U.S. patent application number 13/251910 was filed with the patent office on 2012-04-05 for implantable bioreactor for delivery of paracrine factors.
This patent application is currently assigned to The Johns Hopkins University. Invention is credited to Jason Benkoski, Jeffrey Brinker, George Coles, Gary Gerstenblith, Chao-Wei Hwang, Peter Johnston, Steven P. Schulman, Gordon Tomaselli, Robert G. Weiss.
Application Number | 20120083767 13/251910 |
Document ID | / |
Family ID | 45890420 |
Filed Date | 2012-04-05 |
United States Patent
Application |
20120083767 |
Kind Code |
A1 |
Gerstenblith; Gary ; et
al. |
April 5, 2012 |
IMPLANTABLE BIOREACTOR FOR DELIVERY OF PARACRINE FACTORS
Abstract
An implantable bioreactor containing a barrier which is designed
to allow the release of cell-derived biomolecules, but restricts
the entry of immunologic and other cells, or the egress of the
cells contained within the bioreactor. Two broad classes of
implantable bioreactors are envisioned, encompassing devices for
both systemic delivery of the bio-products and local delivery at
the target tissue. Bioreactors of both classes can be implanted via
surgery, through percutaneous techniques, or other techniques which
effect implantation.
Inventors: |
Gerstenblith; Gary;
(Reisterstown, MD) ; Benkoski; Jason; (Ellicott
City, MD) ; Brinker; Jeffrey; (Baltimore, MD)
; Coles; George; (Baltimore, MD) ; Hwang;
Chao-Wei; (Ellicott City, MD) ; Johnston; Peter;
(Baltimore, MD) ; Tomaselli; Gordon; (Lutherville,
MD) ; Weiss; Robert G.; (Hunt Valley, MD) ;
Schulman; Steven P.; (Reisterstown, MD) |
Assignee: |
The Johns Hopkins
University
Baltimore
MD
|
Family ID: |
45890420 |
Appl. No.: |
13/251910 |
Filed: |
October 3, 2011 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
61388778 |
Oct 1, 2010 |
|
|
|
Current U.S.
Class: |
604/890.1 |
Current CPC
Class: |
A61L 2430/20 20130101;
A61L 27/16 20130101; A61L 2300/42 20130101; C12N 2533/30 20130101;
A61F 2/022 20130101; A61L 27/54 20130101; A61L 29/146 20130101;
A61L 31/16 20130101; A61L 2300/64 20130101; A61L 31/10 20130101;
A61K 9/0024 20130101; A61K 35/28 20130101; A61L 27/3834 20130101;
A61L 29/00 20130101; A61L 2300/30 20130101; A61L 29/16 20130101;
A61L 27/28 20130101; A61L 27/56 20130101; C12N 5/0663 20130101;
A61K 35/545 20130101; A61L 29/005 20130101; A61L 31/145 20130101;
A61L 31/146 20130101 |
Class at
Publication: |
604/890.1 |
International
Class: |
A61M 5/00 20060101
A61M005/00 |
Claims
1. An implantable bioreactor comprising a housing comprising cells
which can produce paracrine factors; wherein the housing comprises
a barrier that shields the enclosed cells from immunological attack
and permits the transfer of paracrine factors out of the
housing.
2. The implantable bioreactor of claim 1, wherein the cells are
stem cells.
3. The implantable bioreactor of claim 2, wherein the stem cells
are selected from the group consisting of embryonic stem cells,
mesenchymal stem cells, whole bone marrow stem cells,
adipose-derived stem cells, myocardial derived stem cells and
endothelial progenitor stem cells.
4. The implantable bioreactor of claim 1, wherein said housing is
selected from the group consisting of a pouch, a semi-permeable
membrane, a polymer matrix, a matrix gel and a microfabricated
cellular enclosure.
5. The implantable bioreactor of claim 1, wherein said housing is
composed of cellulosic and/or synthetic materials.
6. The implantable bioreactor of claim 5, wherein said cellulosic
material is cellulose acetate.
7. The implantable bioreactor of claim 5, wherein said synthetic
material is selected from the group consisting of polysulfone,
polyamide, polacrylonitrile, copolymers thereof,
polymethylmetacrylate, polytetrafluroethylene and derivatives
thereof, silicon carbide and micro-mechanical porous silicon
diaphragms.
8. The implantable bioreactor of claim 1, wherein said bioreactor
is a pouch-based bioreactor.
9. The implantable bioreactor of claim 8, wherein the pouch-based
bioreactor is mounted on a wire, a catheter or as part of another
implantable device.
10. The implantable bioreactor of claim 4, wherein the polymer
matrix is selected from the group consisting of polyethylene
glycol, hyaluronic acid, chitosan, dextran, collagen and
self-assembling oligopeptides.
11. The implantable bioreactor of claim 4, wherein the matrix gel
is a cross-linked hydrogel.
12. The implantable bioreactor of claim 11, wherein the
cross-linked hydrogel is selected from the group consisting of a
N-hydroxysuccinimide ester/amine conjugate, an isocyanate/amine
conjugate, an epoxy/amine conjugate, an isothiocyanate/amine
conjugate, an alcohol/glutamate conjugate, and a thiol/maleimide
conjugate.
13. The implantable bioreactor of claim 1, wherein the bioreactor
is adhered to a medical device.
14. The implantable bioreactor of claim 4, wherein the housing is a
microfabricated cellular enclosure.
15. The implantable bioreactor of claim 14, wherein the
microfabricated cellular enclosure comprises a silicon-permeable
membrane.
16. The implantable bioreactor of claim 14, wherein the
microfabricated cellular enclosure is adhered to a medical
device.
17. A method of making an implantable bioreactor comprising
combining a housing with cells capable of producing paracrine
factors.
18. The method of claim 17, wherein the cells are selected from the
group consisting of embryonic stem cells, mesenchymal stem cells,
adipose-derived stem cells, whole bone marrow stem cells,
myocardial derived stem cells and endothelial progenitor stem
cells.
