U.S. patent application number 15/892481 was filed with the patent office on 2019-02-14 for wet-electrospun biodegradable scaffold and uses therefor.
The applicant listed for this patent is University of Pittsburgh - Of the Commonwealth System of Higher Education. Invention is credited to Nicholas J. Amoroso, Kazuro Lee Fujimoto, Ryotaro Hashizume, Yi Hong, William R. Wagner.
Application Number | 20190046689 15/892481 |
Document ID | / |
Family ID | 45004418 |
Filed Date | 2019-02-14 |
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United States Patent
Application |
20190046689 |
Kind Code |
A1 |
Hong; Yi ; et al. |
February 14, 2019 |
Wet-Electrospun Biodegradable Scaffold and Uses Therefor
Abstract
A cell growth matrix is provided that comprises a biodegradable
elastomeric polymer electrodeposited concurrently with a sprayed or
electrosprayed liquid that is a physiological solution or which
comprises a mammalian blood product such as serum, plasma or
platelet rich plasma. The matrix is useful as a cell-growth matrix
and for repair of a tissue in a mammal, for instance by
implantation in a mammal at a site in need of repair, such as in an
abdominal wall.
Inventors: |
Hong; Yi; (Pittsburgh,
PA) ; Amoroso; Nicholas J.; (Pittsburgh, PA) ;
Fujimoto; Kazuro Lee; (Pittsburgh, PA) ; Hashizume;
Ryotaro; (Pittsburgh, PA) ; Wagner; William R.;
(Gibsonia, PA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
University of Pittsburgh - Of the Commonwealth System of Higher
Education |
Pittsburgh |
PA |
US |
|
|
Family ID: |
45004418 |
Appl. No.: |
15/892481 |
Filed: |
February 9, 2018 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13698757 |
Apr 2, 2014 |
9901659 |
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PCT/US2011/038332 |
May 27, 2011 |
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15892481 |
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61349031 |
May 27, 2010 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61L 27/3616 20130101;
A61L 27/3691 20130101; C12N 2533/30 20130101; A61L 27/54 20130101;
C12N 5/0068 20130101; A61L 27/58 20130101; A61L 27/18 20130101 |
International
Class: |
A61L 27/36 20060101
A61L027/36; A61L 27/58 20060101 A61L027/58; C12N 5/00 20060101
C12N005/00; A61L 27/18 20060101 A61L027/18; A61L 27/54 20060101
A61L027/54 |
Goverment Interests
STATEMENT REGARDING FEDERAL FUNDING
[0002] This invention was made with government support under Grant
No. W81XWH-08-2-0032 awarded by the Armed Forces Institute for
Regenerative Medicine, and Grant No. R01 HL-068816 awarded by the
National Institutes of Health. The government has certain rights in
the invention.
Claims
1. A method of repairing damaged or defective tissue in a mammal,
comprising: implanting in the mammal at the site of the damaged or
defective tissue a cell-free matrix comprising: a biodegradable
elastomeric polymer, comprising: a polylactide; a
poly(lactide-co-glycolide); a poly(lactide-co-caprolactone); a
poly(L-lactide-co-caprolactone); a polyglycolic acid; a
poly(lactide-co-glycolide); a poly(DL-lactide-co-glycolide); a
poly(L-lactide-co-DL-lactide); a polyhydroxybutyrate; a
polyhydroxyvalerate; a polydioxanone; a polyglactin; a
polycaprolactone; a polycarbonate; a polyglyconate; a
poly(trimethylene carbonate); a poly(glycolide-co-trimethylene
carbonate); a poly(glycolide-co-trimethylene
carbonate-co-dioxanone); a triblock copolymer comprising
polycaprolactone; a triblock copolymer comprising
polycaprolactone-b-polyethylene, and glycol-b-polycaprolactone; a
polycarbonate urethane; a poly(ester carbonate)urethane urea; or a
poly(carbonate)urethane urea, electrodeposited concurrently with a
sprayed or electrosprayed liquid that comprises a cell-free
mammalian blood product, thereby resulting in repair of the damaged
or defective tissue.
2. The method of claim 1, wherein the tissue is muscle tissue and
implantation of the cell-free matrix results in infiltration and
growth of muscle cells, and mature vascular formation in the
matrix, and thereby repairing the damaged or defective tissue.
3. The method of claim 1 wherein the polymer comprises: a
polylactide; a poly(lactide-co-glycolide); a
poly(lactide-co-caprolactone); a poly(L-lactide-co-caprolactone); a
polyglycolic acid; a poly(lactide-co-glycolide); a
poly(DL-lactide-co-glycolide); a poly(L-lactide-co-DL-lactide); a
polyhydroxybutyrate; a polyhydroxyvalerate; a polydioxanone; a
polyglactin; a polycaprolactone; a polycarbonate; a polyglyconate;
a poly(trimethylene carbonate); a poly(glycolide-co-trimethylene
carbonate); a poly(glycolide-co-trimethylene
carbonate-co-dioxanone); a triblock copolymer comprising
polycaprolactone; or a triblock copolymer comprising
polycaprolactone-b-polyethylene, and glycol-b-polycaprolactone, and
the tissue is abdominal wall tissue and implantation of the
cell-free matrix results in infiltration and growth of skeletal
muscle cells, and mature vascular formation in the matrix, and
thereby repairing the damaged or defective abdominal wall
tissue.
4. The method of claim 1, wherein the sprayed or electrosprayed
liquid comprises one or more of serum, plasma and platelet rich
plasma.
5. The method of claim 1, wherein the mammal is a human.
6. The method of claim 1, wherein the sprayed or electro sprayed
liquid comprises from 1% to 100% v/v of a mammalian blood
product.
7. The method of claim 1, wherein the sprayed or electro sprayed
liquid comprises 100% serum, plasma or platelet rich plasma.
8. The method of claim 1, wherein the sprayed or electro sprayed
liquid comprises from 10% to 30% v/v serum, plasma or platelet rich
plasma.
9. The method of claim 1, wherein the sprayed or electro sprayed
liquid comprises 20% v/v serum, plasma or platelet rich plasma.
10. The method of claim 1, wherein the blood product is allogeneic
or autologous with respect to the mammal.
11. The method of claim 1, wherein the sprayed or electrosprayed
liquid comprises a platelet rich plasma.
12. The method of claim 1, wherein the matrix is electrospun.
13. The method of claim 1, wherein the cell-free matrix further
comprises at least one active agent selected from the group
consisting of an anti-inflammatory, an immunosuppressant, an
antiangiogenic composition, an antiproliferative composition, an
antibody, a composition that acts on an immunophilin, a taxane, a
nitric oxide donor, an antibiotic, a growth factor, and
combinations thereof.
14. A method of making a cell-free matrix for tissue repair,
comprising electrodepositing a polymer composition comprising: a
polylactide; a poly(lactide-co-glycolide); a
poly(lactide-co-caprolactone); a poly(L-lactide-co-caprolactone); a
polyglycolic acid; a poly(lactide-co-glycolide); a
poly(DL-lactide-co-glycolide); a poly(L-lactide-co-DL-lactide); a
polyhydroxybutyrate; a polyhydroxyvalerate; a polydioxanone; a
polyglactin; a polycaprolactone; a polycarbonate; a polyglyconate;
a poly(trimethylene carbonate); a poly(glycolide-co-trimethylene
carbonate); a poly(glycolide-co-trimethylene
carbonate-co-dioxanone); a triblock copolymer comprising
polycaprolactone; a triblock copolymer comprising
polycaprolactone-b-polyethylene, and glycol-b-polycaprolactone; a
polycarbonate urethane; a poly(ester carbonate)urethane urea; or a
poly(carbonate)urethane urea, and concurrently electrospraying a
liquid comprising a cell-free mammalian blood product onto the
electrodeposited polymer.
15. The method of claim 14, wherein the polymer composition is
electrospun.
16. The method of claim 14, wherein the polymer composition further
comprises collagen and/or elastin.
17. The method of claim 14, wherein the cell-free matrix further
comprises at least one active agent selected from the group
consisting of an anti-inflammatory, an immunosuppressant, an
antiangiogenic composition, an antiproliferative composition, an
antibody, a composition that acts on an immunophilin, a taxane, a
nitric oxide donor, an antibiotic, a growth factor, and
combinations thereof.
18. The method of claim 14, wherein the sprayed or electrosprayed
liquid comprises one or more of serum, plasma and platelet rich
plasma.
19. The method of claim 14, wherein the sprayed or electrosprayed
liquid comprises from 1% to 100% v/v of a mammalian blood
product.
20. The method of claim 14, wherein the sprayed or electrosprayed
liquid comprises a human blood product.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of U.S. patent
application Ser. No. 13/698,757, filed Apr. 2, 2014, which is a
National Stage of International Patent Application No.
PCT/US2011/038332, filed May 27, 2011, which claims the benefit of
U.S. Provisional Patent Application No. 61/349,031, filed May 27,
2010, each of which is herein incorporated by reference in its
entirety.
[0003] Employing prosthetic materials in abdominal wall repair is
commonplace, with clear benefits over repair by direct tissue
apposition for larger defects. Synthetic, non-biodegradable
biomaterials such as polypropylene and polyester meshes, as well as
expanded polytetrafluoroethylene (ePTFE) are widely employed for
this purpose with few complications. Complications associated with
placement of these materials, although not commonly encountered,
include seromas and fistulas, chronic patient discomfort, surgical
site infections and decreasing abdominal wall compliance that can
infringe upon the patient's physical activity. These complications
are generally more common in the settings of massive ventral
hernia, contaminated fields, and emergency surgery. An array of
biologic prosthetic materials from both allogenic and xenogeneic
tissue sources, processed with and without chemical crosslinking,
have been utilized in an effort to address some of the limits
associated with synthetic materials. The potential benefits of
biologic materials include improved infection resistance, host
tissue ingrowth, and less adhesion formation. However, the
downsides of these materials include concerns with mechanical
failure, higher costs, and greater difficulty in tailoring physical
properties, which can lead to mechanical property mismatch at the
native tissue interface with the implant.
SUMMARY
[0004] Although a variety of materials are currently used for
abdominal wall repair, general complications encountered include
herniation, infection, and mechanical mismatch with native tissue.
An approach wherein a degradable synthetic material is ultimately
replaced by tissue mechanically approximating the native state
could obviate these complications. Provided herein are matrices
comprising a biodegradable elastomeric polymer electrodeposited
concurrently with an electrosprayed liquid comprising either a
mammalian blood product, such as serum, plasma or platelet rich
plasma or a physiological solution, such as water, normal (NS,
e.g., 0.9% wt.) saline or phosphate-buffered saline (PBS). The
biodegradable elastomeric polymer is poly(ester urethane)urea
(PEUU), poly(ether ester urethane)urea (PEEUU), poly(ester
carbonate)urethane urea (PECUU) or poly(carbonate)urethane urea
(PCUU) in certain embodiments. These matrices show superior
cellular infiltration as compared to similar materials that are
dry-electrodeposited or other materials, such as expanded
polytetrafluoroethylene (ePTFE). These matrices can be used for any
tissue repair or reconstruction by implanting the matrix at a site
in a patient in need of repair, including, without limitation:
skin, blood vessel(s), muscle, esophagus, trachea, stomach,
intestine, rectum or bladder.
[0005] Thus, a method of repairing a defect in a tissue, such as an
abdominal wall or heart of a mammal, such as a human, but including
any mammal, such as a dog, cat, horse, etc. is provided. The method
comprises implanting a matrix of a wet-electrodeposited
biodegradable, elastomeric polymer, such as poly(ester urethane)
urea (PEUU), poly(ether ester urethane)urea (PEEUU), poly(ester
carbonate)urethane urea (PECUU) and poly(carbonate)urethane urea
(PCUU) nonwoven matrix in a mammal. The matrix is electrodeposited
as it is electrosprayed with either a blood product-containing
liquid, such as a serum-, plasma- or platelet rich
plasma-containing liquid or a physiological solution, such as
water, NS or PBS. Also provided is method of growing cells (in
vitro or in vivo). The method comprises contacting cells with a
matrix of a wet-electrodeposited biodegradable, elastomeric
polymer, such as as poly(ester urethane) urea (PEUU), poly(ether
ester urethane)urea (PEEUU), poly(ester carbonate)urethane urea
(PECUU) and poly(carbonate)urethane urea (PCUU) nonwoven matrix.
The matrix is electrodeposited as it is electrosprayed with a blood
product-containing liquid, such as a serum-, plasma- or platelet
rich plasma-containing liquid.
[0006] A method of making a matrix useful, e.g., as a cell growth
matrix for muscle repair, such as for repair of a defect in an
abdominal wall of a mammal, is provided which comprises wet
electrodepositing a biodegradable, elastomeric polymer concurrently
with electrospraying a blood product-containing liquid, such as a
serum-, plasma- or platelet rich plasma-containing liquid, onto the
electrodeposited polymer to produce a wet-electrodeposited matrix.
