U.S. patent application number 14/567474 was filed with the patent office on 2015-04-09 for novel biodegradable elastomeric scaffold for tissue engineering and light scattering fingerprinting methods for testing the same.
The applicant listed for this patent is NORTHWESTERN UNIVERSITY. Invention is credited to Guillermo Ameer, Antonio Webb, Jian Yang.
Application Number | 20150099853 14/567474 |
Document ID | / |
Family ID | 34381088 |
Filed Date | 2015-04-09 |
United States Patent
Application |
20150099853 |
Kind Code |
A1 |
Ameer; Guillermo ; et
al. |
April 9, 2015 |
Novel Biodegradable Elastomeric Scaffold for Tissue Engineering and
Light Scattering Fingerprinting Methods for Testing the Same
Abstract
The present invention is directed to a novel biocompatible
polymer that may be used in tissue engineering. More specifically,
the specification describes methods and compositions for making and
using citric acid copolymers.
Inventors: |
Ameer; Guillermo; (Chicago,
IL) ; Yang; Jian; (Arlington, TX) ; Webb;
Antonio; (Chicago, IL) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
NORTHWESTERN UNIVERSITY |
Evanston |
IL |
US |
|
|
Family ID: |
34381088 |
Appl. No.: |
14/567474 |
Filed: |
December 11, 2014 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13596529 |
Aug 28, 2012 |
8911720 |
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14567474 |
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12370312 |
Feb 12, 2009 |
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13596529 |
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10945354 |
Sep 20, 2004 |
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12370312 |
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60503943 |
Sep 19, 2003 |
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60556642 |
Mar 26, 2004 |
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Current U.S.
Class: |
525/437 |
Current CPC
Class: |
C08G 63/685 20130101;
C08G 63/40 20130101; C08G 63/06 20130101; C08G 63/66 20130101; C08G
63/914 20130101; C08G 63/16 20130101 |
Class at
Publication: |
525/437 |
International
Class: |
C08G 63/91 20060101
C08G063/91 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] This invention was made with government support under Grant
No. R21HL71921-02 awarded by the National Institutes of Health. The
government has certain rights in the invention.
Claims
1-14. (canceled)
15. A method of producing a thermoset elastomer comprising
crosslinked polyesters having the formula: ##STR00001## wherein R
is selected from a hydrogen or a poly(diol citrate), and wherein
each A is a linear aliphatic diol monomer; comprising the steps of:
(a) polymerizing citric acid and said linear aliphatic diol monomer
to produce poly(diol citrate) pre-polymers; and (b) crosslinking
the poly(diol citrate) pre-polymers to produce the elastomeric
network.
16. The method of claim 15, wherein the crosslinking step is done
at lower temperature than the polymerizing step.
17. The method of claim 16, wherein the polymerizing step is done
at 120-140.degree..
18. The method of claim 16, wherein the crosslinking step is done
at about 60-120.degree. C.
19. The method of claim 15, wherein each A is O(CH2).sub.xO, and X
is between 2 and 20.
20. The method of claim 19, wherein X is between 4 and 18.
21. The method of claim 20, wherein the linear aliphatic diol
monomer is 1,8-ocantediol, 1,10-decanediol, or
1,12-dodecanediol.
22. The method of claim 15, wherein each linear aliphatic diol
monomer is the same.
23. A thermoset elastomer produced by the method of claim 15.
24. The method of claim 15, wherein the crosslinking step is
performed under vacuum.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of U.S. patent
application Ser. No. 13/596,529, filed Aug. 28, 2012, which is a
continuation of U.S. patent application Ser. No. 12/370,312, filed
Feb. 12, 2009, which is a divisional of U.S. patent application
Ser. No. 10/945,354, filed Sep. 20, 2004, which claims the benefit
U.S. Provisional Patent Application No. 60/503,943 filed Sep. 19,
2003 and U.S. Patent Application No. 60/556,642, filed Mar. 26,
2004, each of which is incorporated by reference in its
entirety.
FIELD OF THE INVENTION
[0003] The present invention is generally directed to a substrate
used for tissue engineering. The substrate is a biodegradable
elastomeric polymer. Methods and compositions for testing and using
the same are disclosed.
BACKGROUND OF THE RELATED ART
[0004] The field of tissue engineering has slowly emerged within
the past 2 decades, driven primarily by the large demand for
replacement of diseased or damaged tissue [1]. Tissue engineering
presents enormous challenges and opportunities for materials
science from the perspective of both materials design and materials
processing [2]. Successful tissue regeneration must go beyond
reproducing shape and structure to restore biological and
mechanical function and long-term integration with surround native
tissues [3]. Tissue engineering requires the use of a three
dimensional scaffold for cells to grow and differentiate
properly.
[0005] Generally, the ideal cell scaffold in tissue engineering
should be biocompatible and biodegradable, promote cellular
interaction and tissue development, and possess proper mechanical
and physical properties. The cell scaffolds are implanted in a
mechanically dynamic environment in the body; the scaffold must
sustain and recover from various deformations without mechanical
irritations to the surrounding tissues. The properties of scaffolds
should resemble those of the extracellular matrix (ECM), a soft,
tough, and elastomeric proteinaceous network that provides
mechanical stability and structural integrity to tissues and organs
[4].
[0006] Mechanical stimuli play an important role in the development
of tissues. In vascular engineering, for example, the extent to
which the initial compliance may affect the long-term function of
the graft remains controversial[5]. It has long been realized the
fibrous tissue formation within and surrounding and implanted
vascular graft would compromise graft compliance. Compliance
mismatch between the grafts and host vessel may contribute to the
development of incomplete endothelialization and myointimal
hyperplasia at the anastomotic regions. Hence, elastomeric
materials are attractive in tissue engineering especially in soft
tissue engineering such as vascular, ligament, and meniscus
engineering [6].