19. The method of claim 17, wherein the housing is selected from
the group consisting of a pouch, a semi-permeable membrane, a
polymer matrix, a matrix gel and a microfabricated cellular
enclosure.
20. The method of claim 17, wherein the implantable bioreactor is
adhered to a medical device.
21. A medical device comprising a bioreactor comprising a housing
comprising cells which can produce paracrine factors; wherein the
housing comprises a barrier that shields the enclosed cells from
immunological attack and permits the transfer of paracrine factors
out of the housing.
22. The medical device of claim 21, wherein the medical device is a
stent, intra-aortic balloon pump, ventricular assist device,
prosthetic valve, prosthetic valve clip, prosthetic valve ring,
thrombus filter, pacemaker, defibrillator, pacemaker or
defibrillator wire, septal occluder, atrial appendage device,
pulmonary artery catheter, venous catheter or arterial
catheter.
23. The medical device of claim 22, wherein the bioreactor housing
is selected from the group consisting of a pouch, a semi-permeable
membrane, a polymer matrix, a matrix gel and a microfabricated
cellular enclosure.
24. The medical device of claim 23, wherein the matrix gel is a
cross-linked hydrogel.
25. The medical device of claim 23, wherein the medical device is a
stent and the bioreactor housing is a matrix gel.
26. The medical device of claim 25 wherein the matrix gel is a
cross-linked hydrogel.
27. The medical device of claim 23, wherein the medical device is a
stent and the bioreactor housing is a microfabricated cellular
enclosure.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application claims benefit of U.S. Provisional Patent
Application No. 61/388,778, filed Oct. 1, 2010, the contents of
which are incorporated by reference in its entirety.
BACKGROUND
[0002] Heart muscle damage is the final common pathway for most
forms of cardiovascular disease and when extensive can impair
quality of life and shorten survival. The most common cause is
obstruction in coronary arteries, but the heart as well as other
organs can be damaged by trauma, toxins, and infections. There are
no currently approved therapies to generate new myocardium, or
heart muscle. Stem cells have been administered intravenously (Hare
et al. J Am Coll Cardiol. Dec. 8, 2009; 54(24):2277-2286), via
infusion into the coronary arteries (Abdel-Latif et al. Arch Intern
Med. May 28, 2007; 167(10):989-997, Johnston et al. Circulation.
Sep. 22, 2009; 120(12):1075-1083) and by injection into the heart
muscle itself (Williams et al. Circ Res. Apr. 1, 2011;
108(7):792-6). The results in terms of improved muscle function,
however, have been very limited, possibly because of early
"washout" (short duration of viable cells at the target site
because of blood flow), the hostile environment into which the
cells are delivered, and immunologic attack (Chavakis et al.
Circulation; Jan. 19, 2010; 121(2):325-335, Terrovitis et al. Circ
Res. Feb. 19, 2010; 106(3):479-494). Biomolecules released by the
stem cells, called paracrine factors, may be responsible, in part,
for their benefit (Gnecchi et al. Nat Med. April 2005;
11(4):367-368, Gencchi et al. Circ Res. Nov. 21, 2008;
103(11):1204-1219, Chimenti et al. Circ Res. Mar. 19, 2010;
106(5):971-980).
SUMMARY
[0003] An embodiment of the invention relates to an implantable
bioreactor comprising a housing comprising cells which produce
paracrine factors in situ; wherein the housing comprises a barrier
that shields the enclosed cells from immunological attack and
permits the transfer of paracrine factors out of the housing. The
implantable bioreactor can be for systemic or local delivery of
paracrine factors. The implantable bioreactor housing can be in the
form of a pouch, semi-permeable membrane, a cellular microenclosure
or a matrix gel. The bioreactor can be optionally adhered to a
medical device.
BRIEF DESCRIPTION OF DRAWINGS
[0004] FIG. 1: Miniature bioreactor for in vitro paracrine factor
production and cell viability experiments.
[0005] FIG. 2: Simplified catheter-based temporary stem cell
bioreactor.
[0006] FIG. 3: Catheter-based temporary stem cell bioreactor.
[0007] FIG. 4: Microfabricated silicon cell enclosure.
[0008] FIG. 5: Microfabricated cellular enclosure mounted to a
vascular stent.
[0009] FIG. 6: Microfabricated cellular enclosure bonded to
coronary guide wire.
[0010] FIG. 7: Hydrogel coated stent encapsulating stem cells
within a crosslinked hydrophilic polymer.
[0011] FIG. 8: Final solid fraction of the hydrogel plotted versus
the initial solid fraction. Higher final solid fractions are
associated with greater crosslink densities.
[0012] FIG. 9: Final organic solid fraction plotted against the
molecular weight of the polyethylene glycol precursor.
DETAILED DESCRIPTION
[0013] 1000131 Clinical trials using intra-coronary and
intra-myocardial injection of stem cells in an attempt to heal and
regenerate infarcted myocardium have produced modest results to
date (Chavakis et al. Circulation; Jan. 19, 2010; 121(2):325-335,
Terrovitis et al. Circ Res. Feb. 19, 2010; 106(3):479-494).
Potential reasons for the modest results may be related to
inadequate levels of paracrine factors from the cells, which may be
due to poor retention of cells due to cell death, removal via
immunologic mechanisms, and simple "washout" following delivery,
leaving cells with only a brief opportunity to exert beneficial
effects. Embodiments of the implantable bioreactor disclosed herein
can solve the problems of diffusion or washout by providing
adequate, prolonged delivery of paracrine factors secreted from the
bioreactor while protecting the contents of the bioreactor from
immunologic clearance in an enclosed housing. In an embodiment, the
housing allows for the passage of nutrients, which are smaller than
immune cells, into the reactor. The invention includes in one
embodiment, a minimally invasive percutaneous bioreactor and, in
another embodiment, an implantable device either of which can
adequately produce and release paracrine factors. The bioreactor
can also be used to promote healing and regeneration by release of
paracrine factors in any other tissue or organ.