The polymer may be a as poly(ester urethane) urea (PEUU),
poly(ether ester urethane)urea (PEEUU), poly(ester
carbonate)urethane urea (PECUU) and poly(carbonate)urethane urea
(PCUU) matrix. In all embodiments of the methods and matrices
described herein, the matrix may be isotropic or anisotropic and is
a non-woven. Also provided is a use of a nonwoven matrix comprising
a biodegradable elastomeric polymer electrodeposited concurrently
with a sprayed or electrosprayed liquid comprising a mammalian
blood product in the manufacture of a cell growth matrix, where the
biodegradable elastomeric polymer and the sprayed or electrosprayed
liquid are any biodegradable elastomeric polymer and sprayed or
electrosprayed liquid as described herein, for example in relation
to the methods described above.
BRIEF DESCRIPTION OF THE DRAWINGS
[0007] FIG. 1. Schematic of wet electrospinning where PEUU is
electrospun with concurrent electrospraying of cell culture medium
(A). Electron micrographs of dry ePEUU (B) and wet ePEUU (C). Scale
bar: 20 .mu.m.
[0008] FIG. 2. (A) Macroscopic appearance of implanted ePTFE (a)
dry ePEUU (b) and wet ePEUU (c) 8 weeks after surgical
implantation. Scale bar: 10 mm. (B) Wall thickness of patches prior
to implant and after 8 weeks in vivo, as well as thickness of the
native rat abdominal wall. *p<0.01 compared with ePTFE group at
same time point.
[0009] FIG. 3. Representative cross-section mosaic images of
implanted ePTFE (A and B), dry ePEUU (C and D), and wet ePEUU (E
and F). The upper row is from 4 week explants (A, C, and E) and the
lower row from 8 week explants (B, D, and F). Within each box,
staining for the upper image is with H&E, and for the lower
image with Masson's trichrome. Scale bar: 1 mm.
[0010] FIG. 4. High magnification micrographs of dry ePEUU with
H&E staining for 4 (A) and 8 (B) week explanted specimens, as
well as wet ePEUU explanted patches at 4 (C) and 8 (D) weeks.
Arrows in wet ePEUU micrographs indicate remnant ePEUU in the
specimen. Scale bar: 50 .mu.m.
[0011] FIGS. 5A and 5B. (FIG. 5A) Hematoxylin and eosin (H&E)
and nuclear staining for each implanted material at each time
point. (FIG. 5B) Quantification of nuclei as a measure of cellular
infiltration. Scale bar: 200 mm *p<0.01 compared with both ePTFE
group and dry ePEUU group at each time point.
[0012] FIG. 6. Immunostaining for CD68 and .alpha.-smooth muscle
actin for wet ePEUU patches explanted after 8 weeks. Nuclear
staining is blue, .alpha.-smooth muscle actin is green, and
anti-CD68 for macrophage labeling is red. (A) At the edge of the
patch, the boundary between the implanted material and native
abdominal wall is indicated with white broken line. (B) Inner
portion of the wet ePEUU. Scale bar=100 .mu.m.
[0013] FIGS. 7A and 7B. FIG. 7A: immunohistochemical staining for
vWF in dry ePEUU explanted patches at 4 (A) and 8 weeks (B), and
wet ePEUU at 4 (C) and 8 weeks (D). With double staining, vWF(+)
structures (red) are seen to be surrounded by .alpha.-smooth muscle
actin positive cells (green, E), implying vascular ducts.
Quantification of vWF labeled structures is summarized in (FIG.
7B). Scale bar: 50 .mu.m. *p<0.05 compared with dry ePEUU group
at 8 week point.
[0014] FIG. 8. Immunostaining for collagen type I, collagen type
III, elastin, and .alpha.-smooth muscle actin of ePTFE, dry ePEUU,
and wet ePEUU explanted patches 8 weeks after implantation. * on
ePTFE micrographs indicates nonspecific antibody binding. Scale
bar: 100 .mu.m.
[0015] FIG. 9. Collagen protein concentration in explanted
abdominal patch region at 8 weeks, standardized by tissue wet
weight. *p<0.05.
[0016] FIG. 10. High magnifications of wet ePEUU Masson's trichrome
staining at 4 (A) and 8 (B) week time point, collagen type I (C and
D), collagen type III (E and F), and elastin (G and H). Scale bar:
50 .mu.m.
[0017] FIG. 11. Immunostaining for .alpha.SMA, calponin,
SM22.alpha., and h-caldesmon in 8 week wet ePEUU patches.
.alpha.SMA (+) cells co-localize with calponin, SM22.alpha., and
h-caldesmon. Scale bar: 50 .mu.m.
[0018] FIGS. 12A and 12B. Biaxial stress-stretch curves of native
abdominal wall tissue (FIG. 12A) and for patch materials 4 and 8
weeks post-implant (FIG. 12B). Under equal planar biaxial tension,
normal abdominal wall tissue exhibits a high degree of anisotropy,
with the circumferential axis being markedly stiffer than the
longitudinal axis. Wet ePEUU remodels in vivo to approach this
pattern at 8 weeks to a much greater degree than the controls.
DETAILED DESCRIPTION
[0019] The use of numerical values in the various ranges specified
in this application, unless expressly indicated otherwise, are
stated as approximations as though the minimum and maximum values
within the stated ranges are both preceded by the word "about". In
this manner, slight variations above and below the stated ranges
can be used to achieve substantially the same results as values
within the ranges. Also, unless indicated otherwise, the disclosure
of these ranges is intended as a continuous range including every
value between the minimum and maximum values. For definitions
provided herein, those definitions refer to word forms, cognates
and grammatical variants of those words or phrases.
[0020] A method of repairing a defect in a tissue and a method of
growing tissue are provided. The methods can be used for repairing
a defect in a tissue, such as an abdominal wall or heart, of a
mammal, such as a human, but including any mammal, such as a dog,
cat, horse, etc. is provided. The method comprises implanting a
wet-electrodeposited biodegradable elastomeric polymer, such as a
as poly(ester urethane) urea (PEUU), poly(ether ester urethane)urea
(PEEUU), poly(ester carbonate)urethane urea (PECUU) and
poly(carbonate)urethane urea (PCUU) in a mammal, typically, but not
exclusively at a site in need of repair. A method of making an
implantable matrix useful, e.g., for repair of a defect in an
abdominal wall of a mammal is provided which comprises wet
electrospinning a biodegradable polymer, such as a as poly(ester
urethane) urea (PEUU), poly(ether ester urethane)urea (PEEUU),
poly(ester carbonate)urethane urea (PECUU) and
poly(carbonate)urethane urea (PCUU), nonwoven matrix in the
presence of a sprayed or electrosprayed liquid comprising serum.
The nonwoven matrix described above and throughout may be used in
the manufacture of a cell growth matrix. In all instances, unless
otherwise specified, the serum can be xenogeneic, allogeneic,
syngeneic or autogenic to the mammal.
[0021] Also described herein are methods of making an implantable
matrix for muscle repair, comprising electrodepositing a
biodegradable elastomeric polymer and concurrently electrospraying
a liquid comprising a mammalian blood product onto the
electrodeposited polymer. The biodegradable elastomeric polymer and
liquid comprising a mammalian blood product may be any
biodegradable elastomeric polymer and liquid comprising a mammalian
blood described herein.
[0022] As used herein, the terms "comprising," "comprise" or
"comprised," and variations thereof, are meant to be open ended.
The terms "a" and "an" are intended to refer to one or more.
[0023] As used herein, the term "patient" or "subject" refers to
members of the animal kingdom including but not limited to human
beings and "mammal" refers to all mammals, including, but not
limited to human beings.
[0024] A biodegradable polymer is "biocompatible" in that the
polymer and degradation products thereof are substantially
non-toxic to cells or organisms, including non-carcinogenic and
non-immunogenic, and are cleared or otherwise degraded in a
biological system, such as an organism (patient) without
substantial toxic effect. Non-limiting examples of degradation
mechanisms within a biological system include chemical reactions,
hydrolysis reactions, and enzymatic cleavage.
[0025] As used herein, the term "polymer composition" is a
composition comprising one or more polymers. As a class, "polymers"
includes, without limitation, homopolymers, heteropolymers,
co-polymers, block polymers, block co-polymers and can be both
natural and synthetic. Homopolymers contain one type of building
block, or monomer, whereas co-polymers contain more than one type
of monomer.
[0026] A polymer "comprises" or is "derived from" a stated monomer
if that monomer is incorporated into the polymer. Thus, the
incorporated monomer that the polymer comprises is not the same as
the monomer prior to incorporation into a polymer, in that at the
very least, certain terminal groups are incorporated into the
polymer backbone. A polymer is said to comprise a specific type of
linkage if that linkage is present in the polymer.
[0027] As described herein, a "fiber" an elongated, slender,
thread-like and/or filamentous structure. A "matrix" is any two- or
three-dimensional arrangement of elements (e.g., fibers), either
ordered (e.g., in a woven or non-woven mesh) or randomly-arranged
(as is typical with a mat of fibers typically produced by
electrospinning) and can be isotropic or anisotropic.
[0028] By "biodegradable or "bioerodible", it is meant that a
polymer, once implanted and placed in contact with bodily fluids
and tissues, will degrade either partially or completely through
chemical reactions with the bodily fluids and/or tissues, typically
and often preferably over a time period of hours, days, weeks or
months. Non-limiting examples of such chemical reactions include
acid/base reactions, hydrolysis reactions, and enzymatic cleavage.
The biodegradation rate of the polymer matrix may be manipulated,
optimized or otherwise adjusted so that the matrix degrades over a
useful time period. The polymer or polymers typically will be
selected so that it degrades in situ over a time period to optimize
mechanical conditioning of the tissue. For instance, in the case of
abdominal wall repair, it is desirable that the matrix dissolves
over at least a week and preferably longer. More importantly, the
matrix would have to retain its supportive capacity until tissue
remodeling occurs, such as for at least 2-8 weeks, or longer.
[0029] A number of biocompatible, biodegradable elastomeric
(co)polymers are known and have been established as useful in
preparing cell growth matrices, including biodegradable poly(ester
urethane) urea (PEUU), poly(ether ester urethane)urea (PEEUU),
poly(ester carbonate)urethane urea (PECUU) and
poly(carbonate)urethane urea (PCUU). In general, useful copolymers
comprise monomers derived from alpha-hydroxy acids including
polylactide, poly(lactide-co-glycolide),
poly(L-lactide-co-caprolactone), polyglycolic acid,
poly(dl-lactide-co-glycolide), poly(l-lactide-co-dl-lactide);
monomers derived from esters including polyhydroxybutyrate,
polyhydroxyvalerate, polydioxanone and polyglactin; monomers
derived from lactones including polycaprolactone; monomers derived
from carbonates including polycarbonate, polyglyconate,
poly(glycolide-co-trimethylene carbonate),
poly(glycolide-co-trimethylene carbonate-co-dioxanone); monomers
joined through urethane linkages, including polyurethane,
poly(ester urethane) urea elastomer. In one example, as described
in US Patent Publication No. 20080268019, incorporated herein by
reference for its technical disclosure, the polymer composition
comprises a synthetic polymeric component and a biological
polymeric component. The synthetic polymeric component may comprise
a thermoplastic biodegradable elastomer, and/or the biological
polymeric component may comprise an extracellular matrix-derived
material. The synthetic polymeric component may comprise one or
both of a poly(ester urethane) urea elastomer and a poly(ether
ester urethane urea) elastomer. The elastomer may comprise a
diamine, such as putrescine or lysine ethyl ester, or a diol. The
elastomer may comprise a polycaprolactone or a polycaprolactone
diol, such as a triblock copolymer comprising polycaprolactone or a
polycaprolactone-b-polyethylene glycol-b-polycaprolactone triblock
copolymer. The elastomer can be functionalized with an
adhesion-promoting peptide, such as the oligopeptide RGD. In one
non-limiting embodiment, the elastomer comprises an isocyanate
derivative, a polycaprolactone diol, and a diamine chain extender,
which may comprise a ratio of isocyanate
derivative:polycaprolactone diol:diamine chain extender of about
2:1:1. In another non-limiting embodiment, the elastomer comprises
an isocyanate derivative, a triblock copolymer comprising
polycaprolactone, and a diamine chain extender in which the ratio
of isocyanate derivative:triblock copolymer:diamine chain extender
optionally is about 2:1:1. For example, as described in the
Examples below, the molar feed ratio for preparation of PEUU was
polycaprolactone:diisocyanatobutane:putrescine=1:2:1, and the molar
feed ratio for synthesis of PEEUU,
polycaprolactone-PEG-polycaprolactone
copolymer:diisocyanatobutane:putrescine=1:2:1.
[0030] Other useful copolymers include: polylactide, polyglycolide,
poly(lactide-co-glycolide), polycaprolactone,
poly(lactide-co-caprolactone), poly(trimethylene carbonate) based
polymers, polyhydroxybutyrate and its copolymer, polydioxanone,
poly(ester carbonate urethane) urea, poly(carbonate urethane) urea,
polycarbonate urethane, or polyester urethane. In another
embodiment, natural polymers are included in the polymer
composition, including gelatin, collagen, chitosan, hyaluronic
acid, etc.