[0007] Current elastomers in tissue engineering can be categorized
as naturally derived materials and synthetic polymers. Naturally
derived materials such as collagen and elastin must be isolated
from human, animal or plant tissue. This process typically results
in a high cost and large batch to batch variations. These materials
also exhibit a limited range of physical properties and immune
response is always a concern [7][8]. Typical synthetic elastomeric
materials include poly(4-hydroxybutyrate) (P4HB), polyurethane
(PU), polycarpolactone (PCL), poly(glycerol-sebacate) (PGS) [4] and
so on. PHB has a much higher modulus (stiffer) and much lower sfum
to failure compared the normal soft tissues. PU has been
investigated extensively as elastomeric materials for vascular
grafts. One major concern about PU, however, is the potential
carcinogenic effect of its degradation products. A statement issued
by FDA suggested that the implanted PU foam might degrade and form
2,4-toluene diamine, which has been shown to cause liver cancer in
laboratory animals [6]. PCL is a semi-crystalline linear resorbable
aliphatic elastomeric polyester.
[0008] The Food and Drug Administration (FDA) has approved a number
of medical and drug delivery devices made by PCL. However,
applications of PCL might be limited because degradation and
resorption of PCL are considerably slow due to its hydrophobic
character and high crystallinity. The hydrophobic surface also has
impacts to the cell attachment on PCL [9]. PGS is a newly developed
elastomer which exhibits good mechanical properties and
biocompatibility. High temperature and high vacuum, however, are
needed for the polymer synthesis. [10]
SUMMARY OF THE INVENTION
[0009] The present invention is directed to a novel biocompatible
elastomeric polymer that may be used in tissue engineering. More
specifically, the specification describes methods and compositions
for making and using citric acid copolymers. In certain
embodiments, there is provided a composition comprising a citric
acid polyester having a linear aliphatic dihydroxy monomer or
linear diol; and citric acid. In specific embodiments, a linear
diol comprises between about 2 and about 20 carbons. While in
certain embodiments, all the linear aliphatic dihydroxy monomers of
a polymer the same linear diol, other embodiments contemplate
different linear diols. A particularly preferred linear diol is
1,8-octanediol. In other embodiments, the linear aliphatic
dihydroxy monomer may be 1,10-decanediol. The diol also may be an
unsaturated diol, e.g., tetradeca-2,12-diene-1,14-diol, or other
diols including macromonomer diols such as polyethylene oxide, and
Nmethyldiethanoamine (MDEA). This family of elastomers is named as
poly(diol citrate). In particularly preferred embodiments, the
composition of the invention is dihydroxy poly 1,8-octanediol
co-citric acid. Poly(diol citrate) can also form hybrids with other
materials like hydroxyapatite to form elastomeric composites.
[0010] Another aspect of the invention contemplates a substrate
that may be formulated for tissue culture and/or tissue engineering
wherein the substrate is made of a citric acid polymer as described
herein. In preferred embodiments, the substrate may further
comprise a surface modification that allows cellular attachment.
Preferably, the polymer of the invention employed as cell/tissue
culture substrate is biodegradable. Preferably, the polymer also is
biocompatible. The "biocompatible" is intended to encompass a
polymer that may be implanted in vivo or alternatively may be used
for the growth of cells that may be implanted in vivo without
producing an adverse reaction, such as an immunological response or
otherwise altering the morphology of the cells grown thereon to
render the cells incompatible with being implanted in vivo or used
to model an in vivo organ.
[0011] Also contemplated herein is a method of producing engineered
tissue, comprising providing a biodegradable citric acid polymer of
the present invention as a scaffold for the growth of cells and
culturing cells of said tissue on the scaffold. In preferred
methods, the polymer is poly 1,8-octanediol-co-citric acid, or a
derivative thereof. In specific embodiments, the cells are selected
from the group consisting of endothelial cells, ligament tissue,
muscle cells, bone cells, cartilage cells. In other preferred
embodiments, the tissue engineering method comprises growing the
cells on the scaffold in a bioreactor.
[0012] Other features and advantages of the invention will become
apparent from the following detailed description. It should be
understood, however, that the detailed description and the specific
examples, while indicating preferred embodiments of the invention,
are given by way of illustration only, because various changes and
modifications within the spirit and scope of the invention will
become apparent to those skilled in the art from this detailed
description.
BRIEF DESCRIPTION OF THE FIGURES
[0013] FIG. 1 is a schematic representation of the synthesis of
poly(1,8-octanediol-co-citric acid)
[0014] FIG. 2 is an FTIR spectrum of POC
[0015] FIG. 3 is a graph depicting stress-strain curves of POC
under different reaction conditions
[0016] FIG. 4 is a comparison of the stress-strain curves of POC,
PDC, PDDC, PDDCPEO400, POCM and POC-HA.
[0017] FIG. 5 is a graph depicting DSC thermograms of POC
[0018] FIG. 6 is a graph depicting the contact angle to water vs.
time curve of POC.
[0019] FIG. 7 is a graph depicting the degradation of POC
synthesized under different conditions after incubated in PBS at
37.degree. C. for 6 weeks.
[0020] FIG. 8 is a graph depicting weight loss in alkali solution
(0.1 M sodium hydroxide aqueous) of POC with or without 5% (monomer
mole ratio) glycerol.
[0021] FIG. 9 is a photomicrograph (.times.100) of human aortic
smooth muscle cells on POC at different culture times: A) 1 hour B)
5 hours C) 24 hours and D) 8 days.
[0022] FIG. 10 is a graph depicting the results of an
MTT-tetrazonium assay of human aortic smooth muscle cells on POC,
PLLA (Mw=300,000), and tissue culture polystyrene (TCPS). Formosan
absorbance is expressed as a function of culture time.
[0023] FIG. 11 is a photomicrograph (.times.100) (A, B, C, and D)
and SEM pictures (E and F) of human aortic endothelial cells on POC
at different culture times: A) 1 hour; B) 24 hours: C) 4 days; D) 6
days; E) and F) 6 days.
[0024] FIG. 12 is a photomicrograph (.times.100) of human aortic
smooth muscle cells (A) and human aortic endothelial cells (B) on
PDC.
[0025] FIG. 13 is a photograph depicting porous and non-porous tube
scaffold and sponge scaffold made by POC.
[0026] FIGS. 14A and 14B shows graphs depicting the results of wet
mechanical tests for POC and PDDC under different conditions. FIG.
14A shows tensile strength and FIG. 14B shows elongation.
[0027] FIG. 15 is a schematic drawing depicting a biphasic
scaffold.