[0014] As used herein, "bioreactor" refers to a collection of
cells, in a housing, capable of producing and releasing paracrine
factors. As used herein, "paracrine factors" are diffusible
components produced by one cell to affect another cell. The
diffusible components can be any protein, growth factor,
biomolecule, nutrient or fluid produced by the cells housed in the
bioreactor. As used herein, "cells" include any cell capable of
producing and releasing paracrine factors. As used herein, "cells"
can include pancreatic beta cells, endothelial cells, myocardial
cells, and fibroblasts, as well as genetically altered cells. As
used herein, "stem cells" include, but are not limited to,
embryonic and adult stem cells. Embryonic stem cells include,
without limitation, totipotent, pluripotent and multipotent stem
cells, and adult stem cells include, without limitation,
mesenchymal stem cells, adipose-derived stem cells, whole
bone-marrow derived stem cells, myocardial derived stem cells and
endothelial progenitor stem cells. Combinations of stem cells and
these other cell types are also contemplated.
[0015] Various embodiments of the invention include (a) an
implantable bioreactor which enhances recovery of injured
myocardium and other tissue utilizing a stem cell strategy; (b) a
percutaneously implantable bioreactor, which also includes an
easily retrievable percutaneous bioreactor allowing removal once an
intended treatment period is complete; (c) a temporary implantable
device that releases paracrine factors, which are generated de novo
by stem cells and/or other cell types; (d) a permanent implantable
device that releases paracrine factors, which are generated de novo
by stem cells and/or other cell types; (e) a bioreactor implanted
via a standard vascular sheath; (f) an implantable bioreactor which
locally releases paracrine factors in the target tissue; (g) an
implantable bioreactor which contains a barrier with pores which
allow the release of cell-derived biomolecules, but not large
enough to allow the entry of immunologic and other cells, or the
egress of the stem cells and/or other cell types; (h) an
implantable bioreactor which contains a barrier composed of the
material described herein and which is designed to allow the
release of cell-derived biomolecules, but not large enough to allow
the entry of immunologic and other cells, or the egress of the stem
cells and/or other cell types; and (i) an implantable bioreactor
which systemically releases paracrine factors.
[0016] Disclosed herein are two classes of implantable bioreactors.
Class I includes bioreactors designed for systemic delivery of
produced bio-products. Class II includes bioreactors designed for
local delivery of produced bio-products at the target tissue. Both
classes encompass embodiments for implantation via open surgery,
percutaneous techniques, or any other technique to effect
implantation. Both classes, in various embodiments, can adhere to a
medical device.
[0017] Class I implantable bioreactor for systemic delivery. In one
embodiment, the systemic delivery implantable bioreactor comprises
an enclosure housing stem cells and/or other cell types which
produce and release paracrine factors. The enclosure comprises, in
one embodiment, a physical enclosure fabricated with a semi-porous
membrane, or in another embodiment fabricated with a microporous
polymer matrix encapsulating the cells. The enclosure includes
micropores impermeable to cells, but of sufficient size to allow
free permeation of fluid so that biomolecules, waste and nutrients
can be transferred efficiently and without hindrance. In one
embodiment, the implantable bioreactor is deployed in the
intravascular space (such as central veins and large arteries), but
implantation into any other body cavity, tissue, blood vessel, or
organ is also contemplated. In one embodiment, the implantable
bioreactor is temporarily implanted and retrieved later. In another
embodiment, the implantable bioreactor remains indefinitely as a
permanent implant.
[0018] In another embodiment of a systemic delivery bioreactor, the
implantable bioreactor comprises a membranous pouch which contains
within its lumen stem cells and other cell types and/or media to
enhance the viability of the stem cells and which is constructed of
a semi-permeable membrane which shields the cells from immunologic
attack and allows the release of the paracrine factors and other
biomolecules. In various embodiments, the membranous pouch can be
(a) stand-alone, in which case it can be surgically implanted, or
(b) mounted on a wire or catheter, in which case it can be
percutaneously implanted, or (c) mounted as part of another
implantable device, in which case it can be implanted along with
the other device. In an embodiment, the membranous pouch can be
pre-filled with its intended contents or, if attached to a
catheter, filled and potentially emptied and re-filled during and
after implantation. When percutaneously implanted, the device can
be passed via a vascular sheath (as illustrated in FIG. 2).
[0019] Stand-alone Membranous Pouch-based bioreactor. In an
embodiment, the bioreactor is a stand-alone membranous pouch
designed to be surgically implanted. Material composition: The
pouch housing is made of a semi-permeable membrane with a
pre-defined molecular weight cut-off designed to effectively
restrict movement of cells, but allow free transfer of paracrine
factors, nutrients, and waste. The membrane can be composed of a
wide spectrum of cellulosic (such as cellulose acetate) and
synthetic materials (such as polysulfone, polyamide,
polyacrylonitrile, and their copolymers, polymethylmetacrylate,
polytetrafluroethylene and their various derivatives, silicon
carbide, and micro-machined porous silicon diaphragms, among
others). Coatings: As long as membrane porosity is not disturbed,
the interior surface of the pouch can be coated with molecules
which enhance stem cell attachment and function. The exterior of
the pouch can be coated with anticoagulants to minimize thrombosis,
and/or other substances to improve deliverability. This adjustment
of coatings is well known to those of ordinary skill in the art.
Geometry: The physical shape of the pouch can be designed to
incorporate surface undulations and crevices to maximize surface
area for mass transfer. Bioreactor contents: Any number of stem
cells and/or other cell types or cells with and/or without genetic
alterations can be used in the device based, in part, on the type
of injury and organ that is being targeted for regeneration. The
cells can be used stand-alone or bathed in a media containing any
number of proteins, growth factors, or other molecules.