[0031] According to a non-limiting embodiment, the polymer
composition comprises one or both of a collagen and an elastin.
Collagen is a common ECM component and typically is degraded in
vivo at a rate faster than many synthetic bioerodible polymers.
Therefore, manipulation of collagen content in the polymer
composition can be used as a method of modifying bierosion rates in
vivo. Collagen may be present in the polymer composition in any
useful range, including, without limitation, from about 2% wt. to
about 95% wt., but more typically in the range of from about 25%
wt. to about 75% wt., inclusive of all ranges and points
therebetween, including from about 40% wt. to about 75% wt.,
including about 75% wt. and about 42.3% wt. Elastin may be
incorporated into the polymer composition in order to provide
increased elasticity. Use of elastin can permit slight
circumferential expansion of the restrictive matrix in order to
assist the tubular tissue, such as a vein, adapt to its new
function, such as an arterial use. Elastin may be present in the
polymer composition in any useful range, including without
limitation, from about 2% wt. to about 50% wt., inclusive of all
ranges and points therebetween, including from about 40% wt. and
about 42.3% wt., inclusive of all integers and all points
therebetween and equivalents thereof. In one non-limiting
embodiment, collagen and elastin are present in approximately equal
amounts in the polymer composition, In another embodiment, the sum
of the collagen and elastin content in the polymer composition is
in any useful range, including, without limitation, from about 2%
wt. to about 95% wt., but more typically in the range of from about
25% wt. to about 75% wt., inclusive of all ranges and points
therebetween, including from about 40% wt. to about 75% wt.,
including about 75% wt. and about 42.3% wt.
[0032] The biological polymeric component is useful for promoting
cell growth on the elastomeric scaffold, extracting appropriate
host cells for construction, remodeling, and/or enhancement of
biocompatibility. In one non-limiting embodiment, the biological
polymeric component comprises and includes an extracellular
matrix-derived material. As used herein, the terms "extracellular
matrix" and "ECM" refer to a complex mixture of structural and
functional biomolecules and/or biomacromolecules including, but not
limited to, structural proteins, specialized proteins,
proteoglycans, glycosaminoglycans, and growth factors that surround
and support cells within mammalian tissues.
[0033] Generally, any type of extracellular matrix (ECM) can be
used to prepare the biological, ECM-derived polymeric component of
the biodegradable elastomeric scaffold (for example and without
limitation, see U.S. Pat. Nos. 4,902,508; 4,956,178; 5,281,422;
5,352,463; 5,372,821; 5,554,389; 5,573,784; 5,645,860; 5,771,969;
5,753,267; 5,762,966; 5,866,414; 6,099,567; 6,485,723; 6,576,265;
6,579,538; 6,696,270; 6,783,776; 6,793,939; 6,849,273; 6,852,339;
6,861,074; 6,887,495; 6,890,562; 6,890,563; 6,890,564; and
6,893,666; each of which is incorporated by reference in its
entirety for its technical disclosure). By "ECM-derived material"
it is meant a composition that is prepared from a natural ECM or
from an in vitro source wherein the ECM is produced by cultured
cells and comprises one or more polymeric components (constituents)
of native ECM.
[0034] According to one non-limiting example of the ECM-derived
material, ECM is isolated from a vertebrate animal, for example,
from a warm blooded mammalian vertebrate animal including, but not
limited to, human, monkey, pig, cow, sheep, etc. The ECM may be
derived from any organ or tissue, including without limitation,
urinary bladder, intestine, liver, heart, esophagus, spleen,
stomach and dermis. The ECM can comprise any portion or tissue
obtained from an organ, including, for example and without
limitation, submucosa, epithelial basement membrane, tunica
propria, etc. In one non-limiting embodiment, the ECM is isolated
from urinary bladder, which may or may not include the basement
membrane. In another non-limiting embodiment, the ECM includes at
least a portion of the basement membrane. In certain non-limiting
embodiments, the material that serves as the biological component
of the scaffold consists primarily (e.g., greater than 70%, 80%, or
90%) of ECM. In another non-limiting embodiment, the biodegradable
elastomeric scaffold may contain at least 50% ECM, at least 60%
ECM, at least 70% ECM, and at least 80% ECM. In yet another
non-limiting embodiment, the biodegradable elastomeric scaffold
comprises at least 10% ECM. The ECM material may or may not retain
some of the cellular elements that comprised the original tissue
such as capillary endothelial cells or fibrocytes. The type of ECM
used in the scaffold can vary depending on the intended cell types
to be recruited during wound healing or tissue regeneration, the
native tissue architecture of the tissue organ to be replaced, the
availability of the tissue source of ECM, or other factors that
affect the quality of the final scaffold and the possibility of
manufacturing the scaffold. For example and without limitation, the
ECM may contain both a basement membrane surface and a non-basement
membrane surface, which would be useful for promoting the
reconstruction of tissue such as the urinary bladder, esophagus, or
blood vessel all of which have a basement membrane and non-basement
membrane component.
[0035] Commercially available ECM preparations can also be used as
the biological polymeric component of the scaffold. In one
non-limiting embodiment, the ECM is derived from small intestinal
submucosa or SIS. Commercially available preparations include, but
are not limited to, Surgisis.TM., Surgisis-ES.TM., Stratasis.TM.,
and Stratasis-ES.TM. (Cook Urological Inc.; Indianapolis, Ind.) and
GraftPatch.TM. (Organogenesis Inc.; Canton, Mass.). In another
non-limiting embodiment, the ECM is derived from dermis.
Commercially available preparations include, but are not limited to
Pelvicol.TM. (sold as Permacol.TM. in Europe; Bard, Covington,
Ga.), Repliform.TM. (Microvasive; Boston, Mass.) and Alloderm.TM.
(LifeCell; Branchburg, N.J.). In another embodiment, the ECM is
derived from urinary bladder. Commercially available preparations
include, but are not limited to UBM (Acell Corporation; Jessup,
Md.).
[0036] The matrix may take many different forms. In certain
non-limiting embodiments, the matrix s formed as a flexible fabric
that can be sewn directly on to the site to be treated, as in the
abdominal wall replacement methods described herein. The matrix can
be affixed in place at the site of implantation or affixed using a
medically acceptable adhesive. In one non-limiting embodiment, the
scaffold is substantially planar (having much greater dimension in
two dimensions and a substantially smaller dimension in a third,
comparable to bandages, gauze, and other substantially flexible,
flat items). In another non-limiting embodiment, the matrix is
tubular, and can be attached by anastomosis to tissue such as
intestine or blood vessel tissue. The matrix can be
electrodeposited in two or more steps, with different polymers
being deposited at different locations and/or electrosprayed with
different media at different times, for instance containing
different growth factors, biological components, etc. to generate a
3-dimensional structure that may be used for partial or complete
organ replacement.
[0037] The biodegradable elastomeric scaffold can also have
three-dimensional shapes useful for treating wounds and tissue
deficiencies, such as plugs, rings, wires, cylinders, tubes, or
disks. A useful range of thickness for the biodegradable
elastomeric scaffold is between from about 10 .mu.m (micrometers or
microns (.mu.)) to about 3.5 cm, including increments therebetween,
including, without limitation from about 10 .mu.m to about 50
.mu.m, 50 .mu.m to 3.5 cm, 100 .mu.m to 3.0 cm, and between 300
.mu.m and 2.5 cm.
[0038] In certain non-limiting embodiments, the formation and
initial processing of the synthetic polymeric component and the
biological polymeric component are separate. For example, the
synthesis and dissolution of the synthetic polymeric component may
involve solvents that would adversely affect the desirable
biological properties of the biological polymeric component. By
performing the synthesis and initial processing of the synthetic
polymeric component separately from the corresponding synthesis and
initial processing steps of the biological polymeric component, it
is possible to substantially protect the biological polymeric
component against degradation that it would otherwise face when
exposed to the solvents used in the synthesis and processing the
synthetic polymeric component. In certain non-limiting embodiments,
the synthetic polymeric component and biological polymeric
component are dispersed in different solvents and subsequently
combined (e.g., by electrospraying the biological component onto
the polymer matrix as it is being electrodeposited). For example,
the electrosprayed media may comprise dissolved/solubilized
collagen, elastin, ECM-derived material or other biopolymers,
though concentrations or Mw of the biopolymer might need to be
minimized in order to keep the electrosprayed solution viscosity
within tolerances.
[0039] A poly(ester urethane) urea elastomer (PEUU) may be made
from polycaprolactonediol (MW 2,000) and 1,4-diisocyanatobutane,
with a diamine, such as putrescine as the chain extender. A
suitable PEUU polymer may be made by a two-step polymerization
process whereby polycaprolactone diol (Mw 2,000),
1,4-diisocyanatobutane, and putrescine are combined in a 1:2:1
molar ratio though virtually any molar feed ratio may suffice so
long as the molar ratio of each monomer component is >0. In one
embodiment, the molar feed ratio of polycaprolactone diol plus
putrescine is equal to that of diisocyanatobutane. In the first
polymerization step, a 15 wt % solution of 1,4-diisocyanatobutane
in DMSO is stirred continuously with a 25 wt % solution of diol in
DMSO. In the second step, stannous octoate is added and the mixture
is allowed to react at 75.degree. C. for 3 hours, with the addition
of triethylamine to aid dissolution. A poly(ether ester urethane)
urea elastomer (PEEUU) may be made by reacting
polycaprolactone-b-polyethylene glycol-b-polycaprolactone triblock
copolymers with 1,4-diisocyanatobutane and putrescine. In a
preferred embodiment, PEEUU is obtained by a two-step reaction
using a 2:1:1 reactant stoichiometry of
1,4-diisocyanatobutane:triblock copolymer:putrescine. In the first
polymerization step, a 15 wt % solution of 1,4-diisocyanatobutane
in DMSO is stirred continuously with a 25 wt % solution of triblock
compolymer diol in DMSO. In the second step, stannous octoate is
added and the mixture is allowed to react at 75.degree. C. for 3
hours. The reaction mixture is then cooled to room temperature and
allowed to continue for 18 h. The PEEUU polymer solution is then
precipitated with distilled water and the wet polymer is immersed
in isopropanol for 3 days to remove unreacted monomer and dried
under vacuum.
[0040] Poly(ester carbonate)urethane urea (PECUU) and
poly(carbonate)urethane urea (PCUU) are described, for example, in
Hong et al. (Tailoring the degradation kinetics of poly(ester
carbonate urethane)urea thermoplastic elastomers for tissue
engineering scaffolds B iomaterials,
doi:10.1016/j.biomaterials.2010.02.005). Poly(ester carbonate
urethane)urea (PECUU) is synthesized, for example using a blended
soft segment of polycaprolactone (PCL) and poly(1,6-hexamethylene
carbonate) (PHC) and a hard segment of 1,4-diisocyanatobutane (BDI)
with chain extension by putrescine. Different molar ratios of PCL
and PHC can be used to achieve different physical characteristics.
Putrescine is used as a chain extender by a two-step solvent
synthesis method. In one example, the (PCL+PHC):BDI:putrescine
molar ratio is defined as 1:2:1. Variable molar ratios of PCL and
PHC (e.g., PCL/PHC ratios of 100/0 (yielding a PEUU), 75/25, 50/50,
25/75 and 0/100 (yielding a PCUU)) are completely dissolved in DMSO
in a 3-neck flask with argon protection and then BDI is added to
the solution, following 4 drops of Sn(Oct).sub.2. The flask is
placed in an oil bath at 70.degree. C. After 3 h, the prepolymer
solution is cooled at room temperature and then a putrescine/DMSO
solution is added dropwise into the agitated solution. The final
polymer solution concentration is controlled to be approximately 4%
(w/v). Then the flask is than placed in an oil bath and kept at
70.degree. C. overnight. The polymer is precipitated in an excess
volume of cool deionized water and then dried in a vacuum at
60.degree. C. for 3 days. The polyurethane ureas synthesized from
the different PCL/PHC molar ratios defined above are referred to as
PEUU, PECUU 75/25, PECUU 50/50, PECUU 25/75 and PCUU, respectively.
In practice, the yields of all final products using this method is
approximately 95%.
[0041] The electrodeposited matrices described herein are "wet
electrodeposited" meaning a liquid, such as a serum-containing
liquid, is deposited as the polymer is electrodeposited. One method
would be to spray the liquid onto a rotating mandrel at the same
time the polymer is deposited. In one embodiment, the liquid is
electrosprayed onto the matrix in substantially the same manner as
the polymer is electrospun, the only difference being the deposited
liquid is less viscous than the polymer, and the potential
difference is such that droplets, rather than fibers are deposited.