[0028] FIGS. 16A-D show SEM pictures of A) a cross section of a POC
biphasic scaffold; B) the pore structure of the porous phase; C)
human aortic smooth muscles cells on the porous phase of
co-cultured biphasic scaffold; D) human aortic endothelial cells on
the lumen of co-cultured biphasic scaffold.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
A. New Biodegradable Elastomeric Polymers
[0029] Described in the present specification are a family of novel
biodegradable elastomeric polymers comprising a polyester network
of citric acid copolymerized with a linear aliphatic di-OH monomer
in which the number of carbon atoms ranges from 2 to 20. Polymer
synthesis conditions vary from mild conditions, even at low
temperature (less than 100.degree. C.) and no vacuum, to tough
conditions (high temperature and high vacuum) according the
requirements for the materials properties. By changing the
synthesis conditions (including, but not limited to,
post-polymerization temperature, time, vacuum, the initial monomer
molar ratio, and the di-OH monomer chain length) the mechanical
properties of the polymer can be modulated over a wide range. This
series of polymers exhibit a soft, tough, biodegradable,
hydrophilic properties and excellent biocompatibility in vitro.
[0030] The polymers of the present invention comprise a linear,
aliphatic diol and citric acid.
[0031] In preferred embodiments, the linear, aliphatic diol is
1,8-octanediol. However, it should be understood that this is
merely an exemplary linear, aliphatic diol. Those of skill are
aware of other aliphatic alcohols that will be useful in
polycondensation reactions to produce poly citric acid polymers.
Exemplary such aliphatic diols include any diols of between about 2
carbons and about 20 carbons. While the diols are preferably
aliphatic, linear, unsaturated diols, with the hydroxyl moiety
being present at the Cl and Cx position (where x is the terminal
carbon of the diol), it is contemplated that the diol may be an
unsaturated diol in which the aliphatic chain contains one or more
double bonds. The preferred identity for the linear, aliphatic diol
in one embodiment is 1,8, octanediol, however it may be any other
aliphatic alcohols. While in specific embodiments, the linear,
aliphatic diols of the polymer are the same diol, e.g.,
1,8-octanediol, it should be understood that linear, aliphatic
diols of the polymer may have different carbon lengths. For
example, linear, aliphatic diols of the polymer may be 2, 3, 4, 5,
6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20 or more
carbons in length. Exemplary methods for the polycondensation of
the citric acid with the linear diols are provided herein below in
the Examples.
[0032] The polymers of present invention may be utilized to form
hybrids with other materials to form elastomeric composites. In
those embodiments where the other materials are used, the other
materials can be in-organic materials, polymers with any kind of
forms such as powder, fiber, and films. The other materials can
also be elastomeric or non-elastomeric. In a particularly
embodiment, the elastomeric composite can be a hybrid of the
polymers of present invention with hydroxyapatite (POC-HA).
[0033] The polymers of the present invention may be useful both as
substrata for the growth and propagation of tissues cells that may
be seeded on the substrata and also as implantable devices. In
those embodiments where the polymers are used as bioimplantable
devices, the substrate may be formulated into a shape suitable for
implantation. For example, as described in U.S. Pat. No. 6,620,203
(incorporated herein by reference), it may be desirable to produce
prosthetic organ tissue for implantation into an animal, such as
e.g., testicular tissue described in the U.S. Pat. No. 6,620,203.
Other organs for which tissue implantation patches may be generated
include, but are not limited to skin tissue for skin grafts,
myocardial tissue, bone tissue for bone regeneration, testicular
tissue, endothelial cells, blood vessels, and any other cells from
which a tissue patch may be generated. Thus, those of skill in the
art would understand that the aforementioned organs/cells are
merely exemplary organs/cell types and it should be understood that
cells from any organ may be seeded onto the biocompatible polymers
of the invention to produce useful tissue for implantation and/or
study.
[0034] The cells that may be seeded onto the polymers of the
present invention may be derived from commercially available cell
lines, or alternatively may be primary cells, which can be isolated
from a given tissue by disaggregating an appropriate organ or
tissue which is to serve as the source of the cells being grown.
This may be readily accomplished using techniques known to those
skilled in the art. Such techniques include disaggregation through
the use of mechanically forces either alone or in combination with
digestive enzymes and/or chelating agents that weaken cell-cell
connections between neighboring cells to make it possible to
disperse the tissue into a suspension of individual cells without
appreciable cell breakage. Enzymatic dissociation can be
accomplished by mincing the tissue and treating the minced tissue
with any of a number of digestive enzymes either alone or in
combination. Digestive enzymes include but are not limited to
trypsin, chymotrypsin, collagenase, elastase, and/or hyaluronidase,
Dnase, pronase, etc. Mechanical disruption can also be accomplished
by a number of methods including, but not limited to the use of
grinders, blenders, sieves, homogenizers, pressure cells, or
sonicators to name but a few. For a review of tissue disaggregation
techniques, see Freshney, Culture of Animal Cells. A Manual of
Basic Technique, 2d Ed., A. R. Liss, Inc., New York, 1987, Ch. 9,
pp. 107-126.
[0035] Once the primary cells are disaggregated, the cells are
separated into individual cell types using techniques known to
those of skill in the art. For a review of clonal selection and
cell separation techniques, see Freshney, Culture of Animal Cells.
A Manual of Basic Techniques, 2d Ed., A. R. Liss, Inc., New York,
1987, Ch. 11 and 12, pp. 137-168. Media and buffer conditions for
growth of the cells will depend on the type of cell and such
conditions are known to those of skill in the art.
[0036] In certain embodiments, it is contemplated that the cells
attached to the biocompatible polymeric substrates of the invention
are grown in bioreactors. A bioreactor may be of any class, size or
have any one or number of desired features, depending on the
product to be achieved. Different types of bioreactors include tank
bioreactors, immobilized cell bioreactors, hollow fiber and
membrane bioreactors as well as digesters. There are three classes
of immobilized bioreactors, which allow cells to be grown: membrane
bioreactors, filter or mesh bioreactors, and carrier particle
systems. Membrane bioreactors grow the cells on or behind a
permeable membrane, allowing the nutrients to leave the cell, while
preventing the cells from escaping. Filter or mesh bioreactors grow
the cells on an open mesh of an inert material, allowing the
culture medium to flow past, while preventing the cells from
escaping. Carrier particle systems grow the cells on something very
small, such as small nylon or gelatin beads. The bioreactor can be
a fluidized bed or a solid bed. Other types of bioreactors include
pond reactors and tower fermentors. Any of these bioreactors may be
used in the present application for regenerating/engineering
tissues on the citric acid polymers of the present invention.
