U.S. patent application number 12/072232 was filed with the patent office on 2008-06-19 for tissue scaffold having aligned fibrils, apparatus and method for producing the same, and artificial tissue and methods of use thereof.
This patent application is currently assigned to University of South Carolina. Invention is credited to Edie C. Goldsmith, Richard L. Goodwin, C. Michael Gore, Louis Terracio, Michael J. Yost.
Application Number | 20080147199 12/072232 |
Document ID | / |
Family ID | 33568590 |
Filed Date | 2008-06-19 |
United States Patent
Application |
20080147199 |
Kind Code |
A1 |
Yost; Michael J. ; et
al. |
June 19, 2008 |
Tissue scaffold having aligned fibrils, apparatus and method for
producing the same, and artificial tissue and methods of use
thereof
Abstract
A tubular tissue scaffold is described which comprises a tube
having a wall, wherein the wall includes biopolymer fibrils that
are aligned in a helical pattern around the longitudinal axis of
the tube where the pitch of the helical pattern changes with the
radial position in the tube wall. The scaffold is capable of
directing the morphological pattern of attached and growing cells
to form a helical pattern around the tube walls. Additionally, an
apparatus for producing such a tubular tissue scaffold is
disclosed, the apparatus comprising a biopolymer gel dispersion
feed pump that is operably connected to a tube-forming device
having an exit port, where the tube-forming device is capable of
producing a tube from the gel dispersion while providing an angular
shear force across the wall of the tube, and a liquid bath located
to receive the tubular tissue scaffold from the tube-forming
device. A method for producing the tubular tissue scaffolds is also
disclosed. Also, artificial tissue comprising living cells attached
to a tubular tissue scaffold as described herein is disclosed.
Methods for using the artificial tissue are also disclosed.
Inventors: |
Yost; Michael J.;
(Lexington, SC) ; Gore; C. Michael; (West
Columbia, SC) ; Terracio; Louis; (New York, NY)
; Goodwin; Richard L.; (Columbia, SC) ; Goldsmith;
Edie C.; (Lexington, SC) |
Correspondence
Address: |
Marcia T. Greci
17th Floor, 1320 Main Street
Columbia
SC
29201
US
|
Assignee: |
University of South
Carolina
Columbia
SC
|
Family ID: |
33568590 |
Appl. No.: |
12/072232 |
Filed: |
February 25, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10861664 |
Jun 4, 2004 |
7338517 |
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12072232 |
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60475680 |
Jun 4, 2003 |
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60475866 |
Jun 4, 2003 |
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60475986 |
Jun 4, 2003 |
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Current U.S.
Class: |
623/23.72 ;
435/29; 435/395 |
Current CPC
Class: |
B29C 48/33 20190201;
C12M 21/08 20130101; B29C 48/09 20190201; C12M 25/14 20130101 |
Class at
Publication: |
623/23.72 ;
435/29; 435/395 |
International
Class: |
A61F 2/02 20060101
A61F002/02; C12Q 1/02 20060101 C12Q001/02; C12N 5/02 20060101
C12N005/02 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] The U.S. Government has a paid-up license in this invention
and the right in limited circumstances to require the patent owner
to license others on reasonable terms as provided for by the terms
of grant no. 1 K25 HL67097 awarded by the U.S. National Institutes
of Health of the Department of Health and Human Services.
Claims
1. A method for preconditioning a tubular artificial tissue
scaffold comprising seeding a tubular artificial tissue scaffold
having aligned biofibrils with living cells and culturing the cells
in the presence of media containing at least one growth factor and
under conditions where the tubular artificial tissue scaffold is
subjected to stretch and pressure pulse of controlled frequency and
amplitude.
2. A preconditioned artificial tissue comprising living cells that
are attached to a tubular artificial tissue scaffold having aligned
biofibrils, wherein the cells have been cultured in the presence of
media containing at least one growth factor and under conditions
where the tubular artificial tissue scaffold is subjected to
stretch and pressure pulse of controlled frequency and
amplitude.
3. A method of treatment comprising implanting in the body of a
subject artificial tissue comprising living cells attached to a
tubular tissue scaffold comprising a tube having a wall, wherein
the wall includes biopolymer fibrils that are aligned in a helical
pattern around the longitudinal axis of the tube where the pitch of
the helical pattern changes with the radial position in the tube
wall.
4. The method according to claim 3, wherein the biopolymer is
selected from the group consisting of collagen, fibronectin,
laminin, elastin, fibrin, proteoglycans, hyaluronan, and
combinations thereof.
5. The method according to claim 4, wherein the collagen is
selected from the group consisting of type I collagen, type II
collagen, type III collagen, type V collagen, type XI collagen, and
combinations thereof.
6. The method according to claim 3, wherein the living cells are
introduced to the luminal and outer surfaces of the substrate.
7. The method according to claim 3, wherein the living cells align
along the helical pattern of the biopolymer fibrils.
8. The method according to claim 3, wherein the living cells are
selected from the group consisting of myocyte precursor cells,
cardiac myocytes, skeletal myocytes, satellite cells, fibroblasts,
cardiac fibroblasts, chondrocytes, osteoblasts, endothelial cells,
epithelial cells, embryonic stem cells, hematopoetic stem cells,
neuronal cells, mesenchymal stem cells, anchorage-dependent cell
precursors, and combinations thereof.
9. The method according to claim 8, wherein the living cells
originate from the subject receiving treatment.
10. The method according to claim 3, wherein the living cells are
cultured in vitro for a period of time necessary for the cells to
degrade the original scaffold biopolymer and regenerate a matrix of
secreted extracellular matrix proteins.
11. The method according to claim 3, wherein the living cells
establish intercellular and extracellular connections.
12. The method according to claim 11, wherein the living cells are
contractile.
13. The method according to claim 12, wherein the contractile cells
are cardiac myocytes.
14. The method according to claim 13, wherein the cardiac myocytes
contract synchronously along the helical pattern of the biopolymer
fibrils to pump fluid through the lumen of the tubular tissue
scaffold.
15. The method according to claim 12, wherein the contractile
cardiac myocyte-containing artificial tissue is preconditioned by
the application of mechanical stress before being implanted in the
subject.
16. The method according to claim 3, wherein treatment is for an
injury, disease, or disorder selected from the group consisting of
congestive heart failure, dilated cardiomyopathy, hypertrophic
cardiomyopathy, infiltrative cardiomyopathy, ischemic heart
disease, heart attack, heart failure, coronary artery disease,
atherosclerosis, hypertension, chronic renal disease,
cerebrovascular disease, carotid artery disease, and peripheral
vascular disease.
17. The method according to claim 12, wherein the rate of
contraction is controlled by a pacing device.
18. The method according to claim 3, wherein the tubular artificial
tissue is used to repair or replace a tube-shaped organ.
19. The method according to claim 18, wherein the tube-shaped organ
is selected from the group consisting of blood vessels, coronary
arteries, renal arteries, ureters, fallopian tubes, and nerve fiber
conduits.
20. The method according to claim 3, wherein the tubular tissue
scaffold is split longitudinally and opened into a sheet prior to
attaching living cells to the tissue scaffold.
21. The method according to claim 20, wherein the sheet is coated
with a layer of biopolymer fibrils, wherein the biopolymer fibrils
polymerize and are oriented in the direction in which they are
applied.
22. The method according to claim 21, wherein the living cells are
selected from the group consisting of myocyte precursor cells,
cardiac myocytes, skeletal myocytes, satellite cells, fibroblasts,
cardiac fibroblasts, chondrocytes, osteoblasts, endothelial cells,
epithelial cells, embryonic stem cells, hematopoetic stem cells,
neuronal cells, mesenchymal stem cells, anchorage-dependent cell
precursors, and combinations thereof.
23. The method according to claim 22, wherein the living cells are
selected from the group consisting of myocyte precursor cells,
cardiac myocytes, skeletal myocytes, satellite cells, fibroblasts,
and combinations thereof.
24. The method according to claim 22, wherein the living cells
originate from the subject receiving treatment.
25. The method according to claim 20, wherein treatment is for
hernia, heart attack, congenital heart defects, skin burns, organ
damage, and muscle damage.
26. A method of identifying the effects of a pharmaceutical
composition on cell function comprising administering said
pharmaceutical composition in vitro to artificial tissue comprising
living cells attached to a tubular tissue scaffold comprising a
tube having a wall, wherein the wall includes biopolymer fibrils
that are aligned in a helical pattern around the longitudinal axis
of the tube where the pitch of the helical pattern changes with the
radial position in the tube wall.
27. The method according to claim 26, wherein the living cells are
contractile cardiac myocytes.
28. The method according to claim 27, further comprising
determining the effects of the pharmaceutical composition on the
living cells.
29. The method according to claim 26, wherein the tubular
artificial tissue has been split longitudinally and opened into a
sheet.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a divisional application of co-pending
U.S. Non-Provisional patent application Ser. No. 10,861,664, filed
Jun. 4, 2004, which is related to and claims the priority benefit
of U.S. Provisional Application Ser. Nos. 60/475,680; 60/475,866;
and 60/475,986, filed Jun. 4, 2003, and each of which is
incorporated herein by reference in its entirety.
BACKGROUND OF THE INVENTION
[0003] (1) Field of the Invention
[0004] The present invention relates to a tubular tissue scaffold,
and, in particular, to a tubular tissue scaffold having aligned
biopolymer fibrils, for use in tissue engineering applications; an
apparatus and method of producing a tubular tissue scaffold having
aligned biopolymer fibrils; and artificial tissue, and methods of
use thereof, comprising living cells attached to a tubular tissue
scaffold having aligned biopolymer fibrils.
[0005] (2) Description of the Related Art
[0006] The National Science Foundation defines tissue engineering
as "the application of principles and methods of engineering and
life sciences to obtain a fundamental understanding of
structure-function relationships in novel and pathological
mammalian tissues and the development of biological substitutes to
restore, maintain, or improve [tissue] function." See Shalak, R.
and Fox, J. eds., Tissue Engineering, Proceedings of a Workshop
held at Granlibakken, Lake Tahoe, Calif., Feb. 26-29, 1988, New
York: Alan Liss (1988). In the last decade, over $3.5 billion
dollars has been invested worldwide in tissue engineering research.
More than 70 start-up companies or businesses having a combined
annual expenditure of over $600 million dollars now participate in
a significant engineering and scientific effort toward developing
alternative sources of transplant materials through in vitro tissue
engineering.
[0007] Several aspects of creating an engineered tissue make it a
daunting task. One of the most difficult challenges is directing
the behavior of specialized cells outside of the body to mimic the
normal, endogenous phenotype those cells exhibit in vivo.
Additionally, in order for an engineered tissue to be tolerated
upon implantation, the material that provides the scaffolding for
the cells must meet several important criteria. The material must
be biocompatible, so as not to be toxic or injurious, and not cause
immunological rejection. Also, the material must be biodegradable,
by having the capability of being broken down into innocuous
products in the body. Because cells respond biologically to the
substrate on which they adhere, the materials that provide the
growth surface for engineered tissues must promote cell growth.
Further, the scaffolding material should be replaced by
extracellular matrix components secreted by the grafted cells as
the scaffold is broken down in the body. Additionally, the material
should allow cells to grow and function as they would in vivo.
[0008] Initially, researchers adapted synthetic degradable
polyesters that had been used in surgical materials since the early
1970s to construct scaffold materials for use in tissue
engineering. These degradable polyesters include polyglycolide and
polylactide, as well as the more recently developed polymer,
polylactide coglycolide. However, those degradable polyesters
tended to be inflexible, and their degradation in vivo has been
associated with adverse tissue reactions. These shortcomings have
led to the development of a host of new synthetic polymers, for
example, polyhydroxybutyrate and copolymers of hydroxybutyrate with
hydroxyvalerate. See Amass, W. et al., Polymer Int, 47:89-144
(1998).