In one embodiment, the liquid is serum in normal saline, PBS, cell
culture medium or a balanced salt solution, optionally comprising
other additives. In the example below, the electrosprayed medium is
Dulbecco's Modified Eagle Medium (DMEM) with 10% fetal bovine serum
(FBS), 10% horse serum, 1% penicillin/streptomycin, and 0.5% chick
embryo extract. As can be recognized by those of ordinary skill in
the relevant arts, there are a multitude of salt solutions,
buffered salt solutions, media, media supplements, active agents,
such as antibiotics, growth factors and cytokine, biopolymers and
ECM-derived material that would serve equally as a substitute for
the electrosprayed serum-containing liquids described in the
examples below. Compositions that are do not include blood products
are referred to herein as physiological solutions, which are
biocompatible, aqueous solutions, including salt solutions and
blood-product-free medium, though blood products can be added to
the physiological solutions. Other potentially useful media
include, without limitation: DMEM, MEM, RPMI 1640, F10, OptiMEM,
serum-free media, EMEM, EBM-2, F12, IMDM, and Media 199 (available,
e.g., from Invitrogen). Salt solutions may be used instead of
media, such as, without limitation: saline, normal saline
(approximately 0.9% (w/v)), Dulbecco's phosphate-buffered salines,
Hanks' balanced salt solutions, phosphate buffered salined or
Earle's balanced salt solutions. Media supplements include, without
limitation: HEPES, Calcium chloride, or sodium bicarbonate.
Antibiotics include, without limitaiton: actinomycin D, ampicillin,
carbenicillin, cefotaxime, fosmidomycin, gentamycin, kanmycin,
neomycin, penicillin streptomycin, polymyxin B and streptomycin.
Mixtures of more than one media, supplement, or antibiotic can also
be used.
[0042] According to one embodiment, the electrosprayed liquid
comprises one or more xenogeneic, allogeneic, isogeneic, syngeneic
or autologous blood products, such as serum, plasma or
platelet-rich plasma. "Serum" is a cell-free, fibrinogen-free blood
fraction. In one non-limiting embodiment, an aliquot of a patient's
blood is removed and serum is prepared from the blood by allowing
the blood to clot and removing the clotted material and cellular
material, typically by first "ringing" the sample, and then by
centrifugation. Plasma is made by centrifuging a tube of fresh
blood containing an anti-coagulant in a centrifuge until the blood
cells fall to the bottom of the tube. Platelet Rich Plasma is
defined as a volume of the plasma fraction of autologous blood
having a platelet concentration above baseline. (See, generally,
Sampson et al. Curr Rev Musculoskelet Med. 2008 December;
1(3-4):165-74). One method of preparing platelet-rich plasma is by
density-gradient centrifugation and collection of the buffy coat. A
device, such as the Biomet Biologics GPS III device can be used to
obtain a platelet rich plasma fraction. Platelet-rich buffy coat
preparations can be mixed with plasma, serum, saline, PBS or any
suitable salt, buffer, media, etc.
[0043] Stabilizing compositions, such as stabilizing proteins may
be included in the electrosprayed liquid composition. Likewise
viscosity enhancers, including, without limitation: polymeric
compounds may also be added.
[0044] Allogeneic blood fractions, such as one or more of serum,
plasma or platelet-rich plasma, may be used. An electrospray liquid
to be concurrently electrosprayed during electrodeposition of the
polymer component of the matrices described herein may comprise
blood fraction (e.g., serum, plasma or platelet-rich plasma, or
mixtures thereof) concentrations ranging from approximately 1% to
100%, including any increment therebetween, such as 1%, 5%, 10%,
20%, 30%, 40%, 50%, 60%, 70%, 80%, 90% and 100% and any increment
therebetween. In one embodiment, the electrospray liquid comprises
from 5%-25% autologous or allogeneic blood product(s), and in
another embodiment, 20%. When the blood fraction(s) is not 100% of
the electrospray liquid, the electrospray liquid will comprise a
suitable aqueous liquid, such as water, normal saline, PBS, or a
cell culture medium as described above. As described elsewhere
herein, the electrospray liquid also may comprise antibiotics,
buffers, active agents, growth factors, cytokines, biopolymers, ECM
derived material etc. in appropriate concentrations.
[0045] Because the electrodeposition/electrospinning process may be
controlled either manually or by computer, different ratios of
polymer and electrosprayed liquid may be deposited in different
layers of the matrix. For example, the ration of liquid to polymer
may increase in regions of the matrix where it is desirable to get
increased cell infiltration, though too much liquid could lead to
polymer delamination.
[0046] The biodegradable polymers useful herein also can be, and
are preferably elastomeric. Generally, an elastomeric polymer that
has properties similar to that of the soft tissue to be replaced or
repaired is appropriate. For example, in certain embodiments, the
polymers used to make the device is distensible. Non-limiting
examples of suitable polymers include those that have a breaking
strain of from 100% to 1700%, more preferably between 200% and
800%, and even more preferably between 325% and 600%. In
particularly preferred embodiments, the breaking strain of the
polymer is between 5% and 50%, more preferably between 10% and 40%,
and even more preferably between 20% and 30%. Further, it is often
useful to select polymers with tensile strengths of from 10 kPa-30
MPa, more preferably from 5-25 MPa, and even more preferably
between 8 and 20 MPa. In certain embodiments, the initial modulus
is between 10 kPa to 100 MPa, more preferably between 10 and 90
MPa, and even more preferably between 20 and 70 MPa.
[0047] In certain embodiments, the polymers used herein also
release therapeutic agents when they degrade within the patient's
body. For example, the individual building blocks of the polymers
may be chosen such that the building blocks themselves provide a
therapeutic benefit when released in situ through the degradation
process. In one particularly preferred embodiment, one of the
polymer building blocks is putrescine, which has been implicated as
a substance that causes cell growth and cell differentiation.
[0048] In other embodiments, at least one therapeutic agent is
added to the bioerodible matrix or fibers, for instance during
electrospraying or mixed with the liquid copolymer before
electrodeposition. Useful therapeutic agents include any substance
that can be coated on, attached, absorbed, adsorbed, embedded or
otherwise associated with the bioerodible fibers that would provide
a therapeutic benefit to a patient. Therapeutic agent may be
blended with the polymer while the polymer is being processed. For
example, the therapeutic agent may be dissolved in a solvent (e.g.,
DMSO) and added to the polymer blend during processing. In another
embodiment, the therapeutic agent is mixed with a carrier polymer
(for example and without limitation, a polyethylene glycol hydrogel
or polylactic-glycolic acid microparticles) which is subsequently
processed with the elastomeric polymer. By blending the therapeutic
agent with a carrier polymer or the elastomeric polymer itself, the
rate of release of the therapeutic agent may be controlled by the
rate of polymer degradation. In one embodiment, a bioerodible
hydrogel comprising an active agent or cells is applied to the
bioerodible fibers after they are applied to a surface of a tubular
tissue.
[0049] Additionally, other active agents that may be incorporated
into the bioerodible matrix or fibers include, without limitation,
anti-inflammatories, such as, without limitation, NSAIDs
(non-steroidal anti-inflammatory drugs) such as salicylic acid,
indomethacin, sodium indomethacin trihydrate, salicylamide,
naproxen, colchicine, fenoprofen, sulindac, diflunisal, diclofenac,
indoprofen sodium salicylamide, antiinflammatory cytokines, and
antiinflammatory proteins or steroidal anti-inflammatory agents),
antibiotics, anticlotting factors such as heparin, Pebac,
enoxaprin, aspirin, hirudin, plavix, bivalirudin, prasugrel,
idraparinux, warfarin, coumadin, clopidogrel, PPACK, GGACK, tissue
plasminogen activator, urokinase, and streptokinase; growth
factors. Other active agents include, without limitation: (1)
immunosuppressants; glucocorticoids such as hydrocortisone,
betamethisone, dexamethasone, flumethasone, isoflupredone,
methylpred-nisolone, prednisone, prednisolone, and triamcinolone
acetonide; (2) antiangiogenics such as fluorouracil, paclitaxel,
doxorubicin, cisplatin, methotrexate, cyclophosphamide, etoposide,
pegaptanib, lucentis, tryptophanyl-tRNA synthetase, retaane, CA4P,
AdPEDF, VEGF-TRAP-EYE, AG-103958, Avastin, JSM6427, TG100801, ATG3,
OT-551, endostatin, thalidomide, becacizumab, neovastat; (3)
antiproliferatives such as sirolimus, paclitaxel, perillyl alcohol,
farnesyl transferase inhibitors, FPTIII, L744, antiproliferative
factor, Van 10/4, doxorubicin, 5-FU, Daunomycin, Mitomycin,
dexamethasone, azathioprine, chlorambucil, cyclophosphamide,
methotrexate, mofetil, vasoactive intestinal polypeptide, and
PACAP; (4) antibodies; (5) drugs acting on immunophilins, such as
cyclosporine, zotarolimus, everolimus, tacrolimus and sirolimus
(rapamycin), interferons, TNF binding proteins; (6) taxanes, such
as paclitaxel and docetaxel; statins, such as atorvastatin,
lovastatin, simvastatin, pravastatin, fluvastatin and rosuvastatin;
(7) nitric oxide donors or precursors, such as, without limitation,
Angeli's Salt, L-Arginine, Free Base, Diethylamine NONOate,
Diethylamine NONOate/AM, Glyco-SNAP-1, Glyco-SNAP-2,
(.+-.)-S-Nitroso-N-acetylpenicillamine, S-Nitrosoglutathione,
NOC-5, NOC-7, NOC-9, NOC-12, NOC-18, NOR-1, NOR-3, SIN-1,
Hydrochloride, Sodium Nitroprusside, Dihydrate, Spermine NONOate,
Streptozotocin; and (8) antibiotics, such as, without limitation:
acyclovir, afloxacin, ampicillin, amphotericin B, atovaquone,
azithromycin, ciprofloxacin, clarithromycin, clindamycin,
clofazimine, dapsone, diclazaril, doxycycline, erythromycin,
ethambutol, fluconazole, fluoroquinolones, foscarnet, ganciclovir,
gentamicin, iatroconazole, isoniazid, ketoconazole, levofloxacin,
lincomycin, miconazole, neomycin, norfloxacin, ofloxacin,
paromomycin, penicillin, pentamidine, polymixin B, pyrazinamide,
pyrimethamine, rifabutin, rifampin, sparfloxacin, streptomycin,
sulfadiazine, tetracycline, tobramycin, trifluorouridine,
trimethoprim sulphate, Zn-pyrithione, and silver salts such as
chloride, bromide, iodide and periodate. Growth factors and/or
cytokines, such as, without limitation: basic fibroblast growth
factor (bFGF), acidic fibroblast growth factor (aFGF), vascular
endothelial growth factor (VEGF), hepatocyte growth factor (HGF),
insulin-like growth factors 1 and 2 (IGF-1 and IGF-2), platelet
derived growth factor (PDGF), stromal derived factor 1 alpha (SDF-1
alpha), nerve growth factor (NGF), ciliary neurotrophic factor
(CNTF), neurotrophin-3, neurotrophin-4, neurotrophin-5,
pleiotrophin protein (neurite growth-promoting factor 1), midkine
protein (neurite growth-promoting factor 2), brain-derived
neurotrophic factor (BDNF), tumor angiogenesis factor (TAF),
corticotrophin releasing factor (CRF), transforming growth factors
.quadrature..alpha. and .beta. (TGF-.alpha. and TGF-.beta.),
interleukin-8 (IL-8), granulocyte-macrophage colony stimulating
factor (GM-CSF), interleukins, and interferons. Any combination of
the active agents described above amy be combined.
[0050] Although the matrix described herein can be implanted
without cells with success, cells may be micro-integrated within
the bioerodible matrix or fibers using a variety of methods. For
example, the matrix may be submersed in an appropriate growth
medium for the cells of interest, and then directly exposed to the
cells. The cells are allowed to proliferate on the surface and
interstices of the matrix. The matrix is then removed from the
growth medium, washed if necessary, and implanted.
[0051] In another embodiment, the cells of interest are added to
the liquid that is sprayed or electrosprayed onto the bioerodible
matrix while the matrix is being formed by electrospinning. This
method is particularly suitable when a highly cellularized tissue
engineered construct is desired. In one embodiment, pressure
spraying (i.e., spraying cells from a nozzle under pressure) is
used to deposit the cells. In another, the cells are electrosprayed
onto the non-woven mesh during electrospinning. As described
herein, electrospraying involves subjecting a cell-containing
solution with an appropriate viscosity and concentration to an
electric field sufficient to produce a spray of small charged
droplets of solution that contain cells. As a non-limiting example,
the solution comprising the cells at 10 kV is sprayed onto a target
charged at -15 kV while PEUU or PEEUU is electrodeposited.
[0052] The cells that may be incorporated on or into the
bioerodible matrix include stem cells, progenitor (precursor)
cells, smooth muscle cells, skeletal myoblasts, myocardial cells,
endothelial cells, endothelial progenitor cells, bone-marrow
derived mesenchymal cells and genetically modified cells. In
certain embodiments, the genetically modified cells are capable of
expressing a therapeutic substance, such as a growth factor.
