U.S. patent application number 11/704039 was filed with the patent office on 2007-09-06 for functionalizing implantable devices with a poly (diol co-citrate) polymer.
This patent application is currently assigned to NORTHWESTERN UNIVERSITY. Invention is credited to Guillermo Ameer, Jian Yang.
Application Number | 20070208420 11/704039 |
Document ID | / |
Family ID | 38345814 |
Filed Date | 2007-09-06 |
United States Patent
Application |
20070208420 |
Kind Code |
A1 |
Ameer; Guillermo ; et
al. |
September 6, 2007 |
Functionalizing implantable devices with a poly (diol co-citrate)
polymer
Abstract
The present invention is directed to a novel poly(diol
citrates)-based coating for implantable devices. More specifically,
the specification describes methods and compositions for making and
using implantable devices coated with citric acid copolymers or
citric acid copolymers impregnated with therapeutic compositions
and/or cells.
Inventors: |
Ameer; Guillermo; (Chicago,
IL) ; Yang; Jian; (Arlington, TX) |
Correspondence
Address: |
MARSHALL, GERSTEIN & BORUN LLP
233 S. WACKER DRIVE, SUITE 6300
SEARS TOWER
CHICAGO
IL
60606
US
|
Assignee: |
NORTHWESTERN UNIVERSITY
Evanston
IL
|
Family ID: |
38345814 |
Appl. No.: |
11/704039 |
Filed: |
February 8, 2007 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60771348 |
Feb 8, 2006 |
|
|
|
Current U.S.
Class: |
623/1.41 ;
427/2.25; 427/240; 435/402; 623/1.46 |
Current CPC
Class: |
A61P 7/02 20180101; C12N
2533/40 20130101; A61P 31/04 20180101; C12N 5/0692 20130101; Y10T
428/31736 20150401; A61L 27/34 20130101; A61L 27/34 20130101; A61L
27/58 20130101; A61L 27/3804 20130101; A61L 27/507 20130101; A61L
27/48 20130101; Y10T 428/31663 20150401; A61L 27/16 20130101; Y10T
428/31565 20150401; Y10T 428/31681 20150401; Y10T 428/3154
20150401; A61L 33/068 20130101; Y10T 428/31786 20150401; Y10T
428/3179 20150401; Y10T 428/31797 20150401; Y10T 428/31507
20150401; A61L 27/16 20130101; A61L 33/068 20130101; A61L 31/10
20130101; C08L 27/18 20130101; C08L 67/04 20130101; C08L 67/04
20130101 |
Class at
Publication: |
623/001.41 ;
623/001.46; 435/402; 427/002.25; 427/240 |
International
Class: |
A61F 2/06 20060101
A61F002/06; A61F 2/02 20060101 A61F002/02; C12N 5/08 20060101
C12N005/08; A61L 33/00 20060101 A61L033/00; B05D 3/00 20060101
B05D003/00 |
Claims
1. An implantable medical device wherein at least one surface of
said has deposited thereon a coating comprising a citric acid
polyester having the generic formula (A-B-C)n, wherein A is a
linear aliphatic dihydroxy monomer; B is citric acid, C is a linear
aliphatic dihydroxy monomer, and n is an integer greater than
1.
2. The implantable medical device of claim 1, wherein A is a linear
diol comprising between about 2 and about 20 carbons.
3. The implantable medical device claim 1, wherein C is a linear
diol comprising between about 2 and about 20 carbons.
4. The implantable medical device claim 1, wherein both A and C are
the same linear diol.
5. The implantable medical device of claim 4, wherein said linear
diol is 1, 8, octanediol.
6. The implantable medical device of claim 1, wherein A and C are
different linear diols.
7. The implantable medical device of claim 5, wherein said linear
aliphatic dihydroxy poly 1,8-octanediol co-citric acid.
8. The implantable medical device of claim 5, wherein said linear
aliphatic dihydroxy poly 1,10-decanediol co-citric acid.
9. The implantable medical device wherein said device is made of a
base material that comprises one or more materials selected from
the group consisting of: stainless steel, tantalum, titanium,
nitinol, gold, platinum, inconel, iridium, silver, tungsten, a
biocompatible metal, carbon, carbon fiber, cellulose acetate,
cellulose nitrate, silicone, polyethylene terephthalate,
polyurethane, polyamide, polyester, polyorthoester, polyanhydride,
polyether sulfone, polycarbonate, polypropylene, high molecular
weight polyethylene, polytetrafluoroethylene, a biocompatible
polymeric material, polylactic acid, polyglycolic acid, a
polyanhydride, polycaprolactone, polyhydroxybutyrate valerate, a
biodegradable polymer, a protein, an extracellular matrix
component, collagen, fibrin, a biologic agent, PEBAX, polyethylene,
irradiated polyethylene or a suitable mixture, copolymer, or alloy
of any of these.
10. The implantable medical device of claim 1, wherein said base
material is ePTFE.
11. The implantable medical device of claim 10, wherein said
medical device has improved properties for use in blood
vessels.
12. The implantable medical device of claim 11, wherein said device
is less thrombogenic than a similar device prepared from ePTFE that
lacks said coating.
13. The implantable medical device of claim 11, wherein said device
has increased properties of endothelialization and cell adherence
as compared to a similar device prepared from ePTFE that lacks said
coating.
15. The implantable medical device of claim 1, wherein said coating
comprises a bioactive agent wherein the bioactive agent is selected
from the group consisting of: an antisense nucleotide, a thrombin
inhibitor, an antithrombogenic agent, a tissue plasminogen
activator, a thrombolytic agent, a fibrinolytic agent, a vasospasm
inhibitor, a calcium channel blocker, a nitrate, a nitric oxide
promoter, a vasodilator, an antimicrobial agent, an antibiotic, an
antiplatelet agent, an antimitotic, a microtubule inhibitor, an
actin inhibitor, a remodeling inhibitor, an agent for molecular
genetic intervention, a cell cycle inhibitor, an inhibitor of the
surface glycoprotein receptor, an antimetabolite, an
antiproliferative agent, an anti-cancer chemotherapeutic agent, an
anti-inflammatory steroid, an immunosuppressive agent, an
antibiotic, a radiotherapeutic agent, iodine-containing compounds,
barium-containing compounds, a heavy metal functioning as a
radiopaque agent, a peptide, a protein, an enzyme, an extracellular
matrix component, a cellular component, a biologic agent, an
angiotensin converting enzyme (ACE) inhibitor, ascorbic acid, a
free radical scavenger, an iron chelator, an antioxidant, a
radiolabelled form or other radiolabelled form of any of the
foregoing, or a mixture of any of these.
16. The implantable medical device of claim 1, wherein said coating
comprises cells are selected from the group consisting of
endothelial cells, ligament tissue, muscle cells, bone cells,
cartilage cells.
17. The implantable medical device of claim 1, wherein said coating
comprises endothelial cells.
18. The implantable medical device of claim 1, wherein said coating
comprises smooth muscle cells.
19. The implantable medical device of claim 1, wherein said coating
further comprises a second polymer selected from the group
consisting of poly(hydroxyvalerate), poly(lactide-co-glycolide),
poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate),
polyorthoester, polyanhydride, poly(glycolic acid),
poly(glycolide), poly(L-lactic acid), poly(L-lactide),
poly(D,L-lactic acid), poly(D,L-lactide), poly(caprolactone),
poly(trimethylene carbonate), and polyester amide
20. A method of producing engineered tissue, comprising: a.
preparing an implantable medical device of any of claims 1-19; and
b. culturing cells of said tissue on the scaffold of part (a).
21. A method of decreasing the thrombogenicity of a graft
comprising coating said graft with a coating comprising a citric
acid polyester having the generic formula (A-B-C)n, wherein A is a
linear aliphatic dihydroxy monomer; B is citric acid, C is a linear
aliphatic dihydroxy monomer, and n is an integer greater than 1
wherein said graft is less thrombogenic than a similar graft that
does not comprise said coating.
22. A method of preparing a small diameter blood vessel graft from
ePTFE, comprising coating said graft with a coating comprising a
citric acid polyester having the generic formula (A-B-C)n, wherein
A is a linear aliphatic dihydroxy monomer; B is citric acid, C is a
linear aliphatic dihydroxy monomer, and n is an integer greater
than 1 wherein said graft is less thrombogenic than a similar graft
that does not comprise said coating.
23. The method of claim 22, wherein said coating is deposited on
the inner lumen surface of the small diameter blood vessel graft,
on the outer surface of the small blood vessel graft or on both the
inner lumen surface and the outer surface of the small diameter
blood vessel graft.
24. The method of claim 22, wherein said graft further comprises a
layer of endothelial cells deposited on the surface of said
coating.
25. The method of any of claims 22-24, wherein said graft further
comprises an addition therapeutic agent impregnated into said
coating.
26. The method of claim 25, wherein said therapeutic agent is
selected from the group consisting of an antisense nucleotide, a
thrombin inhibitor, an antithrombogenic agent, a tissue plasminogen
activator, a thrombolytic agent, a fibrinolytic agent, a vasospasm
inhibitor, a calcium channel blocker, a nitrate, a nitric oxide
promoter, a vasodilator, an antimicrobial agent, an antibiotic, an
antiplatelet agent, an antimitotic, a microtubule inhibitor, an
actin inhibitor, a remodeling inhibitor, an agent for molecular
genetic intervention, a cell cycle inhibitor, an inhibitor of the
surface glycoprotein receptor, an antimetabolite, an
antiproliferative agent, an anti-cancer chemotherapeutic agent, an
anti-inflammatory steroid, an immunosuppressive agent, an
antibiotic, a radiotherapeutic agent, iodine-containing compounds,
barium-containing compounds, a heavy metal functioning as a
radiopaque agent, a peptide, a protein, an enzyme, an extracellular
matrix component, a cellular component, a biologic agent, an
angiotensin converting enzyme (ACE) inhibitor, ascorbic acid, a
free radical scavenger, an iron chelator, an antioxidant, a
radiolabelled form or other radiolabelled form of any of the
foregoing, or a mixture of any of these.
27. The method of claim 25, wherein said therapeutic agent is an
anti-inflammatory agent.
28. The method of claim 22 wherein said small diameter blood vessel
has an improved property selected from the group consisting of
blood flow, vessel tone, platelet activation, adhesion and
aggregation, leukocyte adhesion, SMC migration and SMC
proliferation as compared to a small diameter blood vessel prepared
from a graft that does not comprise said coating.