[0009] In animals, collagens make up a majority of endogenous
scaffolding materials. They are the most commonly occurring
proteins in the human body and they play a central role in the
formation of extracellular matrix. Collagens are triple-helical
structural proteins. It is this triple-helical structure that gives
collagens the strength and stability that are central to their
physiological role in the structure and support of the tissues in
the body. Although there are over twenty types of mammalian
collagens, collagen types I, II, III, V, and XI make up the fibrous
collagens. Type I collagen molecules polymerize into fibrils which
closely associate in a parallel fashion to form fibers with
enormous tensile strength, which are found in skin, tendon, bone
and dentin. Type II is the major collagen found in cartilage, where
the fibrils are randomly oriented to impart both stiffness and
compressibility to the proteoglycan matrix. Type III collagen is
found in skin, muscle, and vascular structures, frequently together
with type I collagen.
[0010] Collagen has been used successfully in several tissue
engineering applications. As a copolymer with glycoaminoglycans
such as chondroitin 6-sulfate, collagen has been utilized as an
artificial skin scaffold to induce regeneration in vivo for the
treatment of burn injury since the early 1980s. See Burke, J. F.,
et al., Ann Surg, 194:413-28 (1981); Yannis, I. V., et al.,
Science, 215:174-6 (1982). Collagen has also been used to form
anti-adhesion barriers for use on surgical wounds. See U.S. Pat.
Nos. 5,201,745 and 6,391,939 to Tayot et al.; U.S. Pat. No.
6,451,032 to Ory et al. Zilla et al., in U.S. Pat. No. 6,554,857,
describe the use of collagen, among other materials, as a component
in a concentric multilayer ingrowth matrix that can have a tubular
form.
[0011] However, some problems remain to be solved in the use of
collagen as a scaffolding material for applications requiring
structural and mechanical stability, such as for vascular
prosthetics. This is at least partly due to an inability to isolate
collagen possessing the physical properties required to maintain
necessary mechanical integrity of a scaffold, as it is remodeled in
vivo, for use in, for example, cardiovascular indications.
Additionally, in order for a tubular construct to mimic endogenous
components of the cardiovascular system, it must promote the proper
growth, orientation, association, and function of specialized cell
types.
[0012] Despite significant work in the field of tissue engineering
and the numerous synthetic biomaterials that have been developed in
the last decade, there is still a need for improved scaffolding
materials for use in specialized applications. It would be useful,
therefore, to provide a tissue scaffold in the form of a tube,
comprising a biopolymer which promotes maintenance of an in vivo
cell phenotype and, particularly, a tubular tissue scaffold that
was non-toxic, biologically degradable in vivo, and causes little
or no immune reaction in a host. It would also be useful to provide
an apparatus and method for the production of such a tubular tissue
scaffold. Also, despite the advances in biomaterials research, and
the elucidation of the molecular biology of cell behavior and
cell:matrix interactions, the gap between in vitro engineered
tissue and biologically functional implantable organs remains
significant. Therefore, it would be useful to provide artificial
tissue that can act as a functional prosthetic. It would also be
useful if the artificial tissue could be produced in the form of a
tube, utilizing the tissue scaffolding described herein. This
structural configuration would be particularly useful in
cardiovascular applications.
SUMMARY OF THE INVENTION
[0013] Briefly, therefore, the present invention is directed to a
novel tubular tissue scaffold comprising a tube having a wall,
wherein the wall includes biopolymer fibrils that are aligned in a
helical pattern around the longitudinal axis of the tube where the
pitch of the helical pattern changes with the radial position in
the tube wall.
[0014] The present invention is also directed to a novel apparatus
for producing a tubular tissue scaffold having aligned biopolymer
fibrils, the apparatus comprising a biopolymer gel dispersion feed
pump that is operably connected to a tube-forming device having an
exit port, where the tube-forming device is capable of producing a
tube from the gel dispersion while providing an angular shear force
across the wall of the tube, and a liquid bath located to receive
the tubular tissue scaffold from the tube-forming device.
[0015] The present invention is also directed to a novel method of
producing a tubular tissue scaffold, the method comprising:
[0016] providing a gel dispersion comprising a biopolymer;
[0017] feeding the gel dispersion to a tube-forming device that is
capable of producing a tube from the gel dispersion while providing
a radial shear force across the wall of the tube and having a gas
channel connecting a gas source with the luminal space of the
tubular tissue scaffold as it exits the tube-forming device;
[0018] forming the gel dispersion into a tube; and
[0019] causing the gel dispersion to solidify, thereby forming a
tubular tissue scaffold comprising a tube wall having biopolymer
fibrils that are aligned in a helical pattern around the
longitudinal axis of the tube and where the pitch of the helical
pattern changes with the radial position in the tube wall.
[0020] The present invention is also directed to novel artificial
tissue comprising living cells attached to a tubular tissue
scaffold comprising a tube having a wall, wherein the wall includes
biopolymer fibrils that are aligned in a helical pattern around the
longitudinal axis of the tube where the pitch of the helical
pattern changes with the radial position in the tube wall.
[0021] The present invention is also directed to a novel method of
preconditioning artificial tissue for implantation into the body of
a subject, the method comprising: seeding a tubular artificial
tissue scaffold having aligned biofibrils with living cells and
culturing the cells in the presence of media containing at least
one growth factor and under conditions where the tubular artificial
tissue scaffold is subjected to stretch and pressure pulse of
controlled frequency and amplitude.
[0022] The present invention also includes a novel preconditioned
artificial tissue comprising living cells that are attached to a
tubular artificial tissue scaffold having aligned biofibrils,
wherein the cells have been cultured in the presence of media
containing at least one growth factor and under conditions where
the tubular artificial tissue scaffold is subjected to stretch and
pressure pulse of controlled frequency and amplitude.
[0023] Additionally, the present invention provides a method of
treatment comprising implanting in the body of a subject artificial
tissue comprising living cells attached to a tubular tissue
scaffold comprising a tube having a wall, wherein the wall includes
biopolymer fibrils that are aligned in a helical pattern around the
longitudinal axis of the tube where the pitch of the helical
pattern changes with the radial position in the tube wall.
[0024] The present invention also encompasses a method of
identifying the effects of a pharmaceutical composition on cell
function comprising administering said pharmaceutical composition
in vitro to artificial tissue comprising living cells attached to a
tubular tissue scaffold comprising a tube having a wall, wherein
the wall includes biopolymer fibrils that are aligned in a helical
pattern around the longitudinal axis of the tube where the pitch of
the helical pattern changes with the radial position in the tube
wall.
[0025] Among the several advantages found to be achieved by the
present invention, therefore, may be noted the provision of a
tubular tissue scaffold having sufficient structural strength to
withstand pressure; the provision of a scaffold having a
composition that is biologically degradable in vivo and will result
in a minimal immunological response from a host; the provision of a
tissue scaffold having a structural composition that allows the
penetration of cells and provides cells the requisite signals to
develop an in vivo functional phenotype; the provision of an
apparatus and a method for the production of such a tubular tissue
scaffold, and; the provision of artificial tissue comprising cells
attached to a tissue scaffold of biopolymer fibrils that is
non-immunogenic, has a construction that mimics that of cardiac
tissue, and has improved structural integrity.
BRIEF DESCRIPTION OF THE DRAWINGS
[0026] FIG. 1 is an illustration showing a partial cross-sectional
cutaway view of a tubular tissue scaffold of the present invention
illustrating helical alignment patterns of different direction of
biopolymer fibrils at the luminal surface (left-hand pitch) and
outside surface (right-hand pitch) of the tube (A), and also
showing three side views of tubular tissue scaffolds showing that
the pitch of the helical alignment pattern of fibrils at the
luminal surface (dotted line having an angle with the longitudinal
axis of .alpha..sub.L) and outside surface (solid lines having an
angle with the longitudinal axis of .alpha..sub.O) of the tube can
be approximately equal, but opposite in direction (B), or unequal
in pitch, but opposite in direction (C), and include the case where
the biopolymer fibrils are aligned at zero pitch
(.alpha..sub.O=0.degree.), at the outside wall in this case, while
the biopolymer fibrils at another location in the tube wall (here,
the luminal surface of the tube wall) are aligned in a right-hand
pitch of angle .alpha..sub.L (D); and
[0027] FIG. 2 shows scanning electron micrographs of the luminal
and outside surfaces of a tubular scaffold of the present invention
in which the longitudinal axis (along the center of the tube) is
roughly vertical between the two images and where panel (A) shows
the alignment of the collagen fibrils on the outside of the tube
wall and panel (B) shows the reverse pitch orientation of the
collagen fibrils on the luminal surface of the tube wall;
[0028] FIG. 3 is a sectional view illustrating the major parts of a
counter rotating cone extruder of the present invention;
[0029] FIG. 4 is a schematic illustration of the major parts of an
apparatus for producing a tubular tissue scaffold of the present
invention;
[0030] FIG. 5 shows a laser scanning confocal micrograph (A) of
cardiac myocytes on a collagen scaffold of the present invention,
stained with antibodies to F-actin and connexin 43, illustrating
expression of connexin 43 along the sides and ends of the myocytes.
Panels (B) and (C) are a stereo pair of laser scanning confocal
micrographs of cardiac myocytes on a collagen scaffold of the
present invention stained with an antibody to F-actin (Alexa 488
phalloidin). Multiple layers of cells with aligned fibrils can be
seen when viewing the images with stereo viewing equipment;
[0031] FIG. 6 shows transmission electron micrographs of cardiac
myocytes on a collagen scaffold of the present invention. The
micrographs demonstrate the presence of Z-bands (Z), aligned
microfibrils (F), numerous mitochondria (M), cell:cell junctions
(arrows in inset A), and interaction with collagen in the tube wall
(arrows in inset B);
[0032] FIG. 7 shows a representative electrical signal obtained
from a tubular scaffold of the present invention seeded with
neonatal cardiac myocytes in panel (A), and Fast Fourier transform
analysis of the electrical data in panel (B); and
[0033] FIG. 8 shows a schematic illustration of a system including
a pulsatile internal flow stretch bioreactor wherein the tubular
tissue scaffold (tissue engineered tube) is attached at each end to
supply and effluent tubing and media is fed through the tubing. A
basal level of pressure is provided due to the height of the media
reservoir and pulsed flow to cyclically stretch the cells. The
bioreactor stretches the tubular scaffold both radially and
longitudinally.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0034] In accordance with the present invention, it has been found
that a tubular tissue scaffold can be produced that comprises a
tube wall which includes biopolymer fibrils that are aligned in a
helical pattern around the longitudinal axis of the tube where the
pitch of the helical pattern changes with the radial position in
the tube wall, using a novel apparatus and method of producing such
a tubular tissue scaffold. It has also been found that artificial
tissue comprising living cells attached to the novel tubular tissue
scaffold described herein can be produced and used to treat a
variety of disease conditions by implanting the artificial tissue
in the body of a subject.
[0035] The novel scaffold is unique in that it is not a very soft
gel with little mechanical integrity, nor is it a tough highly
crosslinked material resembling leather. In addition, toxic or
potentially cytotoxic chemicals are not required to crosslink the
scaffold material. Rather, the novel scaffold is made as an
extruded tube with aligned biopolymer fibrils that impart
biochemical and biomechanical information to cells that attach to
them, which instruct the cells to adopt an in vivo-like
phenotype.
[0036] The novel tissue scaffold provides the advantages that it
has sufficient structural strength to maintain its tubular
conformation without a support such as a mandrel, and can withstand
at least about 12 mmHg internal pressure and in preferred forms can
withstand pressures of at least about 280 mmHg, or more. The
scaffold has a composition that will result in a minimal
immunological response from a host; and the scaffold has a
structural composition that allows penetration by living cells and
provides cells the requisite signals to develop an in vivo
functional phenotype.
[0037] From a mechanical standpoint, the properties provided by
biopolymers that form fibrils give the scaffold a suitable modulus,
suitable flexural rigidity, and a surface that is not too
hydrophobic or too hydrophilic to support the cells and fluid
necessary to form a tissue. Moreover, the biopolymers are
biodegradable, nontoxic and support the maintenance and synthesis
of new tissue. Furthermore, the present scaffold material will
biodegrade at a controlled rate to allow new tissue to invade.
[0038] Pore size and distribution have been reported as important
parameters for scaffold characterization and efficacy. See
Dagalakis, N. et al., J. Bio Med Matl Res. 14(4):511 (1980), and
Zeltinger, J. S. et al, Tissue Engineering. 7(5):557 (2001). The
present tissue scaffold material is porous, with most pores between
1 and 10 .mu.m in diameter. Although the average size of cells such
as neonatal rat cardiac myocytes, for example, is 15-18 .mu.m, it
is believed that pore size is not a limiting factor that prevents
myocyte penetration into the present scaffold material, since cells
can remodel and create an appropriate extracellular matrix material
if placed in the right environment. Pore size in the novel tissue
scaffold material is discussed further below.