Examples of suitable growth factors include angiogenic or
neurotrophic factor, which optionally may be obtained using
recombinant techniques. Non-limiting examples of growth factors
include basic fibroblast growth factor (bFGF), acidic fibroblast
growth factor (aFGF), vascular endothelial growth factor (VEGF),
hepatocyte growth factor (HGF), insulin-like growth factors (IGF),
transforming growth factor-beta pleiotrophin protein, midkine
protein. In one preferred embodiment, the growth factor is
IGF-1.
[0053] Electrodeposition, for instance electrospinning, is used
herein for depositing substantially uniform fibers onto such a
surface. Electrospinning is a form of electrodeposition that
permits fabrication of scaffolds that resemble the scale and
fibrous nature of the native extracellular matrix (ECM). The ECM is
composed of fibers, pores, and other surface features at the
sub-micron and nanometer size scale. Such features directly impact
cellular interactions with synthetic materials such as migration
and orientation. Electrospinning also permits fabrication of
oriented fibers to result in scaffolds with inherent anisotropy.
These aligned scaffolds can influence cellular growth, morphology
and ECM production. For example, Xu et al. found smooth muscle cell
(SMC) alignment with poly(L-lactide-co-.epsilon.-caprolactone)
fibers (Xu C. Y., et al., "Aligned biodegradable nanofibrous
structure: a potential for blood vessel engineering", Biomaterials
2004 (25) 877-86.) and Lee et al. submitted aligned
non-biodegradable polyurethane to mechanical stimulation and found
cells cultured on aligned scaffolds produced more ECM than those on
randomly organized scaffolds (Lee C. H, et al., "Nanofiber
alignment and direction of mechanical strain affect the ECM
production of human ACL fibroblast", Biomaterials 2005 (26)
1261-1270).
[0054] Generally, electrodeposition involves placing a
polymer-containing fluid (e.g, a polymer solution, a polymer
suspension, or a polymer melt) in a reservoir equipped with a small
orifice, such as a needle or pipette tip and a metering pump. One
electrode of a high voltage source is also placed in electrical
contact with the polymer-containing fluid or orifice, while the
other electrode is placed in electrical contact with a target
(typically a collector screen or rotating mandrel). During
electrodeposition, e.g., electrospinning, the polymer-containing
fluid is charged by the application of high voltage to the solution
or orifice (e.g., about 3-15 kV) and then forced through the small
orifice by the metering pump that provides steady flow. While the
polymer-containing fluid at the orifice normally would have a
hemispherical shape due to surface tension, the application of the
high voltage causes the otherwise hemispherically shaped
polymer-containing fluid at the orifice to elongate to form a
conical shape known as a Taylor cone. With sufficiently high
voltage applied to the polymer-containing fluid and/or orifice, the
repulsive electrostatic force of the charged polymer-containing
fluid overcomes the surface tension and a charged jet of fluid is
ejected from the tip of the Taylor cone and accelerated towards the
target, which typically is biased between -2 to -10 kV. Optionally,
a focusing ring with an applied bias (e.g., 1-10 kV) can be used to
direct the trajectory of the charged jet of polymer-containing
fluid. As the charged jet of fluid travels towards the biased
target, it undergoes a complicated whipping and bending motion. If
the fluid is a polymer solution or suspension, the solvent
typically evaporates during mid-flight, leaving behind a polymer
fiber on the biased target. If the fluid is a polymer melt, the
molten polymer cools and solidifies in mid-flight and is collected
as a polymer fiber on the biased target. As the polymer fibers
accumulate on the biased target, a non-woven, porous mesh (matrix)
is formed on the biased target.
[0055] The properties of electrospun elastomeric matrices can be
tailored by varying the electrospinning conditions. For example,
when the biased target is relatively close to the orifice, the
resulting electrospun mesh tends to contain unevenly thick fibers,
such that some areas of the fiber have a "bead-like" appearance.
However, as the biased target is moved further away from the
orifice, the fibers of the non-woven mesh tend to be more uniform
in thickness. Moreover, the biased target can be moved relative to
the orifice. In certain embodiments, the biased target is moved
back and forth in a regular, periodic fashion, such that fibers of
the non-woven mesh are substantially parallel to each other. When
this is the case, the resulting non-woven mesh may have a higher
resistance to strain in the direction parallel to the fibers,
compared to the direction perpendicular to the fibers. In other
embodiments, the biased target is moved randomly relative to the
orifice, so that the resistance to strain in the plane of the
non-woven mesh is isotropic. The target can also be a rotating
mandrel. In this case, the properties of the non-woven mesh may be
changed by varying the speed of rotation. The properties of the
electrospun elastomeric scaffold may also be varied by changing the
magnitude of the voltages applied to the electrospinning system. In
one particularly preferred embodiment, the electrospinning
apparatus includes an orifice biased to 12 kV, a target biased to
-7 kV, and a focusing ring biased to 3 kV. Moreover, a useful
orifice diameter is 0.047'' (I.D.) and a useful target distance is
about 23 cm. A useful range of high-voltage to be applied to a
polymer suspension or melt is from 0.5-30 kV, more preferably 5-25
kV, even more preferably 10-15 kV.
[0056] Electrospinning may be performed using two or more nozzles,
wherein each nozzle is a source of a different polymer solution.
The nozzles may be biased with different biases or the same bias in
order to tailor the physical and chemical properties of the
resulting non-woven polymeric mesh. Additionally, many different
targets may be used. In addition to a flat, plate-like target, a
mandrel may be used as a target.
[0057] When the electrospinning is to be performed using a polymer
suspension, the concentration of the polymeric component in the
suspension can also be varied to modify the physical properties of
the elastomeric scaffold. For example, when the polymeric component
is present at relatively low concentration, the resulting fibers of
the electrospun non-woven mesh have a smaller diameter than when
the polymeric component is present at relatively high
concentration. Without any intention to be limited by this theory,
it is believed that lower concentration solutions have a lower
viscosity, leading to faster flow through the orifice to produce
thinner fibers. One skilled in the art can adjust polymer
concentrations to obtain fibers of desired characteristics. Useful
ranges of concentrations for the polymer component include from
about 1% wt. to about 15% wt., from about 4% wt. to about 10% wt.
and from about 6% wt. to about 8% wt.
[0058] In electrospinning, polymer fibers are deposited about the
circumference of a mandrel and to generate a planar or
substantially planar structure, the electrodeposited mat/matrix is
cut substantially in the direction of the rotational axis of the
mandrel, or in any manner to generate a useful topology. In use,
more than one electrospun mats/matrices can be attached by any
useful means, such as by "sewing" using sutures, heat annealing,
chemical annealing/cross-linking, etc, though it should be
recognized that the method of attaching the two or more
mats/matrices would have to be strong enough for the end use, e.g.,
to resist rupture/herniation.
[0059] Thickness of the matrix can be controlled by either
adjusting the viscosity of the polymer composition to be deposited
and/or adjusting duration of the electrospinning Use of more
viscous polymer composition may result in thicker fibers, requiring
less time to deposit a matrix of a desired thickness. Use of a less
viscous polymer composition may result in thinner fibers, requiring
increased deposition time to deposit a matrix of a desired
thickness. The thickness of the matrix and fibers within the matrix
affects the speed of bioerosion of the matrix. These parameters are
optimized, depending on the end-use of the matrix, to achieve a
desired or optimal physiological effect.
[0060] Although any form of spraying is expected to be effective,
the liquid, e.g. cell growth media, may be electrosprayed.
Electrospraying is done essentially concurrently with the
electrodeposition of polymer fibers and is conducted in an
essentially identical manner. That is, a potential is created
between the liquid and the target and the liquid is drawn towards
the target. Useful voltage ranges for media range from 0 kV to -10
kV for the target and from 1 kV to 20 kV for the media, and
preferably from -2 kV to -6 kV for the target and from 5 kV to 10
kV for the media. Because the liquid, such as cell growth media, is
less viscous that polymer solutions, the liquid is dispersed onto
the target as droplets rather than as a fiber. As indicated below,
the essentially concurrent spraying of the electrodeposited or
electrospun matrix with fiber deposition results in a structure
with significantly different physical properties as well as
capacity to act as an effective implant for abdominal wall
repair.
[0061] In use, the nonwoven structures described herein are
surgically-implanted. Methods of implanting the materials are known
in the art and are a matter of using standard surgical techniques,
such as for suturing a patch in place about an abdominal wall or
cardiac defect. Methods of implanting are described in the examples
below. For cardiac implantation, exemplary methods are described in
Fujimoto et al. (An Elastic, Biodegradable Cardiac Patch Induces
Contractile Smooth Muscle An Elastic, Biodegradable Cardiac Patch
Induces Contractile Smooth Muscle Infarction, J. Am. Coll. Cardiol.
published online May 24, 2007; doi:10.1016/j.jacc.2007.02.050).
Thus, in one embodiment, a patch of the nonwoven material is
sutured about an infarction on the wall of the heart.
Example 1: Morphological and Mechanical Characteristics of the
Reconstructed Rat Abdominal Wall Following Use of a Wet Electrospun
Biodegradable Polyurethane Elastomer Scaffold
[0062] Although a variety of materials are currently used for
abdominal wall repair, general complications encountered include
herniation, infection, and mechanical mismatch with native tissue.
An approach wherein a degradable synthetic material is ultimately
replaced by tissue mechanically approximating the native state
could obviate these complications. We report here on the generation
of biodegradable scaffolds for abdominal wall replacement using a
wet electrospinning technique in which fibers of a biodegradable
elastomer, poly(ester urethane)urea (PEUU), were concurrently
deposited with electrosprayed serum-based culture medium. Wet
electrospun PEUU (wet ePEUU) was found to exhibit markedly
different mechanical behavior and to possess an altered
microstructure relative to dry processed ePEUU. In a rat model for
abdominal wall replacement, wet ePEUU scaffolds (1.times.2.5 cm)
provided a healing result that developed toward approximating
physiologic mechanical behavior at 8 weeks. An extensive cellular
infiltrate possessing contractile smooth muscle markers was
observed together with extensive extracellular matrix (collagens,
elastin) elaboration. Control implants of dry ePEUU and expanded
polytetrafluoroethylene did not experience substantial cellular
infiltration and did not take on the native mechanical anisotropy
of the rat abdominal wall. These results illustrate the markedly
different in vivo behavior observed with this newly reported wet
electrospinning process, offering a potentially useful refinement
of an increasingly common biomaterial processing technique.
INTRODUCTION
[0063] A tissue engineering approach employing an implanted
degradable synthetic material designed to adequately function
throughout a period of tissue ingrowth and scaffold remodeling and
to result in tissues which mechanically approximate the native
tissue would represent a regenerative approach likely to reduce the
complications seen with current replacement materials, particularly
in the application areas with higher complication rates mentioned
above. Toward this end, we hypothesized that generation of an
engineered tissue based upon an elastic biodegradable synthetic
material, electrospun poly(ester urethane) urea (ePEUU) designed to
better mimic tissue passive mechanical properties prior to
implantation would result in improved outcomes in the
reconstruction of the abdominal wall and other sites of fascia
reconstruction (Courtney T, et al. Biomaterials 2006; 27:3631-8). A
concern, however, is that cellular migration into the ePEUU might
not proceed in a timely fashion (Stankus J J, et al. J Biomater Sci
Polym Ed 2008; 19:635-52) and that a new processing methodology
might be required to facilitate the scaffold remodeling
process.
[0064] In an effort to address the limited cellular infiltration
and remodeling of electrospun scaffolds that might be candidates
for abdominal wall replacement, we report here on the development
of a "wet" electrospinning process in which electrospun PEUU fibers
were concurrently deposited onto a collection mandrel with
electrosprayed serum-supplemented culture medium. Abdominal wall
patches generated using both wet and the traditional "dry"
electrospinning processes with PEUU were evaluated in vitro and in
vivo in a rat abdominal wall replacement model with an emphasis on
evaluating the cellular remodeling process and changes in tensile
mechanical properties under estimated physiological stress levels
were evaluated. For control purposes, ePTFE patches were implanted
and similarly evaluated. The results illustrate the markedly
different in vivo behavior observed with wet versus dry
electrospinning, offering a potentially useful refinement of an
increasingly common biomaterial processing technique.
[0065] Materials and Methods
[0066] Scaffold Fabrication
[0067] Poly(ester urethane) urea (PEUU) was synthesized from
polycaprolactone diol (Mn=2000, Sigma), 1,4-diisocyanatobutane
(Sigma) and putrescine (Sigma) according to previously described
methods (Guan J, et al. J Biomed Mater Res 2002; 61:493-503, also
describing physical characteristics of the product).
[0068] In further detail, PEUU and PEEUU are synthesized as
follows. The molar feed ratio for preparation of PEUU was
polycaprolactone:diisocyanatobutane:putrescine=1:2:1, and the molar
feed ratio for synthesis of PEEUU,
polycaprolactone-PEG-polycaprolactone copolymer:
diisocyanatobutane:putrescine=1:2:1.
