetermined time from one day to 3
weeks, depending on the temperature, to achieve the
Poly(1,6-hexanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
[0065] Synthesis of Poly(1,12-dodecanediol-co-citric acid) PDDC. In
a typical experiment, 19.212 g citric acid and 20.234 g
1,12-dodecanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,12-dodecanediol-co-citric acid). Nitrogen was introduced
into the reaction system before the polymer was taken out from
reaction system.
[0066] Synthesis of Poly(1,8-octanediol-co-citric acid-co-glycerol)
In a typical experiment (Poly(1,8-octanediol-co-citric acid-co-1%
glycerol), 23.0544 g citric acid, 16.5154 g 1,8-octanediol and
0.2167 g glycerol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for another hour
at 140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,8-octanediol-co-citric acid-co-1% glycerol). Nitrogen was
introduced into the reaction system before the polymer was taken
out from reaction system.
[0067] Synthesis of Poly(1,8-octanediol-citric acid-co-polyethylene
oxide). In a typical experiment, 38.424 g citric acid, 14.623 g
1,8-octanediol and 40 g polyethylene oxide with molecular weight
400 (PEO400)(100 g PEO1000 and 200 g PEO2000 respectively) (molar
ratio: citric acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 135.degree. C.
The mixture was stirred for 2 hours at 135.degree. C. to get the
pre-polymer. Nitrogen was vented throughout the above procedures.
The pre-polymer was post-polymerized at 120.degree. C. under vacuum
for predetermined time from one day to 3 dyas to achieve the
Poly(1,8-octanediol-citric acid-co-polyethylene oxide). Nitrogen
was introduced into the reaction system before the polymer was
taken out from reaction system. The molar ratios can be altered to
achieve a series of polymers with different properties.
[0068] Synthesis of Poly(1,12-dodecanediol-citric
acid-co-polyethylene oxide). In a typical experiment, 38.424 g
citric acid, 20.234 g 1,12-dodecanediol and 40 g polyethylene oxide
with molecular weight 400 (PEO400)(100 g PEO1000 and 200 g PEO2000
respectively) (molar ratio: citric
acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a 250 ml or 500
ml three-neck round-bottom flask, fitted with an inlet adapter and
an outlet adapter. The mixture was melted within 15 min by stirring
at 160-165.degree. C. in silicon oil bath, and then the temperature
of the system was lowered to 120.degree. C. The mixture was stirred
for half an hour at 120.degree. C. to get the pre-polymer. Nitrogen
was vented throughout the above procedures. The pre-polymer was
post-polymerized at 120.degree. C. under vacuum for predetermined
time from one day to 3 days to achieve the
Poly(1,12-dodecanediol-citric acid-co-polyethylene oxide). Nitrogen
was introduced into the reaction system before the polymer was
taken out from reaction system. The molar ratios can be altered to
achieve a series of polymers with different properties.
[0069] Synthesis of Poly(1,8-octanediol-citric
acid-co-N-methyldiethanoamine) POCM. In a typical experiment,
38.424 g citric acid, 26.321 g 1,8-octanediol and 2.3832 g
N-methyldiethanoamine (MDEA) (molar ratio: citric
acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a 250 ml or 500
ml three-neck round-bottom flask, fitted with an inlet adapter and
an outlet adapter. The mixture was melted within 15 min by stirring
at 160-165.degree. C. in silicon oil bath, and then the temperature
of the system was lowered to 13520.degree. C. The mixture was
stirred for half an hour at 120.degree. C. to get the pre-polymer.
Nitrogen was vented throughout the above procedures. The
pre-polymer was post-polymerized at 80.degree. C. for 6 hours,
120.degree. C. for 4 hours without vacuum and then 120.degree. C.
for 14 hours under vacuum to achieve the Poly(1,8-octanediol-citric
acid-co-N-methyldiethanoamine). Nitrogen was introduced into the
reaction system before the polymer was taken out from reaction
system. The molar ratios can be altered to citric
acid/1,8-octanediol/MDEA=1/0.95/0.05.
[0070] Synthesis of Poly(1,12-dodecanediol-citric
acid-co-N-methyldiethanoamine) PDDCM. In a typical experiment,
38.424 g citric acid, 36.421 g 1,12-dodecanediol and 2.3832 g
N-methyldiethanoamine (MDEA) (molar ratio: citric
acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a 250 ml or 500
ml three-neck round-bottom flask, fitted with an inlet adapter and
an outlet adapter. The mixture was melted within 15 min by stirring
at 160-165.degree. C. in a silicon oil bath, and then the
temperature of the system was lowered to 120.degree. C. The mixture
was stirred for half an hour at 120.degree. C. to get the
pre-polymer. Nitrogen was vented throughout the above procedures.
The pre-polymer was post-polymerized at 80.degree. C. for 6 hours,
120.degree. C. for 4 hours without vacuum and then 120.degree. C.
for 14 hours under vacuum to achieve the
Poly(1,12-dodecanediol-citric acid-co-N-methyldiethanoamine).
Nitrogen was introduced into the reaction system before the polymer
was taken out from reaction system. The molar ratios can be altered
to citric acid/1,12-dodecanediol/MDEA=1/0.95/0.05.
