Note: Descriptions are shown in the official language in which they were submitted.
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ENGINEERED EXTRACELLULAR MATRICES CONTROL STEM CELL
BEHAVIOR
10 FIELD OF THE INVENTION
This invention relates to the preparation of a collagen based matrix for
culturing and differentiating stem cells and progenitor cells and the use of
such
compositions as tissue graft constructs.
BACKGROUND
The interaction of cells with their extracellular matrix (ECM) as it
occurs in vivo plays a crucial role in the organization, homeostasis, and
function of
tissues and organs. Continuous communication between cells and their
surrounding
ECM environment orchestrates critical processes such as the acquisition and
maintenance of differentiated phenotypes during embryo genesis, the
development of
form (morphogenesis), angiogenesis, wound healing, and even tumor metastasis.
Both biochemical and biophysical signals from the ECM modulate fundamental
cellular activities including adhesion, migration, proliferation, differential
gene
expression, and programmed cell death.
In turn, the cell can modify its ECM environment by modulating the
synthesis and degradation of specific matrix components. The realization of
the
significance of cell-ECM interaction has led to a renewed interest in
characterizing
ECM constituents and the basic mechanisms of cell-ECM interaction.
Tissue culture allows the study in vitro of animal cell behavior in an
investigator-controlled physiochemical environment. Presumably cultured cells
function best (i.e., proliferate and perform their natural in vivo functions)
when
cultured on substrates that closely mimic their natural environment.
Currently,
studies in vitro of cellular function are limited by the availability of cell
growth
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substrates that present the appropriate physiological environment for
proliferation and
development of the cultured cells. Complex scaffolds representing combinations
of
ECM components in a natural or processed form are commercially available, such
as
Human Extracellular Matrix (Becton Dickinson) and MATRIGEL . However, none
of the existing scaffolds have been prepared under conditions that regulate
the
polymerization of the scaffold in a controlled manner so as to produce a
composition
having mechanical properties and a predetermined 3D microstructure of collagen
fibrils and/or soluble ECM components that optimizes cell-substrate
interactions to
yield predictable and reproducible cellular outcomes. Applicants have
discovered that
the physical state of an ECM scaffold and not just its molecular composition
should
be considered in the design of new and improved scaffolds.
As reported herein, modifying the conditions used to form a collagen
based matrix from a solubilized collagen solution allows for the controlled
alteration
of the micro-structural and subsequent mechanical properties of the resulting
ECM
scaffold. Furthermore, the micro-structural and mechanical properties of the
ECM
scaffold directly impact fundamental cell behavior including survival,
adhesion,
proliferation, migration and differentiation of cells cultured within the
scaffold.
Basement membrane tissues and submucosal material harvested from
warm blooded vertebrates have shown great promise as unique graft materials
for
inducing the repair of damaged or diseased tissues in vivo, and for supporting
fundamental cell behavior (e.g., cell proliferation, growth, maturation,
differentiation,
migration, adhesion, gene expression, apoptosis and other cell behaviors) of
cell
populations in vitro. Submucosal material can be extracted or fluidized to
provide
enriched extracts that can be utilized as additives for tissue culture media,
or
polymerized to form collagen based scaffolds, to promote in vitro cell growth
and
proliferation.
As a tissue graft, submucosal tissue undergoes remodeling and induces
the growth of endogenous tissues upon implantation into a host. Numerous
studies
have shown that submucosal tissue is capable of inducing host tissue
proliferation,
remodeling and regeneration of tissue structures following implantation in a
number
of in vivo environments, including the urinary tract, the body wall, tendons,
ligaments,
bone, cardiovascular tissues and other vascular tissues, and the central
nervous
system. Upon implantation of the submucosal tissues, cellular infiltration and
a rapid
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neovascularization are observed and the submucosa materials are remodeled into
host
replacement tissue with site-specific structural and functional properties.
Accordingly, submucosa tissue can be used as a tissue graft construct,
for example, in its native form, in its fluidized form, in the form of an
extract, or as
components extracted from submucosa tissue and subsequently purified. The
fluidized forms of vertebrate submucosa tissue can be gelled to form a semi-
solid
composition that can be implanted as a tissue graft construct or utilized as a
cell
culture substrate. As a tissue graft material, the fluidized form can be
injected, or
delivered using other methods, to living tissues to enhance tissue remodeling.
Furthermore, the fluidized form can be modified, or can be combined with
specific
proteins, growth factors, drugs, plasmids, vectors, or other therapeutic
agents for
controlling the enhancement of tissue remodeling at the site of injection.
Moreover,
the fluidized, solubilized form can be combined with primary cells or cell
lines prior
to injection to further enhance the remodeling properties that result in the
repair or
replacement of diseased or damaged tissues.
Because the molecular forces that orchestrate the self assembly of
soluble, monomeric collagen into higher ordered structures are weak their
assembly
can easily turn into an unstructured aggregation of misfolded proteins. In the
literature, there are known methods for isolating collagen from a variety of
tissues,
e.g., placenta and animal tails and using the isolated material to
reconstitute
collagenous matrices. These known methods rely on the protein's intrinsic
ability to
retain its secondary structure during protein isolation and assume that, for
instance,
the alpha helix will retain its helical structure throughout. The end result,
even with a
homogenous biochemical composition, can be a heterogeneous secondary
structure.
Controlling the assembly of the constituting monomers into tertiary or
quaternary
multimeric arrangements is very hard to achieve under such conditions. One
embodiment of the present invention is directed to controlling the
polymerization of a
composition comprising solubilized collagen to form a collagen based scaffold
that
has the requisite microstructure and composition to allow for the expansion,
differentiation and/or clonal isolation of stem cells in a highly reproducible
and
predictable manner.
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SUMMARY
The present invention relates to compositions comprising a three
dimensional matrix that is formed to have the requisite composition and
microstructure to enhance the proliferation and /or differentiation of stem
cells or
progenitor cells cultured within such a matrix. In accordance with one
embodiment
an improved method for culturing stem cells is provided. The method comprises
preparing a solubilized collagen composition from a source of collagen, adding
cells
to the solubilized collagen composition and polymerizing the collagen
composition
under controlled conditions to provide a matrix formed from collagen fibrils
and
having the desired microstructure. In one embodiment cells are added to the
collagen
based matrix at a cell density within two orders of magnitude of the minimum
cell
number required to maintain cell viability, and the cells are cultured under
conditions
suitable for proliferation of the cells. In one embodiment the three
dimensional
matrix has a fibril area fraction (defined as the percent area of the total
area occupied
by fibrils in a cross-sectional surface of the matrix; providing an estimate
of fibril
density) of about 8% to about 26% and an elastic or linear modulus (defined by
the
slope of the linear region of the stress-strain curve) of about 0.5 to about
401cPa. In
one embodiment the three dimensional matrix is farther provided with an
exogenous
source of glucose and calcium chloride.
In accordance with one embodiment, stem cell seeded engineered
purified collagen based matrices are used as novel compositions for inducing
the
repair of damaged or disease tissues in vivo. In one embodiment the tissue
graft
construct comprises an engineered purified collagen based matrix, wherein the
matrix
is formed by contacting purified collagen with hydrochloric acid to produce a
solubilized collagen composition and subsequently polymerizing the solubilized
collagen composition under controlled conditions and in the presence of a
population
of cells to produce the engineered purified collagen based matrix containing
cells
entrapped within the matrix. In one embodiment the population of cells
comprises
stem cells initially added to the composition at a density of less than 105
cells per
milliliter, or the progeny of such stem cells. In one embodiment the stem cell
seeded
engineered purified collagen based matrices are implanted into a host without
culturing the seeded stem cells in vitro. In another embodiment the stem cell
seeded
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engineered purified collagen based matrix is further incubated under
conditions suitable for
inducing the proliferation and/or differentiation of the seeded stem cells.
In another embodiment the stem cells are added to the engineered purified
collagen based matrices at densities of less than 103 cells per milliliter and
the cells are
cultured under conditions that are minimally permissive for stem cell
functionality. These
conditions result in the production of localized populations of stem cells and
thus allow for the
isolation of clonal populations of stem cells. Accordingly, in one embodiment,
a method of
isolating clonal populations of individual stem cells is provided. The method
comprises the
steps of contacting a collagen based matrix with a low density of stem cells
wherein said
collagen matrix is formed by contacting a source of collagen with HC1 to
prepare a solubilized
collagen composition, polymerizing the solubilized collagen composition using
a final
collagen concentration of 1.0 to 3.0 mg/ml, at a pH of about 6.5 to about 7Ø
In one
embodiment the initial seeded population of stem cells ranges from about 10 to
about 103 cells
per milliliter. The seeded stem cells are cultured under conditions suitable
for proliferation of
the cells and individual populations of stem cells are isolated.
Specific aspects of the invention include:
- a composition for supporting stem cells, said composition comprising a
synthetic collagen-based matrix comprised of collagen fibrils; and a
population of stem cells,
wherein the matrix has an elastic or linear modulus of 0.5 kPa to 40 kPa;
- a method for culturing stem cells in vitro, said method comprising providing
a solubilized collagen composition; adding the stem cells at a cell density
within two orders of
magnitude of the minimum cell number required to maintain cell viability;
polymerizing the
collagen composition to form a matrix comprising the cells within a collagen
fibril network,
wherein the matrix has an elastic or linear modulus of 0.5 kPa to 40 kPa; and
providing
conditions conducive to cell growth;
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- a tissue graft construct comprising a synthetic collagen-based matrix
comprising collagen fibrils; and a population of stem cells, wherein said
synthetic collagen-
based matrix has a fibril area fraction of 7.7% to 25%, has an elastic or
linear modulus of 0.5
to 40 kPa and further comprises exogenously added glucose and calcium
chloride;
- use of the construct as described herein for enhancing the repair of tissues
at a
site in need of repair in a warm blooded vertebrate;
- a method of isolating a clonal population of individual stem cells, said
method comprising the steps of contacting a collagen matrix with a seed
population of stem
cells, wherein the density of the seed population ranges from 10 to 103 cells
per milliliter of
culture medium, wherein said collagen matrix is formed by contacting a source
of collagen
with hydrochloric acid to prepare a solubilized collagen composition;
polymerizing the
solubilized collagen composition using a collagen concentration of 1.0 to 3.0
mg/ml, at a pH
of 6.5 to 7.0, and culturing said seed population of stem cells and isolating
the clonal
population of stem cells;
- use of a mixture comprising stem cells and solubilized collagen for
enhancing
the repair of tissues in a warm blooded vertebrate, wherein said mixture is
for injection into a
desired site of said vertebrate, wherein the mixture is produced in vitro by
adding the stem
cells at a cell density within two orders of magnitude of the minimum cell
number required to
maintain cell viability, to a solubilized collagen composition, and wherein
upon polymerizing
the collagen composition in vivo, said mixture forms a collagen-based matrix
comprising the
stem cells entrapped within a collagen fibril network; and
- use of a collagen-based matrix comprising stem cells entrapped within a
collagen fibril network, for enhancing the repair of tissues in a warm blooded
vertebrate,
wherein said collagen-based matrix is for injection into a desired site of
said vertebrate, and
wherein the collagen-based matrix is produced in vitro by: adding the stem
cells at a cell
density within two orders of magnitude of the minimum cell number required to
maintain cell
viability, to a solubilized collagen composition to produce a mixture, and
inducing
polymerization of the mixture.
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BRIEF DESCRIPTION OF THE DRAWINGS
Figs. 1A-1F present data showing the effect of various parameters on the
stiffness (elastic or linear modulus) of the formed matrix. Fig. lA represents
the effect of
polymerization temperature on a matrix formed from a solubilized collagen
composition
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concentration on a matrix formed from a solubilized ECM collagen composition
in
1X PBS at 37 C.
Figs. 2A & 2B represent a series of graphs showing the quantification
of fibril area fraction (Fig. 2A) and fibril diameter distribution (Fig. 2B)
based upon
and compared for neonatal human dermal fibroblasts (NHDFs) seeded within 3D
15 Figs. 4A-4D represent a series of images depicting cell
contractility
and matrix remodeling by individual NHDFs resident within type I collagen (1.5
mg/ml) ECMs prepared with type III collagen concentrations of 0.25 mg/ml
(Figs. 4A
and 4B) and 0.75 mg/ml (Figs. 4C and 4D). Figs. 4A and 4C represent 2D
projections of confocal reflection image stacks showing changes to NHDF
Fig. 5 represents a graph depicting contractility and matrix remodeling
within engineered ECMs. NHDFs were grown within engineered ECMs in which the
Fig. 6 represents a graph showing that points of maximum local
deformation or strain induced within a 3D tissue construct, by low passage
neonatal
human dermal fibroblasts, occurred at distances further from the cell than for
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engineered ECMS prepared with lower amounts of type III collagen. NHDFs were
grown within engineered ECMs in which the type I collagen concentration was
kept
constant at 1.5 mg/ml and the amount of type III collagen was either 0.25
mg/ml or
0.75 mg/ml.