29. A method of preparing a small diameter blood vessel or similar
graft, from ePTFE comprising modifying the lumen of said small
diameter blood vessel by coating the small diameter blood vessel
through a spin-shearing method wherein said method comprises: a.
coating a glass-rod with said citric acid polymer by use of a
mechanical stirrer; b. rotating said coated glass rod at a low
speed of about 300 rpm; c. contacting said rotating glass rod with
ePTFE or other material use to form the small? d. shearing the
lumen of the graft by rotating the graft in a direction counter to
the direction in which it is spinning in step (a); e. optionally,
repeating steps a to d.
Description
[0001] The present application claims benefit of U.S. Provisional
Application No. 60/771,348 filed Feb. 8, 2006. The entire text of
the aforementioned application is incorporated herein by
reference
FIELD OF THE INVENTION
[0002] The present invention describes methods and compositions for
coating implantable devices with a polymer to improve the long-term
biocompatibility and/or patency of the device.
BACKGROUND
[0003] Implantable medical devices are used in the treatment and
assessment of a variety of medical conditions. Such devices may be
introduced into the body for a short period of time or may be
placed therein permanently and have been used for the treatment of
diseased, injured, or deformed body vessels. In cases where
malfunctioning body vessels have reduced inner diameters, there is
usually reduced flow of vital fluid or gas through the vessels and
in extreme cases, the vessels often are occluded. Implantable
medical devices have proven useful to open and/or expand, or to
otherwise treat such obstructed or constricted vessels. These
devices often are placed inside the vessel for a period of time and
serve to mechanically support the inside of the malfunctioning
vessel to keep the vessel open or patent.
[0004] The coated implantable medical device may be partly or
completely placed into the esophagus, trachea, colon, biliary
tract, urinary tract, vascular system or other location within a
human or veterinary patient. Many treatments of the vascular or
other systems involve the insertion of a stent, a catheter, a
balloon, a wire guide, a cannula or the like into such a location
within a human or veterinary patient. A stent is most simply
considered as a cylinder of relatively short length which opens a
body passage or lumen or which maintains a body passage or lumen in
an open condition. In addition, balloons such as angioplasty or
dilation balloons are expanded to open a body passage or vessel
lumen, thereby causing potential trauma or injury to the expanded
passage or vessel.
[0005] While implantable medical devices have gained widespread
use, they do have attendant drawback. For example, introduction of
a stent into the vascular system of a patient may cause the blood
vessel walls into which the stent is being placed to become
disrupted or injured. Healing of the injury site will involve clot
formation (i.e., thrombosis), thereby causing stenosis (i.e.,
vessel closing) of the blood vessel. Moreover, if the medical
device is left within the patient for an extended period of time,
thrombus often forms on the device itself, again causing stenosis.
As a result, the patient is placed at risk of a variety of
complications, including heart attack, pulmonary embolism, and
stroke. Thus, the use of such a medical device can entail the risk
of precisely the problems that its use was intended to
ameliorate.
[0006] Another site of injury during implantation of medical
devices is the tissue at and beyond the ends of the implanted
stent. Regardless of the cause of the trauma or injury to the
vessel wall, the tissue will react such as with smooth muscle cell
proliferation and the like thereby creating an adverse reaction and
subsequent closure or stenosis of the vessel.
[0007] Another way in which blood vessels undergo stenosis is
through disease. Probably the most common disease causing stenosis
of blood vessels is atherosclerosis. Indeed, atherosclerotic
vascular disease, in the form of coronary artery and peripheral
vascular disease remains the leading cause of mortality in the
United States. [1] Many medical devices and therapeutic methods are
known for the treatment of atherosclerotic disease. Autogenous
veins are the first choice of treatment for vein grafts because of
their long-term patency especially in below knee anastomosis.
However, for many patients suitable vein grafts are not available.
[2] Allografts are in short supply and carry the risk of poor
healing characterized by slow wall lysis, compaction and loss of
elastic tissue, ulceration, mural thrombosis, and calcification.
[3] Large-diameter (>6 mm inner diameter) blood vessels could be
replaced by using non-degradable polymeric materials such as Dacron
(polyethylene terephthalate) and ePTFE (expanded
polytetrafluorotethylene).
[0008] Unfortunately, Dacron and ePTFE are not applicable to
small-diameter (.ltoreq.6 mm inner diameter) blood vessels (SDBV),
especially in locations below the knee. Synthetic materials trigger
inflammatory responses and activate platelets and leukocytes,
initiating thrombogenesis and intimal hyperplasia. Poor patency is
problematic due in part to incomplete endothelialization,
thrombosis, and intimal hyperplasia particularly at the distal
anastomosis. [10, 11]
[0009] Tissue engineering is an emerging alternative which utilizes
biodegradable scaffolds seeded with cells to reconstruct lost
tissues or organs. Significant progress has been made for in vitro
regeneration of SDBV, [4-6] however, there is still a long way to
go before tissue engineered blood vessel substitutes are approved
by Food and Drug Adminstration (FDA). [7] Modification of existing
vascular grafts to improve performance remains a viable option with
room for innovation. It is well known that synthetic grafts do not
spontaneously endothelialize in humans. In addition, the highly
hydrophobic surfaces of ePTFE grafts limit endothelial cell
adhesion. Various modifications have been proposed to either
stimulate in vivo graft endothelialization or improve the retention
of in vitro seeded endothelial cells when they are exposed to
physiological blood flow. [12]. These modifications included
coating or immobilization of endothelial-specific adhesion ligands
such as collagen, albumin, thrombomodulin, gelatin, fibronectin,
collagen-elastin matrices, dipyridamole, fibrin glue, heparin, and
peptides (such as RGD, REDV). [13-17] However, issues of stability
of the coating, transmission of pathogens, and high costs remain.
[18] Plasma treatment is a convenient and widely used method for
modifying the surface of materials without altering their bulk
properties, [19-23] typically used to confer hydrophilicity to a
surface. However, without the inclusion of a polymerizable agent,
the plasma-induced effects are temporary and difficult to control.
Using plasma to polymerize compounds such as ethylene to form a
coating on materials can provide a modified surface that is stable.
[24] Nevertheless, the resulting film is normally non-degradable,
and long-term biocompatibility and the increased compliance
mismatch of the modified grafts are a concern.
[0010] It would be desirable to develop implantable medical devices
and methods for reliably delivering suitable therapeutic and
diagnostic agents, drugs and other bioactive materials directly
into a body portion during or following a medical procedure, so as
to treat or prevent the abrupt closure and/or restenosis of a body
portion such as a passage, lumen or blood vessel.
SUMMARY OF THE INVENTION
[0011] The present invention is directed to a novel poly(diol
citrates)-based coating for implantable devices. More specifically,
the specification describes methods and compositions for making and
using implantable devices coated with citric acid copolymers or
citric acid copolymers impregnated with therapeutic compositions
and/or cells.
[0012] More particularly, the present invention provides an
implantable medical device wherein at least one surface of said
device has deposited thereon a coating comprising a citric acid
polyester having the generic formula (A-B-C)n, wherein A is a
linear aliphatic dihydroxy monomer; B is citric acid, C is a linear
aliphatic dihydroxy monomer, and n is an integer greater than
1.
[0013] Preferably, in the citric acid polymer A is a linear diol
comprising between about 2 and about 20 carbons. In other
embodiments, C is a linear diol comprising between about 2 and
about 20 carbons. In some embodiments, both A and C are the same
linear diol. In other embodiments A and C are different linear
diols. In specific embodiments a preferred linear diol is
1,8-octanediol. In specific embodiments, the linear diol is
aliphatic dihydroxy poly 1,8-octanediol co-citric acid. In other
embodiments, the linear aliphatic dihydroxy poly 1,10-decanediol
co-citric acid.
[0014] The base material from which the implantable device is
prepared may be any base material that is typically used in medical
devices and in prosthetic materials. Exemplary implantable devices
may be prepared from one or more materials selected from the group
consisting of: stainless steel, tantalum, titanium, nitinol, gold,
platinum, inconel, iridium, silver, tungsten, a biocompatible
metal, carbon, carbon fiber, cellulose acetate, cellulose nitrate,
silicone, polyethylene terephthalate, polyurethane, polyamide,
polyester, polyorthoester, polyanhydride, polyether sulfone,
polycarbonate, polypropylene, high molecular weight polyethylene,
polytetrafluoroethylene, a biocompatible polymeric material,
polylactic acid, polyglycolic acid, a polyanhydride,
polycaprolactone, polyhydroxybutyrate valerate, a biodegradable
polymer, a protein, an extracellular matrix component, collagen,
fibrin, a biologic agent, PEBAX, polyethylene, irradiated
polyethylene or a suitable mixture, copolymer, or alloy of any of
these. Preferably, the base material is ePTFE.
[0015] In specific preferred embodiments compositions of the
present invention are such that they can be used to improve the
medical device to give it improved properties for use in blood
vessels or any other use for which the base material of the
implantable device is used. In particular embodiments, the
implantable medical device having the coating of the invention is
rendered less thrombogenic than a similar device prepared from
ePTFE that lacks said coating.
[0016] In preferred embodiments, the coating also may comprise a
therapeutic or other agent for delivery to an in vivo site. In
specific embodiments, such a therapeutic or other agent may be a
bioactive agent wherein the bioactive agent is selected from the
group consisting of: an antisense nucleotide, a thrombin inhibitor,
an antithrombogenic agent, a tissue plasminogen activator, a
thrombolytic agent, a fibrinolytic agent, a vasospasm inhibitor, a
calcium channel blocker, a nitrate, a nitric oxide promoter, a
vasodilator, an antimicrobial agent, an antibiotic, an antiplatelet
agent, an antimitotic, a microtubule inhibitor, an actin inhibitor,
a remodeling inhibitor, an agent for molecular genetic
intervention, a cell cycle inhibitor, an inhibitor of the surface
glycoprotein receptor, an antimetabolite, an antiproliferative
agent, an anti-cancer chemotherapeutic agent, an anti-inflammatory
steroid, an immunosuppressive agent, an antibiotic, a
radiotherapeutic agent, iodine-containing compounds,
barium-containing compounds, a heavy metal functioning as a
radiopaque agent, a peptide, a protein, an enzyme, an extracellular
matrix component, a cellular component, a biologic agent, an
angiotensin converting enzyme (ACE) inhibitor, ascorbic acid, a
free radical scavenger, an iron chelator, an antioxidant, a
radiolabelled form or other radiolabelled form of any of the
foregoing, or a mixture of any of these.