EXAMPLE 2
[0071] Biodegradable Elastomeric Composites Made from Biodegradable
Elastomeric Polymers
[0072] Example 1 describes the production of PDC as well as a
number of other poly(diol)citrate polymers. In the present Example,
there are provided teachings of how to further strengthen and
stiffen the composite by incorporating nanoparticles into the
elastomeric polymer matrix [2-4, 18]. The nanoparticles act as
additional crosslink points to reinforce the network chains and in
general, the increase in mechanical properties is inversely
proportional to the nanoparticle diameter [2]. The exemplary
polymers chosen for the nanophase were poly(L-lactic acid) (PLLA)
and poly(lactic-co-glycolic acid) (PLGA). These polymers were
chosen as they are rigid and strong and have been used in many
tissue engineering applications [19]. Furthermore, the rate of
degradation could be tailored to match that of the surrounding
elastomeric matrix. Poly(L-lactic acid) has a degradation time of
greater than two years while poly(glycolic acid) has a degradation
time of 1-2 months [20]. By changing the ratio of lactic to
glycolic acid, the degradation rate could be varied from fast (1-2
months) to slow (>2 years). For tissue engineering, the rate of
degradation of the polymer scaffold should match that of tissue
regrowth.
[0073] Materials & Methods
[0074] Poly(diol citrate) Pre-polymer Synthesis: Citric acid and
1,10-decanediol were melted under a flow of nitrogen gas by
stirring at 165.degree. C. in a silicon oil bath and then stirred
for another hour at 140.degree. C. to create a pre-polymer
solution, poly(1,10-decanediol-co-citrate) (PDC). The mechanism of
PDC synthesis is shown in FIG. (1) below. For nanoscaffold
composite preparation, the PDC prepolymer was dissolved in ethanol
to a concentration of 30%. For nanoparticle composite preparation,
PDC was dissolved in 1,4-dioxane to a concentration of 25%.
[0075] Nanoscaffold Synthesis: Poly(L-lactic acid) (PLLA)
nanoscaffolds were prepared using a thermally induced gelation,
solvent exchange, and freeze drying method [21]. PLLA
(Mw.about.50,000, polydispersity.about.1.8) was dissolved in
tetrahydrofuran (THF) to make solutions with concentrations of
2.5%, 5%, and 10%. The PLLA solution was cast into a Teflon Petri
dish and immediately placed into a -80.degree. C. freezer for 1
hour to induce gelation. The gel was then transferred to a
-20.degree. C. freezer for 1 hour. After 1 hour, the dish
containing the gel was removed from the -20.degree. C. freezer and
immersed in a large volume of water for solvent exchange. The water
was changed 3 times per day for two days ensure complete removal of
solvent. The gel was removed from the water, blotted dry with
filter paper, and freeze-dried for 2 days using a Freezone-6
lyophilizer (Labconco, Kansas City, Mo.).
[0076] Nanoscaffold Composite Synthesis: Freeze-dried nanoscaffolds
were transferred to a Teflon Petri dish and PDC prepolymer in
ethanol was poured over the nanoscaffold. The dish was subjected to
repeated vacuum/repressurization cycles to ensure the prepolymer
filled the nanopores. Excess prepolymer was removed from the dish
and the nanocomposite polymerized at 60.degree. C. for 4 hours.
Following this, the PDC prepolymer coating and
vacuum/repressurization cycles was repeated one additional time to
ensure that the pores were completely filled with polymer. After
the final coating, the composite was polymerized at 120.degree. C.
for 1 day without vacuum followed by 120.degree. C. for 1 day with
vacuum.
[0077] Characterization of Nanoscaffolds: The density and porosity
of the scaffold was measured using a method based on Archimedes'
principle [24]. Briefly, a density bottle was filled with ethanol
and weighed once equilibrated to 30.degree. C. (W.sub.1) in a water
bath. The polymer scaffold of mass W.sub.s was then immersed in the
density bottle and the air bubbles removed from the pores under
vacuum. The bottle was supplemented with ethanol and weighed once
again (W.sub.2). The scaffold saturated with ethanol was removed
from the bottle and the bottle weighed one final time (W.sub.3).
The density and porosity of the scaffolds were calculated using the
following equations where V.sub.p is the volume of the scaffold
pores, V.sub.s is the volume of the scaffold skeleton, .rho..sub.s
is the scaffold density, and .epsilon. is the porosity. V p = ( W 2
- W 3 - W s ) .rho. e ( 1 ) V s = ( W 1 - W 2 + W s ) .rho. e ( 2 )
.rho. s = W s V s = W s .times. .rho. e ( W 1 - W 2 + W s ) ( 3 ) =
V p ( V p + V s ) = ( W 2 - W 3 - W s ) ( W 1 - W 3 ) ( 4 )
##EQU1##
[0078] The morphology of the nanofibrous matrices and nanoscaffold
composites were studied using an LEO 1525 scanning electron
microscope (Zeiss, Thornwood, N.Y.). SEM images were analyzed using
image analysis software (Image-Pro.RTM. Plus V.4.0, Silver Spring,
Md.) to obtain the pore size and fiber diameter data. Over fifty
measurements were averaged to get the mean value and standard
deviation of the pore size.