[0020] Catheter-mounted Membranous Pouch-based bioreactor. In an
embodiment the bioreactor is a pouch bioreactor, as described
above, which is attached to catheter tubing with ports connecting
the pouch lumen to the catheter exterior, allowing infusion,
sampling, and circulation of cells and media. In another
embodiment, multiple lumens within the housing of the catheter can
provide additional options for continuous circulation of stem cells
within the pouch, and an open distal port can be used as an
intravenous line or for central venous pressure measurement (as
illustrated in FIG. 3). Alternatively, the pouch bioreactor can be
mounted directly on a wire.
[0021] The catheter housing can made of polyvinyl chloride tubing
(or any other suitable biocompatible material) composed of a
multitude of lumens, proximal access ports and distal apertures,
similar to standard multi-lumen central venous catheters. These
ports can be used for removal and/or replacement of cells, media,
or other substances which may promote maintenance of the cells
and/or enhance their function. Coatings: The surface of the
catheter upon which the bioreactor is mounted and the portion of
which is inside of the bioreactor can be coated with molecules
which enhance stem cell attachment and function. Some of these
molecules include polylysine, fibronectin, or other proteins with
Arg-Gly-Asp sequences. This can be accomplished by first oxidizing
the catheter surface in a plasma reactor or with chemical oxidizing
agents such as potassium permanganate, then reacting the surface
with the appropriate molecular functional group. The exterior of
the catheter can also be spray- or dip-coated with an
anti-coagulant (such as heparin) to minimize thrombus formation
during and following implantation. Air-filled guiding balloon: A
small balloon (.about.1 cm in diameter) can be placed near the tip
of the catheter, which upon filling with air, can assist in guiding
intravascular placement. Guidewire directability is provided by
incorporation of a channel for the introduction of a steerable and
removable guidewire within the device. Bioreactor contents: Any
number of stem cells and/or other cell types or cells with and/or
without genetic alterations can be used in the device based on the
type of injury and organ that is being targeted for regeneration.
The cells could be used stand-alone or bathed in a media containing
any number of proteins, growth factors, or other molecules. The
device would also allow for slow infusion, intermittent recycling
or continuous circulation of cells, cell-conditioned media,
concentrated paracrine factors, or any other fluids that are
determined to have a beneficial effect. Removability: The
catheter-based device can be easily removed when desired. The pouch
housing is first evacuated by withdrawing its contents through the
infusion port. Vacuum is then created allowing the pouch to
collapse to a small profile facilitating removal. The entire device
can then be pulled out of the body. The vascular sheath is removed
and manual compression or vascular closure devices, if needed, can
be used to achieve hemostasis.
[0022] Secondary device-mounted pouch-based bioreactor. In an
embodiment, pouch bioreactors as described above can be attached to
various other secondary devices and implanted with the device. The
pouch bioreactors can be miniaturized as needed to attach to the
secondary device. In the cardiovascular arena, secondary devices
include, but are not limited to, stents, intra-aortic balloon
pumps, percutaneous and surgically implanted ventricular assist
devices, percutaneous and surgically implanted prosthetic valves
and valve clips or rings, endovascular grafts, thrombus filters,
pacemaker or defibrillator surfaces or leads, septal occluders,
atrial appendage closure devices, pulmonary artery catheters,
venous catheters, and arterial catheters, among others. Outside the
cardiovascular arena, any implantable device conferring access to a
target tissue is a possible candidate for attaching a pouch-based
bioreactor. These variations are evident to one of ordinary skill
in the art.
[0023] Encapsulant-based cell enclosure. In an embodiment, an
implantable encapsulant-based bioreactor consists of a matrix
encapsulating desired cells and media coated on a surface. The
matrix can be (a) coated directly to a tissue surface or (b) coated
on a separate implantable device. The matrix can be pre-formed and
stored for later use, or stored in its individual component
reactants and then prepared on site when needed. The matrix can be
composed of a number of polymers including but not limited to
polyethylene glycol (PEG), hyaluronic acid, chitosan, dextran,
collagen or self-assembling oligopeptides. Factors such as the
arginine-glycine-aspartic acid (RGD) oligopeptide can also be
incorporated within the matrix to assist in stem cell adhesion and
enhance proliferation and function. The matrix used will be porous
enough to allow free transfer of growth factors, nutrients and
wastes, while restricting mobility of stem cells. At its lower
dimensional limit, single cells can be encapsulated using this
method. For example, a hydrogel matrix, which is typically 90%
water, is porous enough to allow free transfer of growth factors
and other substances, yet will restrict mobility of stem cells.
Also arising from the high water content are the excellent
anti-fouling properties of the free surface.
[0024] Direct tissue coating. In an embodiment, matrices
encapsulating desired cells and media are directly coated onto a
tissue surface. This can be performed via either direct surgical
exposure of the target tissue, or through percutaneous injection
into the target tissue or into a surface or potential space (e.g.,
pericardial sac, peritoneal space) around the target tissue.
[0025] Coating onto an implantable device. In an embodiment,
matrices encapsulating desired cells and media are applied to a
secondary implantable device. Application of the matrix to the
device can be via dip-coating, spray-coating, spin-coating, or any
other method that achieves adequate adhesion of the matrix to the
implantable device. Coating may be performed long a priori or
immediately before device implantation. In the cardiovascular
arena, possible secondary implantable devices include, but are not
limited to the pouch-based bioreactor, a microfabricated cellular
enclosure, stents, intra-aortic balloon pumps, percutaneous and
surgically implanted ventricular assist devices, percutaneous and
surgically implanted prosthetic valves and valve clips or rings,
endovascular grafts, thrombus filters, pacemaker or defibrillator
surfaces or leads, septal occluders, atrial appendage closure
devices, pulmonary artery catheters, venous catheters, and arterial
catheters, among others. Outside the cardiovascular arena, any
implantable device conferring access to a target tissue is a
possible candidate for coating with the matrices.
[0026] Class II implantable bioreactor for local delivery. Local
delivery of paracrine factors confers dual advantages of direct
targeting of diseased tissue and reduction or prevention of any
systemic side effects. In an embodiment, an implanted device is
coated with cells using cell-specific ligands or antibodies as cell
anchors. In another embodiment, a cell micro-enclosure, constructed
using micro-electro-mechanical systems (MEMS) fabrication
technology, is disclosed, for implantation in small target tissues,
such as the intracoronary space. In another embodiment,
miniaturized forms of the bioreactors of Class I can be utilized
for local delivery when the bioreactor is deployed at the
appropriate target location.