[0017] In preferred embodiments, the coating acts as an
extracellular matrix to support growth of cells. It may be
impregnated with specific factors that facilitate growth of cells,
e. g., growth factors, cytokines, chemokines and the like.
[0018] In other preferred embodiments, the coating comprises cells
selected from the group consisting of endothelial cells, ligament
tissue, muscle cells, bone cells, cartilage cells. In preferred
embodiments, the coating comprises endothelial cells. In other
preferred embodiments, the coating comprises smooth muscle cells.
Preferably, the coating will support the growth of cells in vivo
such that ultimately those cells are able to form part of the
tissue site at which the medical device is implanted.
[0019] In other preferred embodiments, the coating further
comprises a second polymer selected from the group consisting of
poly(hydroxyvalerate), poly(lactide-co-glycolide),
poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate),
polyorthoester, polyanhydride, poly(glycolic acid),
poly(glycolide), poly(L-lactic acid), poly(L-lactide),
poly(D,L-lactic acid), poly(D,L-lactide), poly(caprolactone),
poly(trimethylene carbonate), and polyester amide.
[0020] Also contemplated by the invention is a method of producing
engineered tissue, comprising preparing an implantable medical
device of the invention as outlined above; and culturing cells of
said tissue on the scaffold. The scaffold may be biphasic.
[0021] The invention provides methods of decreasing the
thrombogenicity of a graft comprising coating said graft with a
coating comprising a citric acid polyester having the generic
formula (A-B-C)n, wherein A is a linear aliphatic dihydroxy
monomer; B is citric acid, C is a linear aliphatic dihydroxy
monomer, and n is an integer greater than 1 wherein said graft is
less thrombogenic than a similar graft that does not comprise said
coating.
[0022] In the methods of the invention the coating is preferably
deposited on the inner lumen surface of the small diameter blood
vessel graft, on the outer surface of the small blood vessel graft
or on both the inner lumen surface and the outer surface of the
small diameter blood vessel graft. In other preferred embodiments,
the graft further comprises a layer of endothelial cells deposited
on the surface of said coating. In some preferred methods, the
graft further comprises an addition therapeutic agent impregnated
into said coating. The therapeutic agent may be selected from the
group consisting of an antisense nucleotide, a thrombin inhibitor,
an antithrombogenic agent, a tissue plasminogen activator, a
thrombolytic agent, a fibrinolytic agent, a vasospasm inhibitor, a
calcium channel blocker, a nitrate, a nitric oxide promoter, a
vasodilator, an antimicrobial agent, an antibiotic, an antiplatelet
agent, an antimitotic, a microtubule inhibitor, an actin inhibitor,
a remodeling inhibitor, an agent for molecular genetic
intervention, a cell cycle inhibitor, an inhibitor of the surface
glycoprotein receptor, an antimetabolite, an antiproliferative
agent, an anti-cancer chemotherapeutic agent, an anti-inflammatory
steroid, an immunosuppressive agent, an antibiotic, a
radiotherapeutic agent, iodine-containing compounds,
barium-containing compounds, a heavy metal functioning as a
radiopaque agent, a peptide, a protein, an enzyme, an extracellular
matrix component, a cellular component, a biologic agent, an
angiotensin converting enzyme (ACE) inhibitor, ascorbic acid, a
free radical scavenger, an iron chelator, an antioxidant, a
radiolabelled form or other radiolabelled form of any of the
foregoing, or a mixture of any of these. In specific embodiments,
the therapeutic agent is an anti-inflammatory agent.
[0023] In preferred embodiments, the implantable device is one
which is formulated for use as small diameter blood vessel and the
small diameter blood vessel has an improved property selected from
the group consisting of blood flow, vessel tone, platelet
activiation, adhesion, and aggregation, leukocyte adhesion, SMC
migration and SMC proliferation as compared to a small diameter
blood vessel prepared from a graft that does not comprise said
coating.
[0024] Other features and advantages of the invention will become
apparent from the following detailed description. It should be
understood, however, that the detailed description and the specific
examples, while indicating preferred embodiments of the invention,
are given by way of illustration only, because various changes and
modifications within the spirit and scope of the invention will
become apparent to those skilled in the art from this detailed
description.
BRIEF DESCRIPTION OF THE DRAWINGS
[0025] The following drawings form part of the present
specification and are included to further illustrate aspects of the
present invention. The invention may be better understood by
reference to the drawings in combination with the detailed
description of the specific embodiments presented herein.
[0026] FIG. 1 is a schematic representation of the synthesis of
poly(1,8-octanediol-co-citric acid).
[0027] FIG. 2 shows SEM images of the graft lumen. A) untreated
ePTFE graft; B) POC-ePTFE graft with 1 POC coating (POC-ePTFE 1C);
C) POC-ePTFE graft with 3 coatings (POC-ePTFE 3C); D) POC-ePTFE
graft with 6 coatings (POC-ePTFE 6C). Scale bar=50 .mu.m.
[0028] FIG. 3 shows the effect of POC coating on the compliance of
the ePTFE graft. Compliance of untreated ePTFE grafts (control) and
POC-ePTFE grafts with: a) 1 coating, b) 3 coatings and c) 6
coatings. * P>0.05 vs. control; # P<0.05 vs. control.
[0029] FIG. 4 shows the surface FTIR analysis of coated and
uncoated grafts. A) untreated ePTFE and B) ePTFE with 3 coatings of
POC.
[0030] FIG. 5 shows XPS analysis of the grafts. A) control; and B)
POC-ePTFE with 3 coatings.
[0031] FIG. 6 shows the effect of POC coating on water-in-air
contact angle. A) untreated ePTFE; and B) POC-ePTFE graft with 3
coatings.
[0032] FIG. 7 shows a comparison of Platelet adhesion on various
type of materials.
[0033] FIG. 8 shows a plasma re-calcification clotting profiles.
Platelet poor plasma on TCP (blue), PLGA (80:20)(red), and POC
(green). Clotting is significantly delayed on POC relative to TCP
and PLGA. Purple marks are PPP on TCP with no calcium.
Mean.+-.S.D., N=5.
[0034] FIG. 9 shows In vivo assessment in pigs of the foreign body
reaction to the grafts. A) Intraoperative photo of the implanted
grafts. (B) contrast-enhanced MR angiogram demonstrating patency of
both grafts at 7 days, I=ligated native common iliac arteries. C)
MAC 387 stain for new macrophages and granulocyte (arrowheads) on
POC-ePTFE graft; D) MAC 387 stain for new macrophages and
granulocyte (arrowheads) on control ePTFE graft; L=lumen of the
graft. Scale bars=50 .mu.m. E) and (F): SEM of the lumen of the
grafts. Scale bars=100 .mu.m.
[0035] FIG. 10 shows representative SEM photographs of endothelial
progenitor cells (EPCs) that were cultured on ePTFE graft (A and B)
and POC-ePTFE graft (C and D) for 10 days. Scale bar: A and C 1 mm;
B and D 50 .mu.m.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0036] While there are numerous technologies available for the
treatment and prevention of disorders resulting from occluded or
damaged vessels that are based on the use of implantable medical
devices, there is a need for improved implantable devices. It would
be desirable to produce such implantable devices that have improved
biocompatibility, cell adhesion properties, endothelialization as
well as to decrease thrombogenicity. Further, it would be desirable
to limit the systemic exposure of a subject to the bioactive
materials that are being used in the treatment regimen,
particularly when the agent is a toxic or chemotherapeutic agent
being delivered to a specific site of action through an intravenous
catheter (which itself has the advantage of reducing the amount of
agent needed for successful treatment), by preventing stenosis both
along the catheter and at the catheter tip. It would be desirable
to similarly improve other therapies. Of course, it would also be
desirable to avoid degradation of the agent, drug or bioactive
material during its incorporation on or into any such device. It
would further be highly desirable to develop a method for coating
an implantable medical device with a drug, therapeutic agent,
diagnostic agent or other bioactive material which entailed a
minimum number of steps, thereby reducing the ultimate cost of
treating the patient. It would be desirable to deliver the
bioactive material without causing additional problems with a poor
biocompatible carrier or containment material. Finally, it would be
highly desirable to develop a method for coating an implantable
medical device with a drug, therapeutic agent, diagnostic agent or
other bioactive material which could be carried out in such a way
as to minimize any environmental or personal risks or
inconveniences associated with the manufacture of the device.
[0037] Tissue engineering is an emerging alternative which utilizes
biodegradable scaffolds seeded with cells to reconstruct lost
tissues or organs. Significant progress has been made for in vitro
regeneration of small diameter blood vessels (SDBV), [4-6] however,
there is still a long way to go before tissue engineered blood
vessel substitutes are approved by Food and Drug Administration
(FDA). [7] Modification of existing vascular grafts to improve
performance remains a viable option with room for innovation. The
intrinsic hydrophobicity of ePTFE grafts limit the ability of
endothelial cell to adhere to them.
[0038] By way of explanation, a functional endothelium provides a
continuous thromboresistant layer between blood and the blood
vessel wall. It also controls blood flow and vessel tone, platelet
activation, adhesion and aggregation, leukocyte adhesion, and SMC
migration and proliferation. [25] Mammalian cells maintained in
vitro require nutrients and growth factors but also require an
appropriate substratum upon which to attach and spread. The
intrinsic hydrophobicity of ePTFE grafts limit the ability of
endothelial cell to adhere to them. Previous studies have confirmed
that ECs are able to secrete extracellular matrix (ECM). [8,9] To
date, no vascular graft has successfully achieved long-term patency
when used for SDBV applications. The synthetic graft market is
currently dominated by ePTFE and Dacron; therefore, we are
proposing a novel approach which could significantly reduce the
thrombogenicity of ePTFE grafts, thereby facilitating their use in
SDBV applications. As of yet, no one has investigated the use of a
synthetic biodegradable scaffold as a means to create a more stable
and functional endothelial cell monolayer on ePTFE grafts.
[0039] The present invention describes implantable medical devices
suitable for use inside or outside anatomical structures, such as
body vessels (e.g., vessels of the vasculature like blood vessels).