[0079] Nanoparticle Preparation: Poly(lactic-co-glycolic acid)
(PLGA) and Poly(L-lactic acid) nanoparticles were prepared using a
modified spontaneous emulsion and solvent diffusion method [23,
25]. PLGA or PLLA was dissolved in a 60:40 acetone/ethanol mixture
to form solutions with concentrations ranging from 1-4%. Poly(vinyl
alcohol) (PVA) (80% hydrolyzed) was dissolved in ml of water to
create a 4% (w/v) solution. The PLGA in solvent mixture was added
dropwise to the PVA in water solution (1 part PLGA, 4 parts PVA)
under high speed magnetic stirring to form nanoparticles. The
nanoparticles in solution were concentrated to a volume of 300 ml
using an Amicon 8400 ultrafiltration cell with polyethersulfone
membrane (Mw.about.500,000) (Millipore, Billerica, Mass.) and then
centrifuged at 48,400 g to pellet the nanoparticles. The
supernatant was discarded and the tube washed several times with
water before resuspending the particles in water. The mechanism of
nanoparticle formation is shown in FIG. (2).
[0080] Nanoparticle Composite Synthesis: To create a nanocomposite,
the nanoparticles were added to the prepolymer under magnetic
stirring and the mixture quickly cast into a glass Petri dish. The
amount of nanoparticles added was varied to create composites with
5% and 10% nanoparticles by weight. Water and 1,4-dioxane were
removed by freeze drying for 1 week using a Freezone-6 lyophilizer
(Labconco, Kansas City, Mo.). After removal of solvent and water,
the dish transferred to an 80.degree. C. oven for 3 days to finish
polymerization.
[0081] Characterization of Nanoparticles: Recovery of nanoparticles
after centrifugation was assessed by measuring the dry mass of the
particles and comparing it to the mass of polymer used in the
formulation. Nanoparticle diameters were assessed using dynamic
light scattering. The morphology of the nanoparticles and
nanocomposites were studied using an LEO 1525 scanning electron
microscope (Zeiss, Thornwood, N.Y.). The amount of PVA on the
surface of the nanoparticles was determined using the Alleman
method [26]. Briefly, 200 mg of freeze dried nanoparticles were
dissolved in 25 ml of chloroform with sonication for 30 minutes.
The organic solution was filtered using a membrane filter and the
residual PVA on the filter washed with a small volume of
chloroform. The filter was dried under atmospheric pressure and
then transferred to a 250 ml beaker. One hundred grams of purified
water was added to the flask and heated to dissolve the PVA. After
removing the filter and cooling, the weight of the PVA solution was
adjusted to 100 g. Fifteen milliliters of a boric acid solution (4%
w/w) and 3 ml of an iodine solution (1.27% (w/w) iodine and 2.5%
(w/w) potassium iodide in purified water) were added 5 ml of the
PVA solution. The volume was adjusted to 50 ml with water and the
absorbance assayed spectrophotometrically (.lamda.=620, 690
nm).
[0082] Mechanical Characterization of Polymers: Tensile tests using
ASTM D412a with a modified sample size (26 mm.times.4 mm.times.1.5
mm) were performed on an Instron 5544 mechanical tester (Instron,
Canton, Mass.) equipped with a 500N load cell. From these tests the
ultimate tensile strength, Young's modulus, and elongation at break
were recorded.
[0083] Statistical methods: Data are expressed as means.+-.standard
deviation. The statistical significance between two sets of data
was calculated using two-tail Student's t-test. Analysis of
variance (ANOVA) and post-hoc analysis using the Tukey and
Bonferoni tests were used to determine significant differences
among three or more means. Data were taken to be significant, when
a P-value of 0.05 or less was obtained.
[0084] Results & Discussion
Nanoscaffold Results
[0085] Three different nanoscaffolds were produced with 2.5%, 5%,
or 10% PLLA. By varying the concentration of PLLA in
tetrahydrofuran, the characteristics of the nanoscaffold could be
adjusted. Table (1) shows the characteristics of the PLLA
nanoscaffolds before the pores were filled with PDC. With
increasing PLLA concentration, the density and fiber diameter
increase, while the porosity and pore size decrease. A lower
porosity decreases the amount of poly(diol citrate) macro-phase
that can be incorporated into the composite. However, a lower
porosity and greater fiber size result in a greater proportion of
PLLA in the final nanocomposite. TABLE-US-00001 TABLE (1)
Nanoscaffold Characteristics 2.5% PLLA 5% PLLA 10% PLLA Fiber
Diameter (nm) 124.3 .+-. 27.3 246.3 .+-. 69.1 206.7 .+-. 70.4
Density (g/cm.sup.3) 0.033 .+-. 0.007 0.063 .+-. 0.004 0.117 .+-.
0.006 Porosity (%) 97.37 .+-. 0.57 94.95 .+-. 0.32 90.56 .+-. 0.50
Pore Size (nm) 706.6 .+-. 266.1 643.8 .+-. 233.5 617.2 .+-.