[0039] The present tubular tissue scaffolds provide biopolymer
fibrils that are aligned in the tube wall in a helical fashion
around the longitudinal axis of the tube. It is important, however,
that in the present scaffolds the pitch of the helical pattern of
the aligned fibrils changes as the radial location in the tube wall
progresses from the inside (luminal) surface of the tube wall to
the outside (or, outer) surface of the tube wall. In some
embodiments, this pattern of fibril alignment can be characterized
as being in a "corkscrew" pattern at the luminal wall of the tube
and changing to a "counter-corkscrew" pattern at the outside wall
of the tube. In other words, if fibril alignment on the luminal
surface is in a right-hand helical pitch, the fibril alignment on
the outside surface of the tube would be in a left-hand pitch, or
vice-versa. This pattern mimics the extracellular matrix pattern of
heart tissue, and, in one embodiment, allow the construction of
more biologically-similar vascular constructs and heart constructs.
By way of example, when the tissue scaffold is used to support the
growth of neonatal cardiac myocytes, the myocytes can be made to
constrict simultaneously, thereby twisting the tube while reducing
its diameter, and thereby "wringing" fluid from the tube in a
motion similar to that of a biological heart.
[0040] The present tubular tissue scaffold differs from previous
constructs, such as those described in U.S. Pat. No. 6,540,780 to
Zilla et al., which describes synthetic vascular grafts with
helically oriented ingrowth channels within the tube wall, and U.S.
Pat. No. 6,554,857 to Zilla et al., in that the present tubular
tissue scaffolds comprise biopolymer fibrils that are aligned in a
helical pattern within the tube wall. In Zilla et al. '780, only
synthetic polymers are used, and both of these patents describe
constructs where it is the ingrowth channels that are aligned
within the tube wall, rather than biopolymer fibrils. In order to
obtain the present tubular scaffolds, it is necessary to form the
tube wall in a manner that imposes an angular shear across the tube
wall as it is being formed, as is described below. The present
tubular scaffolds, in fact, are substantially free of helically
oriented channels within the tube walls.
[0041] When discussing the geometry of a tubular scaffold herein,
the axis along the center of a cylindrical tube will be referred to
as the longitudinal axis. The axis that is perpendicular to the
longitudinal axis and runs outward from the center of the tube in a
direction that is perpendicular to the tube wall is referred to as
the radial axis. The helical pattern of the aligned biopolymer
fibrils within the tube wall is described in terms of the "pitch"
of the helix. As in the description of screw threads, for example,
the pitch of the helix can be "right hand" or "left hand". The
pitch of the helical pattern formed by the aligned fibrils is
described in terms of the angle (.alpha.) between a tangent to the
helix and a projection of the longitudinal axis of the tube (as
shown, for example, in FIG. 1, of the present specification). The
pitch of the helical pattern of the biopolymer fibers at the
luminal (internal) wall of the tube is designated as .alpha..sub.L
and the pitch of the helical pattern of the biopolymer fibers at
the outside wall of the tube is designated as .alpha..sub.O.
[0042] When it is said that the biopolymer fibrils are "aligned",
it is meant that most of the fibrils in the same radial plane in a
tube wall run roughly parallel to each other. It is not meant that
every fibril must be parallel to every other fibril in the plane,
but that a general alignment pattern must be discernable. Such
fibril alignment is shown, for example, in FIG. 2. The fibrils of
the present tubular tissue scaffold are preferably not isolated
from each other in the wall of the tube, but, rather, are
associated in sheets or areas containing dozens, if not hundreds or
thousands of fibrils that are adjacent or touching and are
associated and a part of the helical pattern of alignment.
[0043] The term "fibrils" is used herein to describe an association
of several biopolymer molecules into a structure that appears
fibrous with suitable magnification, as is typical for collagen and
other selected biopolymers of biological origin. There is no
particular size limit to the fibrils of the present invention, and
both fibrils and small fibers are included in the term.
[0044] As used herein, the term "biopolymer" refers to a natural or
synthetic polymer that is biologically compatible. Polymers that
are biologically compatible are those which can be implanted into a
living vertebrate subject without triggering a severe adverse
immune reaction. Examples of biopolymers that can be used in the
present invention include, but are not limited to, collagen,
fibronectin, laminin, elastin, fibrin, proteoglycans, hyaluronan,
and combinations thereof. In some embodiments, the biopolymer is a
collagen, selected from the group consisting of type I collagen,
type II collagen, type III collagen, type V collagen, type XI
collagen, and combinations thereof. In one embodiment, the collagen
is selected from the group consisting of type I collagen, type III
collagen, and combinations thereof. In preferred embodiments, the
biopolymer is type I collagen.
[0045] Biopolymers that have not been isolated or purified to some
degree from their natural sources, however, are not included in the
scope of the present invention. In other words, the present tubular
tissue scaffold is not meant to include natural vessels, arteries,
or other natural biological tubular structure.
[0046] Type I collagen is the most prevalent structural
extracellular matrix protein in the human body. Collagen has the
property of being able to direct cell behavior by signaling the
cells to modify their growth, differentiation, intercellular
contacts, and production of molecules such as collagen and other
extracellular matrix proteins and cytokines. Additionally, cells
are capable of recognizing and correctly modifying the collagen
matrix to conform to their cellular requirements. These features
allow cells introduced to a collagen scaffold to mimic their normal
in vivo phenotype and organization. By way of example, cells such
as myocytes introduced to the surfaces of the tubular scaffold
would readily attach to the collagen, remodel the matrix, and
develop intercellular connections.
[0047] The tubular tissue scaffold of the present invention
comprises a tube having defined dimensions such as an outside
diameter, a luminal diameter, and a wall thickness. The novel
tubular tissue scaffold can have any dimensions and is not limited
to any particular diameter or wall thickness.
[0048] In one embodiment, the outside diameter of the tubular
tissue scaffold is between about 0.1 millimeter and about 100
millimeters. In some embodiments, the outside diameter is between
about 0.5 millimeters and about 50 millimeters, or between about 1
and about 10 millimeters, or between about 4 and about 10
millimeters. In preferred embodiments, the outside diameter is
about 5 millimeters.
[0049] The luminal diameter of the tubular tissue scaffold is
smaller than the outside diameter of the tube and may be between
about 0.1 millimeters and about 49 millimeters. In one embodiment,
the luminal diameter is between about 0.4 millimeters and about 49
millimeters. In preferred embodiments, the luminal diameter is
between about 4 millimeters and about 5 millimeters.
[0050] The thickness of the tube wall, along with the size of the
lumen, determines the rate of diffusion of nutrients that are
critical for cell growth from the outside of the tube to the
luminal surface. The tube wall can be of any thickness that will
provide properties that are suitable for the intended application.
In one embodiment, the tube wall thickness is between about 0.05
millimeters and about 10 millimeters. In another embodiment, the
wall thickness is between about 0.1 millimeters and about 5
millimeters. In preferred embodiments, the wall thickness is
between about 0.1 millimeters and about 1 millimeter.
[0051] As has been described briefly above, the tube wall of the
tubular tissue scaffold comprises a helical alignment of biopolymer
fibrils, the alignment having a certain pitch in relation to the
longitudinal axis of the tube. A purpose of the particular
alignment of biopolymer fibrils of the present invention is to
allow for the generation of a pumping, or "wringing" type of
mechanical force in a tubular scaffold construct, given the
introduction of the proper cells to the scaffold.
[0052] In the present invention, the pitch of the helical pattern
of biopolymer fibrils on the luminal surface can be between about 0
degrees and about 89 degrees and the pitch of the helical pattern
on the outer surface is between about 0 degrees and about 89
degrees, where the pitch of the fibrils at the luminal surface is
different from the pitch of the fibrils at the outer surface of the
tube. In one embodiment, the pitch of the helical pattern of the
fibrils at the luminal surface is between about 18 degrees and
about 62 degrees and the pitch of the helical pattern of the
fibrils at the outer surface is between about 18 degrees and about
62 degrees, where the pitch of the fibrils at the luminal surface
is different than the pitch of the helical pattern of the
biopolymer fibrils at the outer surface. In preferred embodiments,
the pitch of the helical pattern of the biopolymer fibrils at the
luminal surface is between about 26 degrees and about 60 degrees
and the pitch of the helical pattern of the biopolymer fibrils at
the outer surface of the tube is between about 26 degrees and about
60 degrees, where the pitch of the fibrils at the luminal surface
is different from the pitch of the fibrils at the outer surface. In
another embodiment, the pitch of the helical pattern of the
biopolymer fibrils changes in a linear manner through the thickness
of the tube wall, and the change can be from a right-hand pitch to
a left-hand pitch, or vice versa.
[0053] In one embodiment of the present invention, the tubular
tissue scaffold has pores. One feature of pore size is the effect
that it has on the ability of cells to invade the matrix. In
preferred embodiments of the present invention in which a
biopolymer, such as collagen, for example, is used, cells
introduced to the scaffold are able to remodel the material as they
would in vivo, enabling the cells to infiltrate the matrix. Pore
size must also provide sufficient permeability for the diffusion of
nutrients from the media in in vitro culture conditions. In
preferred embodiments of the present invention, the size of the
pores is from about 1 micron to about 20 microns. More preferably,
the size of the pores is from about 2 microns to about 10 microns.
Unless explicitly described to the contrary herein, pore size is
expressed in terms of number average pore size.
[0054] In some embodiments of the tubular tissue scaffold of the
present invention, the luminal surface and outer surface of the
tube can support cell attachment and growth. In preferred
embodiments, the tissue scaffold can direct the morphology of
attached cells to align along the helical pattern of the
biopolymer.
[0055] In addition to growth and attachment of cells at the luminal
surface and/or the outer surface of the tube, it is sometimes
desirable that growing cells penetrate and grow into the tube wall.
In other words, that growing cells grow throughout the depth or
thickness of the tissue scaffold material. An advantage of the
present tissue scaffold is that cells are able to penetrate and
grow throughout the entire thickness of the tube wall, and in
preferred embodiments, the growing cells align with the general
alignment of the biopolymer fibrils throughout the thickness of the
wall. This provides for a construct of artificial tissue scaffold
having changing alignment of biopolymer fibrils throughout the tube
wall thickness and also having growing cells throughout the tube
wall thickness that match the alignment of the biopolymer fibrils.
This structure permits closer matching of the physical structure of
the artificial tissue with that of normal cardiovascular
tissue.
[0056] In one embodiment, the tissue scaffold of the present
invention can be treated with UV radiation to crosslink the
biopolymer fibrils making up the tissue scaffold to increase the
strength of the scaffold, as measured by increased burst pressure.
As used herein, the term "burst pressure" refers to the maximum
amount of fluid pressure which may be applied to the tubular tissue
scaffold internally without causing a rupture. Treatment with
radiation will be further described below, but the inventors have
found that treatment of tubular tissue scaffolds with UV radiation
increases the strength of the artificial tissue scaffold. When the
artificial tissue scaffold is in the form of a tube, its strength
can easily be measured and expressed in terms of the "burst
pressure". As used herein, the terms burst pressure mean the
internal pressure at which a tube bursts. Treatment of a present
tubular artificial tissue scaffold with UV radiation at a
wavelength that is from about 250 to about 280 nm and an energy
density of from about 100 to about 1000 .mu.w/cm.sup.2, is
preferred. By way of example, treatment with UV radiation at a
wavelength of 254 nm and an energy density of 500 .mu.w/cm.sup.2
for 140 minutes increased the burst pressure of the tubes to 280 mm
Hg, as compared to untreated tissue scaffolds which have an average
burst pressure of about 125 mm Hg.
[0057] In preferred embodiments, a tubular tissue scaffold of the
present invention has a burst pressure of at least about 100 mm Hg,
more preferably at least about 200 mm Hg, even more preferably at
least about 250 mm Hg, and most preferably at least about 280 mm
Hg.
[0058] In another embodiment, the present invention is directed to
an apparatus for use in the production of a tubular tissue scaffold
having aligned biopolymer fibrils, as described herein.