[0069] PEUU synthesis: Poly(ester urethane)urea (PEUU) was
synthesized from PCL diol, and BDI using putrescine as a chain
extender by a two-step solvent synthesis method. The
PCL:BDI:putrescine molar ratio was defined as 1:2:1. Briefly, PCL
diol were completely dissolved in DMSO in a 3-neck flask with argon
protection and then BDI was added to the solution, following 4
drops of Sn(Oct).sub.2. The flask was placed in an oil bath at
70.degree. C. After 3 h, the prepolymer solution was cooled at room
temperature and then a putrescine/DMSO solution was added dropwise
into the agitated solution. The final polymer solution
concentration was controlled to be approximately 4% (w/v). Then the
flask was placed in the oil bath and kept at 70.degree. C.
overnight. The polymer was precipitated in an excess volume of cool
deionized water and then dried in a vacuum at 60.degree. C. for 3
d. The yields of all final products were approximately 95%.
[0070] For the current study, a wet electrospun PEUU (wet ePEUU)
was fabricated by a combination of electrospinning and
electrospraying (Stankus J J, et al. Biomaterials 2007; 28:2738-46
and Stella J A, et al. Biomaterials 2008; 29:3228-36). Cell culture
medium (Dulbecco's Modified Eagle Medium (DMEM, Invitrogen) with
10% fetal bovine serum (FBS, GIBCO), 10% horse serum (GIBCO) and 1%
penicillin/streptomycin (GIBCO), and 0.5% chick embryo extract
(GIBCO)) was fed by a syringe pump at 0.2 mL/min into a sterilized
capillary (1.2 mm inner diameter) charged at 7 kV and suspended 4
cm above the target mandrel (6 mm diameter). Concurrently, PEUU in
hexafluoroisopropanol solution (12%, w/v) was fed at 1.5 mL/h from
a capillary, charged at 12 kV and perpendicularly located 20 cm
from the target mandrel. The mandrel was charged at -4 kV and
rotated at 250 rpm (8 cm/s tangential velocity) while translating
back and forth 8 cm along the x-axis at 0.15 cm/s (FIG. 1). As a
control, a dry electrospun PEUU sheet (dry ePEUU) was prepared
using only polymer electrospinning (without media electrospraying)
using the same parameters described above.
[0071] Animal Study
[0072] Adult female syngeneic Lewis rats (Harlan Sprague Dawley
Inc.) 10-12 weeks old, weighing 200-250 g were used for the
abdominal wall reconstruction procedure. Rats were anesthetized
with the inhalation of 1.25-2.5% isoflurane with 100% oxygen. The
abdomen was shaved and prepared with povidone-iodine solution.
Procedures were performed in a sterile environment on a heating
blanket. The surgical procedure was based on the method previously
reported by Lai et al. (J Surg Res 2005; 127:85-92). Briefly, a
skin incision along the linea alba, 3.5 cm in length, was made from
2 cm caudal of the xiphoid process. A surgical defect (1.times.2.5
cm) involving all of the layers of the abdominal wall including the
fascia, rectus muscle, and parietal peritoneum (with the exception
of the skin and subcutaneous soft tissue) was then created. This
anatomic defect was then subsequently repaired by one of three
types of patches selected randomly. The patches (1.times.2.5 cm,
400 .mu.m thick) were sutured to the abdominal fascia by a
continuous 7-0 polypropylene suture without overlap between muscles
and patches in direct contact with the subcutaneous tissue and the
peritoneal viscera. Skin closure was obtained over the patch by
double-layer suturing. The animals were allowed to recover from
anesthesia and returned to their cages. For postoperative analgesic
treatment, 0.1 mg/kg of buprenorphine was administered
subcutaneously 2 times per day for 3 days after surgery.
[0073] The prosthetic materials used in this study were 1) dry
ePEUU, 2) wet ePEUU and 3) expanded polytetrafluoroethylene (ePTFE,
Impra-Bard, Tempe, Ariz.) as a control. Laminar ePTFE prosthesis
was chosen as a control in the study because a laminar prosthesis,
not reticular, is ideal in the case in which the prosthesis has to
be placed in direct contact with viscera (Belton J M, et al. J
Laparoendosc Adv Surg Tech A 2007; 17:160-6) Both ePEUU patches
were processed to 0.4 mm in thickness, matching the ePTFE sheet
thickness employed. Both ePEUU patch groups were oriented so that
the circumferential direction of the mandrel was aligned with the
circumferential direction of the animal and that the axial
direction of the mandrel was aligned with the longitudinal axis
upon their implantation.
[0074] For each group, the implanted samples were surgically
retrieved at 4 and 8 weeks post-implantation (n=7 per group per
time point). At retrieval, animals were euthanized by isoflurane
(5%) inhalation and the abdominal wall was circumferentially
incised to expose the peritoneal cavity and the repair site.
Representative specimens were photographed in situ for later review
and comparisons. The patches were explanted by cutting along an
apron border approximately 5 mm from the original suture line.
Subsequently, a 1.times.1 cm square shape was cut from each
retrieved sample (not including the suture line) and was used for
mechanical characterization of implanted materials. Thickness was
measured in these retrieved samples with a dial thickness gauge
(L.S. Starrett Co.). The remainder of the retrieved sample from all
animals was processed for the histological examination and collagen
assay.
[0075] Histology and Immunohistochemistry
[0076] Hematoxylin and eosin staining (H&E) and
immunohistochemical staining were performed as previously described
(Fujimoto K L, et al. J Am Coll Cardiol 2007; 49:2292-300). The
samples used for histology were fixed in 4% phosphate buffered
paraformaldehyde for 4 h, followed by immersion in 30% sucrose
solution for at least 2 days. The samples were frozen and serially
cryosectioned into 8 .mu.m-thick specimens and processed for
H&E and immunohistochemical evaluation. To assess the
extracellular matrix, sections were stained with the Masson's
modified IMEB trichrome stain kit (IMEB Inc.). Sections for
immunohistochemistry were reacted with primary antibodies against
collagen type I (monoclonal 1:100, Abcam), collagen type III
(monoclonal 1:400, Abcam), and elastin (polyclonal 1:100 Abcam). A
polyclonal antibody against von Willebrand factor (vWF; 1:200,
Abcam) was used to identify endothelial cells. A monoclonal
antibody against CD68 (1:100, AbD Serotec) was used to identify
macrophages. Nuclei were stained with 4',6-diamidino-2-phenyindole,
DAPI (1:10,000, Sigma). A monoclonal antibody against alpha-smooth
muscle actin (.alpha.SMA; 1:200, Abcam), a monoclonal antibody
against calponin (1:200, Abcam), a polyclonal antibody against
SM22.alpha. (1:50, Abcam), and a monoclonal antibody against 150
kDa high molecular weight caldesmon (h-caldesmon; 1:200, Abcam)
were used to identify smooth muscle cell antigens, and a monoclonal
antibody against alpha sarcomeric actin (1:200, Abcam) and alpha
sarcomeric actinin (1:200, Sigma) for skeletal muscle cells. Slides
were examined with an Olympus IX51 microscope and images captured
using DP2-BSW software (Olympus America Inc.). For each retrieved
sample, 10 different microscopic fields at 400.times. magnification
for nuclei count and 10 different fields at 100.times. for vWF
positive structures were photographed. To determine quantity of
cellular infiltration into the materials, the number of nuclei was
measured using a digital image analyzer (Image J, National
Institutes of Health, Bethesda, Md.). Capillaries were identified
as tubular structures positively stained for vWF.
[0077] Collagen Assay
[0078] Collagen levels in retrieved patches were measured by using
the Sircol collagen assay kit (Accurate Chemical and Scientific
Corp.), as described previously by DuBay et al. (DuBay D A, et al.
Surgery 2005; 137:463-71). The approximately 100 mg (wet weight)
samples of abdominal wall patches without apron tissue were weighed
and mechanically dissolved with scissors and sonicator in protease
inhibitor cocktail (Sigma-Aldrich) and 0.5 M acetic acid solvent (1
mL to 100 mg wet tissue weight). The samples were then stirred
overnight at 4.degree. C. and centrifuged at 16,000.times.g for 60
min. Sircol dye reagent (1.0 mL) was added to 10 .mu.l of
supernatant from each sample followed by placement in a mechanical
shaker at room temperature for 30 min. The samples were centrifuged
at 16,000.times.g for 10 min and the supernatant was removed.
Sodium hydroxide (0.5 M, 1.0 mL) was added to the collagen-bound
dye pellet to release the bound dye into solution. Aliquots of each
sample (200 .mu.L) were transferred to the wells of a 96-well plate
and the optical density was measured at 540 nm. Results were
normalized as mg collagen/g wet tissue.
[0079] Biaxial Mechanical Property Measurements
[0080] Biaxial mechanical testing was performed for patches prior
to implantation, for native tissues removed during the implantation
procedure, and for retrieved samples at each time point (4 and 8
weeks) using a method previously described (Sacks M S. J Elasticity
2000; 61:199-246). Samples were prepared for testing through
immersion into Ringer's solution (82 mM NaCl, 60 mM KCl, 2 mM
CaCl.sub.2, 10 mM Trizma-HCl, 10 mM Trizma-base, 11 mM dextrose)
supplemented with verapamil (0.5 mM) and ethylene glycol
tetraacetic acid (EGTA, 0.5 mM) for 1 h to obtain complete muscle
fiber relaxation. Samples were then trimmed to achieve 10.times.10
mm sections for testing. A biaxial mechanical testing approach was
employed as follows. Samples were tested in physiological saline
solution at room temperature using a Lagrangian membrane tension
(T, force/unit length) controlled protocol designed to apply
constant and equal biaxial tension to the sample up to a maximum of
200 N/m. This value was chosen based on preliminary results from
our laboratory that indicated this was the maximum tension that the
native tissue could reliably withstand without incurring damage.
Using thin slices of polypropylene suture (Ethicon) affixed to the
sample to form four small markers of w1 mm in diameter in the
central region used to compute local strains as well as the
deformation gradient tensor F. From F, the axial stretches
.lamda..sub.CD=F.sub.11 and .lamda..sub.LD=F.sub.22 were determined
(CD=circumferential direction, LD=longitudinal direction). Two
equi-biaxial tension protocols containing 10 cycles each were
performed. The first protocol was used to precondition the sample;
data were recorded from the final cycle of the second protocol.
Post-processing was performed using a preconditioned free-float
reference state image.
[0081] Statistical Analysis
[0082] Statistical evaluation was performed using Prism version
4.0c (GraphPad Software Inc.). Results are listed as
mean.+-.standard error of the mean. One-way analysis of variance
(ANOVA) followed by Tukey-Kramer multiple comparison testing was
applied where multiple comparisons were made. The Mann-Whitney U
test was used for the comparison of vWF positive structures count
among the wet group since these data were not normally distributed.
Statistical analysis for mechanical characterization was performed
by using one-way ANOVA to compare the maximum stretches observed in
each sample. Differences were considered to be statistically
significant at p<0.05.
[0083] Results
[0084] Material Characteristics
[0085] Wet PEUU sheet microstructure prepared using the
electrospinning/electrospray method is seen in FIG. 1, as is the
dry PEUU sheet. Macroscopically the wet PEUU was moist upon cutting
and clearly retained culture medium within the scaffold
interstices. While both dry and wet PEUU microstructures exhibited
continuous polymer fibers without bead formation, the wet PEUU
fibers qualitatively exhibited a greater degree of looping and more
tortuosity. These structural features were consistently observed
across the scaffolds generated. The average diameter of wet PEUU
fibers (1549.+-.270 nm) was found to be larger than that for dry
PEUU fibers (824.+-.47 nm) (p<0.05).
[0086] Postoperative Course and Gross Observations
[0087] No abnormal behavior indicating pain or distress, or
abnormal weight gain or loss was observed after surgery. No
herniation at the repair site of the abdominal wall was observed in
any of the rats during the study (FIG. 2A, a-c). Additionally, no
group showed any adhesions to the visceral organs at the site of
the implanted patch except for slight omental tissue adhesion which
was seen consistently across groups. Although the pre-implantation
materials were 400 .mu.m thick, both wet and dry ePEUU patched
areas were 0.9-1 mm thick at 8 week, almost the same as native rat
abdominal wall, while ePTFE showed no apparent change in thickness
(FIG. 2B). Considering that the thickness measured includes the
thickness of neoperitoneum formed, this result suggests that
neo-formation of tissue beneath and above ePTFE was minimal.
[0088] Histology and Immunochemistry
[0089] H&E and Masson's Trichrome Staining
[0090] Under light microscopy, all of the prosthetic materials
showed layered fibrous tissue surrounding the materials at each
time point (FIG. 3). Almost no tissue ingrowth was observed in
ePTFE, while for wet ePEUU polymer degradation was accompanied by
collagenous fiber deposition. Collagenous fiber deposition
indicated by blue staining in the Masson's trichrome images of FIG.