Figs. 7A & 7B represent a series of graphs depicting data regarding the
proliferation of low passage human dermal fibroblasts when grown within a 3D
ECM
format consisting of type I collagen ECMs prepared within increasing amounts
of
type III collagen (see Fig. 7A). NHDFs exposed to 2D ECM surface coatings
representing the same biochemical compositions and collagen type I/III ratios
showed
no significant changes in proliferative response (see Fig. 7B). Selected
relationships
showing statistically significant differences (p<0.05) are indicated with
symbols (*,
**,=,o).
Fig. 8 represents a bar chart showing differences in the expression of
select tissue-specific genes by multi-potential bone marrow derived
mesenchymal
cells grown on standard 2D plastic and within 3D ECM microenvironments of
increased fibril density and stiffness (elastic or linear modulus). Gene
expression
patterns for mesenchymal cells cultured within a given 2D or 3D format was
also
modulated by changing the composition of the culture medium.
Fig. 9 is a schematic representation of general cell behavior of multi-
potential bone marrow derived mesenchymal cells when cultured within 3D
matrices
that differ in collagen concentration to provide an ECM microenvironment
characterized by increased fibril density and stiffness (elastic or linear
modulus).
Points of arrow indicate low frequency events and wide ends of arrows indicate
high
frequency events.
DETAILED DESCRIPTION
DEFINITIONS
As used herein, the term "stem cell" refers to an unspecialized cell
from an embryo, fetus, or adult that is capable of self-replication or self-
renewal and
can develop into specialized cell types of a variety of tissues and organs.
The term as
used herein, unless further specified, encompasses totipotent cells (those
cells having
the capacity to differentiate into extra-embryonic membranes and tissues, the
embryo,
and all post-embryonic tissues and organs), pluripotent cells (those cells
that can
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differentiate into cells derived from any of the three germ layers), and
multipotent
cells (those cells having the capacity to differentiate into a limited range
of
differentiated cell types).
As used herein the term "progenitor cell" refers to a stem cell with
more specialization and less differentiation potential than a totipotent stem
cell. For
example, progenitor cells include unipotential cells (those cells having the
capacity to
differentiate along a single cell lineage).
As used herein, the term "lyophilized" relates to the removal of water
from a composition, typically by freeze-drying under a vacuum. However,
lyophilization can be performed by any method known to the skilled artisan and
the
method is not limited to freeze-drying under a vacuum. Typically, the
lyophilized
tissue is lyophilized to dryness, and in one embodiment the water content of
the
lyophilized tissue is below detectable levels.
As used herein "solubilized collagen composition" refers to a
composition that comprises collagen in a predominantly soluble monomeric form
(for
example wherein less than 20% of the collagen is insoluble, denatured, or
assembled
in higher ordered structures).
As used herein "solubilized extracellular matrix composition" refers to
a naturally occurring extracellular matrix that has been treated, for example,
with an
acid to reduce the molecular weight of at least some of the components of the
extracellular matrix and to produce a composition wherein at least some of the
components of the extracellular matrix have been solubilized from the
extracellular
matrix. The "solubilized extracellular matrix composition" may include
insoluble
components of the extracellular matrix as well as solubilized components.
As used herein the term "collagen-based matrix" refers to extracellular
matrices that comprise collagen. An "engineered purified collagen based
matrix" as
used herein relates to a composition comprising a collagen fibril scaffold
that has
been formed under controlled conditions from a solubilized collagen
composition,
wherein the solubilized collagen composition is prepared from a composition
consisting essentially of collagen. The conditions controlled during the
polymerization reaction include one or more of the following: pH, phosphate
concentration, temperature, buffer composition, ionic strength, and
composition and
concentration of purified collagen components. Similarly, an "engineered
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extracellular matrix" relates to a solubilized extracellular matrix
composition that is
polymerized to form a collagen fibril containing matrix under controlled
conditions,
wherein the controlled conditions include pH, phosphate concentration,
temperature,
buffer composition, ionic strength, and composition and concentration of the
extracellular matrix components which includes both collagen and non-
collagenous
molecules. A "bioactive engineered extracellular matrix" composition refers to
an
engineered extracellular matrix composition that can be polymerized to form a
three
dimensional scaffold that is capable of remodeling tissues in vivo.
As used herein the term "naturally occurring extracellular matrix"
comprises any noncellular material naturally secreted by cells (such as
intestinal
submucosa) isolated in their native configuration with or without naturally
associated
cells.
As used herein the term "submucosal matrices" refers to natural
extracellular matrices, known to be effective for tissue remodeling, that have
been
isolated in their native configuration, including submucosa derived from
vertebrate
intestinal tissue, stomach tissue, bladder tissue, alimentary tissue,
respiratory tissue
and genital tissue.
As used herein the term "exogenous" or "exogenously added"
designates the addition of a new component to a composition, or the
supplementation
of an existing component already present in the composition, using material
from a
source external to the composition.
As used herein "sterilization" or "sterilize" or "sterilized" means
removing unwanted contaminants including, but not limited to, endotoxins,
nucleic
acid contaminants, and infectious agents.
As used herein "stiffness" or elastic or linear modulus" refers to the
fundamental material property defined by the slope linear portion of a stress-
strain
curve that results from conventional mechanical testing protocols.
As used herein, the term "purified" and like terms relate to the isolation
of a molecule or compound in a form that is substantially free from other
components
with which they are naturally associated (e.g., the total amount of
nondesignated
components present in the composition representing less than 5%, or more
typically
less than 1%, of total dry weight).
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As used herein the term "three dimensional purified collagen matrix
(3D matrix)" refers to an engineered purified collagen based matrix, as
defined above,
and the fluid that surrounds the collagen fibril network. A "3D purified
collagen
matrix populated/seeded with cells" further comprises a viable population of
cells
contained within the matrix.
As used herein the term "three dimensional extracellular matrix (3D
ECM)" refers to an engineered extracellular matrix, as defined above, and the
fluid
that surrounds the collagen fibril network. A "3D extracellular matrix
populated/seeded with cells" further comprises a viable population of cells
contained
within the matrix.
As used herein the term "three dimensional matrix (3D matrix)" is a
generic term that is intended to include both "three dimensional purified
collagen
matrices (3D purified collagen matrices)" as well as "three dimensional
extracellular
matrices (3D ECM).
As used herein the term "collagen fibril" refers to a quasi-crystalline,
filamentous structure formed by the self-assembly of soluble trimeric collagen
molecules. The collagen molecules in a collagen fibril typically pack in a
quarter-
staggered pattern giving the fibril a characteristic striated appearance or
banding
pattern along its axis. Solubilized collagen that is assembled in vitro to
form collagen
fibrils exhibit similarities to collagen structures found in vivo (Veis and
George, 1994
Fundamental of interstitial collagen assembly. In: Yurchenco PD, Birk DE, and
Mecham RP (eds.), Extracellular Matrix Assembly and Structure, Academic Press,
Inc., San Diego, pp.15-45.). Within tissues in vivo, collagen fibrils are
organized as
bundles in a hierarchical manner to form fibers. Collagen fibers are further
organized
in a tissue-specific fashion to provide tissues with specific structural-
functional
properties. Collagen fibrils are distinct from the amorphous aggregates or
precipitates
of insoluble collagen that can be formed by dehydrating (e.g., lyophilization)
collagen
suspensions to form porous network scaffolds. Collagen networks formed from
amorphous aggregates, or precipitates of insoluble collagen, have
characteristics that
distinct from those formed from collagen fibrils as defined above.
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EMBODIMENTS
Cell culture scaffolds presenting a more biologically relevant
microenvironment are disclosed. More particularly, these cell culture
scaffolds
comprise three-dimensional matrices/biomaterials that are created from
solubilized
collagen compositions. The solubilized collagen compositions are prepared from
biological sources, such as naturally occurring extracellular matrices,
including for
example submucosal matrices. More particularly, the soluble polymers suitable
for
use in the present invention can be isolated, to varying degrees of purity,
from natural
tissues and include, but are not limited to, type I collagen, type III
collagen, growth
factors and glycosaminoglycans. In one embodiment the solubilized collagen
composition comprises purified type I collagen or a mixture of purified type I
and
type III collagen. When provided with the proper conditions, the solubilized
collagen
composition undergoes polymerization/self assembly to form a three dimensional
scaffold/biomaterial comprised of collagen fibrils. In one embodiment the
soluble
polymers comprise type I collagen monomers, where upon polymerization the
resulting scaffolds represent a composite material comprising insoluble
collagen
fibrils and an interfibrillar fluid component, referred to herein as a three
dimensional
matrix.
An array of scaffolds/biomaterials can be created by varying the
composition of ECM molecules as well as the self-assembly/polymerization
conditions. Surprisingly, applicants have discovered that upon seeding
progenitor
cells or stem cells within engineered purified collagen based matrices
(scaffolds)
representing different microstructural compositions (e.g., having different
dimensioned and organizations of the collagen fibrils and filaments), distinct
patterns
of cell survival, growth, proliferation, and differentiation are obtained. In
particular,
applicants have discovered that engineered purified collagen based matrices
representing different microstructural compositions (e.g., varied fibril
dimensions
(length, diameter) and densities) will impact the rate of cell proliferation
as well as the
pattern of cellular condensation, aggregation, fusion, and cellular
differentiation
events and their associated time-line. These results are significant because
they
indicate that engineered purified collagen based matrices can be specifically
designed
to foster the proliferation of stem cells while maintaining their precursor or
multi-
potential status. Furthermore, engineered purified collagen based matrices can
be
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designed to direct differentiation of cells down a specific cell lineage (such
as fat,
bone, muscle, or cartilage) to form 3D organotypic tissues (that is
reminiscent of in
vivo tissue structure and function).
In accordance with one embodiment, stem cells and/or progenitor cells
are seeded at relatively low densities on or within the various engineered
purified
collagen based matrices. It is known that, in general, cell behavior is
determined by a
combination of signal inputs arising from soluble factors, biophysical
factors, the
extracellular matrix substrate, and cell-cell interactions. Seeding cells at a
relative
low cell density on or within the collagen based matrices of the present
invention
allows ECM-based signaling to predominate over signals derived from cell-cell
interactions. In accordance with one embodiment, cells are initially seeded on
or
within the engineered purified collagen based matrices at a minimal cell
density that
will allow for cell viability and replication (i.e., the minimal functionality
density).
This minimal functional density can be easily established for the particular
cell type to
be cultured and for the specific culture conditions utilized.
In accordance with one embodiment the stem cells or progenitor cells
are seeded within the collagen based matrix at a cell density substantially
higher than
the minimal functionality density but at a relative low density compared to
standard
cell culture techniques. In one embodiment the cells comprise stem cells,
wherein the
cells are seeded at a density within 3 orders of magnitude of the minimal
functionality
density, in another embodiment stem cells are seeded at a density within 2
orders of
magnitude of the minimal functionality density, and in another embodiment the
stem
cells are seeded at a density within an order of magnitude of the minimal
functionality
density. The stem cells can be seeded at a relatively high density of about 1X
106 to
about 1 X 108 cells/ml, or at a more typical density of about 1X 103 to about
1 X 105
cells/ml. Seeding the cells at the relative high density of about 1X 106 to
about 1 X
108 cells/ml will promote cell to cell interactions over cell to matrix
interactions.
Accordingly, stem cells seeded at relatively high densities will develop into
fat tissue
even when the cells are cultured within 3D matrices of high collagen fibril
density. In
one embodiment stem cells are seeded at a density of less than 5 X 104
cells/ml, more
typically at a density of about 5 X 104 cells/ml. In another embodiment stem
cells are
seeded at a density of less than 1 X 104 cells/ml, in another embodiment stem
cells are
seeded at a density selected from a range of about 1 X 102 to about 5 X 103.