[0027] Stand-alone cell coated device. Any implantable device
(including any of the embodiments described herein) can be coated
with a monolayer of a desired cell type possibly involving the use
of cell-specific ligands or antibodies as anchors. For stem cells,
anchoring ligands and antibodies include, but are not limited to
RGD oligopeptides and those oligopeptides containing the sequence
Ile-Lys-Val-Ala-Val, as well as anti-CD34, and anti-CD31. The
coated device can be prepared prior to implantation by immersing it
in a solution containing the appropriate cell type. In addition,
native cells may attach after implantation. For instance, if the
goal is myocardial healing and regeneration post infarction, an
intracoronary stent could be coated with stem cell-specific ligands
and antibodies and then immersed in a solution containing stem
cells before coronary implantation.
[0028] Cell coat and overcoat. A variation of this embodiment for
cells susceptible to immunologic attack would be to bind the cells
to the desired surfaces using appropriate ligands/antibodies as
described above and, in addition, layer above these cells an
overcoat of a semi-permeable matrix or a semi-permeable membrane.
The semi-permeable matrix or membrane would confer immunologic
isolation while allowing free permeation of bio-molecules, fluids,
nutrients, and waste.
[0029] Microfabricated cellular enclosure. Cell enclosures of
specific dimensions, pre-determined porosity and pore patterns can
be fabricated using microfabrication technology. These enclosures
can be constructed to be far smaller and with far more precise pore
characteristics than available semi-permeable membranes. The pores
can be designed to be impermeable to cells, yet freely permeable to
fluid, nutrients, paracrine factors and/or other bio-products,
wastes and nutrients. The enclosures can be coated with
cell-specific ligands, antibodies or other factors which are well
known to those of ordinary skill in the art to facilitate cell
adhesion and proliferation. The enclosures can be used stand-alone
or attached to a separate implantable device. In one embodiment,
the micro-enclosures are filled with stem cells and/or other cells
and are bonded to a coronary stent for myocardial healing and
regeneration after infarction.
[0030] In one embodiment, microporous cellular enclosures can be
fabricated from silicon wafers. Using photolithographic techniques,
multiple pores of specific diameters (e.g. 0.1 .mu.m to 15 .mu.m)
can be etched into the device layer of a silicon-on-insulator (SOI)
wafer. This is followed by creating cell cavities by etching
multiple large trenches into the handle wafer side. The handle
wafer trenches can be positioned directly beneath the device layer
pores using backside alignment. The buried oxide between the two
silicon surfaces can then be removed, releasing the silicon
semi-permeable membrane. The enclosure can then be coated with cell
specific ligands and antibodies. Finally, stem cells and/or other
cell types, drug, and growth media can then be placed in the
trenches before sealing the cavity. Repeated iterations of this
process can yield porous silicon enclosures (as illustrated in FIG.
4) that can then be bonded to a coronary stent.
[0031] Miniaturized bioreactor. The implantable bioreactors
described herein can all be used for local delivery with
appropriate miniaturization and placement. Pouch-based enclosures
could be miniaturized through micromachining. Thin, uniform
coatings of the matrix-encapsulated cells could be produced using
spray-coating or spin-coating. In another variation, direct
micro-injection of matrix-encapsulated cells into target diseased
tissue can be used as a strategy for local delivery.
EXAMPLES
[0032] The examples presented herein illustrate, but are not
intended to limit, the scope of protection being sought.
Example 1
Stem Cell Viability and Paracrine Factor Release from Miniature
Bioreactors In Vitro
[0033] Methods. To determine whether stem cells can survive in, and
produce and release paracrine factors from the cellulose acetate
semi-permeable membrane used in the implantable pouch-type
bioreactor prototype, we devised an experimental model using
miniature versions of the device chamber. In this model, short
segments of the semi-permeable membrane with 10.sup.6 kDa pores
were fashioned into tubes, and the ends secured with sterile
suture; in addition, limited experiments have been performed to
date using 10.sup.5 kDa pore membrane. One million human
mesenchymal stem cells (MSCs) or bone marrow mononuclear cells
(BMMNCs) were injected into the lumen of the chamber, and then the
entire device was submerged in cell culture media for incubation
(see FIG. 1). Samples of the media outside the device were taken at
24 hrs, 72 hrs, and 7 days for the 10.sup.6 kDa experiments. All
samples underwent measurement of 8 different paracrine factors
(PFs) by Quansys Q-Plex Human Angiogenesis Array (n=3-6). This
micro-array can simultaneously assess for the presence and quantity
of: Vascular Endothelial Growth Factor (VEGF), Hepatocyte Growth
Factor (HGF), Basic Fibroblast Growth Factor (bFGF), Interleukin-8
(IL-8), Tissue Metalloproteinase Inhibitors 1 and 2,
Platelet-Derived Growth Factor-BB (PDGF-BB), and Tissue Necrosis
Factor Alpha (TNF-.alpha.). Multiple 1 mL aliquots of the
conditioned media were collected at stored at -80 C until the time
of analysis. Results of media collected from mini-bioreactors
containing stem cells were compared to those from mini-bioreactors
not containing stem cells media as controls. In addition, at each
time point bioreactors were opened and cells within were removed
for cell viability assessment using trypan blue exclusion
(n=3).
[0034] Results: Paracrine Factor Production. Samples of cell
culture media collected from outside of 10.sup.6 kDa
mini-bioreactors containing MSCs showed substantial production of
VEGF, HGF, IL-8, TIMP-1, and TIMP-2 that increased with time.
Production of bFGF was detected at 24 h, but decreased over time.
The remaining factors, PDGF-BB, and TNF-.alpha. were not
detected.