3, increased at the 8 week time point in the interior of the wet
ePEUU, whereas no staining suggestive of collagenous deposition was
seen within the dry ePEUU patches.
[0091] Cellular Infiltration
[0092] High magnifications of dry and wet ePEUU H&E stained
sections at 4 and 8 weeks point are shown in FIG. 4. In dry ePEUU
some positive labeling is detected that may represent a sparse
cellular infiltrate into the polymer matrix. In contrast, wet ePEUU
exhibited substantial cellular migration near the center of the
patch at 4 weeks surrounding the remaining ePEUU fibers. By 8 weeks
this infiltration is largely complete although regions of fibers
are still clearly visible. For a more specific evaluation of
cellular infiltration FIGS. 5A and 5B shows nuclear staining of
sections from the three patch types at the two explant times and
quantification of nuclear numbers. Virtually no cellular
infiltration was noted in the ePTFE patches at either time. Modest
cellular infiltration occurred in the dry ePEUU group by 8 weeks.
For the wet ePEUU cell numbers were significantly increased
compared to the dry group at both 4 and 8 week time points.
Immunostaining for CD68 in wet ePEUU patches at 8 weeks revealed
macrophage infiltration into the wet ePEUU, that was more
predominant in the edge regions and lighter in the central areas
(FIG. 6).
[0093] Neovascularization
[0094] The vascular density within the patches, as assessed by vWF
immunostaining, is seen in FIGS. 7A and 7B. No positive staining
was seen for either the ePTFE or dry ePEUU at any time point,
whereas the wet ePEUU patches demonstrated sparse vWF positive
structures at 4 weeks, and significantly increased numbers of such
structures at 8 weeks (p<0.05). In many instances these vWF
positive structures were found to be surrounded by .alpha.-smooth
muscle actin positive cells suggesting the presence of smooth
muscle cells. The spatial relationship in the double positive
staining with vWF and .alpha.SMA suggest mature vascular formation
(Bir S C, et al. J Vasc Surg 2009; 50:870-9).
[0095] Extracellular Matrix Deposition
[0096] Immunostaining against collagen types I and III, and elastin
at 8 weeks revealed substantial extracellular matrix (ECM)
component elaboration within wet ePEUU patches, but minimal such
deposition in dry ePEUU patches and none within ePTFE (FIG. 8)
consistent with the cellular infiltration findings. All three
materials had substantial ECM deposition onto the patch exterior
surfaces. With dry ePEUU the ECM deposition within the patch was
associated with the cellular infiltrate near the surfaces. Inside
the wet ePEUU substantial deposition of types I and III collagen
was observed. The collagen assay, performed on the patches
including the surface ECM deposition, showed significantly greater
collagen levels from the wet ePEUU patches than for the other two
patch types at 8 weeks (FIG. 9). Histologically examining the wet
ePEUU patch at 4 versus 8 weeks a trend could be qualitatively
appreciated in ECM deposition increasing with time (FIG. 10).
[0097] Immunostaining for Muscle Cell Markers
[0098] The cellular infiltrate in the 8 week wet ePEUU patches was
also evaluated in terms of immunostaining for muscle cell markers.
Skeletal muscle markers including alpha sarcomeric actin and alpha
sarcomeric actinin were not found in the examined sections.
However, alpha-smooth muscle actin positive cells were found on the
surfaces of both ePEUU patch types and in the cellular infiltrate
for the wet ePEUU patches. In evaluating the presence of
contractile markers for smooth muscle cells, immunostaining for
smooth muscle cell specific antigen calponin was positive for both
wet and dry ePEUU patch surfaces. For the cellular infiltrate in
wet ePEUU, calponin, SM22.alpha., and h-caldesmon were all evident
(FIG. 11). The relative degree of staining for each antigen (-=not
present; +=present; ++=strongly present) is presented in Table
1.
TABLE-US-00001 TABLE 1 Relative degree of immunohistochemical
labeling of cells associated with implanted patches. cal- .alpha.-
.alpha.- .alpha.SMA ponin SM22.alpha. HCALD actinin actin ePTFE
surface - - - - NA NA inner - - - - NA NA dry surface + + - - NA NA
ePEUU inner - - - - NA NA wet surface ++ ++ ++ ++ - - ePEUU inner +
+ + + - - - = not present; + = present; ++ = strongly present; NA =
not assessed
[0099] Biaxial Mechanical Property Measurements
[0100] Native rat abdominal wall tissue was found to possess a very
high degree of anisotropy upon applied equal planar biaxial tension
with the circumferential axis being markedly stiffer than the
longitudinal axis (p<0.0001; FIG. 12A). Prior to implantation,
each patch material exhibited distinct mechanical properties. The
ePTFE patch material was largely isotropic, and very stiff,
deforming less than 5% in either direction under the maximum
applied tension (FIG. 12B). In contrast, dry ePEUU was
significantly more compliant in the circumferential axis
(p<0.001). The addition of sterile media to the ePEUU construct
during fabrication of wet ePEUU served to make both axes more
compliant (p<0.05). Compared to the native tissue, the
circumferential axis was again significantly more compliant
(p<0.001). It was also found that, before implantation, the
ePEUU constructs both displayed an anisotropy (p<0.01) that was
stiffer in the longitudinal (axial) direction relative to the
circumferential direction, in contrast to the native tissue.
Following an implantation period of four weeks, a marked difference
can be seen in the mechanical properties of the elastomeric
constructs (FIG. 12B). Both dry and wet ePEUU became markedly
stiffer and more isotropic (p<0.05). Wet ePEUU remained more
compliant than dry (p<0.05), but only in the longitudinal axis.
The ePTFE showed no change and was again the most stiff and
isotropic. After 8 weeks, wet ePEUU exhibited a reversal from the
implanted anisotropy to exhibit a mechanical behavior that more
resembled native tissue (p<0.001 for 4 versus 8 weeks and for
longitudinal versus circumferential direction at maximum stretch
levels). Neither dry ePEUU nor ePTFE groups changed significantly
between four and eight weeks of implantation.
DISCUSSION
[0101] When repairing large abdominal wall defects, applying simple
surgical closure may lead to increased abdominal pressure and
adversely affect visceral function through the development of
abdominal compartment syndrome (ACS). In trauma patients, injury
itself may lead to ACS, which is treated with damage control
laparotomy to relieve pressure at the expense of an open abdominal
wound. This laparotomy wound is reduced mechanically with time, but
will ultimately still require permanent abdominal wall
reconstruction in the majority of cases. Prosthetic materials are
required to allow repair of abdominal defects to avoid ACS as well
as to avoid hernia recurrence. Desirable abdominal wall prostheses
properties include infection resistance, elasticity and strength,
acceptably low foreign body effect, and facilitation of
vascularized tissue ingrowth without induction of bowel adhesions.
Currently there are no materials that completely address all of
these concerns.
[0102] The role of biomaterial mechanical properties in abdominal
wall repair is well recognized in a general sense, with weak
materials being associated with a risk for herniation and materials
that are too stiff being associated with patient discomfort. The
degree to which these factors come into consideration varies with
the extent of the repair being considered. Abdominal wall
compliance after prosthesis implantation was shown in a small
animal study to be the most important mechanical property for
predicting a low incisional hernia recurrence rate (DuBay D A, et
al. Surgery 2006; 140:14-24). Incisional hernia repair, which may
be associated with increased stiffness in the abdominal wall,
potentially due to fibrosis and atrophy, results in a progressive
mechanical impedance mismatch that increases the transfer of load
forces to the wound healing interface and increases herniation
risk. Moreover, preservation of elasticity would also allow the
abdominal wall to function more naturally in a dynamic fashion,
maintaining its flexibility and avoiding abdominal stiffness and
sites of compliance mismatch. It then follows that special
consideration should be given to the compliance of biomaterials
chosen for this application.
[0103] Here, novel prosthetic materials possessing varied tensile
properties and microstructures were implanted into an abdominal
wall defect model with particular attention to materials that might
better mimic the native tissue behavior.
[0104] Mechanical characteristics of these implanted elastomeric
constructs changed substantially during the implantation period.
This change can be explained through the ingrowth of host tissue
into the construct matrix as well as implanted material
degradation. As ePTFE has been shown to be absent of tissue
ingrowth, it follows that its properties would remain largely
unchanged throughout the course of implantation, as was observed.
FIGS. 3-5, 8 and 9 show the cellular infiltration and ECM
elaboration within the constructs, and the relatively small
quantity of biological material found within dry ePEUU at the time
of explant. This modest deposition may explain the moderate
stiffening effect observed for the dry ePEUU. For wet ePEUU, the
collagen assay and histological results suggested that more
extensive changes in the mechanical results might be expected given
the level of cellular ingrowth which is concomitant with the
macrophage infiltration and ECM elaboration observed. At four
weeks, collagen and elastin were shown to be present, but were
disorganized. By eight weeks, the elaborated ECM and cellular
components had become more extensive, serving as the likely factor
producing the change in anisotropy from what was observed at
implantation to an anisotropy which mimicked native tissue. It is
important to note that the suture line was not included in the
sections that were mechanically tested. This was done to ensure a
proper representation of the explant mechanical properties, rather
than suture strength. Gross examination of the explanted tissue, as
well as the lack of herniation in any animal indicated that breaks
at the anastomosis did not occur.
[0105] In considering a biodegradable material approach to
abdominal wall replacement, the most obvious risk is that the
material will lose strength before sufficient tissue ingrowth and
organization has occurred, thus putting the site of repair at risk
for failure. As discussed above, in the case of dry ePEUU, the
degradation of mechanical properties did not appear to be an issue
in the period of study, and for ePTFE this is not a concern for the
material. For wet ePEUU we were interested in the nature of the
ingrowing tissue, both mechanically and in terms of the cellular
and ECM constituency. If the scaffold were to be replaced by other
than thin and stiff fibrotic scar tissue, this might ultimately be
an improved functional result for the patient in terms of passive
mechanical properties. Both the wet and dry ePEUU implants varied
from ePTFE in terms of the surface tissue that surrounded the
scaffolds, with .alpha.SMA and calponin positive cells being found.
The wet ePEUU material further varied from dry ePEUU in that these
near surface cells were also positive for contractile smooth muscle
cell markers SM22.alpha. and h-caldesmon. In addition, the
extensive cellular infiltrate in the wet ePEUU consistently stained
positive for the array of contractile smooth muscle cell markers
examined by immunohistochemistry. Stressed myofibroblasts could
generate collagen fibers and have actin filaments as a contractile
element, however, D'Addario et al. reported that h-caldesmon in
particular is a specific marker of fully differentiated smooth
muscle (D'Addario S F, et al. J Cutan Pathol 2002; 29:426-9). In
other work with PEUU scaffolds we have observed the contractile
phenotype of smooth muscle cells, confirmed with
immunohistochemical staining as well as electron microscopy, in the
area near degraded PEUU scaffolds placed on cardiac tissue
(Fujimoto K L, et al. J Am Coll Cardiol 2007; 49:2292-300). The
passive mechanical properties of this biological material replacing
the scaffold appeared to better mimic native tissue, although
clearly more temporally extended studies would be needed to examine
the ultimate outcome of the scaffold remodeling process in both wet
and dry ePEUU implants. The active functionality of any smooth
muscle tissue is also unclear. For this report the tissue was
clearly not organized sufficiently to merit such an evaluation.
Whether this nascent tissue might ultimately develop into organized
structures with active mechanical properties is also not clear,
although the goal of this replacement approach currently is to
achieve improved passive mechanical properties.
[0106] From a materials processing perspective, a major aspect of
this investigation was the comparison of traditional dry
electrospinning versus the wet technique wherein cell culture
medium is electrosprayed concurrently with electrospun fiber
deposition. We have previously reported a technique wherein cells
in culture medium are electrosprayed concurrently with PEUU
electrospinning to form a tissue construct that has cells
integrated on a microscale within it (Stankus J J, et al.
Biomaterials 2007; 28:2738-46 and Stankus J J, et al. Biomaterials
2006; 27:735-44). The technique of this example is in a sense a
derivative of that technique with the difference being that here
cells are not utilized. We hypothesized that the so-called wet
electrospinning technique with culture medium would lead to
improved cellular migration since serum factors would be deposited
throughout the scaffold forming process and in early control
experiments we qualitatively noticed that wet electrospinning
resulted in scaffolds with softer mechanical properties and a
distinct morphology with more fiber tortuosity (looping) that might
putatively ease cell migration by more readily locally distended
fibers. It was also considered that the wet electrospinning process
might lead to less inter-fiber bonding, which would also ease cell
migration and contribute to scaffold softening. Quite recently
there has been a brief report in the literature (Yokoyama Y, et al.
Mater Lett 2009; 63:754-6) where another "wet" electrospinning
technique was used. In that method, electrospun fibers were
directly deposited onto fluid surfaces of varying surface tension
to form scaffolds of increased porosity, and as in our method,
apparently increased fiber tortuosity. For the present study, the
specific mechanisms by which wet electrospinning would allow cell
migration were not the focus, rather we sought to investigate
whether this process might indeed result in a different scaffold
remodeling result in an application of clinical relevance,
abdominal wall repair. As mentioned above, in this application area
a regenerative approach that results in a more physiological result
mechanically would be attractive.