362.9
[0086] SEM micrographs of the PLLA nanoscaffolds and resulting
composites are shown in FIG. (3) confirming that the nanopores were
completely filled with PDC.
[0087] The mechanical properties of the composites are shown in
FIG. (4). Increasing the PLLA concentration in the composite
provides additional reinforcement and an increase in the tensile
strength and modulus. The addition of the non-elastic PLLA
nano-phase decreases the elongation at break. However, while the
elongation at break decreased, the materials can still be elongated
to many times there original length before breaking. A
statistically significant difference was found when comparing the
PDC control without PLLA and the PLLA-PDC nanocomposites,
indicating that the mechanical properties can be increased using
nanoscaffold composites.
[0088] Since both PLLA and PDC by themselves have been shown to be
biocompatible, biodegradable, and capable of supporting cell
adhesion and proliferation, it is believed that these composites
will possess the same properties. Therefore it is anticipated that
these materials will be applicable to tissue engineering.
Nanoparticle Results
[0089] Nanoparticles were produced using a spontaneous emulsion and
solvent diffusion method [23, 25]. Both PLGA and PLLA nanoparticles
were produced to allow the possibility of tailoring the degradation
of the particles to that of the surrounding elastomeric matrix. As
noted above the degradation can be adjusted by choice of lactic
acid:glycolic acid ratio. The results of various nanoparticles
produced in 1 gram batches using 2% polymer in solvent solutions
are shown in Table (2) below. TABLE-US-00002 TABLE (2) Size and
Polydispersity of various nanoparticles Recovery Sample Size (nm)
Polydispersity (%) PLGA (85:15) 178.1 0.005 103.84 PLGA (50:50)
190.1 0.005 99.23 PLA 177.4 0.005 98.00 PLLA 183.7 0.045 87.95 PLGA
= poly(lactic-co-glycolic acid) with the ratio of lactic to
glycolic acid designated by (Lactic:Glycolic) PLA = poly(D,L-lactic
acid) PLLA = poly(L-lactic acid)
[0090] The modified spontaneous emulsion and solvent diffusion
method was able to produce nanoparticles from PLGA of various
ratios as well as poly(D,L-lactic acid) and poly(L-lactic acid).
The low polydispersity for all nanoparticles indicates that the
method produces uniform nanoparticles with little aggregation. A
high recovery indicates that this method could potentially be
scaled up to larger batches.
[0091] PLGA nanoparticles of various sizes were also produced by
varying the concentration of polymer in solvent. Results are shown
in Table (3) below. By varying the concentration of polymer in
solvent, the size of the nanoparticles could be controlled. With
rubber elastomers and carbon black, the increase in mechanical
properties is inversely proportional to the size of the
nanoparticles added [2]. Thus it is believed that the increase in
strength and stiffness could be adjusted by changing the size of
the nanoparticles. TABLE-US-00003 TABLE (3) Size of polydispersity
of PLGA nanoparticles Concentration Size (nm) Polydispersity 1%
PLGA (85:15) 166.8 0.033 2% PLGA (85:15) 178.1 0.005 4% PLGA
(85:15) 216.9 0.051
[0092] For nanocomposite fabrication, 5 grams of PLGA (85:15) were
made using a 2% PLGA solution. The diameter, poly(vinyl alcohol)
content, and polydispersity of the nanoparticles used in the
composite fabrication are shown in Table (4). TABLE-US-00004 TABLE
(4) Nanoparticle Characteristics Diameter (nm) 176.5 .+-. 6.8 PVA
Content (%) 13.1 .+-. 2.6 Polydispersity 0.035 .+-. 0.038
[0093] Nanocomposites were fabricated with either 5% (w/w) or 10%
(w/w) PLGA nanoparticles. SEM micrographs of the PLGA nanoparticles
and resulting composites are shown in FIG. (5). FIG. (5A) shows the
dried nanoparticles before incorporation into the PDC matrix. As
can be seen, the nanoparticles are spherical and uniform in size.
FIGS. (5B) and (5C) show the nanocomposite with nanoparticles
embedded in the PDC matrix.
[0094] The mechanical properties of the composites are shown in
FIG. (6). Nanoparticles added to the PDC matrix act as additional
crosslink points and increase the strength and stiffness (Young's
modulus) while decreasing the elongation at break. However, it
should be noted that while the elongation at break decreases, the
nanocomposites still maintain the ability to be elongated to many
times their original length before rupture. A statistically
significant difference was found when comparing the PDC control
without nanoparticles to the PDC-PLGA nanocomposites, indicating
that the strength and stiffness could be increased by incorporating
nanoparticles.
[0095] Although in the present Example nanocomposites were created
with poly(1,10-decanediol citrate), the methods of this Example may
be applied to any elastomer or polymer that can be dissolved in
common solvents. Thus, it is contemplated that the two methods used
to produce nanocomposites could be applicable to other poly(diol
citrates) or other elastomers such as poly(glycerol sebacate).
Nanoscaffold or nanoparticle composites could be created with any
polymer that can be dissolved in common organic solvents such as
ethanol, acetone, 1,4-dioxane, etc. In addition, it has been
demonstrated in other studies that the nanoprecipitation method can
be used to encapsulate drugs and other small molecules [9-12]. This
raises the possibility of creating drug releasing nanocomposites.