The invention is based the principle of coating ePTFE grafts using
our own newly developed coating technique, a spin-shearing method,
with biodegradable elastomeric polymers are used as coating
materials, which could be any other biodegradable elastomers
currently available in the market, such as poly(diol citrates) and
its derivatives, polyurethane and its derivatives, polycarbonate
and its derivatives, polyhydroxyalkanoates (PHAs) and their
derivatives, or aliphatic biodegradable polyester elastomers (such
as poly(glycolide-co-caprolactone)(PGCL), polycaprolactone (PCL),
poly(glycerol sebacate)(PGS)). Particularly preferred materials are
poly(diol citrates) and their derivatives. The biodegradable
elastomeric polymers could be mixed with other materials to modify
the grafts. The coating can be applied on either the lumen or
outside or both inside and outside of grafts.
[0040] In exemplary embodiments, the spin-shearing method is
performed to prepare a small diameter blood vessel or similar graft
from ePTFE, comprising modifying the lumen of said small diameter
blood vessel by coating the small diameter blood vessel through a
spin-shearing method wherein said method comprises coating a
glass-rod with said citric acid polymer by use of a mechanical
stirrer; rotating said coated glass rod at a low speed of about 300
rpm; contacting said rotating glass rod with ePTFE or other
materials used to form the coating; shearing the lumen of the graft
by rotating the graft in a direction counter to the direction in
which it is spinning. These steps optionally are repeated 2, 3, 4,
5, 6, 7, 8 or more times.
[0041] The grafts can contain endothelial cells, stem cells, or
endothelial progenitor cells (EPCs) which can be grown on the graft
to form an endothelium layer on top of the underlying polymer
coating. In addition, or as an alternative, the coating may be
impregnated with a therapeutic agent such as any anticoagulant and
anti-inflammation drugs or any other therapeutic agent that it is
desired to release at the site at which the device is implanted. In
certain embodiments, NO-release agents could be incorporated into
the polymer coating for controlled release.
[0042] The present invention for the first time shows the use of a
novel biodegradable elastomeric polymer, poly(diol citrates) to
modify the ePTFE graft via a convenient method to tissue engineer
an endothelium. A particularly preferred method to accomplish this
is the spin-shearing method. In preferred embodiments, the vascular
graft was coated with a POC elastic polymer on the graft lumen via
an interfacial in-situ polycondensation method. This layer of
biodegradable elastic coating is expected to support attachment,
proliferation and gene expression of the endothelial cells (ECs)
without significantly altering the mechanical properties of the
vessel. This is done by facilitating mechanical interlocking
between secreted extracellular matrix and ePTFE fibrils in the
lumen of the graft. As the polymer degrades, it will be replaced by
an endothelium and its secreted extracellular matrix, improving EC
retention under dynamic flow conditions. In addition, this approach
also allows controlled delivery of anti-inflammatory drugs or other
nitric oxide (NO) releasing compounds by incorporating them in the
post-polymerization reaction resulting in a drug-loaded graft. The
resultant vascular graft could serve better for coronary and
peripheral vascular replacement, and potentially as a viable option
in small-diameter blood vessel applications. This method to modify
the vascular graft can be extended to other medical implants and
devices such as devices prepared for implantation to replace
valves, bones and other prosthetic structures. It is contemplated
that the coatings prepared according to the present invention will
improve the biocompatibility and comfort of use of such devices.
Such devices may be prepared from a base material that comprises
one or more materials selected from the group consisting of:
stainless steel, tantalum, titanium, nitinol, gold, platinum,
inconel, iridium, silver, tungsten, a biocompatible metal, carbon,
carbon fiber, cellulose acetate, cellulose nitrate, silicone,
polyethylene terephthalate, polyurethane, polyamide, polyester,
polyorthoester, polyanhydride, polyether sulfone, polycarbonate,
polypropylene, high molecular weight polyethylene,
polytetrafluoroethylene, a biocompatible polymeric material,
polylactic acid, polyglycolic acid, a polyanhydride,
polycaprolactone, polyhydroxybutyrate valerate, a biodegradable
polymer, a protein, an extracellular matrix component, collagen,
small intestine submucosa (SIS), fibrin, a biologic agent, PEBAX,
polyethylene, irradiated polyethylene or a suitable mixture,
copolymer, or alloy of any of these.
[0043] Compositions of poly(diol citrates) that are used to coat
the implantable medical device comprise a citric acid polyester
having the generic formula (A-B-C)n, wherein A and C could be any
of the diols or any combination of the diols; B could be citric
acid, malic acid or their combinations. The diols include linear or
non-linear aliphatic diols, branched diol, cyclodiol, triol,
heteroatom containing diol (such as N-methyldiethanolamine, MDEA)
and macrodiol or their combinations. Any medical device coated with
any biodegradable elastomers (e.g. poly diol-citric acid,
polyurethanes, polycaprolactone and copolymers thereof) is
contemplated to be within the aspects of the present invention.
[0044] POC has been shown to be compatible (i.e. as per cell
adhesion, proliferation, and differentiation assays) with several
cell types including human and pig endothelial cells, human and pig
smooth muscle cells, bovine chondrocytes, and bovine fibroblasts
[16, 30]. It was also shown to be biocompatible in vivo in a rat
subcutaneous implantation model [15]. In another co-pending
application, the favorable cell adhesion and spreading
characteristics of POC were confirmed in vitro with the use of
primary human osteoblasts.
[0045] Poly(diol citrates) are a family of biodegradable and
biocompatible elastomers that have shown significant potential for
soft tissue engineering, see e.g., U.S. patent application see U.S.
60/721,687 and applications depending therefrom. However, while
those prior compositions are useful in the production of matrices
for tissue culture and implantable tissue patches, those
compositions have not previously been demonstrated for coating
implantable devices to produce devices that will be useful in
implantable devices and in coating prostheses and other implantable
devices. The methods and compositions of the present invention are
directed to the use of poly(diol citrate) based polymers to prepare
improved implantable devices.
[0046] Methods and compositions for preparing POC are described in
detail in PCT/US2004/030631 and U.S. 60/721,687, each of which is
incorporated herein by reference in its entirety. Briefly,
equimolar amounts of citric acid and 1,8-octanediol were stirred
and melted together at 160.degree. C. for 15 minutes, decrease
temperature to 140.degree. C. and continue stirring for 1 hr. The
pre-polymer is purified by precipitation in water and then
post-polymerized at 80.degree. C. for 2 days. The pre-polymer is
soluble in ethanol or 1,4-dioxane, which are less toxic than other
commonly used solvents. The mechanical properties and degradation
rates of the elastomer can be modulated by controlling synthesis
conditions such as crosslinking temperature and time, vacuum,
choice of diol, and initial monomer molar ratio. [28] No permanent
deformation was found during mechanical tests. Other poly(diol
citrates) could also be synthesized using the method described
above. Polycondensation of poly(diol citrates) can be conducted
under no vacuum, no catalyst, and low reaction temperature (under
100.degree. C., such as 60.degree. C., 80.degree. C., even as low
as 37.degree. C.). Catalyst and high temperature could also be
applied if needed.
[0047] Composition of poly(diol citrates) comprise a citric acid
polyester having the generic formula (A-B-C)n, wherein A and C
could be any of the diols or any combination of the diols; B could
be citric acid, malic acid or their combinations. The diols include
aliphalic diols or branched diols, cyclodiols, triols,
heteroatom-containing diols and macrodiols or their
combinations;
[0048] In addition to the POC-based polymers, it is contemplated
that the implantable devices may further be coated with another
biodegradable and biocompatible polymer. Two such polymers are
poly(L-lactic acid) (PLLA) and poly(lactic-co-glycolic acid)
(PLGA). These polymers are rigid and strong and have been used in
many tissue engineering applications. Furthermore, the rate of
degradation could be tailored to match that of the surrounding
elastomeric matrix. Poly(L-lactic acid) has a degradation time of
greater than two years while poly(glycolic acid) has a degradation
time of 1-2 months. By changing the ratio of lactic to glycolic
acid, the degradation rate could be varied from fast (1-2 months)
to slow (>2 years). For tissue engineering, the rate of
degradation of the polymer scaffold should match that of tissue
regrowth.
[0049] The polymer may be a biodegradable polymer or a
non-biodegradable polymer, but preferably is a biodegradable
polymer. Biodegradable polymers include, but are not limited to
collagen, elastin, hyaluronic acid and derivatives, sodium alginate
and derivatives, chitosan and derivatives gelatin, starch,
cellulose polymers (for example methylcellulose,
hydroxypropylcellulose, hydroxypropylmethylcellulose,
carboxymethylcellulose, cellulose acetate phthalate, cellulose
acetate succinate, hydroxypropylmethylcellulose phthalate), casein,
dextran and derivatives, polysaccharides, poly(caprolactone),
fibrinogen, poly(hydroxyl acids), poly(L-lactide) poly(D,L
lactide), poly(D,L-lactide-co-glycolide),
poly(L-lactide-co-glycolide), copolymers of lactic acid and
glycolic acid, copolymers of .epsilon.-caprolactone and lactide,
copolymers of glycolide and .epsilon.-caprolactone, copolymers of
lactide and 1,4-dioxane-2-one, polymers and copolymers that include
one or more of the residue units of the monomers D-lactide,
L-lactide, D,L-lactide, glycolide, .epsilon.-caprolactone,
trimethylene carbonate, 1,4-dioxane-2-one or 1,5-dioxepan-2-one,
poly(glycolide), poly(hydroxybutyrate), poly(alkylcarbonate) and
poly(orthoesters), polyesters, poly(hydroxyvaleric acid),
polydioxanone, poly(ethylene terephthalate), poly(malic acid),
poly(tartronic acid), polyanhydrides, polyphosphazenes, poly(amino
acids). The biodegradable polymers used herein may be copolymers of
the above polymers as well as blends and combinations of the above
polymers. (see generally, Illum, L., Davids, S. S. (eds.) "Polymers
in Controlled Drug Delivery" Wright, Bristol, 1987; Arshady, J.
Controlled Release 17:1-22, 1991; Pitt, Int. J. Phar. 59:173-196,
1990; Holland et al., J. Controlled Release 4:155-0180, 1986).