[0107] The findings of slow cellular infiltration into dry ePEUU
were expected. In subcutaneous implantations comparing dry ePEUU
with dry ePEUU blended to a varying degree with a urinary bladder
derived ECM, we found little degradation of dry ePEUU, presumably
due to both slow polymer hydrolysis and the lack of macrophage
access to interior fibers due to the tight fiber format (Stankus J
J, et al. J Biomater Sci Polym Ed 2008; 19:635-52). With ECM
blended scaffolds, degradation was markedly accelerated, attributed
to increased fiber degradation due to the protein blending (Stankus
J J, et al. J Biomater Sci Polym Ed 2008; 19:635-52; El-Kurdi M S,
et al. Biomaterials 2008; 29:3213-20; and Sell S A, et al. Adv Drug
Deliv Rev 2009; 61:1007-19) and thus increased cellular access.
With wet ePEUU processing protein was not incorporated per se into
the fibers, although serum proteins would have likely adsorbed in
the process. These serum factors may have served to encourage
cellular infiltration, and with macrophages, possibly increased
phagocytic activity. It was also observed that fiber diameter was
increased for wet ePEUU. This may have been due to the deposition
process leading to less dense fibers that are not pulled to the
same extent from adhesion point to adhesion point as in dry
electrospinning, due to fiber-fiber sliding in the wet environment.
Such loosely deposited fibers might be less dense or less
crystalline and might hydrolyze more rapidly, but we did not
specifically investigate these potential effects. To separate
scaffold morphology effects from culture medium deposition effects,
a follow up investigation comparing wet ePEUU formed by
electrospraying with a buffered salt solution to electrospraying
with culture medium would be of interest.
[0108] For clinical application one might wish to consider the use
of an allogenic serum solution possibly supplemented with ionic
species, should these prove to be determinant in achieving a
required morphology. The use of serum in this process, with its
array of growth factors and adhesion molecules, would be attractive
versus isolated growth factors (as is done for controlled release
applications) and specific adhesion molecules from both an economic
and regulatory perspective. The general approach of using wet
electrospinning to create an elastic scaffold incorporating a
growth factor rich protein solution may find application in other
clinical areas. Repair of the pelvic floor and fascial replacement
in a variety of other settings such as breast, oral and
maxillofacial reconstructive surgery might benefit from such soft
constructs (Natale F, et al. Curr Opin Urol 2006; 16:407-12 and
Sergent F, et al. Eur J Obstet Gynecol Reprod Biol 2009; 147:106-1)
and postmastectomy reconstruction (Becker S, et al. Plast Reconstr
Surg 2009; 123:1-6).
[0109] Several limitations of the current report should be
mentioned. First, the scaffold remodeling process and mechanical
property changes were only observed over an 8-week period. Although
the wet ePEUU had substantially degraded during this period, the
sustainability of the developing architecture for longer periods is
not clear. Future studies in a larger animal model with longer time
points would better define the clinical potential of this approach.
A larger animal model would also allow the evaluation of more
appropriately sized implants and would better mimic the physical
forces experienced by the human abdomen, although quadrupeds remain
limited for this purpose.
CONCLUSIONS
[0110] A new wet electrospinning technique in which biodegradable
elastomer fibers were concurrently deposited with electrosprayed
culture medium was found to result in markedly different scaffold
mechanical behavior and to experience much greater cellular
infiltration and scaffold remodeling in vivo versus dry electrospun
constructs. In a model for abdominal wall replacement in the rat,
wet ePEUU scaffolds provided a healing result that better
approximated physiologic passive mechanical behavior and where an
extensive cellular infiltrate possessing contractile smooth muscle
markers was observed together with extensive ECM elaboration.
Control implants of dry ePEUU and ePTFE did not experience
substantial cellular infiltration and did not take on the native
mechanical anisotropy of the rat abdominal wall.
Example 2: Preparation of an Autogenic Rat Abdominal Wall Repair
Matrix
[0111] Whole rat blood is obtained from syngeneic rats. Rat serum
is prepared from the whole blood by allowing the blood to clot, and
then removing the clot by ringing and centrifugation. Platelet rich
plasma is prepared as described above. Multiple samples of PEUU and
PEEUU polymer are electrospun onto a mandrel concurrently with
electrospraying of various dilutions of rat serum or platelet rich
plasma in normal saline and PBS to final serum or platelet rich
plasma concentrations of 0%, 10%, 20%, 50%, 75% and 100% (v/v),
essentially as described in Example 1. Each wet-electrospun matrix
is removed from the mandrel by cutting the matrix along its long
axis.
[0112] The resultant 11 matrices (0%, 10%, 20%, 50%, 75% rat serum
or platelet rich plasma in normal saline, 0%, 10%, 20%, 50%, 75%
rat serum or platelet rich plasma in PBS and 100% rat serum or
platelet rich plasma) are implanted in separate, syngeneic rat
abdominal walls as described in Example 1. The rats are sacrificed
at 8 weeks to determine the level of cell infiltration and tissue
regeneration as described in Example 1.
Example 3--Human Abdominal Wall Repair
[0113] Human serum or platelet rich plasma is obtained according to
standard practice substantially as described in Example 2. The
serum or platelet rich plasma is diluted in normal saline to a
final concentration of approximately one of 10%, 20%, 50%, 75%. A
matrix is prepared as described in the previous examples by
electrospraying 100% human serum or platelet rich plasma or 0%,
10%, 20%, 50% or 75% (v/v) human serum or platelet rich plasma in
normal saline, essentially as described above. The resultant matrix
is implanted in the abdominal wall of a human patient in need
thereof, and remains in the patient until it dissolves. Biopsies of
the implant are taken at 8 and 16 weeks post-implantation for
analysis.
Example 4--Human Abdominal Wall Repair
[0114] Serum or platelet rich plasma is obtained from a patient in
need of abdominal wall repair, according to standard practice
substantially as described in Example 2 and above. The serum or
platelet rich plasma is diluted in normal saline to a final
concentration of approximately one of 10%, 20%, 50%, 75% (v/v). A
matrix is prepared as described in the previous examples by
electrospraying 100% human serum or platelet rich plasma or 0%,
10%, 20%, 50% or 75% human serum or platelet rich plasma in normal
saline, essentially as described above. The resultant matrix is
implanted in the abdominal wall of the patient in need thereof, and
remains in the patient until it dissolves. Biopsies of the implant
are taken at 8 and 16 weeks post-implantation for analysis.
Example 5--Human Organ Repair
[0115] Human serum or platelet rich plasma is obtained according to
standard practice substantially as described in Examples 2-4. The
serum or platelet rich plasma is diluted in normal saline to a
final concentration of approximately one of 10%, 20%, 50%, 75%
(v/v). A matrix is prepared as described in the previous examples
by electrospraying 100% human serum or platelet rich plasma or 0%,
10%, 20%, 50% or 75% human serum or platelet rich plasma in normal
saline, essentially as described above, optionally omitting cutting
the matrix off of the mandrel in favor of sliding the matrix off of
the mandrel. The resultant matrix is implanted in one of the skin,
blood vessel(s), muscle, esophagus, trachea, stomach, intestine,
rectum or bladder of the patient to repair damage or a defect at
the site of implantation and remains in the patient until it
dissolves. The serum or platelet rich plasma is obtained from the
patient or one or more other humans. Biopsies of the implant are
taken at 8 and 16 weeks post implantation for analysis.
Example 6--Preparation of Infarcted Model in Rats and Evaluation of
Patch Materials
[0116] All experimental rats (young female Lewis rats (10-week
old)) undergo cardiac muscle infarction. The rats are anesthetized
with isoflurane (5.0% induction and 1.25 to 1.5% maintenance with
100% oxygen). The rats undergo endotracheal intubation and are
mechanically ventilated with a small animal respirator at a
frequency of 60-70/min. Tidal volume is set to 0.6-2.0 ml depending
on body weight of the rat. The rats are placed on a warming pad
(37.degree. C.) in the supine position. The hair on the anterior
chest is trimmed with an electric clipper. The skin leads of
electrocardiography are attached on both forelimbs and on the left
hind limb, and the electrocardiogram is monitored. Before skin
incision, one dose of lidocaine (100 mg/kg) anti-arrhythmia and
cefuroxime (100 mg/kg) antibiotic is administered intramuscularly
for prophylaxis of fatal arrhythmia (left ventricular fibrillation)
after infarction and surgical site infection, respectively.
[0117] The skin of the left chest is sterilized with
povidone-iodine solution. The surgeries are performed using aseptic
techniques with sterile instruments. The rat heart is exposed
through a left thoracotomy (4' inter-costal). After identifying the
left anterior branch of rat coronary artery, it is carefully
ligated with 7-0 prolene suture. If significant arrhythmia is noted
on ECG monitoring, the procedure is held until the rhythm is
normalized. The muscle layer and skin is then closed with 4-0
polyglactin absorbable suture (VICRYL, Ethicon, Inc.). The rats are
observed in the surgical suite until fully recovered from the
anesthesia, and then returned to the housing facility. The first 3
days after surgery, Tramadol (5 mg/kg) analgesia is administered 2
times a day intramuscularly, and cefuroxime (100 mg/kg) antibiotic
is injected intramuscularly once a day. If the animal develops
severe cardiac failure during or after the procedure, it is
immediately euthanized.
[0118] The creation of infarction might cause fatal arrhythmias.
Lidocaine (100 mg/kg) is administered for prophylactic therapy.
These episodes are treated with cardiac compression (squeezing the
heart gently by forceps when a ventricular fibrillation (VF)
happens during open chest. When a VF happens after closing the
chest, pressing the closed chest wall is performed soon after
mechanical ventilation with 100% oxygen. The frequency is 200
times/min.
[0119] Experimental group selection: Two weeks following ligation
of the coronary artery, patch implantation surgery is performed.
Before surgery, animals are screened by echocardiography for
infarct size as estimated by the percentage of scar area (akinetic
or dyskinetic regions) to LV free wall (LVFW) area. Animals with
infarcts greater than 25% of the LVFW are chosen. The rest, with
infarcts less than 25%, are sacrificed. After this selection, all
experimental animals are divided into 5 groups:
[0120] I) Rats implanted with PEUU-PRP patches
[0121] II) Rats implanted with control electrospun PEUU patches
(dry PEUU)
[0122] III) Rats implanted with control electrospun PEUU+PBS
patches
[0123] IV) Rats implanted with control electrospun PEUU+cell
culture media patches
[0124] V) Rats infarcted (no implantation) as a control.
[0125] Patch material is sterilized by exposure to the ultraviolet
light source in a laminar flow hood (Class II A/B3 Biological
Safety Cabinet) for 2 hours. The skin of the left chest of
anesthesized rats (as above) is sterilized with povidone-iodine
solution. The rat heart is exposed through a left thoracotomy. The
infarcted part of rat cardiac muscle is readily identified because
of its color. Covering the infarcted part with material, the
cardiac patch is sutured using 7-0 polypropylene with over-and-over
sutures. The wound is closed as above. For the sham (infarction
control, group V)) surgery, the rats are anesthesized as above and
the skin of the left chest is sterilized with povidone-iodine
solution. The rat heart is exposed through a left thoracotomy.
After observation of the infarct area, no surgical intervention is
performed. The wound is closed as described above.
[0126] For hemodynamic assessment (Simultaneous Tail-cuff blood
pressure measurement and Echocardiography and Pressure-volume loop
measurement), rats are anesthesized as described above. Once the
plane of anesthesia is established, the animal is placed in a spine
position in a custom warmed glass chamber, which includes a
vertical trough. Once it has been affirmed that the animal is
secured and breathing adequately, the chest is shaved and prepared
for echocardiographic examination. Combined blood pressure
measurement (tail-cuff blood pressure system) and echocardiography
are used to take measurements of left ventricular dimensions and
heart inlet and out let blood flow velocities. The probe of the
echocardiography machine is put on the chest wall, and ultrasonic
beam detects the heart motion. This scanning is not invasive.
Hemodynamic measurement is performed at 2 week after 1st surgery
(creation of LV infarction) for experimental group selection, and 4
and 8 weeks after 2nd surgery.
[0127] For pressure-volume loop measurement, each animal is placed
in a supine position under anesthesia at the endpoint (just before
the euthanasia). The right carotid artery is exposed through small
skin incision, and a catheter for real-time pressure-volume
measurement (SPR-838, Millar) is inserted from the right carotid
artery to left ventricle of the heart. When the measurement is
done, the animal is euthanized immediately by potassium solution
injection.
[0128] Having described this invention above, it will be understood
to those of ordinary skill in the art that the same can be
performed within a wide and equivalent range of conditions,
formulations and other parameters without affecting the scope of
the invention or any embodiment thereof.
* * * * *