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As disclosed herein an improved method for culturing stem cells is
provided that uses three dimensional purified collagen based matrices. The
improved
method allows for enhanced proliferation of stem cells as well as better
control over
the differentiation of the cultured cells. In one embodiment the method
comprises the
steps of providing a solubilized collagen composition, adding cells to the
collagen
composition, and polymerizing the solubilized collagen composition to form
collagen
fibrils. The solubilized collagen composition comprises collagen that has been
isolated with or without additional components from natural tissues.
In accordance with one embodiment the solubilized collagen
composition is prepared using purified type I collagen as a starting material.
In one
embodiment collagen, and more particularly type I or type III collagen, that
has been
isolated from tissues is subjected to a final purification step that removes
any reagents
that were used during the isolation steps. In one embodiment the final
purification
step comprises dialyzing the isolated collagen in an aqueous solution, and in
one
embodiment the isolated collagen is dialyzed against a dilute acid solution,
including
for example, hydrochloric acid. In one embodiment the final purification step
comprises dialyzing the isolated collagen against a 0.01 N HC1 solution.
Isolated type
I or isolated type III collagen preparations are commercially available, and
these
commercially available materials are subjected to a further purification step,
including
for example, dialyzing against a dilute (about 0.001 N to about 0.1 N)
hydrochloric
acid solution to produce purified collagen suitable for use for forming 3D
purified
collagen matrices. The dialysate can optionally be filtered and/or centrifuged
to
remove particulate matter. In accordance with one embodiment, the collagen
component of the solubilized collagen composition consists essentially of
purified
collagen, the majority of which are in monomeric form. In a further embodiment
the
composition is formed from purified collagen (the majority of which are in
monomeric form) that is greater than 75% type I collagen, or greater than 90%
type I
collagen. In one embodiment a composition consisting essentially of purified
collagen is dissolved in an acid solution, such as hydrochloric acid to
prepare a
solubilized collagen composition of the desired concentration. In one
embodiment
the purified collagen is dissolved in about 0.001 N to about 0.1 N, from about
0.005 N
to about 0.1 N, from about 0.005 N to about 0.01 N, from about 0.01 N to about
0.1
N, from about 0.05 N to about 0.1 N, from about 0.001 N to about 0.05 N, about
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0.001 N to about 0.01 N, or from about 0.01 N to about 0.05 N hydrochloric
acid
solution.
In another embodiment, a three dimensional purified collagen matrix is
provided, wherein the matrix is formed from a solubilized collagen composition
wherein the collagen components of the solubilized collagen composition
consist
essentially of purified type I and type III collagen. The component fibrils of
such
matrices have been found to have a greater degree of flexibility relative to
the fibrils
of engineered purified collagen matrices that are formed using only type I
collagen.
In one embodiment the matrix comprises type I collagen and type III collagen
in a
ratio of 200:1. The method of forming matrices with fibrils that exhibit a
higher
degree of flexibility comprises the steps of combining in vitro at least 100
ug/ml of
type I collagen with at least 0.5 ug/ml of type III collagen to obtain a total
amount of
collagen, and forming in vitro a three dimensional purified collagen matrix
wherein
the three dimensional matrix has decreased stiffness compared to a 3D matrix
formed
in vitro with type I collagen when the total amount of collagen in the two
matrices is
equivalent.
In another embodiment, a method of preparing an extracellular matrix
composition is provided. The method comprises the steps of combining in vitro
at
least 100 ug/ml of type I collagen with at least 0.5 ug/ml of type III
collagen to obtain
a total amount of collagen, and forming in vitro a three dimensional matrix.
In one
embodiment the type I and type III collagen is dissolved in about 0.001 N to
about 0.1
N, from about 0.005 N to about 0.1 N, from about 0.005 N to about 0.01 N, from
about 0.01 N to about 0.1 N, from about 0.05 N to about 0.1 N, from about
0.001 N to
about 0.05 N, about 0.001 N to about 0.01 N, or from about 0.01 N to about
0.05 N
hydrochloric acid solution either before or after the combining step.
In another embodiment, an extracellular matrix composition for use in
repairing diseased or damaged tissues is provided. The extracellular matrix
composition comprises at least 100 ug/ml of type I collagen and at least 0.5
ug/ml of
type III collagen, wherein the type I collagen to type III collagen ratio is
selected from
the group consisting of 200:1, 100:1, 50:1, 15:1, 10:1, 8:1, 6:1, 5:1, 3:1,
and 2:1, and a
population of cells. The matrix is formed by provided a solubilized collagen
composition comprising type I and type III collagen, in a ratio selected from
the group
consisting of 200:1, 100:1, 50:1, 15:1, 10:1, 8:1, 6:1, 5:1, 3:1, and 2:1,
polymerizing
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the solubilized collagen composition to form collagen fibrils, and adding
cells to the
collagen composition either before or after the polymerization step. In one
embodiment a composition comprising solubilized collagen and stem cells is
injected
into a host and the polymerization of the solubilized collagen composition
occurs in
vivo to form a cell entrapping matrix. Alternatively, the solubilized collagen
composition can be polymerized in vitro and the polymerized matrix, comprising
the
population of cells, can be subsequently injected or implanted in a host. In
another
embodiment the population of cells entrapped within the 3D matrix can be
cultured in
vitro, for a predetermined length of time, to increase cell numbers and/or
induce
differentiation of the cell population prior to implantation into a host. In a
further
embodiment, the population of cells can be cultured in vitro, for a
predetermined
length of time, to increase cell numbers and/or induce differentiation of the
cell
population and the cells can be separated from the matrix and implanted into
the host
in the absence of the polymerized matrix.
In one illustrative embodiment, the engineered purified collagen based
matrix comprises type III collagen in the range of about 0.5% to about 33% of
total
collagen in the matrix. In another illustrative embodiment, the engineered
purified
collagen based matrix comprises type I collagen in the range of about 66% to
about
99.5% of total collagen in the matrix. In yet another illustrative embodiment,
the
type I collagen to type III collagen ratio is in the range of about 2:1 to
about 200:1,
wherein the type I collagen to type III collagen ratio may be selected from
the group
consisting of 200:1, 100:1, 50:1, 15:1, 10:1, 8:1, 6:1, 5:1, 3:1, and 2:1.
In another embodiment, a method of enhancing cell proliferation
within an extracellular matrix composition is provided. The method comprises
the
steps of combining in vitro an amount of type I collagen with an amount of
type III
collagen to obtain a total amount of collagen wherein the ratio of type III
collagen to
type I collagen is at least 1:6, and forming in vitro a three-dimensional
extracellular
matrix wherein the extracellular matrix enhances cell proliferation compared
to an
extracellular matrix formed in vitro with type I collagen wherein the amount
of type I
collagen is equivalent to the total amount of type I collagen in the combining
step. In
yet another embodiment, the method comprises the steps of combining in vitro
at least
3 ug/ml of type I collagen with at least 0.5 ug/ml of type III collagen to
obtain a total
amount of collagen wherein the ratio of type III collagen to type I collagen
is at least
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1:6, and forming in vitro a three-dimensional extracellular matrix wherein the
extracellular matrix enhances cell proliferation compared to an extracellular
matrix
formed in vitro with type I collagen, wherein the amount of type I collagen is
equivalent to the total amount of type I collagen in the combining step.
In another illustrative embodiment, the method of preparing an
engineered purified collagen based matrix comprises combining type I and type
III
collagen wherein the type III collagen is added in the range of about 17% to
about
33% of total collagen in the matrix. In another illustrative embodiment, the
type I
collagen is added in the range of about 66% to about 83% of total collagen in
the
matrix. In yet another illustrative embodiment, the type I collagen to type
III
collagen ratio is in the range of about 6:1 to about 1:1, wherein the type I
collagen to
type III collagen ratio may be selected from the group consisting of 6:1, 5:1,
4:1, 3:1,
2:1, and 1:1.
Applicants have also discovered that the concentration of total collagen
present in solubilized collagen composition will impact the microstructure of
the
matrix, and the behavior of stem cells cultured within a matrix polymerized
from such
a composition (see Fig. 9). 3D matrices can be prepared from solubilized
collagen
compositions having purified collagen concentrations ranging from as little as
0.05
mg/ml to as much as 40mg/ml. Typically the 3D matrices are prepared from
purified
solubilized collagen compositions having a collagen concentration selected
from a
range of about 0.1mg/m1 to about 5.0mg/ml, and in one embodiment about 1.5
mg/ml
to about 3.0 mg/ml. Table 1 summarizes the effect of total collagen
concentration on
the fibril structure of the matrix:
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Table 1: Microstructure and Mechanical Properties of 3D Purified Collagen
Matrices
Collagen Fibril Area Stiffness Fibril Diameter Fibril
Diameter
Concentration Fraction (Linear (confocal (scanning
(Density; %) Modulus; lcPa) reflection electron
microscopy; microscopy,
nm) nm)
0.3 mg/ml, pH 1.54 0.507 418 121
7.4
1 mg/ml, 11.5 1.9 10.7 1.93 446 65
7.4
1.5 mg/ml, pH 12 1.4 8.5 1.65 412.63 76 115.16 23.18
7.4
2 mg/ml, pH 14.8 4.25 16.6 2.68 435 61 80.8 18.3
7.4
3 mg/ml, pH 18.4 1.9 24.3 4.16 430 71
7.4
2 mg/ml, pH 6 1.84 0.701 490 96
2 mg/ml, p117 12.7 1.18 469 73
2 mg/ml, pH 16.6 2.68 435 61
7.4
2 mg/ml, pH 8 22.5 3.65 421 62
2 mg/ml, pH 9 33.0 6.93 392 65
1.5 mg/ml type 21.5 2.6 13.3 1.4 385 72 87 17
I + 0.75 mg/ml
type III
Using the data of Table 1 and assuming a linear relationship between collagen
concentration and the measure properties, predictions of fibril area fraction
and matrix
stiffness can be determined as a function of collagen concentration using the
following equations:
Fibril Area Fraction = 3.6514 X Collagen Concentration + 7.3286
R2 -= 0.9681
Stiffness = 8.1145 X Collagen Concentration - 0.3306
R2 = 0.9304
Prediction of Stiffness as a function of Fibril Diameter (Assumption: fibril
area
fraction does not change; relationship based upon pH data):
Stiffness = -0.2916 X Fibril Diameter + 146.02
R2 = 0.9581 (based upon pH data)
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The 3D matrices formed in accordance with the present disclosure
represent a matrix of collagen fibrils. The fibrils of the matrices are formed
at a fibril
area fraction (density) of about 7.7% to about 25% total volume. In one
embodiment
the 3D matrices have a fibril area fraction of about 12.8% to about 18.3%
total
volume. In another embodiment the 3D matrices have a fibril area fraction of
about
18.5% to about 25% total volume. In one embodiment the 3D matrix has a fibril
area
fraction of about 12.8% to about 18.3% total volume and the fibrils have a
hydrated
diameter of about 350 to about 475nm. In another embodiment the 3D matrix has
a
fibril area fraction of about 18.5% to about 25% total volume and the fibrils
have a
hydrated diameter of about 375 to about 500nm.
Three dimensional matrices having low fibril density and low stiffness
enhance stem cell proliferation with decreased differentiation of the cells.
Accordingly, 3D matrices formed from solubilized collagen compositions having
about 0.1 mg/ml to about 3 mg/ml collagen, and more typically about 0.5 mg/ml
to
about 2.5 mg/ml collagen are utilized to stimulate stem cell proliferation.
The 3D
matrices so formed will have a fibril predicted fibril area fraction (density)
of about
7.7% to about 18.3% total volume and about 9.2% to about 16.5% total volume,
respectively. In one embodiment the 3D matrices are formed from solubilized
collagen compositions having about 3 mg/ml to about 1.5 mg/ml collagen and in
one
embodiment the solubilized collagen compositions have about 2.5, 2.0, 1.5, or
1.0
mg/ml of collagen. Alternatively, higher concentrations of total collagen
present in
the three dimensional matrix leads to differentiation of stem cells.
Accordingly, 3D
matrices (having a fibril area fraction of at least about 18% total volume)
formed from
solubilized collagen compositions having more than about 3 mg/ml are utilized
to
stimulate differentiation of stem cells cultured within the matrix. In one
embodiment
the 3D matrices are formed from solubilized collagen compositions having about
3.2,
3.4, 3.6, 3.8, 4.0, 4.5 or 5.0 mg/ml of collagen, resulting in 3D matrices
having a fibril
area fraction of about 19%, 19.7%, 20.5%, 21.2%, 22%, 23.8% and 25.6% total
volume, respectively.