[0037] Certain tissues that are regenerated by use of the citric
acid polymers of the invention may be encapsulated so as to allow
the release of release of desired biological materials produced by
the cells at the site of implantation, while sequestering the
implanted cells from the surrounding site. Cell encapsulation can
be applied to all cell types secreting a bioactive substance either
naturally or through genetic engineering means. In practice, the
main work has been performed with insulin secreting tissue.
[0038] Encapsulation procedures are most commonly distinguished by
their geometrical appearance, i.e. micro- or macro-capsules.
Typically, in microencapsulation, the cells are sequestered in a
small permselective spherical container, whereas in
macroencapsulation the cells are entrapped in a larger
non-spherical membrane, Lim et al. (U.S. Pat. Nos. 4,409,331 and
4,352,883) discloses the use of microencapsulation methods to
produce biological materials generated by cells in vitro, wherein
the capsules have varying permeabilities depending upon the
biological materials of interest being produced, Wu et al, Int. J.
Pancreatology, 3:91-100 (1988), disclose the transplantation of
insulin-producing, microencapsulated pancreatic islets into
diabetic rats.
[0039] As indicated above, the cells that are seeded on the
polymers of the present invention may be cell lines or primary
cells. In certain preferred embodiments, the cells are genetically
engineered cells that have been modified to express a biologically
active or therapeutically effective protein product. Techniques for
modifying cells to produce the recombinant expression of such
protein products are well known to those of skill in the art.
Example 1
Preparation of Poly(1,8-Octanediol-co-citric acid) (POC)
[0040] In a typical experiment, 19.212 g citric acid and 14.623 g
Octanediol were added to a 250 mL three-neck round-bottom flask,
fitted with an inlet adapter and an outlet adapter. The mixture was
melted within 15 min by stirring at 160-165.degree. C. in silicon
oil bath, and then the temperature of the system was lowered to
140.degree. C. The mixture was stirred for another 1 hr at
140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time (from one day to 3
weeks depending on the temperature, with the lower temperatures
requiring longer times) to achieve the
Poly(1,8-octanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
[0041] Porous scaffolds of POC (tubular and flat sheets) were
prepared via a salt leaching technique. Briefly, sodium chloride
salt was ground up and sieved for particle sizes between 90 and 125
microns. The salt particles are then mixed with the pre-polymer
solution to the desired mass fraction to obtain a corresponding
porosity. Typically, the mass fraction of the salt particles will
result in a similar % porosity.
Example 2
Preparation of Porous Scaffolds of POC
[0042] Porous scaffolds of POC (tubular and flat sheets) were
prepared via a salt leaching technique as follows: POC pre-polymer
was dissolved into dioxane to form 25 wt % solution, and then the
sieved salt (90-120 microns) was added into pre-polymer solution to
serve as a porogen. The resulting slurry was cast into a
poly(tetrafluoroethylene) (PTFE) mold (square and tubular shape).
After solvent evaporation for 72 h, the mold was transferred into a
vacuum oven for post-polymerization. The salt in the resulting
composite was leached out by successive incubations in water
(produced by Milli-Q water purification system every 12 h for a
total 96 h. The resulting porous scaffold was air-dried for 24 hr
and then vacuum dried for another 24 hrs. The resulting scaffold
was stored in a dessicator under vacuum before use. Porous
scaffolds are typically preferred when cells are expected to
migrate through a 3-dimensional space in order to create a tissue
slice. Solid films would be used when a homogenous surface or
substrate for cell growth is required such as an endothelial cell
monolayer within the lumen of a vascular graft.
Example 3
Characterization of POC
[0043] The following Example provides details of methods and
results of characterization of POC.
Methods
[0044] Fourier Transform Infrared (FTIR) Spectroscopy
Measurements.
[0045] Infrared spectra were recorded on a Biorad FTS40 Fourier
transform infrared spectrometer. Sample POC films with thickness of
12-16 microns were prepared from POC solid samples using a
Microtome.
[0046] Mechanical Tests.
[0047] Tensile tests were conducted according to ASTM D412a on an
Instron 5544 mechanical tester equipped with 500 N load cell. The
POC sample size was 26.times.4.times.1.5 mm.
[0048] Differential Scanning Calorimetry (DSC) Measurements.
[0049] Differential scanning calorimetry thermograms were recorded
in the range of -80 to 600.degree. C. on a DSC550 (Instrument
Specialists Inc.) instrument at a heating rate of 10.degree.
C./min.
[0050] In Vitro Degradation.
[0051] The disk specimen (7 mm in diameter, about 1 to 1.5 mm
thickness) was placed in a small container containing 10 ml
phosphate buffer saline (pH 7.4). The container was incubated at
37.degree. C. for various times. After incubation the disk was
washed with water and dried under vacuum for one week. The mass
loss was calculated by comparing the initial mass (W 0) with that a
given time point (WJ), as shown in Eq. (1). Three individual
experiments were performed in triplicate for the degradation test.
The results are presented as means:1::standard deviation (n=3).
Mass loss (%)=[(Wo-W.)/Wo].times.100 (1)
[0052] Alkali Hydrolysis.
[0053] Alkali hydrolysis of the disk specimen (8.5 mm in diameter,
about 1 to 1.5 mm thickness) was conducted in a 0.1 M sodium
hydroxide aqueous solution at 37.degree. C. for various times. The
degree of degradation was estimated from the weight loss expressed
as g/m2, which was calculated by dividing the weight loss by the
total surface area of the disk.
[0054] Cell Culture.
[0055] Human aortic smooth muscle and endothelial cells (Clonetics)
were cultured in a 50 ml culture flask with SmGM-2 and EBM-2
culture medium (Clonetics). Cell culture was maintained in a
water-jacket incubator equilibrated with 5% CO2 at 37.degree. C.