[0059] The apparatus of the present invention can be described with
reference to the figures. In FIG. 4 an apparatus for producing a
tubular tissue scaffold having aligned biopolymer fibrils is
illustrated. The apparatus comprises a gas source (200), which can
be any type of source that can supply gas having a desired
composition. It is preferred that the gas is a mixture of air and
ammonia, and a mixture of air and ammonia in about a 50:50 mixture
by volume is more preferred. A compressed gas cylinder is commonly
used as a gas source, but the apparatus is not limited to such a
source. The gas source (200) is operably connected to provide gas
to a tube-forming device (100) through a conduit (201), and can, if
desired, be operably connected to provide gas to a controlled
atmosphere chamber (600) through a conduit (202). The gas source
(200) can be used to supply gas to the luminal space of the tubular
tissue scaffold (700), and, if desired, to the interior of the
controlled atmosphere chamber (610). In a preferred embodiment, the
controlled atmosphere chamber is filled with a gas mixture
comprising air and ammonia, and a gas mixture comprising a mixture
of about 50:50 air and ammonia by volume is more preferred.
[0060] A biopolymer feed pump (300) is operably connected to feed a
biopolymer gel from a biopolymer gel source to the tube-forming
device (100). The feed pump can be any type of liquid feeding
device. Examples of suitable pumps are centrifugal pumps, gear
pumps, peristaltic pumps, progressing cavity pumps, syringe pumps,
and the like. It is preferred that the pump is one that is capable
of feeding a viscous biopolymer gel (perhaps having a viscosity of
from about 10 centipoise to as much as 1000 centipoise, or even
more), at a metered rate, and under conditions where the
cleanliness, and even sterility, of the biopolymer gel can be
maintained. A syringe pump has been found to be useful for small
scale tube-forming systems.
[0061] The tube-forming device (100) is one that is capable
producing a tube from the gel dispersion while providing an angular
shear force across the gel dispersion as it is formed into the wall
of the tube. The tube-forming device (100) is preferably located so
that the tubular tissue scaffold exiting the device (700) can
easily be transferred to a liquid bath (500) located to receive the
tubular tissue scaffold from the tube-forming device. The liquid
bath contains a liquid that receives and cushions the extruded
tube. The bath also serves to facilitate the solidification of the
tube. In an embodiment, the liquid bath comprises water having
sufficient ammonia absorbed therein to bring the pH of the water to
between about 9 and 11, a pH of about 10 is preferred.
[0062] It is also preferred that the space between the exit of the
tube-forming device (100) and the liquid bath (500) be enclosed to
provide a controlled atmosphere chamber (600). In embodiments where
the tubular tissue scaffold descends from the exit of the
tube-forming device (100) to the liquid bath (500) by force of
gravity, it is desirable that the surface of the liquid in the bath
(510) be located a defined distance (L) below the exit of the
tube-forming device.
[0063] In a preferred embodiment, the surface of the liquid bath is
located between about 0.25 centimeter and about 60 centimeters
below the exit port of the extruder, a distance of between about
2.5 centimeters and about 25 centimeters is more preferred, and a
distance of between about 7.6 centimeters and about 16.2
centimeters is even more preferred.
[0064] In a preferred embodiment, the tube-forming device (100) is
a rotating cone extruder or a rotating disk extruder. In a yet more
preferred embodiment, the tube-forming device is a rotating cone
extruder. In an even more preferred embodiment, the tube-forming
device is a counter-rotating cone extruder. With reference to FIG.
3, the major parts of a counter-rotating cone extruder include a
body (105) with cover plate (104), within which is contained an
external rotating member (120) having a cone-shaped cavity
therethrough and an internal rotating cone (110), which may also be
referred to herein as an "internal member" or internal rotating
member`, and which fits within the cone-shaped cavity of the
external rotating member. The internal cone-shaped member (110) is
connected to and is driven through an internal member drive gear
(112) and the external rotating member (120) is connected to and
driven by an external member drive gear (122). Both the internal
member drive gear (112) and the external member drive gear (122)
can engage and be driven by a pinion gear (130) that is connected
to and driven by a drive shaft from the drive motor (400). The
external rotating member and the internal rotating cone can
terminate near the apex of the cones to form an annular-shaped exit
port (150).
[0065] The external rotating member (120) and the internal rotating
cone (110) can be operably connected to one or more drive motors
(400) that can spin the external rotating member and the internal
rotating cone about a common axis but in opposite directions. In a
preferred embodiment, the external rotating member and the internal
rotating cone are operably connected to the same drive motor. In a
preferred embodiment, the one or more drive motors can be adjusted
to vary the speed at which the external rotating member and the
internal rotating cone rotate. The drive motor(s) can be designed
to rotate at any speed, but it is preferred that the motor(s) be
designed so that the rotational speed of the external rotating
member and the internal rotating cone can be varied between about 1
rpm and 1800 rpm.
[0066] The counter-rotating cone extruder of the present invention
can be smaller than counter-rotating cone extruders that are in
common use and are commercially available. One of the problems
associated with the manufacture of a reasonably priced
counter-rotating cone extruder is the provision of the intricate
bearings needed to retain the alignment and spacing of the external
rotating member and the internal rotating member. It has been found
that bearings can be dispensed with if the external rotating member
and the internal rotating cone of the extruder are constructed of a
durable, bearing-quality polymer. In preferred embodiments, the
external rotating member and the internal rotating cone are
constructed of Delrin.RTM. acetal resin (available from Tri Star
plastics, Reading, Pa.).
[0067] If desired, the annular exit port (150) of the extruder is
interconnected with a gas source via a gas conduit (140) to provide
for the addition of gas to the luminal space of a tubular tissue
scaffold exiting the extruder.
[0068] When the tube-forming device is a rotating-cone extruder or
a counter-rotating cone extruder, the annular-shaped exit port has
an outside diameter and an inside diameter. The difference between
the outside diameter and the inside diameter of the annular space
defines the width of the annular space. The thickness of the wall
of the tubular tissue scaffold is directly related to, but not
equivalent to, the width of the annular space.
[0069] The outside diameter and the inside diameter of the annular
space can be any desired dimension, but it is preferred that the
outside diameter of the annular-shaped exit port is between about
0.5 mm and about 150 mm, and the width of the annular space is
between about 0.1 mm and about 10 mm. It is more preferred that the
outside diameter of the annular-shaped exit port is between about 1
mm and about 20 mm, and the width of the annular space is between
about 0.1 mm and 2 mm. It is yet more preferred that the outside
diameter of the annular-shaped exit port is between about 1 mm and
about 10 mm, and the width of the annular space is between about
0.1 and 1 mm.
[0070] The present invention also includes a method of producing a
tubular tissue scaffold, the method includes the provision of a gel
dispersion comprising a biopolymer.
[0071] In addition to the isolated and purified biopolymer, or
combination of biopolymers, the biopolymer fibrils that form the
tube wall of the present tissue scaffold can be mixed with any
other polymers, or other additives, that are useful for the
formation of, or the performance of, the tubular tissue scaffold.
For example, fillers, dyes, drugs, or any other useful and
pharmacologically acceptable material may be added.
[0072] The biopolymer is prepared to form a gel dispersion prior to
formation of the present tubular tissue scaffold. By way of
example, a gel dispersion of type I collagen can be prepared by: a)
washing bovine hide sequentially in water, water containing
NaHCO.sub.3 and surfactant, and water; b) contacting the hide with
an aqueous solution containing NaHCO.sub.3, Ca(OH).sub.2 and NaHS;
c) washing the hide in water; d) treating the hide with an aqueous
solution of Ca(OH).sub.2; e) rinsing the hide with water and
trimming the hide of any remaining skin tissue and fat; f) placing
the hide in an aqueous salt solution and adding hydrochloric acid
solution until the pH is stable between about 6.0 and 8.0; g)
washing the hide in water; h) placing the hide in an aqueous
solution of acetic acid with or without pepsin; i) mixing and
allowing the hide to swell; j) placing the swollen hide in a mill
and processing into a gel dispersion; k) filtering the gel
dispersion to remove undissolved particles; l) centrifuging the gel
dispersion to remove small undissolved particles; m) adding salt to
the gel dispersion in an amount sufficient to precipitate collagen
from the gel dispersion; n) filtering the collagen precipitate and
resuspending it in deionized water; o) adding a base to bring the
pH of the collagen dispersion to a pH between about 6 and about 8;
p) dialyzing the collagen dispersion against phosphate buffered
saline solution or tris[hydroxymethyl]aminomethane buffer; q)
resuspending the collagen in deionized water; r) centrifuging the
collagen dispersion to concentrate solid collagen gel as a pellet;
and s) resuspending the pellet in aqueous mineral acid or organic
acid. When an organic acid is used, acetic acid is preferred. When
the pellet is resuspended in acid, it is preferred that the
collagen concentration is adjusted to between about 15-35 g/l,
about 20 g/l is more preferred.
[0073] The concentration of the gel dispersion can be adjusted to
contain about 2%-3%, by weight, solids by the addition of water.
The gel dispersion can be fed to a tube-forming device that is
capable producing a tube from the gel dispersion while providing an
angular shear force across the wall of the tube. This device can be
a counter-rotating cone extruder, a counter-rotating disk extruder,
or a counter-rotating cylinder extruder that preferably has a gas
channel connecting a gas source with the luminal space of the
tubular tissue scaffold as it exits the tube-forming device.
[0074] When the terms "angular shear force" are used herein, what
is meant is a shear force that is applied across the wall of the
tube, and in a direction generally perpendicular to both the radial
and longitudinal tubular dimensions. In other words, a shearing
force from the luminal wall of the tube to the outer wall of the
tube, or vice versa, in a circumferential direction--as provided,
for example, be a rotating cone extruder.
[0075] The tube-forming device forms the gel dispersion into a
tube, and the tube is then solidified, thereby forming a tubular
tissue scaffold comprising a biopolymer having fibrils in the tube
wall that are aligned in a helical pattern around the longitudinal
axis of the tube and where the pitch of the helical pattern changes
with the radial position in the tube wall. Often, the pitch of the
helical pattern on the luminal surface of the tube is different
from the pitch of the helical pattern on the outside surface of the
tube.
[0076] The pitch of the helical pattern of the fibrils can be
controlled by control of such variables as the feed rate of the
biopolymer gel dispersion to the tube-forming device, the rate of
shear imposed on the gel as the tube is being formed, and the
degree of drawing or compression of the tube after formation, but
before solidification. In some embodiments, where the tube-forming
device is a counter-rotating cone extruder that has an external
rotating member having a cone-shaped cavity therethrough and an
internal rotating cone which fits within the cone-shaped cavity of
the external rotating member, the external rotating member is
driven to rotate in one direction at a speed of from about 1 to
about 1800 rpm and the internal rotating cone is driven to rotate
in the opposite direction at a speed of from about 1 to about 1800
rpm. In a preferred embodiment, the external rotating member is
driven to rotate in one direction at a speed of from about 150 to
about 900 rpm and the internal rotating cone is driven to rotate in
the opposite direction at a speed of from about 150 to about 900
rpm.
[0077] In a preferred embodiment of the present apparatus, the
biopolymer feed pump (300) can be adjusted to vary the rate at
which biopolymer is fed to the tube-forming device. When the
tube-forming device is an extruder and the liquid bath (500) is
located at a defined distance (L) below the extruder exit port
(150), the tubular tissue scaffold exiting the extruder (700) can
fall into the bath by force of gravity. In this configuration, it
is possible to control the conformation of the tube as it
solidifies by controlling the biopolymer feed rate and the distance
between the exit port of the extruder and the surface of the liquid
in the bath (L). Because the tube exiting the extruder is still an
unsolidified gel dispersion, the tube can be drawn or
compressed--affecting the aligned helical pattern of the
fibrils--or it can collapse, unless certain measures are taken to
control the conformation of the unsolidified tube.