[0035] [000361 Conditioned media samples taken from outside
mini-bioreactors containing BMMNCs showed production and release of
relevant paracrine factors as well, but in a pattern different from
that of the MSCs. BMMNCs showed increasing production of IL-8,
TIMP-1, and TIMP-2 over time, with production of IL-8 substantially
greather than that of the MSCs. TNF-.alpha. was detected as well,
though decreased over time. The factors VEGF, HGF, bFGF, and
PDGF-BB were not detected (see Table below) in conditioned media
taken from experiments using BMMNCs.
TABLE-US-00001 TABLE Stem Cell Viability and Paracrine Factor
Production from Miniature Bioreactors made using 10.sup.6 kDa
Cellulose Acetate Semi-Permeable Membrane. Average Analyte
Concentration Mean (pg/mL) PDGF- Cell Type t n Viability VEGF HGF
bFGF IL-8 BB TIMP-1 TIMP-2 TNF.alpha. MSCs 24 h 6 89.4% 112.6 0.0
60.9 8.6 0.0 7903.2 2118.5 0.0 SD 6.3 0.0 6.3 3.0 0.0 8818.0 595.7
0.0 72 h 6 79.1% 236.5 112.1 31.5 18.2 0.0 3339.3 7799.2 0.0 SD
57.7 25.2 5.3 4.2 0.0 587.7 2191.3 0.0 7 d 6 75.4% 676.4 348.0 29.5
97.5 0.0 89544.8 8165.1 0.0 SD 81.2 61.5 5.2 87.5 0.0 133399.7
2245.8 0.0 Control 24 h 3 n/a 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 (no
cells) SD 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 72 h 3 n/a 0.0 0.0 0.0
0.0 0.0 0.0 0.0 0.0 SD 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 7 d 3 n/a
0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 SD 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0
BMMNCs 24 h 6 78.4% 0.0 0.0 0.0 1497.7 0.0 327.3 0.0 76.9 SD 0.0
0.0 0.0 637.0 0.0 40.6 0.0 41.3 72 h 6 64.2% 0.0 0.0 0.0 4418.5 0.0
1063.1 84.8 49.2 SD 0.0 0.0 0.0 2712.4 0.0 296.6 10.6 13.4 7 d 3
61.5% 0.0 0.0 0.0 6855.7 0.0 1927.9 447.5 23.7 SD 0.0 0.0 0.0 647.5
0.0 201.3 79.1 2.5 Control 24 h 3 n/a 0.0 0.0 0.0 0.0 0.0 0.0 0.0
0.0 (no cells) SD 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 72 h 3 n/a 0.0
0.0 0.0 0.0 0.0 0.0 0.0 0.0 SD 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 7 d
3 n/a 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 SD 0.0 0.0 0.0 0.0 0.0 0.0
0.0 0.0
[0036] Stem Cell Viability. Assessment of MSC and BMMNC viability
in mini-bioreactors made of 10.sup.6 kDa cellulose acetate membrane
at 24 h, 72 h, and 7 d showed substantial cell viability with
89.4%, 79.1%, and 75.4% for MSCs, respectively, and 78.4%, 64.2%,
and 61.5% for BMMNCs (see Table above).
[0037] Summary. These experiments demonstrate that two types of
stem cells, MSCs and BMMNCs, survive for up to seven days in
culture in miniature bioreactor chambers made of the same 10.sup.6
kDa cellulose acetate membrane used for the catheter-based
implantable bioreactor prototype, and that during this time release
paracrine factors (PFs) relevant to angiogenesis and tissue repair.
The amount and type of PFs released by MSCs differ from those of
BMMNCs. It is yet unclear whether this represents an intrinsic
difference in the cell types, or a differential response to growth
within the cellulose acetate membrane. Overall, these results
suggest that the 10.sup.6 kDa cellulose acetate semi-permeable
membrane is compatible with MSC and BMMNC cell survival and PF
release, though there are differences in the combination of PFs
released by the . different cell types.
Example 2
Implantation of Bioreactor Prototype in Farm Pig
[0038] Methods. Bioreactor prototype devices were implanted in two
farm pigs (.about.25kg). Each device was fashioned from 10.sup.6
kDa cellulose acetate semi-permeable membrane secured to an 8F
clinical grade multi-lumen catheter. In both cases a surgical
cut-down procedure was performed over the right neck to expose the
external jugular vein (EJ) after initiation of general anesthesia
and endotracheal intubation. Using fluoroscopic guidance, a 14F
vascular sheath was inserted in the EJ over the wire. The
bioreactor device prototype was then passed through the sheath and
advanced to the junction of the superior vena cava (SVC) and right
atrium (RA). In the first experiment the device was left in place
for 1 hour, after which a suspension of 2.5.times.10.sup.6 human
MSCs was injected into the device. The device was then left in
place for another 90 minutes, during which continuous ECG
monitoring and intermittent fluoroscopic venography were performed,
the latter to determine whether there was any impediment to blood
return to the heart via the EJ/SVC. The device was then removed and
inspected and the animal euthanized using potassium chloride
infusion.
[0039] In the second experiment the bioreactor prototype was again
advanced to the SVC-RA junction, and a suspension of
2.times.10.sup.6 human MSCs were infused into the bioreactor
chamber via the catheter lumen. The bioreactor and vascular sheath
were then sutured in place and the surgical wound closed. The
animal was then allowed to awaken and the device was maintained in
place for 24 hours. The animal was then returned to the surgical
suite and general anesthesia induced. Fluoroscopic venography was
performed as previously and the device removed for inspection. The
animal was then euthanized as above.
[0040] Results. In both experiments, the prototype bioreactor was
well tolerated with no evidence of vascular, hemodynamic, or
arrhythmic compromise. Fluoroscopic venography showed no evidence
of vascular compromise or intra-vascular thrombus formation during
either the 2.5 hours the device was in place during the first
experiment, or after 24 hours in the second experiment. In the
second experiment small amounts of thrombus were found at both ends
of the bioreactor lumen, where the cellulose acetate membrane was
secured to the catheter shaft, but none on the membrane itself. The
addition of MSCs to the bioreactor lumen did not appear to affect
biocompatibility of the device.