Lastly, since PLGA and PDC by themselves have been shown to be
biocompatible and biodegradable, these nanocomposites could readily
be used for tissue engineering applications.
EXAMPLE 2
Biodegradable Poly(diol citrate) Nanocomposite Elastomers for Soft
Tissue Engineering
[0096] At present, synthetic biodegradable polymers commonly used
for scaffolds in tissue engineering have a limited range of
mechanical properties. This limitation is a challenge to in vivo
tissue engineering, as the cell-scaffold construct is expected to
maintain or restore normal tissue biomechanics during new tissue
formation. The biodegradable elastomeric nanocomposite materials
described in the present invention are such that their mechanical
properties can be tailored to meet the requirements of soft tissue
engineering applications. The nanocomposite consists of a
nanofibrous poly(1-lactic acid) (PLLA) nanophase and an elastomeric
poly(diol citrate) macrophase. Incorporation of a PLLA nanophase
provides reinforcement to the poly(diol citrate) as demonstrated by
an increase in tensile strength, modulus, and elongation at break
with minimal permanent deformation. The mechanical properties of
the nanocomposite were altered with the concentration of PLLA,
choice of poly(diol citrate), and polymerization conditions. More
importantly, the tensile mechanical properties compare favorably to
that of human cartilage, ligament, and blood vessel. Furthermore,
the compressive modulus is very similar to that of human and bovine
articular cartilage. These results suggest that poly(diol citrate)
nanocomposite elastomers are promising candidate biomaterials for
soft tissue engineering.
[0097] Materials & Methods
[0098] The following materials and methods are similar to those
already described in Example 1.
[0099] Poly(L-lactic acid) (Mw.about.50,000,
polydispersity.about.1.8) was purchased from Polysciences
(Warrington, Pa.). All other chemicals were purchased from
Sigma-Aldrich (Milwaukee, Wis.).
[0100] Synthesis of poly(diol citrate) pre-polymer. Citric acid (50
g, 0.260 mol) and 1,8-octanediol (38.06 g, 0.260 mol) or
1,10-decanediol (45.36 g, 0.260 mol) were melted under a flow of
nitrogen gas by stirring at 165.degree. C. in a silicon oil bath
and then stirred for another hour at 140.degree. C. to create a
pre-polymer solution of poly(1,8-octanediol-co-citrate) (POC) or
poly(1,10-decanediol-co-citrate) (PDC). For nanocomposite
preparation, the pre-polymer was dissolved in ethanol to a
concentration of 30% (w/v).
[0101] Fabrication of non-porous nanocomposites. The nanocomposite
was fabricated in a two step process. Firstly, a PLLA nanofibrous
network was fabricated via a thermally induced gelation followed by
solvent exchange and freeze drying. [R. Zhang and P. X. Ma, J.
Biomed. Mater. Res., 2000, 52, 430-438; P. Ma and R. Zhang, J.
Biomed. Mater. Res., 1999, 46, 60-72] Secondly, the PLLA
nanofibrous network was impregnated with the PDC or POC pre-polymer
solution. Briefly, PLLA was dissolved in tetrahydrofuran (THF) to
make solutions with concentrations of 5% and 10%. The PLLA solution
was cast into a Teflon Petri dish to a thickness of approximately
1.5 mm and immediately placed into a -80.degree. C. freezer for 1
hour to induce gelation. The gel was then transferred to a
-20.degree. C. freezer for 1 hour. After 1 hour, the dish
containing the gel was removed from the -20.degree. C. freezer and
immersed in a large volume of ethanol for solvent exchange.
[0102] The ethanol was changed 3 times per day for two days to
ensure complete removal of THF. Nanofibrous networks were
transferred to a Teflon Petri dish and poly(diol citrate)
pre-polymer in ethanol was poured over the nanofibrous network. The
dish was subjected to repeated vacuum/repressurization cycles to
ensure thorough penetration of the pre-polymer within the
nanopores. Excess pre-polymer was removed from the dish and the
nanocomposite polymerized at 60.degree. C. for 4 hours. Following
this step, the pre-polymer coating and vacuum/repressurization
cycle was repeated once more. After the final coating, the
composite was polymerized at 80.degree. C. for 3 days without
vacuum or 120.degree. C. for 1 day without vacuum followed by
120.degree. C. for 1 day with vacuum.
[0103] Fabrication of porous nanocomposite scaffolds. Scaffolds
were prepared using solvent casting and salt leaching techniques.
Briefly, sodium chloride was ground using a mortar and pestle and
sieved to obtain particles 90-120 .mu.m in diameter. The salt
particles were mixed with the PLLA in THF at a ratio of 10% PLLA
and 90% salt by mass, packed into a Teflon Petri dish, and
subsequently placed in a -80.degree. C. freezer for 1 hour. The gel
was then transferred to a -20.degree. C. freezer for 1 hour. After
1 hour, the dish containing the gel was removed from the
-20.degree. C. freezer and immersed in a large volume of ethanol
for solvent exchange. The composite was then coated with
pre-polymer and polymerized as described above. After
polymerization to crosslink the polymer, the Petri dish was soaked
in a large volume of water to remove the salt. The water was
changed twice a day for one week to ensure that the salt was
removed. Complete removal of salt was checked by addition of silver
nitrate to the water and observing for precipitation of silver
chloride. The resulting scaffolds were freeze-dried in a Freezone 6
lyophilizer (Labconco, Kansas City, Mo.) and stored in a desiccator
prior to use.