[0078] In order to prevent the tube from collapsing, a gas can be
fed from the gas source to the luminal space of the tubular tissue
scaffold as it exits the extruder. The flow rate of this gas can be
controlled so that it provides an internal pressure inside the tube
sufficient to prevent the collapse of the walls of the tube without
causing undue expansion of the tube that would adversely affect the
helical pattern of the aligned fibrils. When collagen is used as
the biopolymer, it is preferred that the gas comprises a mixture of
air and ammonia gas. Contact of the ammonia with the biopolymer
dispersion causes the pH of the walls of the tube to raise quickly,
thereby facilitating the solidification of the biopolymer gel. In
preferred embodiments, the mixture of air and ammonia is about a
50:50 mixture by volume.
[0079] In order to further facilitate the solidification of the
tube of biopolymer gel, the controlled atmosphere chamber can be
filled with the same gas as is fed to the lumen of the tube and the
outside surface of the tubular tissue scaffold is contacted with a
mixture of air and ammonia gas as it exits the extruder.
Solidification of the tubular tissue scaffold can also be
facilitated by providing that the liquid of the liquid bath is
composed of water having sufficient ammonia dissolved therein to
raise the pH of the bath liquid to about 10.
[0080] As mentioned above, when the liquid bath is located beneath
the exit port of the extruder, the conformation of the tube can be
controlled as it solidifies by controlling the biopolymer feed rate
and the distance between the exit port of the extruder and the
surface of the liquid in the bath (L). This can be accomplished by
feeding the biopolymer gel dispersion to the extruder at a defined
feed rate, and the defined feed rate and the defined distance (L)
are selected so that the gel dispersion solidifies to form a
tubular tissue scaffold comprising a biopolymer having fibrils in
the tube wall that are aligned in a helical pattern around the
longitudinal axis of the tube and where the pitch of the helical
pattern on the luminal surface of the tube is different from the
pitch of the helical pattern on the outside surface of the
tube.
[0081] In an embodiment where the outside diameter of the
annular-shaped exit port is about 5 mm and the width of the annular
space is about 0.5 mm, the defined feed rate can be controlled to
be sufficient to provide a tube extrusion rate of about 150 cm/min.
and the defined distance (L) is between about 10 cm and about 20
cm. The "tube extrusion rate" referred to is the linear rate of
speed at which the tubular tissue scaffold exits the tube-forming
device.
[0082] Although the tubular tissue scaffold can remain in the bath
for any desired length of time, when the bath comprises aqueous
ammonia having a pH of about 10, it is preferred to leave the
tissue scaffold in the bath for about 10 min-20 min, and about 15
min. is more preferred.
[0083] In a preferred embodiment, the step of causing the gel to
solidify can further include immersing the tube in an aqueous
solution containing 0.3% sodium bicarbonate, percent by weight.
[0084] After the gel has solidified, the tube can be sterilized by
exposure to sterilizing radiation. Gamma radiation and/or UV
radiation can be used for this purpose. In some embodiments, the
use of both gamma and UV radiation is preferred.
[0085] The tubular tissue scaffold of the present invention could
be used, for example, as an implanted prosthesis. In and of itself,
the scaffold could act as a growth substrate on which the cells of
the subject in which it is implanted adhere, replicate, and
reconstruct the injured tissue. Additionally, the tubular tissue
scaffold could be seeded with specific cell types in vitro,
cultured, and then implanted in a subject.
[0086] In another embodiment of the present invention, the tubular
tissue scaffold is split along the longitudinal axis and opened to
form a sheet. The sheet contains layers of aligned biopolymer
fibrils, where the direction of the alignment changes in each
successive layer. As with the tubular construct, the sheet could be
seeded in vitro with specific cell types either before or after
splitting. Applications for this type of tissue scaffold could
include any application that requires a sheet-type of tissue,
rather than a tubular structure. Sheet-type scaffold material can
be used, for example, for the preparation of artificial skin to
treat burn injury or surgical patches for internal application.
[0087] In another embodiment, the present invention encompasses
artificial tissue comprising living cells attached to a tubular
tissue scaffold comprising a tube having a wall, wherein the wall
includes biopolymer fibrils that are aligned in a helical pattern
around the longitudinal axis of the tube where the pitch of the
helical pattern changes with the radial position in the tube
wall.
[0088] In an alternative embodiment, when the tubular artificial
tissue scaffold is split to form a sheet-type structure, the
artificial tissue comprises a sheet having a thickness, wherein the
material comprising the sheet includes biopolymer fibrils having an
alignment that changes with the thickness of the sheet.
[0089] In the artificial tissue of the present invention, living
cells may be introduced to the luminal and outer surfaces of the
tubular tissue scaffold. As used herein, the terms "introduced",
"seed", and "seeded" in reference to cells refer to the addition of
cells to the tissue scaffold by providing the cells in a cellular
suspension and supplementing the solution in which the tissue
scaffold is incubating with the cellular suspension. By way of
example, cells are isolated from a given source, such as neonatal
rat hearts, and dispersed in a solution to form a cellular
suspension having a certain cell density. A volume of cellular
suspension is injected into the tubular tissue scaffold using an IV
catheter. The tubes are placed in a rotating wall bioreactor and
the reactor is filled with additional cell suspension. The tubes
and cells are incubated at a rotation rate of 20 rpm with 5%
CO.sub.2 at 37 degrees Celsius for several days.
[0090] In some embodiments, the living cells align along the
helical pattern of the biopolymer. "Aligned along the helical
pattern of the biopolymer", as used herein, means oriented in
parallel with the direction of the biopolymer fibrils. In preferred
embodiments, the living cells establish intercellular and
extracellular connections such as those found in vivo. These
connections can include, for example, intercalated disks between
distal ends of adjacent cardiac myocytes consisting of gap
junctions that mediate electrical signaling between cells, adherens
junctions and desmosomes between adjacent cells (cadherin
interactions), and focal adhesions and hemidesmosomes
(integrin-matrix interactions).
[0091] The living cells in the present invention are selected from
the group consisting of myocyte precursor cells, cardiac myocytes,
skeletal myocytes, satellite cells, fibroblasts, cardiac
fibroblasts, chondrocytes, osteoblasts, endothelial cells,
epithelial cells, embryonic stem cells, hematopoetic stem cells,
neuronal cells, mesenchymal stem cells, anchorage-dependent cell
precursors, and combinations thereof. For example, the living cells
can be a combination of cardiac myocytes and cardiac fibroblasts.
In some embodiments the living cells are contractile. As used
herein the term "contractile" means having the ability to shorten
in length due to mechanical alterations in intracellular structural
proteins. In preferred embodiments, the contractile cells are
cardiac myocytes.
[0092] Neonatal cardiac myocytes maintain some plasticity, as they
undergo hyperplastic growth to reach maturity following birth. On a
planar growth surface, neonatal myocytes exhibit a stellate
morphology unlike that seen in vivo. Given more appropriate growth
substrates in vitro, such as an oriented matrix of collagen or
laminin, neonatal cardiac myocytes will align with and attach to
the matrix, adopt a rod-shaped morphology, and form organized
arrays of myofibrils. See i.e. McDevitt, T. C., et al., J Biomed
Mater Res, 60:472-9 (2002). These cells can also form intercalated
disks to allow for the conduction of electrical signals from one
cell to the next, resulting in coordinated contractile activity
similar to that seen in the intact myocardium of the heart.
However, conventional 2-dimensional culture systems cannot sustain
this activity indefinitely, and cells stop beating after two to
three weeks.
[0093] The 3-dimensional structure of the scaffolding used in the
present invention provides a more physiologically similar context
than 2-dimensional conditions, allowing, for example, grafted
cardiac myocytes to form intercellular and extracellular
connections between successive layers, so that the tube can
contract as an organ. Without being bound by this or any other
theory, the inventors believe that the ability of the ventricle of
the heart to contract in a manner which propels blood out of the
intraventricular space and into the systemic circulatory system may
rely on the unique geometric configuration of the collagen fibrils
in the extracellular matrix to properly align the cells of the
myocardium. Therefore, the purpose of the particular alignment of
biopolymer fibrils of the present invention is to allow for the
generation of the same type of mechanical force, given the
introduction of contractile cells to the scaffold. Accordingly, in
one embodiment of the present invention, the cardiac myocytes
contract synchronously along the helical pattern of the biopolymer,
thereby twisting the tube while reducing its diameter, and thereby
"wringing" fluid from the tube in a motion similar to that of a
biological heart.
[0094] The term "contract synchronously" as used herein in
reference to cardiac myocytes, refers to the ability of an
electrical signal to pass rapidly from cell to cell via gap
junctions to couple the cells so that they contract in unison as a
single functional unit.
[0095] In one embodiment, artificial tissue of the present
invention has living cells attached to two or more tubular tissue
scaffolds, which may preferably be in the form of concentric tubes.
In some embodiments, the living cells attached to individual
tubular tissue scaffolds are of different cell types. By way of
example, in artificial tissue comprising three concentric tubular
tissue structures, the inner tube is seeded with contractile muscle
cells and endothelial cells, the central tube is seeded with
extracellular matrix-producing fibroblasts, and the outer tube is
seeded with contractile muscle cells and endothelial cells.
[0096] In another embodiment, the artificial tissue can be
engineered to contain and release cytokines and other pharmacologic
agents that affect cell proliferation, development, migration,
differentiation, and/or activity, which are appropriate for the
specific application of the tissue. As used herein, such cytokines
and other pharmacologic agents are referred to as "growth factors",
and include, without limitation, epidermal growth factor (EGF),
fibroblast growth factor (FGF), erythropoietin (EPO), hematopoietic
cell growth factor (HCGF), platelet-derived growth factor (PDGF),
stem cell factors, bone morphogenic protein (BMP), fibronectin,
transforming growth factor-beta (TGF-.beta.), and
neurotrophins.
[0097] Growth factors modulate and control the inception, rate, and
cessation of the vital healing events that may be associated with
surgical implantation of engineered tissue. Growth factors are
polypeptides that modulate cellular function and regulate cellular
growth. These peptides are extremely potent and, in very small
quantities, are able to induce a specific cellular response. Three
key growth factors known to be vital to the proper healing of
damaged tissue are vascular endothelial growth factor, platelet
derived growth factor and nerve growth factor.
[0098] Vascular endothelial growth factor, VEGF, is an
extracellular signal protein that acts through the membrane bound
tyrosine kinase receptor, VEGF receptor, to stimulate angiogenesis
in vivo. A shortage of oxygen in practically any type of cell
causes an increase in the intracellular concentration of the active
form of a gene regulatory protein called hypoxia inducible factor I
(HIF-1). HIF-1 stimulates transcription of the VEGF gene (and
others). VEGF is secreted, diffuses through the tissue and acts on
nearby endothelial cells. The response of the endothelial cells to
VEGF includes at least four components. First, the cells produce
proteases to digest their way through the basal lamina of the
parent vessel. Second, the endothelial cells migrate toward the
source of the signal. Third, the cells proliferate. Fourth, the
cells form tubes and differentiate. Thus, endothelial cells create
and line the lumen of the newly formed blood vessels in the tissue.
These neovessels will not persist on their own. They will develop
microaneurysms as well as other abnormalities that eventually
rupture. These vessels rely on the recruitment of pericytes, from
the vasculature, under the influence of signals from the
endothelial cells, to further mature into competent blood vessels
with the addition of vascular smooth muscle cells and extracellular
matrix. The recruitment of pericytes, in particular, depends on
PDGF secreted by endothelial cells.
[0099] Almost every cell in almost every tissue is located within
50-100 .mu.m of a capillary. In the case of wound healing, for
example, there is a burst of capillary growth in the neighborhood
of the damage to satisfy the high metabolic requirements of the
repair process. Local infections and irritations also cause a
proliferation of new capillaries most of which regress and
disappear when the inflammation subsides. In all of these cases,
the invading endothelial cells respond to signals produced by the
tissue that they invade. The signals are complex, but a key part is
played by VEGF.
[0100] Platelet derived growth factor (PDGF) is an extracellular
signal protein that acts through the membrane bound tyrosine kinase
receptors, PDGF receptors .alpha. and .beta., to stimulate the
survival, growth and proliferation of various cell types in vivo.
PDGF stimulates chemotaxis and proliferation of fibroblasts and
smooth muscle cells as well as collagen synthesis and collagenase
activity.