[0050] In particular preferred embodiments, the biodegradable or
resorbable polymer is one that is formed from one or more monomers
selected from the group consisting of lactide, glycolide,
e-caprolactone, trimethylene carbonate, 1,4-dioxan-2-one,
1,5-dioxepan-2-one, 1,4-dioxepan-2-one, hydroxyvalerate, and
hydroxybutyrate. In one aspect, the polymer may include, for
example, a copolymer of a lactide and a glycolide. In another
aspect, the polymer includes a poly(caprolactone). In yet another
aspect, the polymer includes a poly(lactic acid),
poly(L-lactide)/poly(D,-L-Lactide) blends or copolymers of
L-lactide and D,L-lactide. In yet another aspect, the polymer
includes a copolymer of lactide and .epsilon.-caprolactone. In yet
another aspect, the polymer includes a polyester (e.g., a
poly(lactide-co-glycolide). The poly(lactide-co-glycolide) may have
a lactide:glycolide ratio ranges from about 20:80 to about 2:98, a
lactide:glycolide ratio of about 10:90, or a lactide:glycolide
ratio of about 5:95. In one aspect, the poly(lactide-co-glycolide)
is poly(L-lactide-co-glycolide; see e.g., U.S. Pat. No. 6,531,146
and U.S. application No. 2004/0137033.). Other examples of
biodegradable materials include polyglactin, and polyglycolic
acid.
[0051] Representative examples of non-biodegradable compositions
include ethylene-co-vinyl acetate copolymers, acrylic-based and
methacrylic-based polymers (e.g., poly(acrylic acid),
poly(methylacrylic acid), poly(methylmethacrylate),
poly(hydroxyethyl methacrylate), poly(alkylcynoacrylate),
poly(alkyl acrylates), poly(alkyl methacrylates)), polyolefins such
as poly(ethylene) or poly(propylene), polyamides (e.g., nylon 6,6),
poly(urethanes) (e.g., poly(ester urethanes), poly(ether
urethanes), poly(carbonate urethanes), poly(ester-urea)),
polyesters (e.g., PET, polybutyleneterephthalate, and
polyhexyleneterephthalate), olyethers (poly(ethylene oxide),
poly(propylene oxide), poly(ethylene oxide)-poly(propylene oxide)
copolymers, diblock and triblock copolymers, poly(tetramethylene
glycol)), silicone containing polymers and vinyl-based polymers
(polyvinylpyrrolidone, poly(vinyl alcohol), poly(vinyl acetate
phthalate), poly(styrene-co-isobutylene-co-styrene), fluorine
containing polymers (fluoropolymers) such as fluorinated ethylene
propylene (FEP) or polytetrafluoroethylene (e.g., expanded
PTFE).
[0052] The polymers may be combinations of biodegradable and
non-biodegradable polymers. Further examples of polymers that may
be used are either anionic (e.g., alginate, carrageenin, hyaluronic
acid, dextran sulfate, chondroitin sulfate, carboxymethyl dextran,
caboxymethyl cellulose and poly(acrylic acid)), or cationic (e.g.,
chitosan, poly-I-lysine, polyethylenimine, and poly(allyl amine))
(see generally, Dunn et al., J. Applied Polymer Sci. 50:353, 1993;
Cascone et al., J. Materials Sci.: Materials in Medicine 5:770,
1994; Shiraishi et al., Biol. Pharm. Bull. 16:1164, 1993;
Thacharodi and Rao, Int'l J. Pharm. 120:115, 1995; Miyazaki et al.,
Int'l J. Pharm. 118:257, 1995). Preferred polymers (including
copolymers and blends of these polymers) include
poly(ethylene-co-vinyl acetate), poly(carbonate urethanes),
poly(hydroxyl acids) (e.g., poly(D,L-lactic acid) oligomers and
polymers, poly(L-lactic acid) oligomers and polymers, poly(D-lactic
acid) oligomers and polymers, poly(glycolic acid), copolymers of
lactic acid and glycolic acid, copolymers of lactide and glycolide,
poly(caprolactone), copolymers of lactide or glycolide and
.epsilon.-caprolactone), poly(valerolactone), poly(anhydrides),
copolymers prepared from caprolactone and/or lactide and/or
glycolide and/or polyethylene glycol.
[0053] Methods for making POC-PLLA or PLGA or other like composites
are described in U.S. 60/721,687.
[0054] In specific embodiments, the compositions of the invention
(i.e., the compositions that are made up of a poly(diol citrate)
polymer) are used to coat the inner lumen of an implantable device.
Such a coating may be comprised of just the poly(diol citrate) or
the polymer may be impregnated or otherwise loaded with a drug or
other biologically active agent to be delivered (e.g., in a
controlled-release manner) or it may be seeded with cells so that
they can act as cellular tissue patches or tissue grafts.
[0055] In specific embodiments, the POC materials made according to
the methods of the present invention will be useful both as
substrata for the growth and propagation of tissues cells that may
be seeded on the substrata and also as coating on the implantable
devices. In those embodiments where the elastomeric composites are
used in 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 patent. Other organs for
which tissue implantation patches may be generated include, but are
not limited to skin tissue for skin grafts, myocardial tissue, bone
tissue for bone regeneration, testicular tissue, endothelial cells,
blood vessels, and any other cells from which a tissue patch may be
generated. Thus, those of skill in the art would understand that
the aforementioned organs/cells are merely exemplary organs/cell
types and it should be understood that cells from any organ may be
seeded onto the biocompatible elastomeric composites of the
invention to produce useful tissue for implantation and/or
study.
[0056] The cells that may be seeded onto the POC or other 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.
[0057] 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.
[0058] In certain embodiments, it is contemplated that the cells
attached to the biocompatible elastomeric composite substrates of
the invention are grown in bioreactors. A bioreactor may be of any
class, size or have any one or number of desired features,
depending on the product to be achieved. Different types of
bioreactors include tank bioreactors, immobilized cell bioreactors,
hollow fiber and membrane bioreactors as well as digesters. There
are three classes of immobilized bioreactors, which allow cells to
be grown: membrane bioreactors, filter or mesh bioreactors, and
carrier particle systems. Membrane bioreactors grow the cells on or
behind a permeable membrane, allowing the nutrients to leave the
cell, while preventing the cells from escaping. Filter or mesh
bioreactors grow the cells on an open mesh of an inert material,
allowing the culture medium to flow past, while preventing the
cells from escaping. Carrier particle systems grow the cells on
something very small, such as small nylon or gelatin beads. The
bioreactor can be a fluidized bed or a solid bed. Other types of
bioreactors include pond reactors and tower fermentors. Any of
these bioreactors may be used in the present application for
regenerating/engineering tissues on the citric acid based
elastomeric compositions of the present invention.
[0059] Certain tissues that are regenerated by use of the citric
acid based elastomeric composition of the invention may be
encapsulated so as to allow the release of desired biological
materials produced by the cells at the site of implantation, while
sequestering the implanted cells from the surrounding site. Cell
encapsulation can be applied to all cell types secreting a
bioactive substance either naturally or through genetic engineering
means. In practice, the main work has been performed with insulin
secreting tissue.
[0060] 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.
[0061] As indicated above, the cells that are seeded on the
elastomeric composites of the present invention may be cell lines
or primary cells. In certain preferred embodiments, the cells are
genetically engineered cells that have been modified to express a
biologically active or therapeutically effective protein product.
Techniques for modifying cells to produce the recombinant
expression of such protein products are well known to those of
skill in the art. In particular preferred embodiments, the
compositions of the invention may be used to form of a tissue graft
or tissue patch. Endothelial cells are particularly preferred.
Smooth muscle cells also may be used. The cells for the tissue
graft may be an autograft, allograft, biograft, biogenic graft or
xenograft.
[0062] Tissue grafts may be derived from various tissue types.
Representative examples of tissues that may be used to prepare
biografts include, but are not limited to, rectus sheaths,
peritoneum, bladder, pericardium, veins, arteries, diaphragm and
pleura. For such grafts the cells may be endothelial cells,
ligament tissue, muscle cells, bone cells, cartilage cells. Such
cells may be grafted into the compositions of the invention alone
or in combination with a drug or biologically active agent to be
delivered to an in vivo site. For example, such cells for the
biograft may be harvested from a host, loaded with the agent of
interest and then applied in a perivascular manner at the site
where lesions and intimal hyperplasia can develop. Once implanted,
the agent of interest (e.g., paclitaxel and/or rapamycin) is (are)
released from the graft and can penetrate the vessel wall to
prevent the formation of intimal hyperplasia at the treatment site.
In certain embodiments, the biograft may be used as a backing layer
to enclose a composition (e.g., a gel or paste loaded with
anti-scarring agent).
[0063] The patches made of the compositions of the present
invention may be combined with drugs for delivery or therapeutic
agents that can form part of a tissue patch prepared from the
polymers of the invention. For example, the compositions of the
invention may be used to form a mesh or a patch made of the
biodegradable polymeric matrix that conforms to the tissue and
releases the agent (e.g., a therapeutic agent such as a drug or a
diagnostic agent such as a marker, dye or other marker of that will
allow visualization of a diseased state). In preferred examples,
the compositions are fashioned into coating on the surface (inner,
outer or inner and outer surface of an implantable device such as
an ePTFE-based device for use in angioplasty). Such a coating also
may incorporate a drug that can be released in a controlled release
manner. See, e.g., U.S. Pat. No. 6,461,640. Such a drug may be
present in the coating alone, or alternatively, the coating also
may comprise cells that form e.g., an endothelial layer that coats
a surface of the implantable device. The drug may be any
therapeutic agent.
[0064] The coating made of the compositions of the invention may be
impregnated with an antioxidant and/or antimicrobial. See, e.g.,
U.S. Pat. No. 6,572,878. The tissue patch made of the cells, the
ePTFE compound and the coating of the compositions of the invention
may be prepared to be wrapped around a nerve in a canal to reduce
fibroplasia. See, e.g., U.S. Pat. No. 6,106,558. The tissue patch
may be a resorbable collagen membrane that is wrapped around the
spinal chord to inhibit cell adhesions. See, e.g., U.S. Pat. No.
6,221,109. The tissue patch may be used as a dressing to cover a
wound and promote wound healing. See, e.g., U.S. Pat. No.
6,548,728. The compositions of the present invention may be
prepared as a bandage that contains a scar treatment pad with a
layer of silicone elastomer or silicone gel. See, e.g., U.S. Pat.
Nos. 6,284,941 and 5,891,076. The compositions maybe used to
incorporate a biologically active compound. See, e.g., U.S. Pat.
Nos. 6,323,278; 6,166,130; 6,051,648 and 5,874,500.
[0065] Methods for incorporating the biologically active material
onto or into the coating of the present invention include: (a)
affixing (directly or indirectly) to the patch such a biologically
active material (e.g., by either a spraying process or dipping
process as described above, with or without a carrier), (b)
incorporating or impregnating a biologically active material into
the coating made with the composition (e.g., by either a spraying
process or dipping process as described above, with or without a
carrier), (c) by coating the coating made with the composition with
a substance such as a hydrogel which will in turn absorb the
biologically active material, (d) constructing the patch made with
the composition itself with the biologically active material in
either the biodegradable polymer, the poly(diol citrate) polymer,
or in the mixture of the two, or (e) by covalently binding the
biologically active material directly to the surface of the
composition of the invention.