[0041] Summary. These preliminary results suggest that the
prototype bioreactor device is well tolerated in vivo for up to 24
hours with no evidence of vascular, hemodynamic, or arrhythmic
compromise. The presence of a small amount of thrombus at the
attachment points of the bioreactor chamber to the vascular
catheter indicates the need for further optimization; however, the
cellulose acetate membrane itself appears biocompatible.
Example 3
Hydrogel for Stem Cell Encapsulation and Intravascular
Deployment
[0042] Hydrogel Crosslink Chemistry. We describe a method for
immobilizing stem cells on the surfaces of an implantable stent
using a biologically inert hydrogel coating. Polyethylene glycol
(PEG) was chosen to encapsulate the stem cells because of its
biocompatibility, antifouling properties, and permeability to
biomolecules. It is a neutral, water-soluble polymer that forms
hydrogels when crosslinked in the presence of water. Crosslinks
form in our proprietary system through the reaction between
N-hydroxysuccinimide (NHS)-activated esters and amine groups. The
two functional groups react to form an amide bond with the loss of
the NHS group. The reaction rate is greatest at a pH between 7 and
10 and is much slower at a pH below 6. This difference in reaction
rates allows for convenient handling of the liquid hydrogel
precursor at at pH of 6, while raising the pH slightly to
physiologic pH (7.4) results in immediate solidification. To form a
fully crosslinked gel, we mix a PEG molecule with two or more
activated ester groups (NHS-PEG) and a PEG molecule with 2 or more
amine groups (amine-PEG) to generate a 3-dimensional hydrogel
polymer network. The higher the average functionality, the more
rapidly a gel forms for any given number of crosslinks. A hydrogel
is formed by mixing an 8-arm polyethylene glycol molecule with 8
NHS-activated esters with a diamine-functionalized polyethylene
glycol. The average functionality of 5 for this system is chosen to
efficiently generate fully crosslinked gels.
##STR00001##
[0043] This NHS ester/amine reaction was chosen for its low
toxicity, biocompatibility, and the convenience of triggering the
reaction using a mild pH change. Other polyethylene glycol
hydrogels are more commonly formed by a photoinitiated radical
polymerization. Although stem cells have been shown to survive
photopolymerization, the procedure employs UV light to generate
free radicals, both of which are usually used to sterilize samples,
and hence may harm the stem cells.
[0044] Note that other polymerization chemistries are possible,
such as isocyanate/amine, epoxy/amine, isothiocyanate/amine,
alcohol/glutamate, and thiol/maleimide. Although most of these
reaction chemistries may be used in lieu of the current NHS
ester/amine reaction, many have higher toxicity, sub-optimal
reaction kinetics, and do not have the advantage of being able to
"switch" on and off with such a mild pH change.
[0045] Polymer Precursor Concentration. The hydrogels initially
contain between 5% and 50% organic solids by weight, with water
constituting the remainder. After the gel is formed, it expands to
many times its own weight, typically 4.times.-5.times., by soaking
up additional water. Water will continue to absorb until the gel
reaches equilibrium, which is usually around 5%-10% organic solids
by weight. For this system, higher initial solid fractions
facilitate more efficient crosslinking reactions between the
NHS-PEG and diamine-PEG. The greater reaction efficiency can be
seen by the finding that a higher initial solid fraction results in
a higher final solid fraction (FIG. 8). As will be described below,
denser, less swollen films have a higher crosslink density.
[0046] Hydrogel Crosslink Density. Generally speaking, the
equilibrium fraction of organic solids in a hydrogel
(.upsilon..sub.p) increases with increasing crosslink density
(.rho..sub.x, crosslinks per volume, given in mol/L), and decreases
with the molecular weight of the polyethylene glycol between
crosslinks (M.sub.c).
.rho. x = - 1 v _ ( ln ( 1 - .upsilon. p ) + .upsilon. p + .chi.
.upsilon. p 2 .upsilon. p 1 / 3 - .upsilon. p 2 / 2 )
##EQU00001##
[0047] where .upsilon. is the molar volume of water and .chi. is
the Flory-Huggins chi parameter.
.rho. x = 2 f av .rho. M c ##EQU00002##
[0048] where f.sub.av is the average functionality of the monomers,
and .rho. is the density. The same phenomenon is observed for
proprietary gels.
[0049] 1000521 The molecular weight of the polymer precursor is
typically chosen to control the crosslink density. Typical values
range from 1000 g/mol to 100,000 g/mol. Note that high molecular
weight leads to lower crosslink densities and lower stiffness. The
other mitigating factor is the reaction conversion. Low conversions
decrease the crosslink density, increase the effective molecular
weight between crosslinks, and decrease the elastic modulus.
[0050] Cell Adhesion. To improve cell viability, it is desirable to
allow cells to adhere to the hydrogel and pull themselves into a
state of tension. Viability frequently correlates with the
lens-like shape the cells take when they successfully form focal
adhesion sites with the surrounding matrix whereas a lack of
adhesion is characterized by a spherical cell morphology and
generally leads to apoptosis. For our proprietary recipe, the
addition of arginine-glycine-aspartine (RGD) oligopeptides
facilitates the formation of focal adhesion points through the
specific interactions between RGD and integrin. Typical
concentrations ranging from 1-20 mM are used to promote cell
adhesion.
[0051] Our formulation uses cyclo (Arg-Gly-Asp-d-Phe-Lys). The
additional lysine residue has a free amine group, which reacts with
the NHS-PEG to become incorporated into the hydrogel. The cyclic
RGD ring is not strictly required, but it confers greater stability
and selectivity over linear RGD peptides.
##STR00002##
[0052] Stent Adhesive Layer. The final step is to adjust the pH to
a level at which the nucleophilic addition reaction occurs at the
fastest rate. NHS-activated PEG tends to lower the pH to a level at
which the reaction occurs too slowly, so a small amount of base is
typically needed to raise the pH to 7.4, at which the reaction
proceeds rapidly to completion.