[0101] Nerve growth factor (NGF) is an extracellular signal protein
that acts through a membrane bound tyrosine kinase receptor, Trk A,
to stimulate survival and growth of neurons. Cell growth and
division can be controlled by separate extracellular signal
proteins in some cell types. Such independent control may be
particularly important during embryonic development when dramatic
changes in the size of certain cell types can occur. Even in adult
animals, however, growth factors can stimulate cell growth without
affecting cell division. The size of a sympathetic neuron, for
example which has permanently withdrawn from the cell cycle,
depends on the amount of nerve growth factor secreted by the target
cell it innervates. The greater the amount of NGF the neuron has
access to, the larger it becomes. Concentrations of 250 ng/ml of
NGF have been shown to cause migration of neurons through collagen
gels in vivo.
[0102] Therefore, when vascularization and innervation of the
artificial tissue of the present invention is required, the
presence of growth factors such as those described above is
necessary for these processes to proceed in vivo.
[0103] In order to provide a useful amount of one or more growth
factors, and to control their release from the artificial tissue of
the present invention, the tissue scaffolds of the present
invention can be incubated with a growth factor, or combination of
growth factors, and allowed to absorb the growth factor(s).
Alternatively, growth factors may be incorporated into the
biopolymer solution prior to formation of the tubular tissue
scaffold. It should be appreciated however, that the present
invention is not limited by any particular method of treating the
tissue scaffold with a growth factor, and the invention is
applicable to any such method now known or subsequently discovered
or developed. Growth factors useful in the present invention
include, but are not limited to, vascular endothelial growth factor
(VEGF), epidermal growth factor (EGF), fibroblast growth factor
(FGF), erythropoietin (EPO), hematopoietic cell growth factor
(HCGF), platelet-derived growth factor (PDGF), nerve growth factor
(NGF), transforming growth factors .alpha. and .beta. (TGF-.alpha.
and TGF-.beta.), or combinations thereof.
[0104] In one embodiment, the tissue scaffold can be UV irradiated
in order to crosslink the biopolymer fibrils. Such irradiation can
be administered before or after the addition of growth factors. In
one embodiment, the tissue scaffold is irradiated after the
addition of growth factors, but prior to cell seeding in order to
crosslink the biopolymer fibrils after growth factor addition to
the scaffold. While not wishing to be bound by this or any theory,
the degree of crosslinking can control the rate at which growth
factors are released from the tissue. For example, the tissue
scaffolds can be exposed to UV radiation at a wavelength of 254 nm
and an energy density of 500 .mu.w/cm.sup.2 for 140 minutes. When
compared to tissue scaffolds not subjected to UV radiation, tissue
scaffolds treated as described above have a slower rate of release
of growth factors. Therefore, treatment of the tissue scaffolds
with UV radiation can be used to modulate the release of growth
factors once cells have been seeded on the scaffold, in turn
regulating processes such as neovascularization and innervation of
the engineered tissue.
[0105] The present invention also encompasses a method of treatment
comprising implanting in the body of a subject artificial tissue
comprising living cells attached to a tubular tissue scaffold
comprising a tube having a wall, wherein the wall includes
biopolymer fibrils that are aligned in a helical pattern around the
longitudinal axis of the tube where the pitch of the helical
pattern changes with the radial position in the tube wall.
[0106] The term "subject" for purposes of treatment includes any
vertebrate. Preferably, the vertebrate is a human or animal subject
who is in need of treatment for an injury, disease, or disorder of
the type that can be treated by the use of artificial tissue. The
subject is typically a mammal. "Mammal", as that term is used
herein, refers to any animal classified as a mammal, including
humans, domestic and farm animals, and zoo, sports, or pet animals,
such as dogs, horses, cats, cattle, etc. Preferably, the mammal is
a human. Additionally, the term "implanting in the body" refers to
surgically inserting into the subject at the site of the injury,
disease, or disorder being treated.
[0107] When it is said that the present artificial tissue can be
used in treatment for an injury, disease, or disorder of the type
that can be treated by the use of artificial tissue, such treatment
can include, but is not limited to, replacement of vessels such as,
for example, coronary arteries; repair and/or replacement of any
other physiologic tubular structures, such as, for example,
ureters, veins, lymph channels, GI tract components, and the like;
repair of injured bone; repair of damaged nervous tissue as in, for
example, spinal cord injury; or correction of impaired cardiac
function caused by, for example, ischemic heart disease.
[0108] In one embodiment, the artificial tissue used in the method
of treatment has living cells introduced to the luminal and outer
surfaces of the tissue scaffold, and preferably the cells are
aligned along the helical pattern of the biopolymer. The living
cells in the present invention can be selected from the group
consisting of myocyte precursor cells, cardiac myocytes, skeletal
myocytes, satellite cells, fibroblasts, cardiac fibroblasts,
chondrocytes, osteoblasts, endothelial cells, epithelial cells,
embryonic stem cells, hematopoetic stem cells, neuronal cells,
mesenchymal stem cells, anchorage-dependent cell precursors, and
combinations thereof. In some embodiments, the living cells are
selected from the group consisting of cardiac myocytes, skeletal
myocytes, fibroblasts, cardiac fibroblasts, chondrocytes,
osteoblasts, endothelial cells, epithelial cells, embryonic stem
cells, hematopoetic stem cells, neuronal cells, mesenchymal stem
cells and combinations thereof.
[0109] In one embodiment, the living cells of the present invention
originate from the subject receiving treatment. Autologous grafts
are far less susceptible to rejection, as they are not recognized
as foreign and, hence, do not elicit an immune response in the
subject. In preferred embodiments, cells taken from the subject are
introduced to the tissue scaffold in vitro. The cells are cultured
for the period of time necessary for the cells to degrade the
original scaffold biopolymer and regenerate a matrix of secreted
extracellular matrix proteins. Preferably, the cells are cultured
in vitro for a period of time necessary for the cells to degrade
and replace at least 25 percent of the original scaffold. More
preferably, the cells are cultured for a period of time necessary
for the cells to degrade and replace at least 50 percent of the
original scaffold and, more preferably still, a period of time
necessary for the cells to degrade and replace at least 75 percent
of the original scaffold. The remodeled tissue scaffold is then
constructed of proteins produced by the subject's own cells, and is
therefore not recognized as a foreign substance that would induce
an immune response in the subject.
[0110] It is preferred, in the present invention, that the living
cells introduced to the tubular tissue scaffold establish
intercellular and extracellular connections such as those found in
vivo. In one embodiment, these living cells are contractile. In
preferred embodiments the contractile cells are cardiac myocytes.
As mentioned previously, the unique organization of the tissue
scaffold of the present invention is able to guide the formation of
an interconnected, contractile cell system, wherein the cardiac
myocytes contract synchronously along the helical pattern of the
biopolymer in a "wringing" motion to pump fluid through the lumen
of the tubular tissue scaffold. The pumping action of the
artificial tissue can be made to be directional by any one of
several methods. For example, one-way valves can be placed on one
or both sides of a contractile artificial tissue tube that result
in directing the flow of fluid through the tube. Alternatively, one
or more artificial tissue tubes can be located in series and
controlled with one or more pacing devices in a manner that creates
a peristaltic action to force fluid from one end of the tubular
construct to the other.
[0111] Because the artificial tissue is able to act as a pump when
contractile cells are present, this particular embodiment of the
present invention can be used in the treatment of an injury,
disease, or disorder that involves abnormalities in either cardiac
output or vascular tone, or vascular blockage. In some embodiments
of the present invention, these can include congestive heart
failure, dilated cardiomyopathy, hypertrophic cardiomyopathy,
infiltrative cardiomyopathy, ischemic heart disease, heart attack,
heart failure, coronary artery disease, atherosclerosis,
hypertension, chronic renal disease, cerebrovascular disease,
carotid artery disease, and peripheral vascular disease.
[0112] For most in vitro or in vivo uses, the tubular scaffold used
to make the artificial tissue of the present invention should be
sterilized and treated with antibiotic and antifungal agents. By
way of example, the tubular scaffolds can be sterilized by placing
them in sterile Mosconas solution (0.14M NaCl, 0.0027M KCl, 0.012M
NaHCO.sub.3, 4.2.times.10.sup.-5 M NaH.sub.2PO.sub.4, 0.0094M
glucose) and exposing them to gamma radiation or ultraviolet (UV)
radiation or both for up to 4 hours. Following UV treatment, fresh
Marconas solution with 0.01 mg/ml gentamycin, 4 .mu.g/ml
Amphotericin-B, and 10 .mu.g/ml fibronectin can be added to a
culture dish containing the tubular scaffolds and the tubes can be
incubated for 24 hours with 5% CO.sub.2 at 37.degree. Celsius.
[0113] The artificial tissue of the present invention provides a
tubular cell-based prosthetic that would be particularly useful for
repairing or replacing tube-shaped "organs", such as various sized
blood vessels including coronary arteries and renal arteries,
ureters, fallopian tubes, or nerve fiber conduits. If the cells of
the present invention include contractile cells, such as cardiac
myocytes, the tube as a whole can contract in a "wringing" motion,
as dictated by the alignment of the cells along the helical pattern
of the biopolymer fibrils. This type of contractile engineered
construct could be used, for example, as a ventricular assist
device to enhance cardiac contractility in a failing heart. Large
versions of such a contractile tubular artificial tissue could be
implanted near the heart to improve systemic circulation. Smaller
versions could be implanted near vital organs to improve local
perfusion. For example, contractile tubes could be transplanted in
the renal arteries to increase renal perfusion and improve cardiac
performance while unloading the heart. In another example,
contractile tubes could be implanted in the common carotid artery
inferior to the carotid sinus to improve brain perfusion. The
contractile tubes could be used as replacement coronary arteries to
provide both a blood conduit as well as improved local perfusion
during diastole. Preferably, the contractile activity of this type
of construct is controlled with a pacing device. As used herein,
"pacing device" refers to any electrical device that can be used to
maintain a particular rate of contraction, such as, for example,
the implantable cardiac resynchronization device made by Medtronic,
Inc., called InSync.RTM..
[0114] One problem that has limited the application of artificial
tissue, and in particular, artificial muscle tissue such as
replacement cardiac muscle tissue, is that the tissue must be able
to function immediately and accurately in the rigorous environment
of the continuously cyclically-contracting heart. Such tissue
cannot be allowed to significantly degrade and be remodeled in
situ. Currently available scaffolds for tissue engineering undergo
global remodeling once implanted in vivo. In order to avoid or
reduce the degree of remodeling and consequent loss of strength
and/or function of implanted artificial tissue, the present
invention includes a method for preconditioning the novel tubular
artificial tissue scaffold by seeding it with cells, such as
fibroblasts, and stimulating the cells with a combination of matrix
components, growth factors, and mechanical stimulation. Without
being bound by this theory, it is believed that such stimulation
can control fibroblast proliferation, matrix synthesis, matrix
degradation and fibroblast apoptosis, among other parameters that
are important for artificial tissue development.
[0115] When it is said that the preconditioning includes
stimulation of the cells with matrix components, what is meant is
the selection and control of the composition of the present
artificial tissue scaffold and its physical structure, as have been
discussed above. For example, variables such as the type of
biopolymer that is used to produce the fibrils, and how and to what
extent they are aligned, as well as the dimensions of the tubular
scaffold, are included in the selection and control of matrix
components.
[0116] Mechanical signals play an integral role in both directing
myocytes to assume the distinctive cytoarchitectural features
characteristic of differentiated myocytes as well as the
arrangement of individual myocytes into an intact muscle. A primary
developmental force in the heart is mechanical signaling resulting
from contractile force as well as an increase in pressure and
volume. See Terracio, L., et al., In Vitro Cell Dev Biol, 24:53-8
(1988). The regulation, expression, synthesis, and degradation of
various contractile and regulatory proteins, as well as cell size,
are influenced by mechanical stress. Accordingly, in one embodiment
of the present invention, artificial tissue containing certain
cells, such as contractile cardiac myocytes, is preconditioned by
the application of mechanical stress before being implanted in the
subject. What is meant by "preconditioned by the application of
mechanical stress" is permitting cells to become accustomed to
forces, such as stretch, normally found in the cardiovascular
system in vivo, by subjecting the cells to these forces before
introduction into the subject.