[0104] Characterization of nanocomposites. Non-porous samples were
tested for ultimate tensile strength, Young's modulus, and strain
at break. Tensile tests were performed according to ASTM D412a on
an Instron 5544 mechanical tester (Instron, Canton, Mass.) equipped
with a 500N load cell. The size of the samples was (26 mm.times.4
mm.times.1.5 mm) and samples were pulled at a rate of 500 mm/min.
From the stress-strain data, the ultimate tensile strength, Young's
modulus, and elongation at break were calculated. Sample dimensions
were measured before and after the tensile test to assess any
permanent deformation.
[0105] To assess the potential use of the nanocomposites in
cartilage tissue engineering, the compressive modulus of the
salt-leached, porous scaffolds was evaluated in unconfined
compression using a stepwise stress/relaxation test. [J. S.
Jurvelin, M. D. Buschmann and E. B. Hunziker, P. I. Mech. Eng. [H],
2003, 217, 215-219; R. K. Korhonen, M. S. Laasanen, J. Toyras, J.
Rieppo, J. Hirvonen, H. J. Helminen and J. S. Jurvelin, J.
Biomech., 2002, 35, 903-909] Briefly, scaffolds (.about.7 mm
diameter, .about.2 mm thickness) were loaded between two flat
compression fixtures on an Instron 5544 mechanical tester and a
series of fully automated stress relaxation steps (step 10 .mu.m,
velocity 1 .mu.m/s) was repeated up to a 20% strain. A new step was
automatically initiated after 2 minutes to ensure that equilibrium
relaxation was reached. For each step, the complete
time-position-load data were recorded. The compressive modulus was
determined from the linear range (slope) of the stress-strain
curve.
[0106] The porosity of scaffolds was measured using a method based
on Archimedes' principle. [J. Yang, G. Shi, J. Bei, S. Wang, Y.
Cao, Q. Shang, G. Yang and W. Wang, J. Biomed. Mater. Res., 2002,
62, 438-446] The morphology of nanofibrous networks and
nanocomposites was studied using an LEO 1525 scanning electron
microscope (Zeiss, Thornwood, N.Y.). SEM images were analyzed using
image analysis software (Image-Pro.RTM. Plus V.5.0, Silver Spring,
Md.) to obtain the pore size and fiber diameter data. Fifty
measurements were averaged to get the mean value and standard
deviation.
[0107] Data are expressed as means.+-.standard deviation. The
statistical significance between two sets of data was calculated
using two-tail Student's t-test. Analysis of variance (ANOVA) and
post-hoc analysis using the Tukey and Bonferoni tests were used to
determine significant differences among three or more means. Data
were taken to be significant, when a P-value of 0.05 or less was
obtained.
[0108] Results & Discussion
[0109] POC and PDC composites were created using a PLLA nanofibrous
network as a basis. It was hypothesized that a continuous nanofiber
network would provide significant reinforcement to the resulting
composite.
[0110] PLLA nanofibrous network formation. Nanofibrous networks
were produced through dissolution of PLLA in tetrahydrofuran (THF)
followed by gelation at a preset temperature. Nanofibrous networks
were created with PLLA concentrations of 5% or 10%. By varying the
concentration of PLLA in THF, the characteristics of the
nanofibrous network could be adjusted. Table (5) shows the
characteristics of PLLA nanofibrous networks before the pores were
filled with pre-polymer. All networks showed a wide distribution of
nanometer scale pore sizes and fiber diameters with pore sizes
ranging from 250-975 nm and fiber diameters from 100-300 nm.
Increasing the PLLA concentration led to an increase in the
network's density and fiber diameter while decreasing its porosity
and pore size. A lower porosity decreases the amount of poly(diol
citrate) macro-phase that would be incorporated into the final
composite. TABLE-US-00005 TABLE (5) PLLA Nanofibrous Network
Characteristics 2.5% PLLA 5% PLLA 10% PLLA Fiber Diameter (nm)
124.3 .+-. 27.3 246.3 .+-. 69.1 206.7 .+-. 70.4 Density (g
cm.sup.-3) 0.033 .+-. 0.007 0.063 .+-. 0.004 0.117 .+-. 0.006
Porosity (%) 97.37 .+-. 0.57 94.95 .+-. 0.32 90.56 .+-. 0.50 Pore
Size (nm) 706.6 .+-. 266.1 643.8 .+-. 233.5 617.2 .+-. 362.9
[0111] Fabrication of poly(diol citrate)-PLLA nanocomposites. To
create a nanocomposite, the nanopores within the PLLA network were
filled with poly(diol citrate) pre-polymer and polymerized. SEM
micrographs of the PLLA nanofibrous networks and resulting
composites are shown in FIG. 2. The combination of PLLA nanofibrous
networks and poly(diol citrate) elastomers resulted in a composite
with increased mechanical properties. The mechanical properties and
representative stress-strain curves of the composites with
different percentages of PLLA are shown in FIG. (7). The presence
of a PLLA nanofibrous network provided reinforcement to the
poly(diol citrate) as demonstrated by the increased tensile
strength, modulus, and elongation at break. Increases in tensile
strength of 150% over the PDC control and 400% that of PLLA were
seen with a 10% nanocomposite. Similarly, the modulus increased by
1000% compared to PDC without PLLA reinforcement. Although the
modulus of the composite decreased relative to the PLLA network,
the addition of the non-elastic PLLA nano-phase increased the
elongation at break from 200% for the PDC control to 300% for the
10% nanocomposite. In contrast, the elongation at break for PLLA
controls was less than 5%. PDC nanocomposites could be elongated to
a much greater length with little permanent plastic deformation.