[0053] In our system, we designed an adhesive layer that
automatically adjusts the pH from 6 to 7.4. This convenient method
makes it possible to prepare the hydrogel precursor and add stem
cells and/or other cell types when it is a liquid, and then
solidify the gel only after it coats the stent, or another device.
The gel precursor solution containing the stem cells and/or other
cell types may be painted, sprayed, or dip-coated onto the stent or
another secondary device. Gelation typically occurs within ten
minutes of coming in contact with the gel precursor solution.
[0054] Consisting of poly(allylamine) (PAAM) and hexamethylene
diisocyanate (HMDI), the adhesive coating is formed by first
dipping the stent into a 5% solution of PAAM in water. The stent is
then dipped into a 5% solution of HMDI in isopropanol. The HMDI
rapidly reacts with the PAAM to form polyurea crosslinks. Since the
reaction does not proceed with precise stoichiometry, it will
always leave behind a small fraction of amine and isocyanate
groups. To ensure that the excess amine groups do not increase the
pH too strongly, the coated stent is soaked and rinsed in phosphate
buffered saline at a pH of 7.4 until the pH stabilizes at 7.4.
##STR00003##
[0055] The excess of basic amine groups in this adhesive coating
subsequently increase the pH of the hydrogel precursor solution to
initiate the polymerization. These amine groups are free to react
with NHS-PEG in the hydrogel precursor solution. HMDI serves a dual
function as well: it crosslinks the PAAM by forming polyurea bonds,
and any excess isocyanate groups are also free to crosslink with
diamine-PEG in the hydrogel precursor solution.
[0056] Recipe. To make 200 .mu.L of hydrogel, add 50 mg of
N-hydroxysuccinimide-activated polyethylene glycol with 8 activated
ester groups (8 arm NHS-PEG, MW=40 kg/mol) to 65 .mu.L of phosphate
buffered saline (PBS) at pH 7.4. Stir until dissolved. Excess
carboxylic acid groups that are invariably present on the 8 arm
NHS-PEG cause the pH to drop to about 6 or lower. Next, add 60
.mu.L of a 10 mg/mL solution of cyclo
(arginine-glycine-aspartine-d-phenylalanine-lysine) (RGD-lysine),
and 75 .mu.L of a 400 mg/mL solution of polyethylene glycol-diamine
(PEG-diamine, MW=2 kg/mol). Finally, the pH is adjusted with sodium
hydroxide to 7 to drive the polymerization to completion. Gelation
occurs shortly thereafter.
[0057] This mixture provides a hydrogel that is roughly 40% solids
by weight. It gives an RGD concentration of 5 mM, and uses a 1.5:1
ratio of amine groups to activated esters in order to achieve
gelation. Once placed in aqueous buffer solution such as PBS and
allowed to equilibrate, the hydrogel absorbs water and swells to
give a final composition that is between 5% and 10% solids by
weight.
Example 4
Microfabricated Cellular Enclosures: Detailed Materials &
Methods
[0058] Silicon-on-Insulator (SOI) wafers (4'' diameter with a 1-2
micron thick buried silicon oxide layer sandwiched between a
50-micron thick device layer and a 200-250 micron thick handle
layer) were cleaned using Piranha solution
(H.sub.2SO.sub.4/H.sub.2O.sub.2) and a de-ionized water rinse and
blow dried using nitrogen gas.
[0059] The handle layer (which is to contain the cell reservoir)
was processed first. The handle layer was patterned using
photoresist (AZ9260 to a thickness of 14-15 microns cured at
110.degree. C. for 3 to 4 min) so that the silicon can be etched
using a Surface Technology Systems Deep Reactive Ion Etcher (STS
DRIE) system. In addition, the surface of the device layer was
coated with a 1-2 micron thick layer of photoresist (Shipley 1800
cured at 100.degree. C. for 1 min) to provide added structural
strength so that the backside helium flow for the STS DRIE does not
cause rupture. Etching was performed completely through the handle
layer until the buried oxide layer (BOX) etch-stop was reached. The
wafer was then cleaned in acetone at room temperature for 30 to 60
min, sequentially soaked in isopropyl alcohol (IPA) and de-ionized
water, and dried in a convection oven at 95.degree. C. for 20 min.
The drying process minimized the chances of damaging the fragile
device layer.
[0060] The device layer (which is to contain the semi-permeable
diaphragm) was processed next. The device layer was patterned to
specific pore sizes and pore-to-pore dimensions with photoresist
(Shipley SC 1800, 2 microns thick, cured at 100.degree. C. for 2
min), taking care to match the photoresist pattern on the device
layer with the locations of the cell reservoirs on the handle layer
on the back-side. To further structurally protect the device layer
in processing, it was mounted to a secondary supporting silicon
wafer using liquid CrystalBond spun onto the secondary supporting
wafer at 2500 rpm and cured at 100-110.degree. C. for 30-45 sec.
After the bonded wafer was allowed to cool to room-temperature, the
device wafer layer was etched using the STS DRIE system and
intermittently cooled between batches of etching cycles to assure
continued structural integrity.
[0061] At completion of the device wafer etch, individual cellular
micro-enclosures were separated from the wafer. The entire wafer
was over-coated with photo-resist (Shipley SC 1800 series) from
protection, diced on a diamond saw, and soaked in acetone for 24-48
hours to separate the diced micro-enclosures from the secondary
supporting wafer. The individual micro-enclosures were then
thoroughly cleaned in repeated washes with heated acetone,
isopropyl alcohol and de-ionized water, followed by drying in a
convection oven for 15-20 min at 95.degree. C.
[0062] Finally, the intermediate buried oxide layer was removed by
dipping the micro-enclosures in 49% hydrofluoric acid. Cleaning of
the micro-enclosures was then performed in sequential washings in
de-ionized water, isopropyl alcohol, followed by final cleaning and
sterilization in oxygen plasma using a Trion RIE (Reactive Ion
Etcher).
* * * * *