[0117] One method of applying mechanical stimulation to the
artificial tissue scaffold is by culturing artificial cardiac
myocyte-containing tissue in a pulsatile internal flow stretch
bioreactor that mimics the action of a developing cardiovascular
system. Stretch has been shown to up-regulate both matrix
metalloproteinase production as well as the synthesis of new
matrix. It is believed that the pulsatile internal flow and stretch
mimics the action of a developing cardiovascular system.
[0118] An example of a pulsatile internal flow stretch bioreactor
is shown in FIG. 8. In this system, a tubular tissue scaffold,
either before or after the attachment of cells is placed in the
flow pattern of the bioreactor. The media reservoir can be adjusted
to various heights from essentially zero height to a height
equivalent to 10 cm of water pressure, or more. The pulse frequency
and stretch frequency can be adjusted from 0 Hz to 2 Hz. It is
believed that fibroblasts and myocytes that are seeded on the
tubular tissue scaffold will remodel the matrix in response to the
strain imposed by the pulse.
[0119] In addition to the presence of growth factors during the
preconditioning process, other biologically active chemicals, such
as matrix metalloproteinase (MMP) inhibitors can also be added in
order to promote the preconditioning of the artificial tissue.
[0120] In one embodiment, the method for preconditioning the novel
tubular artificial tissue scaffold comprises seeding a tubular
artificial tissue scaffold having aligned biofibrils with living
cells and culturing the cells in the presence of media containing
at least one growth factor and under conditions where the tubular
artificial tissue scaffold is subjected to stretch and pressure
pulse of controlled frequency and amplitude.
[0121] The present invention also includes a novel preconditioned
artificial tissue comprising living cells that are attached to a
tubular artificial tissue scaffold having aligned biofibrils,
wherein the cells have been cultured in the presence of media
containing at least one growth factor and under conditions where
the tubular artificial tissue scaffold is subjected to stretch and
pressure pulse of controlled frequency and amplitude.
[0122] In another embodiment of the present invention, the tubular
artificial tissue is split along the longitudinal axis and opened
to form a sheet. The sheet contains layers of aligned biopolymer
fibrils with attached cells, where the direction of the alignment
changes in each successive layer. Applications for this type of
artificial tissue could include, for example, artificial skin to
treat burn injury or surgical patches for internal application.
Cardiac myocytes seeded on a collagen sheet having the fibril
alignment patterns of the present novel tissue scaffolds could
provide functional myocardial patches to treat infarcted areas of
the heart.
[0123] In one embodiment, the tubular tissue scaffold is split
longitudinally and opened into a sheet prior to attaching living
cells to the tissue scaffold. The sheet can be coated with an
additional layer of biopolymer fibrils such as, for example,
collagen, wherein the biopolymer fibrils polymerize and are
oriented in the direction in which they are applied.
[0124] Living cells useful in this embodiment of the present
invention include, for example, myocyte precursor cells, cardiac
myocytes, skeletal myocytes, satellite cells, fibroblasts, cardiac
fibroblasts, chondrocytes, osteoblasts, endothelial cells,
epithelial cells, embryonic stem cells, hematopoetic stem cells,
neuronal cells, mesenchymal stem cells, anchorage-dependent cell
precursors, and combinations thereof. In preferred embodiments, the
living cells are selected from the group consisting of myocyte
precursor cells, cardiac myocytes, skeletal myocytes, satellite
cells, fibroblasts, and combinations thereof. In one embodiment,
the living cells originate from the subject receiving
treatment.
[0125] The sheet form of the tissue scaffold can be used for
treatment of, for example, hernia, heart attack, congenital heart
defects, tissue missing due to congenital defect, skin burns, organ
damage, and muscle damage.
[0126] In one embodiment, the sheet form of artificial tissue can
be used to repair a tear or defect in a variety of tissues
including, but not limited to, cardiac tissue, skeletal muscle
tissue, epithelial tissue, vascular tissue, nerve tissue, lymphatic
tissue, connective tissue, epidermal tissue, endocrine tissue,
cartilage, and bone.
[0127] By way of example, the sheet form of artificial tissue can
be used to repair a ventral hernia. Collagen tubes are produced as
described herein, and are cut open longitudinally to form sheets. A
thin layer of collagen solution can be streaked on the surface of
the substrates and allowed to polymerize. This procedure results in
a thin layer of collagen fibrils that are arrayed in parallel with
one another along the direction that the collagen solution was
streaked. Myocyte precursor cells such as, for example, satellite
cells, can be introduced to the sheet scaffold and cultured until
the cells differentiate into myocytes and fuse to form mature
myotubes. This artificial skeletal muscle tissue can then be
surgically applied to the damaged area of the abdominal wall to
repair the hernia.
[0128] Additionally, the present invention embraces a method of
identifying the effects of a pharmaceutical composition on cell
function comprising administering said pharmaceutical composition
in vitro to artificial tissue comprising living cells attached to a
tubular tissue scaffold comprising a tube having a wall, wherein
the wall includes biopolymer fibrils that are aligned in a helical
pattern around the longitudinal axis of the tube where the pitch of
the helical pattern changes with the radial position in the tube
wall. Artificial tissue that has been made into sheet form as
described above can also be used for the tissue scaffold for this
method.
[0129] In one embodiment, the method further comprises determining
the effects of a pharmaceutical composition on the cells by
measuring or identifying changes in cell function. This can be
accomplished by many methodologies known to those skilled in the
art including, for example, Western blot analysis, Northern blot
analysis, RT-PCR, immunocytochemical analysis, BrdU labeling, TUNEL
assay, and assays of enzymatic activity. In some embodiments, the
living cells are contractile cardiac myocytes. Accordingly,
measurements of parameters such as isovolumic pressure generation,
length tension, and isometric force generation can be made. By way
of example, the instant artificial tissue containing cardiac
myocytes that contract in culture could be treated with an agent,
such as ephedrine, and any change in contractility of the myocytes
can be measured as described in Example 3. This method provides an
in vitro diagnostic system that can be utilized to rapidly assay
the physiological consequences of administration of a given
pharmaceutical composition on cell function such as, for example,
cardiac contractility.
[0130] The following examples describe preferred embodiments of the
invention. Other embodiments within the scope of the claims herein
will be apparent to one skilled in the art from consideration of
the specification or practice of the invention as disclosed herein.
It is intended that the specification, together with the examples,
be considered to be exemplary only, with the scope and spirit of
the invention being indicated by the claims which follow the
examples. In the examples all percentages are given on a weight
basis unless otherwise indicated.
EXAMPLE 1
[0131] This example illustrates the preparation of a biopolymer gel
comprising type I collagen from a bovine steer hide.
[0132] Collagen was prepared from the hide of an 18-month old
bovine steer by removal of the superficial epidermis including the
hair and follicle pits. To isolate the collagen the hide was cut
into 4.times.6 cm strips and washed with three changes of deionized
(DI) water for 1 hr. each (the second wash contained 0.2%
NaHCO.sub.3). The remaining follicles and noncollagenous proteins
were removed by bathing in a solution of 0.6% NaHCO.sub.3, 2%
Ca(OH).sub.2, and 4.3% NaHS for 30 min at 20.degree. C. After three
washings in DI water, the strips were treated overnight in 2%
Ca(OH).sub.2 at 4.degree. C. The remaining fat and epidermis were
trimmed from the strips the following day. The strips were then
placed in a 2M NaCl solution and neutralized with HCl to a pH of
6.8-7.0. The strips were then washed three times in DI water, cut
into 0.5.times.2 cm pieces, and placed in a solution of 0.5 N
acetic acid with or without pepsin, 1:100 based on hide weight.
After incubating overnight at 4.degree. C., the swollen strips were
mixed with ice and emulsified into a gel dispersion using a
Kitchen-Aid food processor (model # FP500WH; St. Joseph, Md.). The
suspension was centrifuged at 9950.times.g for 35 min to remove
small, unsolubilized particles. Type I collagen was precipitated
from the gel dispersion by mixing with NaCl to a final
concentration of 2M. The collagen was collected by centrifugation
at 9950.times.g for 35 min, resuspended in DI water and neutralized
to a pH of 7.2 using NaOH. The collagen suspension was then
dialyzed vs phosphate buffered saline overnight and vs DI water for
three changes over 24 hr. The resulting collagen gel was then
centrifuged at 9950.times.g for 35 min to remove excess water and
the collagen concentration adjusted to 25 mg/ml by addition of DI
water and the pH adjusted to 2.5-3.0 with concentrated HCl.
EXAMPLE 2
[0133] This example illustrates the production of a collagen
tubular tissue scaffold having aligned fibrils by using a
counter-rotating cone extruder.
[0134] Collagen tubes were produced by loading a collagen gel
dispersion, produced as described in Example 1, into a 60 ml
syringe that was placed into a syringe pump (having an adjustable
feed rate of 0-20 cc/min) that fed the collagen into the feed port
of a counter-rotating cone extruder that was fabricated from Delrin
(Tri Star plastics, Reading, Pa.) according to the illustration
shown in FIG. 1. The drive motor was a 90 volt 0.06 hp DC motor
(Baldor Electric, Ft. Smith AK) with a regenerative feedback drive
and a 5 k.OMEGA. potentiometer for user interface (KB electronics,
Coral Springs Fla.). The motor is coupled to the pinion gear which
drives the outer rotating member and the inner rotating cone drive
gears. The direct drive system can control speeds from 0 to 1800
rpm with a 2:1 reduction from the pinion gear to the drive gears.
The extruder was mounted atop a controlled atmosphere chamber that
consisted of a vented acrylic box that was purged at 1 L/min with a
50/50 mix of anhydrous ammonia and air.
[0135] As the dispersion was fed between the outer and inner
rotating cones, a collagen tube exited through the annular-shaped
exit port and was collected in a deionized water bath. The speed of
the cone rotation and forward flow of the gel dispersion combined
to create a helically aligned orientation pattern of the collagen
fibrils. To maintain a lumen, as tubes are extruded, gas (air or
air and ammonia) is metered between 10 to 60 ml/min through the
tube. Following extrusion the tube was left in a water bath for 1
hr at room temperature followed by 30 min in a 0.3% NaHCO.sub.3
solution and an additional 1 hr. in deionized water.
Determination of the Pitch of the Helical Pattern of the Aligned
Collagen Fibrils:
By Polarized Light Microscopy:
[0136] To determine the pitch of the collagen fibrils at the
outside and luminal walls of the tube, freshly extruded collagen
tubes were fixed in 2.5% glutaraldehyde in phosphate buffered
saline (PBS) for 1 hr at 25.degree. C. The tubes were then stained
with picrosirius red, and then imaged with a Bausch & Lomb
Illuminator optical polarized microscope to examine the
birefringent nature of the collagen fibrils. The lower polarizer
was installed such that light was polarized in a north-south
direction and the upper analyzer was oriented in the east-west
direction.
[0137] Samples were placed on the rotary stage of the polarized
light microscope in an east-west direction with respect to the long
axis of the tube. The stage was rotated from 0.degree. to
180.degree. and the stage angle showing maximum darkness and
brightness recorded. When the collagen fibers were at a 45.degree.
angle between the polarizer and the analyzer, transmitted light was
at its maximum intensity and the collagen orientation could be
directly visualized. When the pitch of the helical pattern
coincided with either the lower polarizer or upper analyzer, the
refracted light was quenched and there was minimal light
transmission. The rotation angle of the stage was recorded at each
interval of maximum and minimum light transmission and the angle of
the fiber array calculated with respect to the long axis of the
tube.
By Scanning Electron Microscopy (SEM):
[0138] To further illustrate the helical orientation of the
collagen fibrils in the tubes, samples were prepared for SEM by the
O-GTA-O-GTA-O method of Takahashi et al., J. Electron Microsc.,
35(3):304 (1986). Extruded collagen tubes were fixed for 2 hr. in
3% glutaraldehyde in 0.1 M sodium cacodylate buffer (pH 7.4),
rinsed in buffer, and immersed in 2% aqueous OsO.sub.4 for 2 hr.
After rinsing, samples were treated with two repetitions of GTA-O
steps: 3.times.4 hr 8% glutaraldehyde/2% tannic acid at 4.degree.
C., rinse, 2 hr. 2% OsO.sub.4. Tubes were then dehydrated in a
graded ethanol, critical point dried, mounted on aluminum stubs and
imaged on a JEOL JSM-6300V at 10 kV.