As reported herein the relative stiffness (elastic or linear modulus) of a
3D matrix can be modified by controlling the relative proportion of type Ito
type III
collagen, the fibril area fraction (density), or the fibril diameter of the
collagen fibrils
in the 3D matrix. In accordance with one embodiment 3D matrices are prepared
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having a relatively low stiffness (elastic or linear modulus) of about 0.48 to
about
24.0 kPa. In one embodiment these matrices are used to propagate stem cells
and
progenitor cells without further differentiation of the cells and/or their
progeny. In
another embodiment 3D matrices are prepared having a relatively high stiffness
of
about 25 to about 40 kPa. In one embodiment these relatively stiffer matrices
are
used to induce the differentiation of stem cells and progenitor cells and/or
their
progeny. In one embodiment a 3D matrix is provided having a relatively low
stiffness
of about 0.48 to about 24.0 kPa and a relatively low fibril area fraction
(density) of
about 7% to about 18% total volume. In an alternative embodiment a 3D matrix
is
provided having a relatively high stiffness of about 25 to about 40 kPa and a
relatively
high fibril area fraction (density) of about 19% to about 26% total volume.
In another embodiment the solubilized collagen composition comprises
collagen monomers isolated from natural tissues, and includes additional
components
that are naturally associated with the native tissues and/or exogenously added
components. In one embodiment various exogenous materials, such as growth
factors
are added to the collagen based matrices of the present invention. In one
embodiment
the solubilized collagen composition represents a solubilized fraction of a
naturally
occurring extracellular matrix, and in one embodiment the naturally occurring
extracellular matrix is a vertebrate submucosal matrix. In one embodiment the
solubilized collagen composition represents a solubilized fraction of
vertebrate
intestinal submucosa.
In other embodiments, acetic acid, formic acid, lactic acid, citric acid,
sulfuric acid, ethanoic acid, carbonic acid, nitric acid, or phosphoric acid
can be used
to solubilize the naturally occurring extracellular matrix (or a purified
lyophilized
collagen composition) to produce a solubilized collagen composition. The
solubilized
collagen composition derived from a naturally occurring extracellular matrix,
such as
vertebrate intestinal submucosa, can then be polymerized to form an engineered
extracellular matrix.
The invention also relates to methods of preparation and compositions
comprising solubilized extracellular matrix components polymerized in vitro
where
the extracellular matrix components are solubilized by other methods known in
the
art. The polymerizing step can be performed under conditions that are
systematically
varied where the conditions are selected from the group consisting of pH,
phosphate
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concentration, temperature, buffer composition, ionic strength, the
extracellular
matrix components in the solubilized extracellular matrix composition, and the
concentration of the extracellular matrix components in the solubilized
extracellular
matrix composition.
In accordance with one embodiment a method of forming a 3D matrix
comprising stem cells is provided. The method comprises the steps of providing
an
acid solubilized purified type I collagen composition. In one embodiment the
collagen composition further comprises type III collagen. In one embodiment
the
purified collagen represents a commercially available isolated preparation of
collagen
that is further subjected to purification, including for example dialyzing
against an
solution of about 0.005 N to about 0.1 N HC1, more typically about 0.01 N HC1.
Typically the solubilized collagen composition comprises purified collagen
that is
suspended in about 0.005 N to about 0.1 N HC1 solution, and in one embodiment
suspended in 0.01N HC1. The solubilized collagen composition is also typically
sterilized using standard techniques including for example contact with
chloroform or
peracetic acid. Stem cells are then added to the solubilized collagen
composition at a
specific density, typically ranging from about 1 X 103 to about 1 X 108. In
one
embodiment the stem cells are added to the solubilized collagen composition at
a
concentration of less than 5 X 104 cells per milliliter, and in one embodiment
the cells
are added at a density of about 10 to about 103 per milliliter. In accordance
with one
embodiment the collagen/cell suspension is then pipetted into a well plate and
allowed to polymerize in a humidified environment at 37 C for approximately 30
minutes. In an alternative embodiment the collagen/cell suspension is injected
into a
host and the composition is polymerized in vivo.
As noted above solubilized collagen compositions can be prepared
from vertebrate submucosal matrices wherein the collagen compositions comprise
additional components besides collagen. Vertebrate submucosal matrices can be
obtained from various sources, including intestinal tissue harvested from
animals
raised for meat production, including, for example, pigs, cattle and sheep or
other
warm-blooded vertebrates. According to one embodiment the solubilized collagen
composition is derived from one or more sources selected from the group
consisting
of intestinal submucosa, stomach submucosa, urinary bladder submucosa, uterine
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submucosa, and any other submucosal material that can be used to remodel
endogenous tissue.
In one embodiment the submucosa comprises the tunica submucosa
delaminated from both the tunica muscularis and at least the luminal portion
of the
tunica mucosa of a warm-blooded vertebrate. Such constructs can be prepared by
mechanically removing the luminal portion of the mucosa and the external
muscle
layers and lysing resident cells with hypotonic washes.
It is known that compositions comprising the tunica submucosa
delaminated from both the tunica muscularis and at least the luminal portion
of the
tunica mucosa of the submucosal tissue of warm-blooded vertebrates can be used
as
tissue graft materials (see, for example, U.S. Patents Nos. 4,902,508 and
5,281,422).
Such submucosal tissue preparations are
characterized by excellent mechanical properties, including high compliance,
high
tensile strength, a high burst pressure point, and tear-resistance.
Submucosa-derived matrices are collagen based biodegradable
matrices comprising highly conserved collagens, glycoproteins, proteoglycans,
and
glycosaminoglycans in their natural configuration and natural concentration.
Such
submucosal material serves as a matrix for the regrowth of endogenous tissues
at the
implantation site (e.g., biological remodeling). The submucosal material
serves as a
rapidly vascularized matrix for support and growth of new endogenous
connective
tissue. Thus, submucosa matrices have been found to be trophic for host cells
of
tissues to which it is attached or otherwise associated in its implanted
environment.
In multiple experiments submucosal tissue has been found to be remodeled
(resorbed
and replaced with auto genous differentiated tissue) to assume the
characterizing
features of the tissue(s) with which it is associated at the site of
implantation or
insertion.
Small intestinal submucosa tissue is an illustrative source of
submucosal tissue for use in this invention. Submucosal tissue can be obtained
from
various sources, for example, intestinal tissue can be harvested from animals
raised
for meat production, including, pigs, cattle and sheep or other warm-blooded
vertebrates. Small intestinal submucosal tissue is a plentiful by-product of
commercial meat production operations and is, thus, a low cost material.
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Suitable intestinal submucosal tissue typically comprises the tunica
submucosa delaminated from both the tunica muscularis and at least the huninal
portion of the timica mucosa, but other tissue constructs can also be used. In
one
illustrative embodiment the intestinal submucosal tissue comprises the tunica
submucosa and basilar portions of the tunica mucosa including the lamina
muscularis
mucosa and the stratum compactum which layers are known to vary in thickness
and
in definition dependent on the source vertebrate species.
The preparation of submucosal tissue is described in U.S. Patent. No.
4,902,508. A
segment of vertebrate intestine, for example, preferably harvested from
porcine, ovine
or bovine species, but not excluding other species, is subjected to abrasion
using a
longitudinal wiping motion to remove the outer layers, comprising smooth
muscle
tissues, and the innermost layer, i.e., the luminal portion of the tunica
mucosa. The
submucosal tissue is rinsed under hypotonic conditions, such as with water or
with
saline under hypotonic conditions, and is optionally sterilized.
The submucosal tissue can be sterilized using conventional
sterilization techniques including glutaraldehyde tanning, formaldehyde
tanning at
acidic pH, propylene oxide or ethylene oxide treatment, gas plasma
sterilization,
gamma radiation, electron beam, and/or peracetic acid sterilization.
Sterilization
techniques which do not adversely affect the structure and biotropic
properties of the
submucosal tissue can be used. An illustrative sterilization technique is
exposing the
submucosal tissue to peracetic acid, 1-4 Mrads gamma irradiation (or 1-2.5
Mrads of
gamma irradiation), ethylene oxide treatment, exposure to chloroform, or gas
plasma
sterilization. The submucosal tissue can be subjected to one or more
sterilization
processes. In illustrative embodiments, the intact extracellular matrix
material can be
sterilized with peracetic acid or the solubilized collagen composition can be
sterilized.
The submucosal tissue can be subjected to one or more sterilization processes.
The
submucosal tissue can be stored in a hydrated or dehydrated state prior to
solubilization in accordance with the invention.
Extracellular matrix-derived tissues other than intestinal submucosa
tissue may be used in accordance with the methods described herein and used as
a
source for preparing solubilized collagen compositions. Methods of preparing
and
treating other extracellular matrix-derived tissues are known to those skilled
in the art
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and may be similar to the methods described above. For example, see U.S.
Patents
Nos. 5,163,955 (pericardial tissue), 5,554,389 (urinary bladder submucosa
tissue),
6,099,567 (stomach submucosa tissues), 6,576,265 (extracellular matrix tissues
generally), 6,793,939 (liver basement membrane tissues), and U.S. patent
application
publication no. US 2005/0019419 Al (liver basement membrane tissues), and
WO 01/45765 (extracellular matrix tissues generally).
The preparation and use of submucosa tissues as graft compositions is also
described in U.S. Patents Nos. 4,902,508, 5,281,422, and 5,275,826.
In one illustrative embodiment, the extracellular matrix material is
solubilized with an acid and the solubilized fraction is recovered for
polymerization to
form the collagen based matrices of the present invention. Typically, prior to
solubilization, the source extracellular matrix material is comminuted by
tearing,
cutting, grinding, or shearing the harvested extracellular matrix material. In
one
illustrative embodiment, the extracellular matrix material can be comminuted
by
shearing in a high-speed blender, or by grinding the extracellular matrix
material in a
frozen or freeze-dried state, and then lyophilizing the material to produce a
powder
having particles ranging in size from about 0.1 mm2 to about 1.0 mm2. The
extracellular matrix material powder can thereafter be hydrated with, for
example,
water or buffered saline to form a fluid or liquid or paste-like consistency.
In one
illustrative embodiment, the extracellular matrix tissue is comminuted by
freezing and
pulverizing under liquid nitrogen in an industrial blender. The preparation of
fluidized forms of the source extracellular matrix material, such as submucosa
tissue,
is described in U.S. Patent No. 5,275,826.
In one illustrative embodiment, an acid, such as hydrochloric acid,
acetic acid, formic acid, sulthric acid, ethanoic acid, carbonic acid, nitric
acid, or
phosphoric acid, is used to solubilize the source extracellular matrix
material. In
various illustrative embodiments, the acidic conditions for solubilization can
include
solubilization at about 0 C to about 60 C, and incubation periods of about 5
minutes
to about 96 hours. In other illustrative embodiments, the concentration of the
acid,
such as hydrochloric acid, can be from about 0.001 N to about 0.1 N, from
about
0.005 N to about 0.1 N, from about 0.01 N to about 0.1 N, from about 0.05 N to
about
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0.1 N, from about 0.001 N to about 0.05 N, about 0.001 N to about 0.01 N, or
from
about 0.01 N to about 0.05 M. However, the solubilization can be conducted at
any
temperature, for any length of time, and at any concentration of acid.
Any of the source extracellular matrix materials described above can
be used and the solubilization step can be performed in the presence of an
acid or in
the presence of an acid and an enzyme. The acid solubilization step results in
a
solubilized extracellular matrix composition that remains bioactive (i.e., is
capable of
polymerizing and remodeling tissues in vivo) after lyophilization.
In one illustrative embodiment, the extracellular matrix material is
treated with one or more enzymes before, during, or after the acid
solubilization step.
For enzymes that are inactive at acidic pH, for example, the extracellular
matrix
material is treated with the enzyme before the acid solubilization step or
after the acid
solubilization step, but under conditions that are not acidic. Enzymatic
digestion of
the extracellular matrix material is conducted under conditions that are
optimal for the
specific enzyme used and under conditions that retain the ability of the
solubilized
components of the extracellular matrix material to polymerize. The
concentration of
the enzyme depends on the specific enzyme used, the amount of extracellular
matrix
material to be digested, the desired time of digestion, and the desired
temperature of
the reaction. In various illustrative embodiments, about 0.01% to about 0.5%
(weight
per volume, such that 0.01% is equivalent to 0.01 g/100 ml) of enzyme is used.