[0066] In specific and preferred embodiments, the poly(diol
citrate) compositions of the present invention are used to coat
devices such as medical stents and the like. For devices that are
coated, the coating process can be performed in such a manner as to
(a) coat only one surface of device with the compositions of the
invention or (b) coating all or parts of the device with the
compositions of the invention.
[0067] The poly(diol citrate)-based coatings or devices coated with
the same may be made sterile either by preparing them under aseptic
environment and/or they may be terminally sterilized using methods
known in the art, such as gamma radiation or electron beam
sterilization methods or a combination of both of these
methods.
[0068] Thus, the therapeutic agent may advantageously be delivered
to adjacent tissues or tissues proximal to the implant site.
Biologically-active agents which may be used alone or in
combination in the implant precursor and implant include, for
example, a medicament, drug, or other suitable biologically-,
physiologically-, or pharmaceutically-active substance which is
capable of providing local or systemic biological, physiological,
or therapeutic effect in the body of the patient. The
biologically-active agent is capable of being released from the
solid implanted matrix into adjacent or surrounding tissue fluids
during biodegradation, bioerosion, or bioresorption of the implant
made from the compositions of the invention.
[0069] Other agents also maybe used in the coating compositions of
the invention. Preferably, such agents are capable of preventing
infection in the host, either systemically or locally at the defect
site, are contemplated as illustrative useful additives. These
additives include anti-inflammatory agents, such as hydrocortisone,
prednisone, and the like, NSAIDS, such as acetaminophen, salicylic
acid, ibuprofen, and the like, selective COX-2 enzyme inhibitors,
antibacterial agents, such as penicillin, erythromycin, polymyxin
B, viomycin, chloromycetin, streptomycins, cefazolin, ampicillin,
azactam, tobramycin, cephalosporins, bacitracin, tetracycline,
doxycycline, gentamycin, quinolines, neomycin, clindamycin,
kanamycin, metronidazole, and the like, antiparasitic agents such
as quinacrine, chloroquine, vidarabine, and the like, antifungal
agents such as nystatin, and the like, antiviricides, particularly
those effective against HIV and hepatitis, and antiviral agents
such as acyclovir, ribarivin, interferons, and the like. Systemic
analgesic agents such as salicylic acid, acetaminophen, ibuprofen,
naproxen, piroxicam, flurbiprofen, morphine, and the like, and
local anaesthetics such as cocaine, lidocaine, bupivacaine,
xylocaine, benzocaine, and the like, also can be used as additives
in the composites.
[0070] Other therapeutic agents include bioactive agents such as
those selected from the group consisting of: an antisense
nucleotide, a thrombin inhibitor, an antithrombogenic agent, a
tissue plasminogen activator, a thrombolytic agent, a fibrinolytic
agent, a vasospasm inhibitor, a calcium channel blocker, a nitrate,
a nitric oxide promoter, a vasodilator, an antimicrobial agent, an
antibiotic, an antiplatelet agent, an antimitotic, a microtubule
inhibitor, an actin inhibitor, a remodeling inhibitor, an agent for
molecular genetic intervention, a cell cycle inhibitor, an
inhibitor of the surface glycoprotein receptor, an antimetabolite,
an antiproliferative agent, an anti-cancer chemotherapeutic agent,
an anti-inflammatory steroid, an immunosuppressive agent, an
antibiotic, a radiotherapeutic agent, iodine-containing compounds,
barium-containing compounds, a heavy metal functioning as a
radiopaque agent, a peptide, a protein, an enzyme, an extracellular
matrix component, a cellular component, a biologic agent, an
angiotensin converting enzyme (ACE) inhibitor, ascorbic acid, a
free radical scavenger, an iron chelator, an antioxidant, a
radiolabelled form or other radiolabelled form of any of the
foregoing, or a mixture of any of these.
[0071] In yet another aspect, the implantable device coatings made
from the compositions of the invention may be used for delivering a
specific therapeutic or other agent to an external portion
(surface) of a body passageway or cavity. Examples of body
passageways include arteries, veins, the heart, the esophagus, the
stomach, the duodenum, the small intestine, the large intestine,
biliary tracts, the ureter, the bladder, the urethra, lacrimal
ducts, the trachea, bronchi, bronchiole, nasal airways, Eustachian
tubes, the external auditory mayal, vas deferens and fallopian
tubes. Examples of cavities include the abdominal cavity, the
buccal cavity, the peritoneal cavity, the pericardial cavity, the
pelvic cavity, perivisceral cavity, pleural cavity and uterine
cavity.
EXAMPLE 1
Biodegradable Elastomeric Polymers
[0072] The coating compositions of the invention are based on
biodegradable elastomeric polymers of poly(diol) citrate molecules.
Such molecules typically comprising a polyester network of citric
acid copolymerized with a linear aliphatic di-OH monomer in which
the number of carbon atoms ranges from 2 to 20. Polymer synthesis
conditions for the preparation of these molecules vary from mild
conditions, even at low temperature (less than 100.degree. C.) and
no vacuum, to tough conditions (high temperature and high vacuum)
according the requirements for the materials properties. By
changing the synthesis conditions (including, but not limited to,
post-polymerization temperature, time, vacuum, the initial monomer
molar ratio, and the di-OH monomer chain length) the mechanical
properties of the polymer can be modulated over a wide range. This
series of polymers exhibit a soft, tough, biodegradable,
hydrophilic properties and excellent biocompatibility in vitro.
[0073] The poly(diol)citrate polymers used herein have a general
structure of: (A-B-C)n
[0074] Where A is a linear, aliphatic diol and C also is a linear
aliphatic diol. B is citric acid. The citric acid co-polymers of
the present invention are made up of multiples of the above
formula, as defined by the integer n, which may be any integer
greater than 1. It is contemplated that n may range from 1 to about
1000 or more. It is particularly contemplated that n may be 1, 2,
3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20,
21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 31, 32, 33, 34, 35, 36,
37, 38, 39, 40, 41, 42, 43, 44, 45, 46, 47, 48, 49, 50, or
more.
[0075] In preferred embodiments, the identity of "A" above is poly
1,10-decanediol and in another preferred embodiment the identity of
A is 1,8-octanediol. However, it should be understood that these
are merely exemplary linear, aliphatic diols. Those of skill are
aware of other aliphatic alcohols that will be useful in
polycondensation reactions to produce citric acid polymers.
Exemplary such aliphatic diols include any diols of between about 2
carbons and about 20 carbons. While the diols are preferably
aliphatic, linear, unsaturated diols, with the hydroxyl moiety
being present at the C.sub.1 and C.sub.x position (where x is the
terminal carbon of the diol), it is contemplated that the diol may
be an unsaturated diol in which the aliphatic chain contains one or
more double bonds. The preferred identity for "C" in one embodiment
is 1,8-octanediol, however as with moiety "A," "C" may be any other
aliphatic alcohols. While in specific embodiments, both A and C are
both the same diol, e.g., 1,8-octanediol, it should be understood
that A and C may have different carbon lengths. For example, A 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, and C may independently be 2, 3, 4,
5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20 or more
carbons in length. Exemplary methods for the polycondensation of
the citric acid with the linear diols are provided in this Example.
These materials are then used as starting materials for the grafts
described in Example 2.
[0076] Synthesis of Poly(1,10-decanediol-co-citric acid) (PDC) In a
typical experiment, 19.212 g citric acid and 17.428 g
1,10-decanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was post-
polymerized at 60.degree. C., 80.degree. C. or 120.degree. C. with
and without vacuum for predetermined time from one day to 3 weeks
depending on the temperature to achieve the
Poly(1,10-decanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from reaction
system.
[0077] Preparation of Poly(1,8-Octanediol-co-citric acid) (POC) In
a typical experiment, 19.212 g citric acid and 14.623 g octanediol
were added to a 250 mL three-neck round-bottom flask, fitted with
an inlet adapter and an outlet adapter. The mixture was melted
within 15 min by stirring at 160-165.degree. C. in silicon oil
bath, and then the temperature of the system was lowered to
140.degree. C.; The mixture was stirred for another 1 hr at
140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was post-
polymerized at 60.degree. C., 80.degree. C. or 120.degree. C. with
and without vacuum for predetermined time (from one day to 3 weeks
depending on the temperature, with the lower temperatures requiring
longer times) to achieve the Poly(1,8-octanediol-co-citric acid).
Nitrogen was introduced into the reaction system before the polymer
was taken out from the reaction system.
[0078] 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 a 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.
[0079] Using similar techniques, porous scaffolds of PDC or other
poly(diol)citrates can be prepared. In other embodiments, biphasic
scaffolds can be prepared. Biphasic scaffolds consist of an outside
porous phase and an inside non-porous phase as depicted in the
schematic drawing shown in FIG. 15 of PCT PCT/US2004/030631,
incorporated herein by reference. 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 the 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, 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 citrate) is a viable strategy
towards the engineering of small diameter blood vessels.
[0081] Synthesis of Poly(1,6-hexanediol-co-citric acid) (PHC). In a
typical experiment, 19.212 g citric acid and 11.817 g
1,6-hexanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in a silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was post-
polymerized at 60.degree. C., 80.degree. C. or 120.degree. C. with
and without vacuum for a predetermined time from one day to 3
weeks, depending on the temperature, to achieve the
Poly(1,6-hexanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from the
reaction system.
[0082] Synthesis of Poly(1,12-dodecanediol-co-citric acid) PDDC. In
a typical experiment, 19.212 g citric acid and 20.234 g
1,12-dodecanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was post-
polymerized at 60.degree. C., 80.degree. C. or 120.degree. C. with
and without vacuum for predetermined time from one day to 3 weeks
depending on the temperature to achieve the
Poly(1,12-dodecanediol-co-citric acid). Nitrogen was introduced
into the reaction system before the polymer was taken out from the
reaction system.
[0083] Synthesis of Poly(1,8-octanediol-co-citric acid-co-glycerol)
In a typical experiment (Poly(1,8-octanediol-co-citric acid-co-1%
glycerol), 23.0544 g citric acid, 16.5154 g 1,8-octanediol and
0.2167 g glycerol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for another hour
at 140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was post-
polymerized at 60.degree. C., 80.degree. C. or 120.degree. C. with
and without vacuum for predetermined time from one day to 3 weeks
depending on the temperature to achieve the
Poly(1,8-octanediol-co-citric acid-co-1% glycerol). Nitrogen was
introduced into the reaction system before the polymer was taken
out from the reaction system.