[0139] The counter rotating action of the extruder cones created a
spiraling, or helical, alignment of the collagen fibrils that had a
uniform direction on each of the outside and luminal walls of the
tube. A section of a tube was split longitudinally down the tube
wall and the tube was opened into a sheet. SEM microscopy of each
side of the sheet is shown in FIG. 2 and clearly shows the
"corkscrew" and "counter corkscrew" orientation of the fibrils at
the surface of the outside and luminal tube walls. In this tube the
pitch of each helix was about 45.degree., and in opposite
directions.
[0140] When the cone rotation rate that varied from 150 rpm to 900
rpm at a fixed extrusion rate of 150 cm/min, the fibril angle of
each array with respect to the long axis of the tube varied between
26.degree. to 62.degree.. Slower rotation speeds resulted in
smaller angles while higher cone rotation speeds resulted in larger
angles.
EXAMPLE 3
[0141] This example illustrates seeding of cardiac myocyte on a
collagen tubular tissue scaffold.
[0142] Collagen tubes were produced as described in Example 2.
Prior to the addition of myocytes, the collagen tubes were
sectioned into 1.25 cm lengths. The small collagen tube sections
were placed in sterile Mosconas solution (0.14M NaCl, 0.0027M KCl,
0.012M NaHCO.sub.3, 4.2.times.10.sup.-5 M NaH.sub.2PO.sub.4,
0.0094M glucose) and exposed to ultraviolet (UV) light for 4 hours
to sterilize. Following UV treatment, fresh Marconas solution with
0.01 mg/ml gentamycin, 4 .mu.g/ml Amphotericin-B, and 10 .mu.g/ml
fibronectin was added to a culture dish containing the collagen
tube sections and incubated for 24 hours with 5% CO.sub.2 at
37.degree. Celsius.
[0143] Isolation of neonatal cardiac myocytes was performed as
described by Simpson et al., J Cell Physiol, 161:89-105 (1994). For
seeding, collagen tube sections were placed in 150 mm culture
dishes and 0.5 ml of a cellular suspension containing
2.times.10.sup.6 cells/ml was injected into the lumen of each tube
using an 18 gauge IV catheter (Surflo Terumo Mol. Corp., Somerset,
N.J.). The tubes were then placed in a Synthecon Rotating Wall
Bioreactor (Synthecon, Houston, Tex.) and the reactor filled with
an additional 2.5 ml of cell suspension. Incubation was at a
rotation rate of 20 rpm with 5% CO.sub.2 at 37.degree. Celsius for
72 hours.
[0144] Following incubation in the bioreactor, tubes were placed in
culture dishes with fresh media containing 4 .mu.g/ml cytosine
.beta.-D-arabino-furanoside. After 72 hours in the bioreactor,
individual myocytes contracted spontaneously. Two further cell
seedings were performed by adding 2.times.10.sup.6 cells per tube
to the tube lumen and outside surface at one week intervals. After
10 days in culture, entire areas of the tube contracted and, after
16 days in culture, synchronous myocyte contraction and forceful
contractions of the entire tube were observed.
Characterization of Cardiac Myocytes in Collagen Tubes:
[0145] For confocal microscopy, tubes were sectioned
longitudinally, fixed and stained for f-actin and connexin 43
(Chemicon MAB3068; 1:1000 dilution) as described by Price, R. L.,
et al., Anat Rec, 245:83-93 (1996); and Angst, B. D., et al., Circ
Res, 80:88-94 (1997). FIG. 5 demonstrates a connexin 43 protein
expression pattern similar to that described in vivo, where
connexin 43 is distributed along the periphery of the cardiac
myocyte and in cell-cell junctions (A). For stereo imaging,
Z-series were collected at 1 .mu.m intervals to a maximum depth of
80 .mu.m. Reconstructed confocal microscopy Z-series through
cardiac myocyte seeded tubes after 4 weeks in culture, shown in
FIGS. 5 (B) and (C), show four to five layers of cardiac myocytes
aligned in parallel with the collagen fibrils in the tube wall.
[0146] Cardiac myocyte seeded tubes were processed for transmission
electron microscopy (TEM) as previously described in Price, R. L.,
et al., Anat Rec, 245:83-93 (1996). TEM images, like those in FIGS.
6 (A) and (B), demonstrate the presence of organized, aligned
myofibrils (F), well-developed sarcomeres (not shown), Z-bands (Z),
numerous mitochondria (M), cell:cell junctions (arrows in inset A),
and attachment to collagen fibrils (arrows in inset B),
characteristic of the cardiac myocyte phenotype seen in vivo.
Determination of Electrical Activity of Cardiac Myocyte Seeded
Tubes:
[0147] Seeded tubes were measured for electrical activity by
attaching two platinum wires at opposite ends of the tube, using a
third wire as a reference in the surrounding media. The wires were
shielded from interference by enclosing all but the distal end of
each wire within a plastic casing. These electrodes were connected
to an amplifier, an analog-to-digital converter, and finally to a
computer where the signal was recorded. One-hundred seconds of
electrical activity was collected for eight different tubes at 100
samples per second. The signal was then processed using a Fast
Fourier transform algorithm to show the electrical activity of
individual tubes. FIG. 7 (A) is a representative electrical signal
recorded from a tube after 4 weeks in culture, and FIG. 7 (B)
illustrates the Fast Fourier transform analysis of the electrical
data, demonstrating a dominant frequency at 3.4 Hz. This
corresponds to the contraction frequency of about 200 beats per
minute obtained by visual observation and counting of
contractions.
EXAMPLE 4
[0148] This example demonstrates the use of the disclosed tissue
scaffold for ventral hernia repair.
[0149] Collagen tubes can be made following the procedure described
in Example 2, or according to any of the methods set forth in this
disclosure. The collagen tubes are cut open longitudinally to form
sheets, and a thin layer of collagen solution is streaked on the
surface of the substrates and allowed to polymerize. This procedure
results in a thin layer of collagen fibrils that are arrayed in
parallel with one another along the direction that the collagen
solution was streaked. Skeletal myoblasts were plated on the
patterned collagen sheets. Over several days, the myoblasts
continued to fuse and a series of uniformly arrayed, densely packed
myotubes with morphology reflective of skeletal muscle
developed.
[0150] The majority of the cells placed on the aligned collagen are
MyoD positive myoblasts that fuse into multinucleate myotubes.
Confocal microscopy of a phalloidin stained culture demonstrates
parallel-aligned skeletal muscle cells with uniformly spaced
sarcomeres. Located between the skeletal muscle cells are a
population of fibroblast-like cells that are a mixture of myoblasts
and fibroblasts that can produce collagen. To increase the
thickness of the aligned skeletal muscle cultures and form a more
tissue-like construct, the collagen sheets with aligned myotubes
are placed in the Rotating Wall Bioreactor (RCCS) (Synthecon), and
additional cells are seeded to achieve multilayering. It is
possible to obtain as many as 40 layers of skeletal muscle cells in
the RCCS.
[0151] In order to use these constructs for repair of ventral
hernia, an experimental model of hernia in a rat is used. After
achievement of general anesthesia, the animal's abdomen is opened
and a 1 cm.sup.2 midline abdominal wall defect is created. At this
point, the engineered skeletal muscle tissue is surgically sewn
into the defect and the abdomen is closed. Topical antimicrobials
are placed on the incision and the animal is allowed to ascend from
anesthesia. Intramuscular analgesics are administered as well as
subcutaneous fluid boluses while the animal recovers. Tissue
integration is monitored post operatively by clinically inspecting
the surgical sight for signs of infection, bleeding, suture
failure, and extrusion of viscera through the rectus abdominous,
indicative of repair failure. Animals are sacrificed at specific
time points and the tissue is evaluated for specific events known
to occur throughout the wound healing process.
[0152] The implanted tissue is evaluated microscopically,
monitoring normal wound healing processes, neovascularization,
innervation, inflammation and infection. Inflammatory cells, tissue
morphology and detection of angiogenesis are investigated using
hematoxylin and eosin (H&E) stains. For this evaluation, 5 mm
diameter punch biopsies of the interface between the native tissue
and the repair material are obtained and serially sectioned. The
sections are evaluated for clinically significant signs of
rejection or infection such as extensive interstitial mononuclear
cell infiltration and edema as well as mild interstitial hemorrhage
and antibody-mediated rejection by the presence of significant
foamy macrophage populations.
[0153] Neovascularization should be apparent on the H&E slides.
Endothelial cells as well as the internal elastic lamina and
vascular smooth muscle cells are identified for arterial vessels.
The tunica intima, tunica media and vasa vasorum are identified for
venous vessels. The presence of red blood cells within the vessels
is a strong indication of their functionality. Collagen
accumulation and location are evaluated using Masson's trichrome
staining. Myoneural junctions should clearly display the myelinated
nerve fiber approaching the skeletal muscle fiber. As the axon
nears the muscle cell, it loses its myelin sheath and continues on
as a non myelinated axon which can be visualized by Sevier-Munger
staining for neural tissues. Motor end-plates should also be
apparent using Gwyn-Heardman staining method described in the Color
Atlas of Histology, 3.sup.rd ed. Baltimore: Lippincott Williams and
Wilkins (2000).
[0154] The tissue at the repair site is also analyzed
immunohistochemically for the temporal and spatial localization of
growth factors VEGF, PDGF, and NGF within and around the repair.
Two sets of the 2 mm biopsies are taken. One set is analyzed for
growth factors by enzyme-linked immunosorbent assay (ELISA) to
quantify the concentration of growth factors within the tissue
(ELISA kits: VEGF and PDGF from R& D systems, NGF from
Chemicon). The second set is fixed with paraformaldehyde,
sectioned, and immunohistochemically stained for VEGF, PDGF and NGF
using commercially available antibodies (R&D systems). The
antibody staining is imaged using a laser scanning confocal
microscope to localize the concentrations of growth factors
spatially within the tissue using the methods of Germani. A., et
al., Am J Path. 163(4):1417-1428 (2003). Briefly, samples are fixed
in 4% paraformaldehyde overnight at 4.degree. C. They are vibratome
sectioned and stained with 488-phalloidin for f-actin, and
immunohistochemically labeled using commercially-available
antibodies. Serial sections are labeled with either VEGF, PDGF or
NGF, one growth factor label per section. Secondary antibodies are
labeled with Texas Red. Immunohistochemical visualization is
performed using a Bio-Rad 1024 ES laser scanning confocal
microscope.
[0155] Additionally, the tissue is evaluated mechanically for
normal wound healing repair strength development. The indenter
force required to puncture the abdominal repair is measured using
an Instron.RTM. force measurement system. Measurements are made on
the abdominal repair by resecting the entire abdominal wall free
from the animal and placing it in ice cold Krebs Ringer solution
with 30 mM 2,3-butanedione monoxime (BDM) to prevent dissection
injury. Samples are mounted in a bath perfused with warm
(37.degree. C.) Krebs Ringer solution oxygenated by bubbling the
solution reservoir with a 95% O.sub.2/5% CO.sub.2 mixture. The
sample is placed in the stainless steel "picture frame" sample
holder which clamps the sample along the edges in a plane parallel
to the base of the Instron.RTM. and perpendicular to the indenter.
A 3 mm diameter spherical indenter is used to impinge on the center
of the repair site. The sample is indented at a rate of 3.0 mm/sec.
The force and displacement of the indenter and the peak force
required to burst through the repair site are measured and
recorded. Samples for mechanical testing are from animals that have
not undergone punch biopsies. Data will be analyzed using ANOVA
followed by pair-wise comparison.
[0156] All references cited in this specification, including
without limitation all papers, publications, patents, patent
applications, presentations, texts, reports, manuscripts,
brochures, books, internet postings, journal articles, periodicals,
and the like, are hereby incorporated by reference into this
specification in their entireties. The discussion of the references
herein is intended merely to summarize the assertions made by their
authors and no admission is made that any reference constitutes
prior art. Applicants reserve the right to challenge the accuracy
and pertinency of the cited references.
[0157] In view of the above, it will be seen that the several
advantages of the invention are achieved and other advantageous
results obtained.
[0158] As various changes could be made in the above methods and
compositions without departing from the scope of the invention, it
is intended that all matter contained in the above description and
shown in the accompanying drawings shall be interpreted as
illustrative and not in a limiting sense.
* * * * *