Exemplary enzymes include pepsin, bromelain, cathepsins, chymotrypsin,
elastase,
papain, plasmin, subtilisin, thrombin, trypsin, matrix metalloproteinases
(e.g.,
stromelysin, elastase), glycosaminoglycan-specific enzymes (e.g.,
chondroitinase,
hyaluronidase, heparinase)and the like, or combinations thereof. The source
extracellular matrix material can be treated with one or more enzymes. In
illustrative
embodiments, the enzyme digestion can be performed at about 2 C to about 37 C.
However, the digestion can be conducted at any temperature, for any length of
time
(e.g., about 5 minutes to about 96 hours), and at any enzyme concentration.
In illustrative embodiments, the ratio of the extracellular matrix
material (hydrated) to total enzyme (weight/weight) ranges from about 25 to
about
2500. If an enzyme is used, it should be removed (e.g., by fractionation) or
inactivated after the desired incubation period for the digestion so as to not
compromise stability of the components in the solubilized extracellular matrix
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composition. Enzymes, such as pepsin for example, can be inactivated with
protease
inhibitors, a shift to neutral pH, a drop in temperature below 0 C, or heat
inactivation,
or a combination of these methods.
In another illustrative embodiment, the source extracellular matrix
material can be extracted in addition to being solubilized with hydrochloric
acid.
Extraction methods for extracellular matrices are known to those skilled in
the art and
are described in detail in U.S. Patent No. 6,375,989.
Illustrative extraction excipients include, for example, chaotropic agents
such as urea,
guanidine, sodium chloride, magnesium chloride, and non-ionic or ionic
surfactants.
In one embodiment, the solubilized collagen composition comprises
soluble and insoluble components, and at least a portion of the insoluble
components
of the solubilized collagen composition can be separated from the soluble
components. For example, the insoluble components can be separated from the
soluble components by centrifugation (e.g., at 12,000 rpm for 20 minutes at 4
C). In
alternative embodiments, other separation techniques known to those skilled in
the
art, such as filtration, can be used.
In accordance with one illustrative embodiment, the solubilized
extracellular matrix composition, prepared with or without the above-described
separation step, is fractionated prior to polymerization. In one illustrative
aspect, the
solubilized extracellular matrix composition can be fractionated by dialysis.
Exemplary molecular weight cut-offs for the dialysis tubing or membrane are
from
about 3,500 to about 12,000 or about 3,500 to about 5,000. In one embodiment,
the
solubilized extracellular matrix composition is dialyzed against an acidic
solution
having a low ionic strength. For example, the solubilized extracellular matrix
composition can be dialyzed against a hydrochloric acid solution, however any
other
acids can be used, including acetic acid, formic acid, citric acid, lactic
acid, sulfuric
acid, ethanoic acid, carbonic acid, nitric acid, or phosphoric acid. In
another example,
the extracellular matrix composition can be dialyzed against water as long as
the pH
is approximately 6 or below.
In various illustrative embodiments, the fractionation, for example by
dialysis, can be performed at about 2 C to about 37 C for about 1 hour to
about 96
hours. In another illustrative embodiment, the concentration of the acid, such
as
acetic acid, hydrochloric acid, formic acid, citric acid, lactic acid,
sulfuric acid,
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ethanoic acid, carbonic acid, nitric acid, or phosphoric acid, against which
the
solubilized extracellular matrix composition is dialyzed, can be from about
0.001 N to
about 0.1 N, from about 0.005 N to about 0.1 N, from about 0.01 N to about 0.1
N,
from about 0.05 N to about 0.1 N, from about 0.001 N to about 0.05 N, about
0.001 N
to about 0.01 N, or from about 0.01 N to about 0.05 N. In one illustrative
embodiment, the solubilized extracellular matrix composition is dialyzed
against 0.01
N HC1. However, the fractionation can be performed at any temperature, for any
length of time, and against any concentration of acid.
In accordance with one embodiment the 3D matrix used for culturing
stem cells comprises a lyophilized, solubilized collagen composition that is
rehydrated prior to contact with the cells. As discussed above, the term
"lyophilized"
means that water is removed from the composition, typically by freeze-drying
under a
vacuum (typically to dryness). In one illustrative aspect, a solubilized
extracellular
matrix composition is lyophilized after solubilization. In another
illustrative aspect,
the solubilized extracellular matrix composition is lyophilized after the
solubilized
portions have been separated from the insoluble portions. In yet another
illustrative
aspect, the solubilized extracellular matrix composition is lyophilized after
a
fractionation step but prior to polymerization. In another illustrative
embodiment, the
polymerized matrix is lyophilized. In one illustrative lyophilization
embodiment, the
solubilized extracellular matrix composition is first frozen, and then placed
under a
vacuum. In another lyophilization embodiment, the solubilized extracellular
matrix
composition is freeze-dried under a vacuum. Any method of lyophilization known
to
the skilled artisan can be used.
In accordance with one embodiment, the solubilized collagen
composition is sterilized before polymerization. In one embodiment the source
of the
solubilized collagen (e.g., a naturally occurring extracellular matrix, or a
lyophilized
purified collagen composition) is sterilized prior to the solubilization step.
Sterilization of the extracellular matrix material can be performed, for
example, as
described in U.S. Patents Nos. 4,902,508 and 6,206,931.
In another embodiment, the solubilized collagen composition is directly
sterilized, for example, with peracetic acid. In one embodiment wherein an
extracellular matrix is solubilized with an acid and the resulting material is
fractionated to isolate a fraction comprising solubilized collagen,
sterilization can be
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carried out either before or after the fractionation step. In another
illustrative
embodiment, the lyophilized composition itself is sterilized before
rehydration, for
example using an e-beam sterilization technique. In yet another illustrative
embodiment, the polymerized matrix formed from the components of the
solubilized
collagen matrix composition is sterilized.
In one illustrative embodiment, the solubilized extracellular matrix
composition is directly sterilized before the fractionation/separation step,
for example,
with peracetic acid or with peracetic acid and ethanol (e.g., by the addition
of 0.18%
peracetic acid and 4.8% ethanol to the solubilized extracellular matrix
composition
before the separation step). In another embodiment, sterilization can be
carried out
during the fractionation step. For example, the solubilized extracellular
matrix
composition can be dialyzed against chloroform, peracetic acid, or a solution
of
peracetic acid and ethanol to disinfect or sterilize the solubilized
extracellular matrix
composition. For example, the solubilized extracellular matrix composition can
be
sterilized by dialysis against a solution of peracetic acid and ethanol (e.g.,
0.18%
peracetic acid and 4.8% ethanol). The chloroform, peracetic acid, or peracetic
acid/ethanol can be removed prior to polymerization of the solubilized
collagen
composition, for example by dialysis against an acid, such as 0.01 N HC1.
If the solubilized collagen composition is lyophilized, the lyophilized
collagen matrix composition can be stored frozen or at room temperature (for
example, at about -80 C to about 25 C). Storage temperatures are selected to
stabilize
the components of the solubilized collagen matrix composition. The
compositions
can be stored for about 1-26 weeks, or longer, hi one illustrative embodiment,
the
storage solvent is hydrochloric acid. As described herein, "storage solvent"
means
the solvent that the solubilized collagen matrix composition is in prior to
and during
lyophilization. For example, hydrochloric acid, or other acids, at
concentrations of
from about 0.001 N to about 0.1 N, from about 0.005 N to about 0.1 N, from
about
0.01 N to about 0.1 N, from about 0.05 N to about 0.1 N, from about 0.001 N to
about
0.05 N, from about 0.001 N to about 0.01 N, or from about 0.01 N to about 0.05
N
can be used as the storage solvent for the lyophilized, solubilized collagen
matrix
composition. Other acids can be used as the storage solvent including acetic
acid,
formic acid, citric acid, lactic acid, sulfuric acid, ethanoic acid, carbonic
acid, nitric
acid, or phosphoric acid, and these acids can be used at any of the above-
described
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concentrations. In one illustrative embodiment, the lyophilizate can be stored
(i.e.,
lyophilized in) an acid, such as acetic acid, at a concentration of from about
0.001 M
to about 0.5 M or from about 0.01 M to about 0.5 M. In another embodiment, the
lyophilizate can be stored in water with a pH of about 6 or below. In other
illustrative
embodiments, lyoproteetants, cryoprotectants, lyophilization accelerators, or
crystallizing excipients (e.g., ethanol, isopropanol, mannitol, trehalose,
maltose,
sucrose, tert-butanol, and Tween 20), or combinations thereof, and the like
can be
present during lyophilization.
In one embodiment, the sterilized, solubilized collagen composition
can be dialyzed against 0.01 N HC1, for example, prior to lyophilization to
remove the
sterilization solution and so that the solubilized extracellular matrix
composition is in
a 0.01 N HCI solution as a storage solvent. Alternatively, the solubilized
extracellular
matrix composition can be dialyzed against acetic acid as the storage solvent,
for
example, prior to lyophilization and can be lyophilized in acetic acid and
redissolved
in HCI or water.
If the solubilized extracellular matrix composition is lyophilized, the
resulting lyophilizate can be redissolved in any solution, but may be
redissolved in an
acidic solution or water. The lyophilizate can be redissolved in, for example,
acetic
acid, hydrochloric acid, formic acid, citric acid, lactic acid, sulfuric acid,
ethanoic
acid, carbonic acid, nitric acid, or phosphoric acid, at any of the above-
described
concentrations, or can be redissolved in water. In one illustrative embodiment
the
lyophilizate is redissolved in 0.01 N HC1. For use in producing engineered
matrices
that can be injected in vivo or used for other purposes in vitro, the
redissolved
lyophilizate can be subjected to varying conditions (e.g., pH, phosphate
concentration,
temperature, buffer composition, ionic strength, and composition and
concentration of
solubilized extracellular matrix composition components (dry weight/m1)) that
result
in polymerization to form engineered extracellular matrices for specific
tissue graft
applications.
Accordingly, in one illustrative embodiment of the method described
herein, a solubilized collagen composition is prepared by enzymatically
treating the
source extracellular matrix material with 0.1% (w/v) pepsin in 0.01 N HC1 to
initially
solubilized the extracellular matrix material, centrifuging the enzymatically
treated
composition at 12,000 rpm for 20 minutes at 4 C to remove insoluble
components,
* Trade-mark
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fractionating the soluble fraction by dialysis against a 0.01 N HC1 solution,
and then
polymerizing the dialyzed fraction.
In another illustrative embodiment, the method does not involve a
fractionation step. In this embodiment, the source extracellular matrix
material is
enzymatically treated with 0.1% (w/v) pepsin in a 0.01 N hydrochloric acid
solution
to produce a solubilized collagen composition, the solubilized composition is
then
centrifuged to remove insoluble components, and then the solubilized fraction
is
polymerized.
In another illustrative embodiment, a solubilized collagen composition
is prepared by grinding source vertebrate submucosa into a powder and
enzymatically
digesting the powderized submucosa with 0.1% w/v pepsin and solubilizing in
0.01 N
HC1 for one to three days at 4 C. Following digestion and solubilization, the
solubilized components of the solubilized submucosa composition are separated
from
the insoluble components by centrifugation at 12,000 rpm at 4 C for 20
minutes. The
supernatant, comprising the soluble components, is recovered and the pellet
containing insoluble components is discarded. The supernatant is then
fractionated by
dialyzing the solubilized submucosa composition against 0.01 N HC1. In one
embodiment, the solubilized submucosa composition is dialyzed against several
changes of 0.01 N hydrochloric acid at 4 C using dialysis membranes having a
molecular weight cut-off of 3500. Thus, the solubilized submucosa composition
is
fractionated to remove components having a molecular weight of less than about
3500. Alternatively, dialysis tubing or membranes having a different molecular
weight cut-off can be used. The fractionated solubilized submucosa composition
is
then polymerized to produce the collagen based matrices of the present
invention.
In accordance with another illustrative embodiment, a solubilized
collagen composition is prepared by grinding vertebrate submucosa into a
powder and
digesting the powderized submucosa composition with 0.1% w/v pepsin and
solubilizing in 0.01 N hydrochloric acid for one to three days at 4 C. The
solubilized
components are then separated from the insoluble components, for example, by
centrifugation at 12,000 rpm at 4 C for 20 minutes. The supernatant,
comprising the
soluble components, is recovered and the pellet containing insoluble
components is
discarded. The non-fractionated solubilized submucosa composition is then
polymerized.