[0084] Synthesis of Poly(1,8-octanediol-citric acid-co-polyethylene
oxide). In a typical experiment, 38.424 g citric acid, 14.623 g
1,8-octanediol and 40 g polyethylene oxide with molecular weight
400 (PEO400)(100 g PEO1000 and 200 g PEO2000 respectively) (molar
ratio:citric acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 135.degree. C.
The mixture was stirred for 2 hours at 135.degree. C. to get the
pre-polymer. Nitrogen was vented throughout the above procedures.
The pre-polymer was post- polymerized at 120.degree. C. under
vacuum for predetermined time from one day to 3 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 the reaction system. The molar ratios can be altered
to achieve a series of polymers with different properties.
[0085] Synthesis of Poly(1,12-dodecanediol-citric
acid-co-polyethylene oxide). In a typical experiment, 38.424 g
citric acid, 20.234 g 1,12-dodecanediol and 40 g polyethylene oxide
with molecular weight 400 (PEO400)(100 g PEO1000 and 200 g PEO2000
respectively) (molar ratio:citric
acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a 250 ml or 500
ml three-neck round-bottom flask, fitted with an inlet adapter and
an outlet adapter. The mixture was melted within 15 min by stirring
at 160-165.degree. C. in silicon oil bath, and then the temperature
of the system was lowered to 120.degree. C. The mixture was stirred
for half an hour at 120.degree. C. to get the pre-polymer. Nitrogen
was vented throughout the above procedures. The pre-polymer was
post- polymerized at 120.degree. C. under vacuum for predetermined
time from one day to 3 days to achieve the
Poly(1,12-dodecanediol-citric acid-co-polyethylene oxide). Nitrogen
was introduced into the reaction system before the polymer was
taken out from the reaction system. The molar ratios can be altered
to achieve a series of polymers with different properties.
[0086] Synthesis of Poly(1,8-octanediol-citric
acid-co-N-methyldiethanoamine) POCM. In a typical experiment,
38.424 g citric acid, 26.321 g 1,8-octanediol and 2.3832 g
N-methyldiethanoamine (MDEA) (molar ratio:citric
acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a 250 ml or 500
ml three-neck round-bottom flask, fitted with an inlet adapter and
an outlet adapter. The mixture was melted within 15 min by stirring
at 160-165.degree. C. in silicon oil bath, and then the temperature
of the system was lowered to 13520.degree. C. The mixture was
stirred for half an hour at 120.degree. C. to get the pre-polymer.
Nitrogen was vented throughout the above procedures. The
pre-polymer was post- polymerized at 80.degree. C. for 6 hours,
120.degree. C. for 4 hours without vacuum and then 120.degree. C.
for 14 hours under vacuum to achieve the Poly(1,8-octanediol-citric
acid-co-N-methyldiethanoamine). Nitrogen was introduced into the
reaction system before the polymer was taken out from the reaction
system. The molar ratios can be altered to citric
acid/1,8-octanediol/MDEA=1/0.95/0.05.
[0087] Synthesis of Poly(1,12-dodecanediol-citric
acid-co-N-methyldiethanoamine) PDDCM. In a typical experiment,
38.424 g citric acid, 36.421 g 1,12-dodecanediol and 2.3832 g
N-methyldiethanoamine (MDEA) (molar ratio:citric
acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a 250 ml or 500
ml three-neck round-bottom flask, fitted with an inlet adapter and
an outlet adapter. The mixture was melted within 15 min by stirring
at 160-165.degree. C. in a silicon oil bath, and then the
temperature of the system was lowered to 120.degree. C. The mixture
was stirred for half an hour at 120.degree. C. to get the
pre-polymer. Nitrogen was vented throughout the above procedures.
The pre-polymer was post- polymerized at 80.degree. C. for 6 hours,
120.degree. C. for 4 hours without vacuum and then 120.degree. C.
for 14 hours under vacuum to achieve the
Poly(1,12-dodecanediol-citric acid-co-N-methyldiethanoamine).
Nitrogen was introduced into the reaction system before the polymer
was taken out from the reaction system. The molar ratios can be
altered to citric acid/1,12-dodecanediol/MDEA=1/0.95/0.05.
EXAMPLE 2
Methods Employed in the Preparation and Characterization of the
Grafts of the Invention
[0088] Surface modification of ePTFE graft with POC: Prior to
modification, standard ePTFE grafts (Gore-Tex.RTM., W.L. Gore &
Associates Inc., Flagstaff Ariz., 6 mm inner diameter) were cleaned
by first soaking under sonication in absolute ethanol then in
acetone followed by vacuum drying. The lumen of ePTFE grafts were
modified by mechanically coating a layer of POC through a
spin-shearing method. Briefly, a 5 mm diameter glass rod was dipped
into 10% of purified pre-POC solution in 1,4-dioxane and inserted
horizontally into the motor of a mechanical stirrer (IKA.RTM.-Werke
GMBH & CO. KG, Eurostar ST P CV PS S1, Staufen, Germany). The
pre-polymer-coated glass rod was spun clockwise at 300 rpm for 2
minutes and a 6-cm-long piece of ePTFE graft was placed
concentrically over the spinning rod. The lumen of the graft was
sheared against the spinning rod for 2 minutes by manually rotating
the graft counterclockwise. The above procedure was considered to
be 1 coating. To change the amount of POC deposited onto the graft
(and therefore the coating thickness), the above procedure was
repeated 3, and 6 times (defined as 3 and 6 coatings) to assess POC
coverage and effects on graft compliance with increasing polymer
content. Followed by air-drying, the pre-POC coated ePTFE graft was
put into an oven for 80.degree. C., 2 days to obtain POC-ePTFE
graft.
[0089] Characterization of POC-ePTFE graft: This Section details
the characterization of exemplary coated implantable device
materials prepared by the invention.
[0090] Fourier transform infrared (FTIR)) analysis. POC-ePTFE graft
and control ePTFE graft were cut open into small piece (1.times.2
cm), modified side up, and placed on the sample stage of a FTIR
spectrometer (Thermo Nicolet Nexus 870, Keck II, NUANCE,
Northwestern University) under a mode of Diffuse reflectance
infrared Fourier transform spectroscopy (DRIFTS) measurement, which
is suited to characterize a rough surface.
[0091] X-ray photoelectron spectroscopy (XPS) analysis. XPS spectra
of the POC-ePTFE and untreated ePTFE graft (control) were acquired
on an Omicron ESCA probe (electron spectroscopy for chemical
analysis, Keck II, NUANCE, Northwestern University) at a power of
300 W. All measurements were taken under vacuum (<3.times.10-9
mbar).
[0092] Scanning electron microscopy (SEM). The surface morphology
of the lumen of POC-ePTFE and control grafts were sputter-coated
with gold and examined under a SEM (Hitachi 3500N, EPIC,
Northwestern University).
[0093] Contact angle measurements. The water-in-air contact angles
were measured at room temperature using the sessile drop
method.sup.20 using a Rame-Hart goniometer and an imaging system
(Rame-Hart Inc., Mouttain Lake, N.J.) after placing water drops on
the surfaces. The contact angles changes over time were monitored.
Four independent measurements at different sites were averaged.
[0094] Compliance measurements. Compliance was defined as diameter
changes percentage for a given pressure change, typically from 80
to 120 mmHg. Measurements were conducted on a custom-made in vitro
closed-loop flow system. Computer controlled pressurization or a
Harvard Apparatus pulsatile blood pump (Harvard Apparatus,
Hollister, Mass.) was used to control pulsation frequency and
pressure ranges. A LED micrometer (Keyence LS7000,
Higashi-Yodogawa, Osaka) was used to non-invasively measure radial
distension. The pressure and distension data over time was recorded
with LabVIEW software (National Instruments, Austin Tex.).
[0095] Hemocompatibility evaluation of POC-ePFE: Platelet adhesion:
The methods used to collect and prepare the platelets used in this
study have been approved by the Institutional Review Board (IRB
Project #1118-001) and the Office for the Protection of Research
Subjects at Northwestern University. Blood was drawn from healthy
adult volunteers by venipuncture into ACD anticoagulant (Acid
Citrate Dextrose, Solution A; BD Franklin Lakes, N.J.).
Platelet-Rich Plasma (PRP) was prepared as previously described
(Grunkemeier, et al, 1998). Briefly, whole blood was centrifuged at
250 g for 15 minutes and the PRP removed. Plasma proteins were
separated from the platelet fraction utilizing size exclusion
chromatography. The columns (Bio-Rad, Hercules, Calif., #732-1010)
were packed with Sepharose 2B (Sigma, St. Louis, Mo., #2B-300) and
equilibrated with platelet-suspending buffer (PSB). The PRP was run
though the column, to allow plasma proteins to bind to the
sepharose. The platelet fraction was collected, counted, and
adjusted to a final concentration of 5.times.10.sup.7/mL in PSB.
Test and control samples were incubated with PRP for 1 hour at
37.degree. C. under static conditions. The suspension was aspirated
and each well rinsed carefully six times with PBS. The number of
adherent platelets was determined by detecting the amount of
lactate dehydrogenase (LDH) using a modification of the methods
described by Tamada et al.sup.30. Adherent platelets were lysed by
incubation with Triton-PSB buffer at 37.degree. C. for 30 minutes.
A colorimetric substrate for LDH (Roche Diagnostics Corporation,
Indianapolis, Ind., 1644793) was added to the solution and
incubated for 20 minutes at 37.degree. C. The reaction was stopped
by the addition of 1N HCl. The O.D. was read at 490 nm with
reference wavelength of 650. Test and control samples were
incubated with the PRP for 1 hour at 37.degree. C. under static
conditions. Following incubation, the suspension was aspirated and
each well was rinsed carefully six times with PBS. Adherent
platelets were fixed using 2.5% glutaraldehyde in PBS for at least
2 hours, dehydrated in a graded series of ethanol and critical
point dried. The samples were then sputter coated with a 7 nm layer
of gold and observed using SEM.