When the cells had grown to confluence, the cells were passaged
using a Subculture Reagent Kit (Clonetics). Polymer films were cut
into small pieces (1.times.2 cml and placed in cell culture dishes
(6 cm in diameter). Polymer films were sterilized in 70% ethanol
and the ethanol was exchanged with an excess amount of
phosphate-buffered saline (PBS). The PBS was removed with a pipette
and then the samples were sterilized UV light for another 30 min. A
5 ml cell suspension with 6.6.times.10.sup.4/ml was added to the
culture dish. The morphology of cell attachment was observed and
photographed with an inverted light microscope (Nikon Eclipse,
TE2000-U) equipped with a Photometrics Coo1SNAP HQ after culturing
a predetermined time. After reaching confluence, the samples were
fixed by 2.5% glutaraldehyde solution and dehydrated sequentially
in 50, 70, 95 and 100% ethanol each for 10 min. The fixed samples
were lyophilized, sputter-coated with gold and examined under
scanning electron microscope (SEM, Hitachi 3500N). Polymer films
were cut into small disks (7 mm in diameter) with the aid of a cork
borer in order to locate the disks into a 96-well tissue culture
plate. PLLA films and Tissue culture polystyrene (TCPS) were used
as control. The samples were sterilized as described above. The
human aortic smooth muscle cells (3.13.times.10.sup.3/well) were
added to the wells. The viability and proliferation of the cells
were determined by MTT assays. The absorbance of produced Formosan
was measured at 570 nm using microplate reader (Tecan,
SAFFIRE).
Results
[0056] Polycondensation of citric acid and 1,8-octanediol yields a
transparent film. The resulting polymer features a small number of
crosslinks and carboxyl and hydroxyl groups directly attached to
the polymer backbone (FIG. 1).
[0057] The typical FTIR spectrum of a POC preparation is shown in
FIG. 2. The intense C.dbd.O stretch at 1,735 cm-1 in FTIR spectrum
confirms the formation of ester bonds. The intense OH stretch at
3,464 cm-1 indicates that the hydroxyl groups are hydrogen
bonded.
[0058] Tensile tests on strips of POC prepared under different
synthetic conditions reveal a stress-strain curve characteristic of
an elastomeric and tough material (FIG. 3). The nonlinear shape of
the tensile stress-strain curve, low modulus and large elongation
ratio is typical for elastomers and resembles those of ligament and
vulcanized rubber [4]. These results further demonstrate that when
the post-polymerization reaction is carried out under lower
temperature, the resulting polymer is more elastic than when it is
performed at higher temperatures. Post polymerization at lower
temperature under vacuum (i.e. 40.degree. C.)) may enable
incorporation of biological molecules within POC without
significant loss of biological activity. Tissue engineering
applications that require significant elasticity and strength such
as for vascular grafts and heart valves may benefit from
post-polymerization at the lower conditions. Tissue engineering
applications that require a more rigid or stiff scaffold such as
cartilage tissue engineering would benefit from post-polymerization
at the higher temperatures.
[0059] The thermal properties of POC were investigated by DSC. From
the thermograms depicted in FIG. 5, no crystallization temperature
and melting temperature are observed and apparent glass transition
temperature (Tg) is observed below 0.degree. C. for POC synthesized
under a variety of conditions. This result shows POC is totally
amorphous at 37.degree. C. similar to the vulcanized rubber. FIG. 5
shows the Tg changes with the synthesis conditions. Increasing
post-polymerization temperature and elongating the treating time
can increase the crosslinking density and then result in the
increase of Tg. The Tg is still significantly below 37.degree. C.,
making the material elastomeric for tissue engineering applications
that require elastomeric scaffolds (i.e. cardiovascular, pulmonary,
ligament tissue engineering). This result also confirms that POC is
a cross-linked polymer. Similar results were observed with PDC.
[0060] FIG. 6 shows the contact angle to water vs. time curve of
POC. The initial contact angle of the POC synthesized under
different conditions is 76.degree. and 84.degree., respectively.
The water drop spread out with the time. The contact angles finally
reach 38.degree. and 44.degree., respectively. Although the initial
contact angle is relatively high, the polymer chains are highly
mobile since POC is a rubber-like and amorphous polymer at room
temperature, and the polar water molecules can induce the polar
groups such as hydroxyl and carboxyl to enrich at the polymer
surface via surface rearrangement. The results show POC is a
hydrophilic polymer. Hydrophilic polymers are expected to promote
endothelial cell adhesion and proliferation as presented in
preliminary data.
[0061] FIG. 7 shows the degradation of POC synthesized under
different conditions after incubation in PBS at 37.degree. C. for 6
weeks. POC synthesized under mild conditions (A) has a faster
degradation rate compared to that of POC synthesized under
relatively tougher conditions (B and C). The degradation rate of
POC (B) is considerably faster than that of POC (C). POC
synthesized under tough conditions features a high cross-linking
degree and the penetration of water molecules into the network
films is difficult because of the smaller network space. This is
the reason why the degradation rate sequence is POC (A)>POC
(B)>POC (C). These results show that POC is degradable polymer.
The degradation rate can be modulated by changing synthesis
conditions.
[0062] In order to achieve better control for the degradation of
"highly cross-linked" POC, a third monomer, glycerol is added in
addition to the citric acid and diol monomer (0-3 mol %, the molar
ratio of carboxyl and hydroxyl group among the three monomers was
maintained as 1/1). Increasing amounts of glycerol will result in
an increased break strength and Young's modulus. The alkali
hydrolysis results show that the addition of glycerol can enhance
the degradation of POC in alkali solution. Glycerol is a
hydrophilic component. Its addition can facilitate the water
penetration into the network films which results in the faster
degradation rate.
[0063] The in vitro biocompatibility of POC was evaluated in order
to investigate the potential application in tissue engineering,
especially for soft tissue engineering such as vascular graft,
ligament, bladder, and cartilage. Human smooth muscle cells and
endothelial cells are chosen as model cells. FIGS. 9 and 11 show
the morphology of both cell types on POC films at different culture
times. The results indicate that POC is a good substrate for
supporting the both cells attachment. Both cells grow promptly and
achieve confluence on POC.
[0064] Cell attachment and growth are also observed on PDC (FIG.
12). MTT assays (an indicator of cell viability) also indicate that
POC is a better substrate for cell growth than PLLA (FIG. 10).