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The present invention encompasses the formation of a solubilized
collagen composition from a complex extracellular matrix material without
purification of the matrix components. However, the components of the
naturally
occurring extracellular matrices can be partially purified and the partially
purified
composition can be used in accordance with the methods described herein to
prepare a
solubilized collagen composition. Purification methods for extracellular
matrix
components are known to those skilled in the art and are described in detail
in U.S.
Patent No. 6,375,989, In accordance with one
embodiment the solubilized collagen composition includes purified type I
collagen or
type I and type HI collagen as the only protein constituents of the
composition.
The solubilized collagen composition can be polymerized under
different conditions to produce a collagen based matrix having the desired
microstrutures and mechanical properties. Polymerization of purified type I
collagen
solutions at different concentrations of collagen affected fibril density
while
maintaining a relatively constant fibril diameter. In addition, both fibril
length and
diameter are affected by altering the pH of the polymerization reaction.
Additional conditions can be varied during the polymerization reaction
to provide engineered purified collagen matrices that have the desired
properties. In
illustrative embodiments, the conditions that can be varied include pH,
phosphate
concentration, temperature, buffer composition, ionic strength, the
extracellular
matrix components in the solubilized extracellular matrix composition, and the
concentration of solubilized extracellular matrix composition components (dry
weight/m1). These conditions result in polymerization of the extracellular
matrix
components to form engineered extracellular matrices with desired
compositional,
microstructural, and mechanical characteristics. Illustratively, these
compositional,
microstructural, and mechanical characteristics can include fibril length,
fibril
diameter, number of fibril-fibril connections, fibril density, fibril
organization, matrix
composition, 3-dimensional shape or form, viscoelastic, tensile, or
compressive
behavior, shear (e.g., failure stress, failure strain, and modulus),
permeability,
swelling, hydration properties (e.g., rate and swelling), and in vivo tissue
remodeling
and bulking properties, and desired in vitro cell responses. The matrices
described
herein have desirable biocompatibility, vascularization, remodeling, and
bulking
properties, among other desirable properties.
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In various illustrative embodiments, qualitative and quantitative
microstructual characteristics of the engineered matrices can be determined by
environmental or cryostage scanning electron microscopy, transmission electron
microscopy, confocal microscopy, second harmonic generation multi-photon
microscopy. In another embodiment, polymerization kinetics may be determined
by
spectrophotometry or time-lapse confocal reflection microscopy. In another
embodiment, tensile, compressive and viscoelastic properties can be determined
by
rheometry or uniaxial tensile testing. In another embodiment, a rat
subcutaneous
injection model can be used to determine remodeling and bulking properties.
All of
these methods are known in the art or are further described in Examples 5-7 or
are
described in Roeder et al., f Biomech. Eng. vol. 124, pp. 214-222 (2002) and
in Pizzo
et al., J. AppL PhysioL, vol. 98, pp. 1-13 (2004).
In accordance with one embodiment, the solubilized collagen
composition is polymerized at a final total collagen concentration of about 1
to about
40 mg/ml, and in one embodiment about Ito about 30 mg/ml, in another
embodiment
about 2 to about 25 mg/ml and in another embodiment about 5 to about 15 mg/ml.
In
one embodiment the final total collagen is selected from a range of about 0.25
to
about 5.0 mg/ml, or in another embodiment the fmal total collagen
concentration is
selected from the range of about 0.5 to about 4.0 mg/ml, and in another
embodiment
the final total collagen concentration is selected from the range of about 1.0
to about
3.0 mg/ml, and in another embodiment the final total collagen concentration is
about
0.3, 0.5, 1.0, 2.0 or 3.0 mg/ml. In other embodiments, the components of the
solubilized extracellular matrix composition are polymerized at final
concentrations
(dry weight/nil) of about 0.25 to about 10 mg/ml, about 0.25 to about 20
mg/ml, about
0.25 to about 30 mg/ml, about 0.25 to about 40 mg/ml, about 0.25 to about 50
mg/ml,
about 0.25 to about 60 mg/ml, or about 0.25 to about 80 mg/ml.
In various illustrative embodiments, the total collagen comprising the
solubilized collagen composition comprises type I and type III collagen,
wherein the
percent range of the type III collagen and type I collagen is selected from
about 17-
33% and about 66-83%, respectively, to achieve various collagen type I/III
ratios.
Examples of percentage ranges of type III collagen and type I collagen,
respectively
that may be used in the matrices include 17% and 83%; 20% and 80%; 25% and
75%;
30% and 70%; and 33% and 66%, respectively. In various illustrative
embodiments,
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the type I collagen to type III collagen ratio may be in the range of about
6:1 to about
1:1. Examples of the type I collagen to type III collagen ratios that may be
used in the
matrices include 6:1, 5:1, 4:1, 3:1, 2:1, 1.5:1, and 1:1.
In various illustrative embodiments, at least 3 ug/ml of type I collagen
is combined with at least 0.5 ug/ml of type III collagen to obtain a total
amount of
collagen. Examples of the amount of type I collagen combined with type III
collagen,
respectively, that may be used in the matrices include 3 ug/ml and 0.5 ug/ml;
1500
ug/ml and 250 ug/ml; 1500 ug/ml and 500 ug/ml; 1500 ug/ml and 750 ug/ml; and
1500 ug/ml and 1500 ug/ml.
In various illustrative embodiments, the conditions for combining type
I collagen and type III collagen can be the same as those described above for
the
method of decreasing stiffiless of an extracellular matrix composition.
Illustratively, the matrix compositions produced by the methods
described herein can be combined, prior to, during, or after polymerization,
with stem
cells or progenitor cells, to further enhance the repair or replacement of
diseased or
damaged tissues. Examples of progenitor cells include those that give rise to
blood
cells, fibroblasts, endothelial cells, epithelial cells, smooth muscle cells,
skeletal
muscle cells, cardiac muscle cells, multi-potential progenitor cells,
pericytes, and
osteogenic cells. The population of progenitor cells can be selected based on
the cell
type of the intended tissue to be repaired. For example, if skin is to be
repaired, the
population of progenitor cells will give rise to non-keratinized epithelial
cells or if
cardiac tissue is to be repaired, the progenitor cells can produce cardiac
muscle cells.
The matrix composition can also be seeded with autogenous cells isolated from
the
patient to be treated. In an alternative embodiment the cells may be
xenogeneic or
allogeneic in nature.
In any of the embodiments described above using purified collagen,
the purified collagen can be sterilized after purification. In yet other
embodiments,
the collagen that is purified can be sterilized before or during the
purification process.
In other embodiments, purified collagen can be sterilized before
polymerization or the
matrix can be sterilized after polymerization.
It has been reported that the use of progenitor or stem cells to treat
damaged tissues (including for example treating myocardial infarction followed
by
heart failure) has demonstrated early evidence of potential utility. However,
recent
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data, has revealed three key issues that significantly limit successful
delivery of
reparative cells to tissues. These are 1.) inefficient and inconsistent local
retention of
cells acutely following injection into tissues [Hou et al., 2005, Circulation,
112:1150-
6]; 2.) limited survival of cells over time following injection into tissues
[Rehman et
al., 2004, Circulation 109: 1292-8]; and 3.) lack of a suitable cellular
microenvironment to modulate differentiation into the desired tissue types
(e.g., either
vascular structures or myocytes in the context of tissue remodeling in
response to
ischemic insult) [Reinlib and Field, 2000, Circulation101: E182-E187].
In accordance with one embodiment a novel cell delivery strategy is
provided that involves the suspension of cells in a liquid-phase, injectable
solubilized
collagen composition that polymerizes in situ to form a three-dimensional (3D)
matrix. The 3D matrix is designed to both entrap cells and provide them with
an
"instructive" microenvironment which promotes cell survival and modulates
their
fate. It is anticipated that the introduction of cells in the presence of a
comparatively
viscous medium (i.e., the solubilized collagen composition, which will
subsequently
assemble in situ shortly after post-injection) will enhance the cells local
retention.
Furthermore, as noted in Examples 12-15, the components of the 3D matrix and
their
microstructural organization play an important role in determining cell fate
with
respect to survival, proliferation, and differentiation. Interestingly, recent
data shows
that a nanofiber microenvironment formed intramyocardially following injection
of a
peptide (8-16 amino acids long) hydrogel (of which the biological signaling
capacity
and degradation properties have yet to be elucidated) resulted in formation of
a
nanofiber microenvironment that promoted endogenous cell recruitment [Davis et
al.,
2005 Circulation 111:442-50]. Furthermore, co-culture of endothelial cells
with
cardiomyocytes within the peptide hydrogel in vitro dramatically decreased
apoptosis
and necrosis of cardiomyocytes [Narmaneva et al., 2004 Circulation 110:962-
968].
As reported herein, the biophysical signals provided by a 3D self-
assembled collagen microenvironment can be used to direct the proliferation
and
differentiation capacity of multi-potential, bone marrow-derived stem cells.
For
example, 3D purified collagen matrices characterized by a relatively high
fibril
density and stiffness supported an increase in clonal growth and enhanced
osteogenesis (bone formation). Collectively, these results demonstrate the
ability to
engineer injectable, self-assembling 3D purified collagen matrices in which
the
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composition, microstructure, and mechanical properties are defined and
systematically varied with discrete outcomes. In general, the biophysical
features of
the 3D matrix, in addition to cellular signaling modalities consisting of
soluble factors
and cell-cell interactions, are determinants of cell fate and represent a new
target for
In accordance with one embodiment a method of enhancing the repair
of damaged, diseased or congenital defective tissues is provided. The method
comprises the steps of suspending a population of cells within a solubilized
collagen
composition, inducing the polymerization of the solubilized collagen
composition,
In one embodiment the solubilized collagen composition comprises
about 0.1 mg/ml to about 3 mg/ml total purified collagen (either type I alone
or a
combination of type I and type III collagen) in about 0.05 to about 0.005N HC1
(and
in one embodiment about 0.01N HC1), about 0.07M to about 0.28M NaC1 (and in
one
(and in one embodiment about 8.1mM Na2HPO4), about 0.7 to about 3.0mM KH2PO4
(and in one embodiment about 1.5mM KH2PO4), about 0.25 to about 1.0mM MgC12
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(and in one embodiment about 0.5mM MgC12), about 2.8mM to about 166mM
glucose, (and in one embodiment about 5mM glucose). Polymerization of the
solubilized collagen composition is induced by the addition of a neutralizing
solution
such as NaOH. For example a NaOH solution can be added to a final
concentration
of 0.01N NaOH. The cells are then added to the composition after the addition
of
neutralizing solution. In accordance with one embodiment a calcium chloride
solution is also added to the solubilized collagen composition. In this
embodiment,
calcium chloride is added to bring the final concentration of CaCl2 in the
solubilized
collagen composition to about 0.4 mM to about 2.0 mM CaC12 (and in one
embodiment about 0.9 mM CaCl2). The composition is then allowed to polymerize
either in vitro or in vivo to form a 3D matrix comprised of collagen fibrils
wherein the
cells are embedded within the 3D matrix. In illustrative embodiments the
polymerization reaction is conducted in a buffered solution using any
biologically
compatible buffer system known to those skilled in the art. For example the
buffer
may be selected from the group consisting of phosphate buffer saline (PBS),
Tris
(hydroxymethyl) aminomethane Hydrochloride (Tris-HC1), 3-(N-Morpholino)
Propanesulfonic Acid (MOPS), piperazine-n,n'-bis (2-ethanesulfonic acid)
(PIPES),
[n-(2-Acetamido)]-2-Aminoethanesulfonic Acid (ACES), N-[2-hydroxyethyl]
piperazine-N'{2-ethanesulfonic acid] (HEPES) and 1,3-
bis[tris(Hydroxymethyl)methylamino]propane (Bis Tris Propane). In one
embodiment the buffer is PBS, Tris or MOPS and in one embodiment the buffer
system is PBS, and more particularly 10X PBS. In accordance with one
embodiment
the 10X PBS buffer at pH 7.4 comprises the following ingredients:
1.37M NaC1
0.027M KC1
0.081M Na2HPO4
0.015M KH2PO4
5mM MgC12
55.5mM glucose
To create 10X PBS buffers of different pH, the ratio of Na2HPO4 and
KH2PO4 is varied. Ionic strength may be adjusted as an independent variable by
varying the molarity of NaCl only.