The amount of permanent deformation after break of PDC scaffolds
was less than 4% (Table 6). For the tensile strength, modulus, and
elongation, a statistically significant difference was found when
comparing the PDC control (no PLLA) and the PLLA-PDC
nanocomposites, confirming that the mechanical properties could be
significantly increased using nanofibrous composite elastomers.
TABLE-US-00006 TABLE (6) Permanent Deformation Sample Deformation
(%) 5% PLLA-PDC 1.31 .+-. 0.14 10% PLLA-PDC 3.27 .+-. 0.83 10%
PLLA-POC 1.88 .+-. 0.60 PDC Control 0.07 .+-. 0.30 POC Control 0.02
.+-. 0.20 5% PLLA Control 0.10 .+-. 0.17 10% PLLA Control 0.52 .+-.
0.52
[0112] The mechanical properties of the composites could be
adjusted by varying the amount of PLLA in the final composite.
Comparing between composites reinforced with different amounts of
PLLA, there is an increase in the mechanical properties with
increasing PLLA concentration from 5% to 10% (FIG. 7). The tensile
strength, Young's modulus, and elongation at break increased by
54%, 75%, and 44%, respectively. The increase in mechanical
properties with the addition of PLLA nanofibrous networks may be
due to mechanical interlocking and/or interactions between the
nanofibers and the elastomer chains.
[0113] With other nanofiber-reinforced composites, good
nanofiber-matrix interaction is necessary to produce composites
with increased mechanical properties. For this reason, researchers
have attempted to functionalize the surface of the nanofibers
within the composite to enhance bonding between the nanofiber and
the surrounding matrix. [Finegan et al., J. Mater. Sci., 2003, 38,
3485-3490; Blond et al., Adv. Funct. Mater., 2006, 16, 1608-1614;
Xu J. Compos. Mater., 2004,38, 1563-1582.] Within the compositions
of the present invention, the high temperatures and release of
water during the polycondensation of citric acid and a linear
aliphatic diol likely partially hydrolyzes PLLA chains on the
surface of the nanofibers yielding free carboxyl groups. These
groups can then participate in ester bond formation with the
poly(diol citrate) backbone, leading to covalent bond
reinforcement. In addition, the use of nanofibrous networks instead
of discontinuous fibers may help increase strength through
mechanical interlocking between the PLLA nanofibers and the
poly(diol citrate) chains.
[0114] The effect of post-polymerization conditions on
nanocomposite mechanical properties was also assessed. Increasing
the polymerization temperature increased the tensile strength and
modulus of the nanocomposite (FIG. 8). However, the elongation at
break decreased significantly, although considerably larger than
that of PLLA only. These results are not surprising as it has been
previously reported that the strength and stiffness of poly(diol
citrates) can be increased while decreasing the elongation at break
through increased polymerization time and/or temperature. [Yang et
al., Biomaterials, 2006, 27, 1889-1898; Yang et al., Adv. Mater.,
2004, 16, 511-516] These results suggest that it is possible to
tailor the mechanical properties of the nanocomposite for specific
soft tissue engineering applications such as cartilage, ligament,
and blood vessel by changing polymerization conditions.
[0115] When compared to native tissue, the tensile mechanical
properties of the highly polymerized 10% PLLA-PDC (tensile
strength: 13.03.+-.3.15 MPa, modulus: 85.13.+-.5.52 MPa) compare
favorably to that of cartilage and ligament (FIG. 8). For instance,
the reported tensile strength of human cartilage and ligament are
3.7-10.5 MPa and 24-112 MPa, respectively.2 Likewise, the reported
Young's modulus of cartilage and ligament are 0.7-15.3 MPa and
65-541 MPa, respectively. Furthermore, the range of tensile
strength values compares favorably to human coronary arteries
(1.4-11.14 MPa). [He et al., Tissue Eng., 2005, 11, 1574-1588] As
noted above, mechanical properties are particularly important when
engineering soft tissues which have specific biomechanical
requirements for successful functional tissue engineering. [Lavik,
E. and R. Langer, Tissue engineering: current state and
perspectives. Appl Microbiol Biotechnol, 2004. 65: p. 1-8;
Griffith, L. G., Polymeric Biomaterials. Acta Mater, 2000. 48: p.