Synthetic materials have attracted many interests as small diameter
grafts. Normally, the synthetic grafts have not produced acceptable
results because of rapid thrombotic buildup in the vessel
lumen[13]. Researchers have been attempting to improve graft
performance by adding an endothelial lining and thus better
mimicking the vessels in the body [14,15]. Failure of grafts was
associated with subintimal hyperplasia and a thrombotic surface,
possibly resulting in part from lack of a confluent layer of
endothelial cells on the graft lumen. Many methods have been
developed for improving the endothelial cell attachment and growth
such as immobilizing cell adhesion peptides (GREDVY) on polymer
surfaces [16], plasma modification using radio frequency glow
discharge [17] and so on. Endothelial cells adherence can be
dramatically increased when the grafts are coated with
extracellular matrix, plasma or fibronectin. Unfortunately for
graft compatibility, coating with fibronectin increases not only
the adhesion of endothelial cells to those surfaces, but of
platelets as well [18]. Optimal adherence has been reported for gas
plasma-treated surfaces with hydrophilicity in the range of
40-60.degree. by Dekker [19] and van Wachem [20]. This effect was
attributed to specific protein adsorption favorable for adhesion,
spreading, and proliferation of endothelial cells, and improved
deposition of endothelial matrix proteins. For POC, the
hydrophilicity is in the above range, which may help the adsorption
of glycoproteins on the polymer surface. The surface-enriched polar
groups such as carboxyl and hydroxyl may facilitate the cell
attachment and growth [21,22]. No additional pre-treatments are
needed and the endothelial cells confluence on POC films can be
achieved in a short time.
Example 4
Synthesis of Poly(1,6-hexanediol-co-citric acid) (PHC)
[0065] In a typical experiment, 19.212 g citric acid and 11.817 g
1,6-hexanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in a silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for a predetermined time from one day to 3
weeks, depending on the temperature, to achieve the
Poly(1,6-hexanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
Example 5
Synthesis of Poly(1,10-decanediol-co-citric acid) (PDC)
[0066] In a typical experiment, 19.212 g citric acid and 17.428 g
1,10-decanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,10-decanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
Example 6
Synthesis of Poly(1,12-dodecanediol-co-citric acid) PDDC
[0067] In a typical experiment, 19.212 g citric acid and 20.234 g
1,12-dodecanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,12-dodecanediol-co-citric acid). Nitrogen was introduced
into the reaction system before the polymer was taken out from
reaction system.
Example 7
Synthesis of Poly(1,8-octanediol-co-citric acid-co-glycerol)
[0068] In a typical experiment (Poly(1,8-octanediol-co-citric
acid-co-1% glycerol), 23.0544 g citric acid, 16.5154 g
1,8-octanediol and 0.2167 g glycerol were added to a 250 ml
three-neck round-bottom flask, fitted with an inlet adapter and an
outlet adapter. The mixture was melted within 15 min by stirring at
160-165.degree. C. in silicon oil bath, and then the temperature of
the system was lowered to 120.degree. C. The mixture was stirred
for another hour at 140.degree. C. to get the pre-polymer. Nitrogen
was vented throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,8-octanediol-co-citric acid-co-1% glycerol). Nitrogen was
introduced into the reaction system before the polymer was taken
out from reaction system.
Example 8
Synthesis of Poly(1,8-octanediol-citric acid-co-polyethylene
oxide)
[0069] In a typical experiment, 38.424 g citric acid, 14.623 g
1,8-octanediol and 40 g polyethylene oxide with molecular weight
400 (PEO400)(100 g PE01000 and 200 g PEO2000 respectively) (molar
ratio: citric acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 135.degree. C.
The mixture was stirred for 2 hours at 135.degree. C. to get the
pre-polymer. Nitrogen was vented throughout the above procedures.
The pre-polymer was post-polymerized at 120.degree. C. under vacuum
for predetermined time from one day to 3 days to achieve the
Poly(1,8-octanediol-citric acid-co-polyethylene oxide). Nitrogen
was introduced into the reaction system before the polymer was
taken out from reaction system. The molar ratios can be altered to
achieve a series of polymers with different properties.
Example 9
Synthesis of Poly(1,12-dodecanediol-citric acid-co-polyethylene
oxide)
[0070] In a typical experiment, 38.424 g citric acid, 20.234 g
1,12-dodecanediol and 40 g polyethylene oxide with molecular weight
400 (PE0400) (100 g PE01000 and 200 g PE02000 respectively) (molar
ratio: citric acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 120.degree. C.
The mixture was stirred for half an hour at 120.degree. C. to get
the pre-polymer. Nitrogen was vented throughout the above
procedures. The pre-polymer was post-polymerized at 120.degree. C.
under vacuum for predetermined time from one day to 3 days to
achieve the Poly(1,12-dodecanediol-citric acid-co-polyethylene
oxide). Nitrogen was introduced into the reaction system before the
polymer was taken out from reaction system. The molar ratios can be
altered to achieve a series of polymers with different
properties.
Example 10
Synthesis of Poly(1,8-octanediol-citric
acid-co-N-methyldiethanoamine) POCM
[0071] In a typical experiment, 38.424 g citric acid, 26.321 g
1,8-octanediol and 2.3832 g N-methyldiethanoamine (MDEA) (molar
ratio: citric acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 120.degree. C.
The mixture was stirred for half an hour at 120.degree. C. to get
the pre-polymer. Nitrogen was vented throughout the above
procedures. The pre-polymer was post-polymerized at 80.degree. C.
for 6 hours, 120.degree. C. for 4 hours without vacuum and then
120.degree. C. for 14 hours under vacuum to achieve the
Poly(1,8-octanediol-citric acid-co-N-methyldiethanoamine). Nitrogen
was introduced into the reaction system before the polymer was
taken out from reaction system. The molar ratios can be altered to
citric acid/1,8-octanediol/MDEA=1/0.95/0.05.
Example 11
Synthesis of Poly(1,12-dodecanediol-citric
acid-co-N-methyldiethanoamine) PDDCM
[0072] In a typical experiment, 38.424 g citric acid, 36.421 g
1,12-dodecanediol and 2.3832 g N-methyldiethanoamine (MDEA) (molar
ratio: citric acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in a silicon oil bath, and
then the temperature of the system was lowered to 120.degree. C.