The polymerization of the solubilized collagen composition is
conducted at a pH selected from the range of about 6.0 to about 9.0, and in
one
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embodiment polymerization is conducted at a pH selected from the range of
about 5.0
to about 11.0 and in one embodiment about 6.0 to about 9.0, and in one
embodiment
polymerization is conducted at a pH selected from the range of about 6.5 to
about 8.5,
in another embodiment polymerization of the solubilized collagen composition
is
conducted at a pH selected from the range of about 7.0 to about 8.0, and in
another
embodiment polymerization of the solubilized collagen composition is conducted
at a
pH selected from the range of about 7.3 to about 7.4.
The ionic strength of the buffered solution is also regulated. In
accordance with one embodiment the ionic strength of the solubilized collagen
composition is selected from a range of about 0.05 to about 1.5 M, in another
embodiment the ionic strength is selected from a range of about 0.10 to about
0.90 M,
in another embodiment the ionic strength is selected from a range of about
0.14 to
about 0.30 M and in another embodiment the ionic strength is selected from a
range
of about 0.14 to about 0.17 M.
In still other illustrative embodiments, the polymerization is conducted
at temperatures selected from the range of about 0 C to about 60 C. In other
embodiments, polymerization is conducted at temperatures above 20 C, and
typically
the polymerization is conducted at a temperature selected from the range of
about
C to about 40 C, and more typically the temperature is selected from the range
of
20 about 30 C to about 40 C. In one embodiment the polymerization is
conducted at
about 37 C.
In yet other embodiments, the phosphate concentration is varied. For
example, in one embodiment, the phosphate concentration is selected from a
range of
about .005 M to about 0.5 M. In another illustrative embodiment, the phosphate
concentration is selected from a range of about 0.01 M to about 0.2 M. In
another
embodiment, the phosphate concentration is selected from a range of about 0.01
M to
about 0.1 M. In another illustrative embodiment, the phosphate concentration
is
selected from a range of about 0.01 M to about 0.03 M. In other illustrative
embodiments, the solubilized collagen composition can be polymerized by, for
example, dialysis against a solution under any of the above-described
conditions (e.g.,
dialysis against PBS at pH 7.4), extrusion or co-extrusion of submucosa
formulations
into a desired buffer, including the buffers described above, or wet-spinning
to form
strands of extracellular matrix material. In one embodiment the strands can be
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formed by extrusion of a solubilized collagen composition through a needle and
can
be air-dried to form threads.
In one embodiment the strands can be formed by extrusion through a
needle and can be air-dried to form fibers or threads of various dimensions.
The
syringe can be adapted with needles or tubing to control the dimensions (e.g.,
diameter) of the fibers or threads. In one embodiment, the extrusion process
involves
polymerization of the solubilized extracellular matrix composition followed by
extrusion into a bath containing water, a buffer, or an organic solvent (e.g.,
ethanol).
In another embodiment, the extrusion process involves coextrusion of the
solubilized
extracellular matrix composition with a polymerization buffer (e.g., the
buffer such as
Tris or phosphate buffers at various concentrations can be varied to control
pH and
ionic strength). In yet another embodiment, the extrusion process involves
extrusion
of the solubilized extracellular matrix composition into a polymerization bath
(e.g.,
the buffer such as Tris or phosphate buffers at various concentrations can be
varied to
control pH and ionic strength). The bath conditions affect polymerization time
and
properties of the fibers or threads, such as mechanical integrity of the
fibers or
threads, fiber dimensions, and the like. In one embodiment the extrusion of a
solubilized collagen composition through a needle is used a method to control
orientation of polymerized fibrils within the fibers. In one embodiment, the
fibers can
be air-dried to create materials that can be crosslinked or woven into three
dimensional meshes or mats that can serve as a substrate, or a component of a
substrate, for culturing cells. In various illustrative embodiments,
engineered
extracellular matrices can be polymerized from the solubilized extracellular
matrix
composition at any step in the above-described methods. For example, the
engineered
matrices can be polymerized from the solubilized extracellular matrix
composition
after the solubilization step or after the separation step, the filtration
step, or the
lyophilization and rehydration steps, if the separation step, the filtration
step, and/or
the lyophilization and rehydration steps are performed.
The engineered matrices can be combined, prior to, during, or after
polymerization, with nutrients, including minerals, amino acids,
pharmaceutical
agents, sugars, peptides, proteins, vitamins (such as ascorbic acid), or
glycoproteins
that facilitate cellular proliferation, such as laminin and fibronectin, or
growth factors
such as epidermal growth factor, platelet-derived growth factor, transforming
growth
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factor beta, or fibroblast growth factor, and glucocorticoids such as
dexamethasone.
In other illustrative embodiments, fibrillogenesis modulators, such as
alcohols,
glycerol, glucose, or polyhydroxylated compounds can be added prior to or
during
polymerization. In accordance with one embodiment, cells can be added to the
solubilized extracellular matrix composition as the last step prior to the
polymerization or after polymerization of the matrix. In another illustrative
embodiment, particulate extracellular matrix compositions can be added to the
solubilized extracellular matrix composition and can enhance in vivo bulking
capacity. In other illustrative embodiments, cross-linking agents, such as
carbodiimides, aldehydes, lysl-oxidase, N-hydroxysuccinimide esters, imido
esters,
hydrazides, and maleimides, and the like can be added before, during, or after
polymerization.
Hyaluronic acid (HA) is a glycosaminoglycan found naturally within
the extracellular matrix. This mucopolysaccharide is made up of a repetitive
sequence of two modified simple sugars, glucuronic acid and N-acetyl
glucosamine.
HA molecules are negatively charged and typically high in molecular weight
(long in
size). The size and charged nature of this molecule allow it to bind water to
produce a
high viscosity gel. When hyaluronic acid is added to soluble collagen
compositions
and the solubilized collagen compositions are allowed to polymerize, it
appears that
only subtle changes occur to the fibrillar microstructure of the resultant 3D
matrix.
On the other hand, increasing the hyaluronic acid content significantly
affects the
viscous fluid phase of the extracellular matrix, providing it with distinct
mechanical
behavior. Furthermore, the addition of hyaluronic acid to engineered matrices
was
found to modulate the manner by which cells remodel and contract the matrix.
Accordingly, HA content represents a further variable of the present
engineered 3D
matrices that can be manipulated to provide an optimal microenvironment for
cells
cultured within the matrices.
In accordance with one embodiment the engineered purified collagen
based matrices of the present invention can be used as cell culture substrates
that
more accurately mimic the substrates that various cells contact in vivo.
Accordingly,
collagenous based matrices can be designed for specific cell types to mimic
their
native environment. In this manner stem cells or progenitor cells can be
cultured in
vitro without altering the fundamental cell behavior (e.g., cell
proliferation, growth,
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maturation, differentiation, migration, adhesion, gene expression, apoptosis
and other
cell behaviors) of the cells. In another embodiment, the engineered purified
collagen
based matrices of the present invention can be used to expand or differentiate
a cell
population, such a stem cell population (including pluripotent or unipotent
cells),
primary cells, progenitor cells or other eukaryotic cells by seeding the cells
on, or
within, the collagen based matrix and culturing the cells in vitro for a
predetermined
length of time under conditions conducive for that cell type's proliferation
(i.e.,
appropriate nutrients, temperature, pH, etc.). In accordance with one
embodiment
cells are added to the solubilized collagen composition as the last step prior
to the
polymerization of the solubilized collagen composition. The engineered
purified
collagen based matrices of the present invention can be combined with
nutrients,
including minerals, pharmaceutical agents, amino acids, sugars, peptides,
proteins,
vitamins (such as ascorbic acid), or glycoproteins that facilitate cellular
proliferation,
such as laminin and fibronectin and growth factors such as epidermal growth
factor,
platelet-derived growth factor, transforming growth factor beta, or fibroblast
growth
factor, and glucocorticoids such as dexamethasone.
In one example of an embodiment comprising a collagen based matrix
seeded with living cells, a sterilized engineered purified collagen based
matrix may be
seeded with living cells and packaged in an appropriate medium for the cell
type used.
For example, a cell culture medium comprising Dulbecco's Modified Eagles
Medium
(DMEM) can be used with standard additives such as non-essential amino acids,
glucose, ascorbic acid, sodium pyruvate, fungicides, antibiotics, etc., in
concentrations deemed appropriate for cell type, shipping conditions, etc.
The cell seeded engineered purified collagen based matrices of the
present invention can be used simply for culturing cells in vitro, or the
composition
can be implanted or injected as a tissue graft construct to enhance the repair
of
damaged or diseased tissue. In one embodiment an improved tissue graft
construct is
provided wherein the construct comprises a 3D purified collagen based matrix
and a
population of cells. The 3D purified collagen based matrix is formed from a
solubilized collagen composition wherein the solubilized composition is formed
by
contacting a source of purified collagen with an acid selected from the group
consisting of hydrochloric acid, acetic acid, formic acid, sulfuric acid,
ethanoic acid,
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carbonic acid, nitric acid, or phosphoric acid. The solubilized collagen
composition is
then polymerized as described above to form the 3D purified collagen based
matrix.
Cells, and in one embodiment stem cells, are combined with the
collagen based matrix at a low density and can be either added to the
solubilized
collagen composition prior to polymerization, or after formation of the
collagen based
matrix. This initial seeded population of cells can be expanded by incubating
the
composition under conditions suitable for replication of the seeded cells.
Accordingly, cell seeded 3D purified collagen based matrices of the present
invention
comprise a population of cells that consists of, or are the progeny of,
eukaryotic stem
cells initially added to the composition at a low density. In one embodiment a
tissue
graft construct is prepared comprising the 3D purified collagen based matrices
of the
present invention that have been seeded with a low density of cells, wherein
the cells
are cultured within the matrix to expand and/or differentiate the seeded
population of
cells prior to implantation of the graft construct in a host. In one
embodiment the
cells, and more particularly stein cells, are initially seeded within the 3D
purified
collagen matrix at a final concentration of about 10 to about 108 cells per
milliliter,
and in one embodiment at a final concentration of less than 105 cells per
milliliter.
For most cells, cell survival during in vitro culture is known to
decrease as the concentration/density at which the cells are initially seeded
onto a
substrate. Applicants have discovered that using an engineered purified
collagen
based matrix and seeding stem cells at very low densities, clonal populations
of stem
cells can be isolated in a substantially pure form. Typically the isolation of
non-
embryoic stem cells results in the isolation of cells that may differentiate
along
different cell lineage pathways. In accordance with one embodiment of the
present
invention culturing conditions can be selected wherein a decreased seeding
density of
viable pluripotent or multipotent stein cells within an engineered purified
collagen
based matrix leads to clonal growth of cells representing a single cell
lineage. Such
cells can be isolated and transferred to a second engineered purified collagen
based
matrix and conditions can be altered to enhance the proliferation of the
isolated clonal
population of cells. Estimates of optimal cell densities for clonal growth
range from
about 10 cells/ml to about 103 cells/ml and depend upon the specific seeding
efficiencies.
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In accordance with one embodiment a method of isolating clonal
populations of individual stem cells is provided. The method comprises the
steps of
contacting a source of collagen with hydrochloric acid to prepare a
solubilized
collagen composition. The solubilized collagen composition is then polymerized
to
form an engineered purified collagen based matrix. The stem cells are seeded
on or
within the engineered purified collagen based matrix at a low density that
maintains
the functionality of the stem cells but allows for the isolation of clonal
populations of
cells. In accordance with one embodiment the solubilized collagen composition
is
prepared having a type I collagen concentration selected from the range of
about 1.0
to 3.0 mg/ml, and a pH of about 6.5 to about 7.0, wherein the solubilized
collagen
composition further comprises glucose and calcium chloride. In one embodiment
stem cells are seeded at a concentration selected from the range of from about
10 to
about 103 cells per milliliter. In one embodiment the source of collagen used
to
prepare the solubilized collagen composition comprises a purified preparation
of type
I collagen that has been dissolved in a hydrochloric acid solution. In an
alternative
embodiment the source of collagen comprises a hydrochloric acid solubilized
fraction
of a naturally occurring extracellular matrix, such as a submucosal matrix. In
one
embodiment the solubilized collagen composition is prepared from vertebrate
intestinal submucosa. The hydrochloric acid solution used to prepared the
solubilized
collagen composition can be from about 0.005 N to about 0.1 N, from about 0.01
N to
about 0.1 N, from about 0.05 N to about 0.1 N, from about 0.001 N to about
0.05 N,
about 0.001 N to about 0.01 N, or from about 0.01 N to about 0.05 N HC1.