263-277]
[0116] Although the mechanical properties of highly polymerized
PDC-PLLA compare favorably to many soft tissues, it may be possible
to further tailor the properties for specific applications through
the selection of linear aliphatic diols of different carbon length.
For this reason, nanocomposites of POC with PLLA were also
fabricated. The mechanical properties of nanocomposites using POC
and PDC with the same concentration of PLLA and identical
polymerization conditions are shown in FIG. (9). For both
materials, there is a statistically significant increase in the
mechanical properties as compared to the poly(diol citrate)
controls. Comparing between the composites, there is a
statistically significant increase in Young's modulus when using
the shorter chain 1,8-octanediol monomer relative to the longer
1,10-decanediol monomer. This finding is in agreement with
published results for poly(diol citrates), as a shorter chain diol
leads to a lower molecular weight between crosslinks and thus a
greater crosslink density and increase in stiffness. [Yang et al.,
Biomaterials, 2006, 27, 1889-1898; Yang et al., Adv. Mater., 2004,
16, 511-516] In addition, a longer chain diol can be elongated to a
much greater length prior to break. [Yang et al., Biomaterials,
2006, 27, 1889-1898; Yang et al., Adv. Mater., 2004, 16,
511-516]
[0117] Since poly(diol citrates) and PLLA by themselves have been
shown to be biocompatible, biodegradable, and capable of supporting
cell adhesion and proliferation, it was believed that these
composites would possess similar properties. Therefore it was
contemplated that these materials could be used as scaffolds for
tissue engineering. Porous nanocomposites were created through the
incorporation of salt particles during the gelation step. [Zhang
and P. X. Ma, J. Biomed. Mater. Res., 2000, 52, 430-438] After
gelation and solvent exchange with ethanol, the nanocomposite was
created by impregnating the scaffold with poly(diol citrate)
pre-polymer followed by polymerization and salt removal. The
porosity of the nanofibrous scaffold composites is shown in Table
(7). Both the 5% and 10% nanocomposites exhibited an open and
interconnected pore structure with high porosity (FIG. 10). This
feature is beneficial for tissue engineering applications as a
three dimensional highly porous structure is required to support
cell attachment, proliferation, and extracellular matrix synthesis.
More importantly, the scaffold should possess mechanical properties
that match the host tissue at the site of implantation. In
relevance to cartilage tissue engineering, the mechanical
properties of porous nanocomposites were tested in unconfined
compression using a stepwise stress relaxation method commonly used
with cartilage. [Jurvelin, P. I. Mech. Eng. [H], 2003, 217,
215-219; Korhonen, J. Biomech., 2002, 35, 903-909]. The compressive
modulus increased with increasing PLLA concentration in the
nanocomposite (FIG. 11). For the 10% PLLA-PDC concentration, the
compressive modulus was 439.+-.106 kPa, similar to that of human
(581.+-.168 kPa) and bovine (310.+-.180 kPa) articular cartilage.
[Jurvelin, P. I. Mech. Eng. [H], 2003, 217, 215-219; Korhonen, J.
Biomech., 2002, 35, 903-909]. In addition, a statistically
significant difference was found when comparing the PDC control
scaffolds to the PLLA-PDC nanofibrous scaffolds. TABLE-US-00007
TABLE (7) Porosity of nanofibrous scaffold composites Sample
Porosity (%) 10% PLLA-PDC 83.14 .+-. 4.67 5% PLLA-PDC 88.27 .+-.
4.47 PDC Control 92.50 .+-. 3.57
[0118] Although the degradation rate of PLLA may be too slow for
most tissue engineering applications, it has been shown that
nanofibrous PLLA scaffolds degrade at a significantly faster rate
than solid walled counterparts.37 While degradation was not
examined in this study, it is believed that the degradation rate of
the nanocomposite scaffolds would be increased compared to
conventional PLLA microfiber or salt-leached scaffolds. The high
nanofiber surface area in addition to the acidic degradation
products of the poly(diol citrate) would contribute to faster
observed degradation kinetics. In addition, the rate of degradation
may be adjusted by varying the molecular weight of the PLLA used in
composite fabrication.
[0119] In summary, in the present examples, novel nanocomposite
elastomers were fabricated for potential use in soft tissue
engineering using a poly(diol citrate) elastomeric macrophase and a
nanofibrous network of poly(1-lactic acid). The incorporation of
nanofibrous PLLA to the poly(diol citrate) increased tensile
strength, Young's modulus, and elongation at break with little
permanent deformation. In addition, it was demonstrated that the
mechanical properties could be adjusted by varying the PLLA
concentration in the nanophase, by choice of diol monomer used to
synthesize the poly(diol citrate) elastomeric matrix, or by
adjusting polymerization conditions. The range of mechanical
properties spanned that of human cartilage, ligament, and blood
vessel. This is the first report of a nanocomposite material for
tissue engineering where the macro and nano phases are made from
biodegradable and biocompatible synthetic co-polymers. These
results warrant further study of poly(diol citrate) nanocomposite
elastomers for soft tissue engineering and medical
applications.
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