The mixture was stirred for half an hour at 120.degree. C. to get
the pre-polymer. Nitrogen was vented throughout the above
procedures. The pre-polymer was post-polymerized at 80.degree. C.
for 6 hours, 120.degree. C. for 4 hours without vacuum and then
120.degree. C. for 14 hours under vacuum to achieve the
Poly(1,12-dodecanediol-citric acid-co-N-methyldiethanoamine).
Nitrogen was introduced into the reaction system before the polymer
was taken out from reaction system. The molar ratios can be altered
to citric acid/1,12-dodecanediol/MDEA=1/0.95/0.05.
Example 12
Calcium Modification of Different Polymers
[0073] In a typical experiment, POC and PDDC films or scaffolds
were immersed in a 0.1M CaCl.sub.2 solution for 1 week, rinsed in
mini-Q water and then freeze-dried. The dry samples were stored in
a desiccator before use. In order to evaluate calcium modification
on mechanical properties of POC and PDDC, POC and PDDC was tested
under different treating conditions.
[0074] Mechanical test results show that calcium under wet
conditions, calcium treatment (1 week) may help to maintain
appropriate tensile stress for POC compared to PBS 1 week of
treatment with phosphate buffered saline (PBS). Calcium treatment
has dramatic effects on elongation for POC. After 1 week of calcium
treatment, POC can maintain a similar elongation rate compared to 1
week of PBS (phosphate buffer solution) 1 week treatment. Even
after 1 more week of PBS treatment following 1 week of calcium
treatment, the height elongation rate of POC can be still
maintained. Since PDDC is more hydrophobic than POC, the effects of
calcium treatment on tensile stress and elongation of PDDC is less
than that on POC. This results show that calcium ions chelated by
unreacted carboxyl group of POC and PDDC synthesized under mild
condition (80.degree. C. 2 days) act as crosslinkers to help to
maintain the elasticity and appropriate strength of polymers (FIG.
14).
Example 13
Synthesis of POC-Hydroxyapatite (HA) Composite
[0075] In a typical experiment, 19.212 g citric acid and 14.623 g
Octanediol (molar ratio 1:1) of Citric Acid and 1,8-octanediol was
reacted in a 250 ml three-neck round-bottom flask at 165.degree.
C., forming a pre-polymer solution. Specific amount of HA was then
added to reaction vessel while stirring with a mechanical stirrer.
Acetone was then added until solution liquefied into a slurry
state. Solution was then cast into a Teflon mold and set in a
vacuum oven at 120.degree. C. for 2-4 hours or until the acetone
was purged. Film was then incubated without vacuum at 120.degree.
C. overnight allowing the solution to set. Film could then be post
polymerized for various durations depending upon the desired
properties. Mechanical tests on POC-HA (40 wt %) shows tensile
stress is as high as 10.13.+-.0.57 MPa and elongation is
47.78.+-.3.00. Specimen recovery completely after pulling by
mechanical tester.
Example 14
Comparison of the Properties of Different Polymers
[0076] FIG. 4 shows that the mechanical properties of the polymer
can be modulated by choosing different diol monomers. The maximum
elongation ratio for the polymer at break can reach 265:!:10.5%
similar to that of arteries and vein (up to 260%) [10]. The minimum
tensile Young's modulus can reach 1.4:!:0.2 MPa. The Young's
modulus is between those of ligament (KPa scale) [11] and tendon
(GPa scale) [12].
[0077] Similar to the vulcanized rubber, POC, PDC, and PDDC are
thermoset elastomers. In general, thermoset polymers can not be
dissolved in common solvents which adds to the difficulty in making
the polymer into a scaffold for tissue engineering applications.
The present application describes a method to fabricate porous and
non-porous scaffolds which makes it possible to be used in tissue
engineering utilizing the solubility of the pre-polymer in some
solvent such as dioxane, acetone, 1,3-dioxlane, ethanol,
N,N-dimethylformamide. Therefore, this family of polymers is a
potential elastomer in tissue engineering especially in soft tissue
engineering.
Example 15
Further Characterization of Solid Polymeric Materials
[0078] This example is directed to the extent of cross-linking of
the polymeric materials. Current methods to determine the molecular
weight of a polymer include osmotic pressure, light scattering,
ultracentrifugation, solution viscosity, and gel permeation
chromatography measurements. All of these methods normally require
a polymer that can be dissolved in specific solvents..sup.[24]
Crosslinked polymers can not be dissolved in a solvent and their
molecular weight is considered to be infinite. However, a useful
parameter to characterize cross-linked polymers is molecular weight
between cross-links (Mc), which can give a measure of the degree of
cross-linking and therefore some insight into mechanical
properties. According the theory of rubber elasticity, molecular
weight between crosslinks can be calculated using Equation (1)
under some assumptions:.sup.[25]
n = E 0 3 RT = .rho. M c ##EQU00001##
where n represents the number of active network chain segments per
unit volume; Mc represents the molecular weight between cross-links
(mol/m.sup.3); E.sub.0 represents Young's modulus (Pa); R is the
universal gas constant (8.3144 Jmol-1K-1); T is the absolute
temperature (K); .rho. is the elastomer density (g/m.sup.3) as
measured via volume method..sup.[26] From Equation (1), molecular
weight between crosslinks can only be obtained after mechanical
tests and polymer density measurements. Another method for
determining molecular weight between crosslinks for a crosslinked
polymer is by swelling the polymer..sup.[27] Using the swelling
method, molecular weight between crosslinks can be calculated by
Equation (2).
1 M c = 2 M n - v V 1 [ ln ( 1 - v 2 , s ) + v 2 , s + .chi. 1 v 2
, s 2 ] v 2 , s 1 / 3 - v 2 , s 2 ##EQU00002##
where Mc is the number average molecular weight of the linear
polymer chain between cross-links, .upsilon. is the specific volume
of the polymer, V.sub.1 is the molar volume of the swelling agent
and .chi..sub.1 is the Flory-Huggins polymer-solvent interaction
parameter. .upsilon..sub.2s is the equilibrium polymer volume
fraction which can be calculated from a series of weight
measurements.