[0096] Platelet activation: Test and control samples were incubated
with whole blood for 1 hour at 37.degree. C. under static
conditions. The blood was removed and centrifuged at 2000 g for 10
minutes to obtain the Platelet Poor Plasma (PPP). Soluble
P-selectin levels in the plasma were determined using an ELISA kit
(Parameter Human soluble P-selectin Immunoassay, R & D Systems,
Minneapolis, Minn., #BBE 6).
[0097] Aortoiliac bypass graft model in pigs: All procedures and
care were performed in accordance with the regulations of the
animal care and use committee of Northwestern University. Three
male pigs (Yorkshire Landrace, Oak Hill Genetics, Fanning Farms)
weighing 20 to 25 Kg were used in the study. The animals received
pre-operative analgesia with buprenorphine (0.05 mg/kg IM), and
sedation with Acepromazine (0.15 mg/kg IM) and Ketamine (20 mg/kg
IM). After intubation, maintenance anesthesia was conducted with
Isoflurane (0.5-2.0%) delivered with 100% oxygen. A midline
abdominal incision was made in order to expose and dissect both
iliac arteries. Prior to vascular occlusion, the animals received
intravenous heparin (150 units/kg). Bilateral aortoiliac bypass
grafting was performed with a POC-ePTFE graft on one side and an
untreated ePTFE graft (control) on the contralateral side. All
animals received aspirin (325 mg daily) as an antiplatelet therapy
pre and post-operatively. Animals were monitored for the patency of
the grafts via MRI at CAMRI (Center for Advanced MRI, Northwestern
University). Contrast enhanced MR was performed using a
time-resolved T1-weighted gradient echo pulse sequence. A
time-series of 3D contrast-enhanced MR angiograms were acquired
with an injection of 0.1 mmol/kg of a gadolinium based contrast
agent (Magnevist, Berlex, Princeton, N.J.). Following completion of
the angiogram, the bypass grafts was harvested for further analysis
via histology and immunohistochemistry.
[0098] Histology and immunohistochemistry staining of the grafts:
Grafts and adjacent 3-cm segments of attached vessel at each
anastomosis were harvested and fixed in 10% neutral buffered
formalin (Sigma, Milwaukee, Wis.). Sections of each graft were also
fixed in a 2.5% glutaraldehyde solution for observation via SEM.
Formalin-fixed grafts were embedded in paraffin and sectioned at a
distance of 2 mm from the proximal and distal anastomoses.
Five-micron sections were cut and stained with hematoxylin-eosin
(H&E) staining, and macrophage staining using MAC387 antibody
(Serotec, UK).
[0099] Seeding of endothelial progenitor cells on POC-ePTFE grafts:
Prior to cell seeding, pig endothelial progenitor cells (EPCs)
isolated from peripheral blood of pigs were stained with vWF (von
Willerbrand Factor) to confirm their endothelial phenotype.
POC-ePTFE and untreated ePTFE grafts (6 cm in length) were
gas-sterilized via exposure to ethylene oxide and placed in a
culture dish (150.times.15 mm, Becton Dickinson, N.Y.). An
suspension of EPCs was injected into the lumen of the grafts at a
density of 3.9.times.0.sup.6 cells/ml and incubated at 37.degree.
C. for 60 min. The graft was rotated 180.degree., and an additional
suspension of EPCs injected into the lumen of the grafts. The cell
suspension was incubated in the grafts for 60 min prior to adding
80 ml of fresh culture medium to the each culture dish. After
culture for 1 week, grafts were taken out of culture dishes and
washed with PBS three times. Thereafter the grafts were cut into
three 2 cm-long segments. Each segment was subsequently cut into
two parts. One was fixed with 2.5% glutaraldehyde in PBS for
observation via SEM and the other part was fixed with 10% neutral
buffered formalin for H&E staining and vWF (von Willerbrand
Factor) staining.
EXAMPLE 2
Results and Discussion
[0100] ePTFE grafts are manufactured by a heating, extruding, and
longitudinal stretching at a high strain rate and cracking into a
non-woven porous tube. ePTFE grafts are characteristic of a
node-fibril structure (FIG. 1A) in which non-continuous nodes
connect through fine fibrils. Average internodal distance is about
30 .mu.m for standard ePTFE grafts. ePTFE grafts trigger
inflammation, thrombosis and incomplete endothelium formation, the
major causes for graft failure together with the compliance
mismatch. Many attempts have been tried to modify the lumen of
ePTFE grafts to abolish their anti-adhesion properties and improve
endothelialization including carbon coating to increase surface
electronegativity for reduced thrombus formation.sup.31,32, and
attaching anticoagulant or antithrombotic agents to the
grafts.sup.15,33,34 13-17 through chemical or physical
modifications. However, there are still no satisfactory grafts with
a long term patency available, especially for small diameter blood
vessel replacements that are associated with problems caused by
unsatisfactory antithrombogenicity, weak endothelial adherence, and
always even worse compliance mismatch after
modification.sup.13,35-38.
[0101] Using a newly developed biodegradable elastomer, the lumen
of ePTFE grafts have been modified via a convenient mechanical
spin-shearing method. FIGS. 1B, C and D showed the morphology of
ePTFE lumen after different times of coating with biodegradable
elastomeric POC. The more times the coating was applied, the higher
was the coverage of POC coating on ePTFE lumen. Notably, this
elastic coating could maintain the aligned structure of fibril
between nodes instead of binding the fibrils together when using
other non-elastic polyers such as poly(ethyleneimide) (PEI, data
not shown here). Thus, it is understandable that the coated ePTFE
grafts could maintain their original size instead of shrinking as
can happen when using other non-elastic polymers. The elastic
coating is expected to expand or contract with the nodes and
fibrils under the dynamic physiological fluid flow. The compliance
measurements (FIG. 2) on modified ePTFE grafts confirmed that with
limited coating (<3 times under the proposed coating conditions)
the compliance of the ePTFE grafts were not adversely affected
(POC-ePTFE 1C and POC-ePTFE 3C vs. control ePTFE, P<0.05).
[0102] Surface characterization of ePTFE grafts through surface
FTIR analysis and surface element composition analysis (XPS)
confirmed the successful coating of POC on ePTFE grafts. FIG. 3
showed that the broad peaks centered at 3475 and 3215 cm.sup.-1,
peaks at 2925 and 2850 cm.sup.-1, and peaks within 1690 to 1750
cm.sup.-1 in spectra B were assigned to the hydroxyl group
stretching vibration and .nu..sub.O--H of carboxyl groups, the
--CH2- groups, and carbonyl (C.dbd.O) groups from POC..sup.28 XPS
analysis showed the F/C of control ePTFE was 2.1 which was exactly
in agreement with the previous work.sup.16. There is a peak
appearing at approximately 539 ev, which is assigned to O1s from
POC. A decreased F/C (1.36) in POC-ePTFE occurred with an increased
O/C from 0 to 0.13 compared with control ePTFE.
[0103] FIG. 5 showed that the initial contact angle of ePTFE was
132.5.+-.0.2.degree.. Introduction of POC on ePTFE could
significantly decrease the initial contact angles to
121.2.+-.0.2.degree.. More importantly, The equilibrium
water-in-air contact angles of POC-ePTFE after 75 min of contact
time could reach 38.1.+-.3.9 as compared to 101.7.+-.0.8.degree. of
control ePTFE suggesting that POC-ePTFE have very good wettability,
which is expected to facilitate the cell adhesion..sup.19
[0104] Coating an ePTFE graft with POC could drastically improve
the hemocompatibility of ePTFE grafts. FIG. 6 showed the comparison
of platelet adhesion on various type of materials. POC could
significantly inhibit the platelet adhesion. Since ePTFE is not a
transparent material, the clotting assay was only conducted on POC
films as compared with TCP and PLGA. The results of the clotting
assay (FIG. 7) suggested that clotting time could be significantly
delayed by coating the ePTFE graft with POC. Interestingly, the
clotting delay of POC is significantly longer than the commercially
available biodegradable polymer, PLGA, which is widely used in
tissue engineering.
[0105] The in vivo biocompatibility of POC-ePTFE grafts was
evaluated through a pig iliac artery model. Grafts were explanted
after 1 week implantation. FIGS. 8A and B showed that both
POC-ePTFE and control grafts were open in 3 pigs after 1 week
implantation. Inflammation cells staining (MAC 387) indicated there
are less recruited macrophages in POC-ePTFE grafts compared to
control grafts (FIGS. 8C and D). SEM pictures (FIGS. 8E and F)
indicated that the lumen of POC-ePTFE has less fibrin coagulum,
amorphous platelet-rich and/or white blood cell-rich materials
development compared to control ePTFE. The above results
demonstrated the biocompatibility and hemocompatibility of ePTFE
grafts were greatly improved with POC coating.
[0106] As for the issue of cell sources, harvesting autologous
endothelial cells remains a challenge since the ECs are normally
harvested by sacrificing a healthy blood vessel. Recent studies
evidenced that matures ECs and immature endothelial-like cells
float in the peripheral blood.sup.39-41. Yamashita has reported
embryonic vascular progenitors capable of differentiating into both
endothelial and smooth muscle-like cells in response to different
culturing environment..sup.42 Thus, there is interest in using
autologous EPCs isolated from pig peripheral blood to seed the
grafts for a pig model. FIG. 9 showed the SEM observation of EPCs
seeded grafts. EPCs appeared as patches on control grafts
suggesting an incomplete endothelialization while a totally
confluent endothelium was formed on POC-ePTFE. This indicated that
POC-ePTFE could facilitate the formation of complete
endothelialization. ECs have been confirmed to be able to lay down
extracellular matrix (ECM).sup.8,9.
[0107] Thus, a tissue-engineered endothelium could be formed, which
is expected to be much better than the endothelium formed using
current methods. Firstly, the POC coating provides an excellent
support for ECs adhesion; Secondly, since POC is a biodegradable
polymer, the degrading POC will gradually be replaced by the ECM
produced by ECs; Thirdly, this gradual replacement is expected to
ensure the maintenance of a complete endothelium. Given that the
POC coating is wrapping around the fibrils of the ePTFE lumen just
like the fibrils inserted in the POC films, the ECM produced by ECs
could be induced by the gradual degradation of POC and wrap around
the fibrils eventually. This kind of ECM support would be ideal for
effectively supporting an endothelium.
[0108] In conclusion, the present study has for the first time
demonstrated the use of a novel biodegradable elastomeric polymer
to modify the ePTFE graft via a convenient method to tissue
engineer an endothelium. The long term patency of ePTFE grafts is
expected to be greatly improved. The in vivo long term evaluation
of POC-ePTFE is underway.
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