Example 16
Novel Biphasic Scaffold Design for Blood Vessel Tissue
Engineering
[0079] Biphasic scaffolds consist of outside porous phase and
inside non-porous phase as depicted in the schematic drawing shown
in FIG. 15. The non-porous phase is expected to provide a
continuous surface for EC adhesion and spreading, mechanical
strength, and elasticity to the scaffold. The porous phase will
facilitate the 3-D growth of smooth muscle cells. Biphasic
scaffolds were fabricated via following procedures. Briefly, glass
rods (.about.3 mm diameter) were coated with the pre-polymer
solution and air dried to allow for solvent evaporation. Wall
thickness of the tubes can be controlled by the number of coatings
and the percent pre-polymer in the solution. The pre-coated
pre-polymer was partially post-polymerized under 60.degree. C. for
24 hr; the pre-polymer-coated glass rod is then inserted
concentrically in a tubular mold that contains a salt/pre-polymer
slurry. The pre-polymer/outer-mold/glass rod system is then placed
in an oven for further post-polymerization. After salt-leaching
[4], the biphasic scaffold was then de-molded from the glass rod
and freeze dried. The resulting biphasic scaffold was stored in a
desiccator before use. The same materials or different materials
from the above family of elastomers can be utilized for both phases
of the scaffold. Other biomedical materials widely used in current
research and clinical application such as polylactide (PLA),
polycaprolactone (PCL), poly(lactide-co-glycolide) (PLGA) may also
be utilized for this novel scaffold design.
[0080] The thickness, degradation, and mechanical properties of
inside non-porous phase can be well controlled by choosing various
pre-polymers of this family of elastomers, pre-polymer
concentration, coating times and post-polymerization conditions
(burst pressure can be as high as 2800 mmHg). The degradable porous
phase and non-porous phases are integrated since they are formed
in-situ via post-polymerization. The cell culture experiments shown
in FIG. 16 confirm that both HAEC and HASMC can attach and grow
well in biphasic scaffolds. The results suggest that a biphasic
scaffold design based on poly(diol-co-citrate) is a viable strategy
towards the engineering of small diameter blood vessels.
Example 17
Materials and Methods Employed for Polymer Characterization
[0081] In addition to the materials and methods described above,
the following materials and methods also are exemplary of the
studies performed herein.
Polymer Synthesis
[0082] Preparation of poly(1,8-Octanediol-co-citric acid) (POC)
films: [23] All chemicals were purchased from Sigma-Aldrich
(Milwaukee, Wis.). Equimolar amounts of citric acid and
1,8-octanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet and outlet adapter. The mixture was
melted under a flow of nitrogen gas by stirring at 160.degree.
C.-165.degree. C. in a silicon oil bath, and then the temperature
of the system was lowered to 140.degree. C. The mixture was stirred
for another hour at 140.degree. C. to create the pre-polymer
solution. The pre-polymer was cast in glass dishes and
post-polymerized at 80.degree. C., 120.degree. C. or 140.degree. C.
under vacuum (2 Pa) or no vacuum for times ranging from 1 day to 2
weeks to create POC films with various degrees of cross-linking
Mechanical Tests
[0083] Tensile tests were conducted according to ASTM D412a on an
Instron 5544 mechanical tester equipped with 500 N load cell
(Instron Canton, Mass.). Briefly, a dog-bone-shaped sample
(26.times.4.times.1.5 mm, Length.times.Width.times.Thickness) was
pulled at a rate of 500 mm/min. Values were converted to
stress-strain and a Young's modulus was calculated. 4-6 samples
were measured and averaged.
Molecular Weight Between Crosslinks Measurements
[0084] The molecular weight between crosslinks of POC was
calculated using Equation (1).
Swelling Studies
[0085] Polymers were cut into rectangular strip and the initial
length, width and thickness measured with calipers. The polymers
were then swollen in DMSO at 37.degree. C. overnight to achieve
equilibrium swelling. The equilibrium length, width, and thickness
were measured to determine the change in volume upon swelling.
Results and Discussion
Mechanical Tests and Molecular Weight Measurements of POC
[0086] POC samples of various degrees of cross-linking were
synthesized by reacting the polyfunctional monomer citric acid with
the difunctional monomer 1,8-octanediol under different
post-polymerization conditions and the resulting polymer films were
subjected to mechanical tensile tests and molecular weight between
crosslinks measurements..sup.[25] The results in Table 1 indicate
that increased crosslinking temperatures and time increase the
tensile stress, Young's modulus and the number of active network
chain segment per unit volume (crosslinking density) while
decreasing the molecular weight between crosslinks. Therefore, the
mechanical properties of POC can be well controlled by controlling
polymer network structures via post-polymerization under different
conditions.
TABLE-US-00001 TABLE 1 Mechanical properties, the number of active
network chain segment per unit volume (crosslinking density): n)
and molecular weight between crosslinks (Mc) of POC synthesized
under different conditions Polymerization Young's Tensile n Mc POC
condition Modulus (MPa) Stress (MPa) (mol/m.sup.3) (g/mol) LS1
80.degree. C., no vacuum, 2 days 1.38 .+-. 0.21 1.64 .+-. 0.05
182.59 .+-. 27.78 6874 .+-. 148 LS2 80.degree. C., high vacuum, 2
days 1.72 .+-. 0.45 1.90 .+-. 0.22 227.58 .+-. 59.54 5445 .+-. 116
LS3 120.degree. C., high vacuum, 1 day 2.84 .+-. 0.12 3.62 .+-.
0.32 375.77 .+-. 15.88 3301 .+-. 218 LS4 120.degree. C., high
vacuum, 2 days 3.13 .+-. 0.27 3.66 .+-. 0.61 414.14 .+-. 35.72 2971
.+-. 76 LS5 120.degree. C., high vacuum, 3 days 4.69 .+-. 0.48 5.34
.+-. 0.66 620.68 .+-. 63.51 1857 .+-. 81 LS6 140.degree. C., high
vacuum, 2 days 6.07 .+-. 0.52 5.73 .+-. 1.39 803.14 .+-. 68.80 1516
.+-. 269 LS7 80.degree. C., no vacuum, 5 days 2.21 .+-. 0.17 3.90
.+-. 0.60 292.41 .+-. 22.49 4326 .+-. 68 LS8 80.degree. C., no
vacuum, 14 days 2.24 .+-. 0.09 2.55 .+-. 0.21 296.38 .+-. 11.91
4265 .+-. 33
* * * * *