In any of the embodiments described in this application, the solubilized
collagen composition (purified collagen or extracellular matrix components)
can be
polymerized at final concentrations of collagen (dry weight/m1) of about 5 to
about 10
mg/ml, about 5 to about 30 mg/ml, about 5 to about 50 mg/ml, about 5 to about
100
mg/ml, about 20 to about 50 mg/ml, about 20 to about 60 mg/ml, or about 20 to
about
100 mg/ml. Illustratively, the three-dimensional matrices may contain fibrils
with
specific characteristics, including, but not limited to, a fibril area
fraction (defined as
the percent area of the total area occupied by fibrils in a cross-sectional
surface of the
matrix; i.e., fibril density) of about 7% to about 26%, about 20% to about
30%, about
20% to about 50%, about 20% to about 70%, about 20% to about 100%, about 30%
to
about 50%, about 30% to about 70%, or about 30% to about 100%. In further
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illustrative embodiments, the three-dimensional matrices have an elastic or
linear
modulus (defined by the slope of the linear region of the stress-strain curve
obtained
using conventional mechanical testing protocols; i.e., stiffness) of about 0.5
kPa to
about 40 kPa, about 30 kPa to 100 kPa, about 30 kPa to about 1000 kPa, about
30 kPa
to about 10000 kPa, about 30 kPa to about 70000 kPa, about 100 kPa to 1000
kPa,
about 100 kPa to about 10000 kPa, or about 100 kPa to about 70000 kPa.
In accordance with one embodiment a kit is provided for preparing 3D
matrices that have been optimized for a particular cell that is to be seeded
within the
formed 3D matrix. The kit is provided with purified individual components that
can
be combined to form a solubilized collagen composition that upon
polymerization
foinis a 3D matrix comprised of collagen fibrils that presents an optimal
microenvironrnent for a population of cells. Typically the population of cells
represent cells provided separately from the kit, but in one embodiment the
cells may
also constitute a component of the kit. In one embodiment the cells are
mammalian
cells, including human cells, and in a further embodiment the cells are stem
or
progenitor cells. In accordance with one embodiment a kit is provided
comprising a
solubilized collagen composition and a polymerization composition. In a
further
embodiment the solubilized collagen composition comprises purified type I
collagen
as the sole collagen component. In another embodiment the solubilized collagen
composition comprises purified type I collagen and type III collagen as the
sole
collagen components.
In one embodiment the kit comprises separate vessels, each containing
one of the following components: purified type I collagen, a phosphate buffer
solution, a glucose solution, a calcium chloride solution and a basic
neutralizing
solution. In one embodiment the purified type I collagen of the kit is
provided in a
lyophilized form and the kit is further provided with a solution of HC1 (or
other dilute
acid including for example, acetic acid, formic acid, lactic acid, citric
acid, sulfuric
acid, ethanoic acid, carbonic acid, nitric acid, or phosphoric acid) for
resuspending the
lyophilized collagen. In one embodiment the kit is provided with a solution
comprising a solubilized collagen composition, and in a further embodiment the
solubilized collagen composition comprises a solubilized extracellular matrix
composition. In one embodiment the kit comprises a phosphate buffer solution,
a
glucose solution, a calcium chloride solution, and acid solution, a basic
neutralizing
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solution, a vessel comprising purified type I collagen, and a vessel
comprising
purified type III collagen. In one embodiment the polymerization composition
comprises a phosphate buffer that has a pH of about 7.2 to about 7.6 and the
acid
solution is an HC1 solution comprising about 0.05N to about 0.005N HC1, and in
one
embodiment the acid solution is about 0.01N HC1. In one embodiment the glucose
solution has a concentration selected from the range of about 0.2% to about 5%
w/v
glucose, or about 0.5% to about 3% w/v glucose, and in one embodiment the
glucose
solution is about 1% w/v glucose. In one embodiment the CaC12 solution has a
concentration selected from the range of about 2 mM to about 40.0 mM CaC12 or
about 0.2 mM to about 4.0 mM CaC12, or about 0.2 to about 2mM CaCl2. In one
embodiment the kit is provided with a 10X PBS buffer having a pH of about pH
7.4,
and comprising about 1.37M NaC1, about 0.027M KC1, about 0.081M Na2HPO4,
about 0.015M KH2PO4, about 5mM MgC12 and about 1% w/v glucose.
The kit can further be provided with instructional materials describing
methods for mixing the kit reagents to prepare 3D matrices. In particular, the
instructions materials provide information regarding the final concentrations
and
relative proportions of the matrix components that give optimal
microenvironmental
conditions including fibril microstructure and mechanical properties for a
particular
cell type or for a particular desired result (i.e., clonal expansion of cells,
differentiation, etc.).
The following examples illustrate specific embodiments in further
detail. These examples are provided for illustrative purposes only and should
not be
construed as limiting the invention or the inventive concept in any way.
EXAMPLE 1
Preparation of Lyophilized, Bioactive ECM Compositions from Fractionated
Submucosa Hydrolysates
Small intestinal submucosa is harvested and prepared from freshly
euthanized pigs as previously disclosed in -U.S. Patent. No. 4,956,178.
Intestinal
submucosa is powderized under liquid nitrogen and stored at -80 C prior to
use.
Digestion and solubilization of the material is perfoimed by adding 5 grams of
powdered tissue to each 100 ml of solution containing 0.1% w/v pepsin in 0.01
N
hydrochloric acid and incubating for 72 hours at 4 C. Following the incubation
period, the resulting solubilized composition is centrifuged at 12,000 rpm for
20
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minutes at 4 C and the insoluble pellet is discarded. The supernatant is
dialyzed
against at least ten changes of 0.01 N hydrochloric acid at 4 C (MWCO 3500)
over a
period of at least four days. The solubilized fractionated composition is then
sterilized by dialyzing against 0.18% peracetic acid/4.8% ethyl alcohol for
about two
hours. Dialysis of the composition is continued for at least two more hours,
with
additional changes of sterile 0.01 N hydrochloric acid per day, to eliminate
the
peracetic acid. The contents of the dialysis bags are then lyophilized to
dryness and
stored.
EXAMPLE 2
Preparation of Lyophilized, Bioactive ECM Compositions from Non-Fractionated
Submucosa Hydrolysates
Small intestinal submucosa was harvested and prepared from freshly
euthanized pigs as previously disclosed in U.S. Patent No. 4,956,178.
Intestinal
submucosa was powderized under liquid nitrogen and stored at -80 C prior to
use.
Partial digestion of the material was performed by adding 5 g powdered tissue
to each
100 ml solution containing 0.1% wtv pepsin in 0.01M hydrochloric acid and
digesting
for 72 hours at 4 C. Following partial digestion, the suspension was
centrifuged at
12,000 rpm for 20 minutes at 4 C and the insoluble pellet discarded. The
supernatant
was lyophilized to dryness.
EXAMPLE 3
Preparation of Reconstituted, Bioactive ECM Compositions
Immediately prior to use, lyophilized material from Example 2,
consisting of a mixture of extracellular matrix components, was reconstituted
in 0.01
N HC1. To polymerize the soluble extracellular matrix components into a 3-
dimensional matrix, reconstituted extracellular matrix solutions were diluted
and
brought to a particular pH, ionic strength, and phosphate concentration by the
addition
of a phosphate buffer and concentrated HC1 and NaOH solutions. Polymerization
of
neutralized solutions was then induced by raising the temperature from 4 C to
37 C.
Various polymerization buffers (including, e.g., phosphate buffers) were used
and the
pH of the polymerization reaction was controlled by varying the ratios of mono-
and
dibasic phosphate salts. Ionic strength was varied based on sodium chloride
concentration.
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Type I collagen prepared from calf skin was obtained from Sigma-
Aldrich Corporation, St. Louis, MO, and dissolved in and dialyzed extensively
against
0.01 M hydrochloric acid (HC1) to achieve desired concentrations. Interstitial
ECM
was prepared from porcine small intestinal submucosa (SIS). SIS was powdered
under liquid nitrogen and the powder stirred (5% w/v) into 0.01 N hydrochloric
acid
containing 0.1% (w/v) pepsin for 72 h at 4 C. The suspension was centrifuged
at
12,000 xg for 20 mm at 4 C to remove undissolved tissue particulate and
lyophilized
to dryness. Immediately prior to experimental use, the lyophilized material
was
redissolved in 0.01 N HC1 to achieve desired collagen concentrations. To
polymerize
the soluble collagen or interstitial ECM components into a 3D matrix, each
solution
was diluted and brought to the specified pH, ionic strength, and phosphate
concentration by the addition of a polymerization composition and concentrated
HC1
and NaOH solutions. Polymerization of neutralized solutions was induced by
raising
the temperature from 4 C to 37 C. 'Various polymerization compositions were
used
to make final solutions with the properties shown in Table 2.
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Table 2
Series 1
Collagen formulations SIS formulations
pH I [Pi] [C] pH I [Pi] [C]
6.5 0.16 0.01 1 mg/ml 6.5 0.16 0.01 1
mg/ml
7.0 0.16 0.01 1 mg/ml 7.0 0.16 0.01 1
mg/ml
7.4 0.17 0.01 1 mg/ml 7.4 0.17 0.01 1
mg/ml
8.0 0.17 0.01 1 mg/ml 8.0 0.17 0.01 1
mg/ml
8.5 0.17 0.01 1 mg/ml 8.5 0.17 0.01 1
mg/ml
9.0 0.17 0.01 1 mg/ml 9.0 0.17 0.01 1
mg/ml
Series 2
Collagen formulations SIS formulations
pH I [Pi] [C] pH I [Pi] [C]
7.4 0.06 0.02 1 mg/ml 7.4 0.06 0.02 1
mg/ml
7.4 0.10 0.02 1 mg/ml 7.4 0.30 0.02 1
mg/ml
7.4 0.15 0.02 1 mg/ml 7.4 0.60 0.02 1
mg/m1
7.4 0.20 0.02 1 mg/ml 7.4 0.90 0.02 1
mg/ml
7.4 0.25 0.02 1 mg/ml 7.4 1.20 0.02 1
mg/ml
7.4 1.50 0.02 1 mg/ml
Series 3
Collagen formulations SIS formulations
pH I [Pi] [C] pH I [Pi] [C]
7.4 0.15 0.00 1 mg/ml 7.4 0.3 0.00 1
mg/ml
7.4 0.15 0.01 1 mghnl 7.4 0.3 0.02 1
mg/ml
7.4 0.15 0.02 1 mg/ml 7.4 0.3 0.04 1
mg/ml
7.4 0.15 0.03 1 mg/ml 7.4 0.3 0.06 1
mg/ml
7.4 0.15 0.04 1 mg/ml 7.4 0.3 0.08 1
mg/ml
7.4 0.15 0.05 1 mg/ml 7.4 0.3 0.11 1
mg/ml
Table 2: Engineered ECMs representing purified type I collagen or a complex
mixture
of interstitial ECM components (SIS) were prepared at varied pH (series 1),
ionic
strength (series 2), and phosphate concentration (series 3). [C] represents
collagen
concentration in mg/ml, [Pi] represents phosphate concentration in M, and I
represents
ionic strength in M.
Representative data showing the results of varying the polymerization
temperature, buffer system, pH (using either a phosphate or tris buffer),
ionic
strength, phosphate concentration or concentration of ECM material, on
stiffness
(elastic modulus) of the formed 3D matrix is presented in Figs. 1A-1G. In
summary,
as the polymerization temperature is increased from 4 C up to 37 C, the
polymerization rate and the stiffness of the formed 3D matrix increases. The
effect of
a temperature gradient profile on the microstructural composition of the 3D
matrix
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was also investigated. Polymerizing the matrix using a temperature ramp from
about
4 C to 37 C over 30 minutes was compared to matrices formed using a step
increase
in temperature to 37 C and incubated at that temperature for 30 minutes. The
data
revealed that fibrils formed using a temperature ramp are longer in length and
have
decreased fibril density compared to matrices formed using a single step
increase in
temperature. As the pH of the polymerizing composition is increased, from
about 7.0
up to about pH 9.2, the polymerization rate and the stiffness of the formed 3D
matrix
increases. Buffer selection was found to play a role in determining the
mechanical
properties of the 3D matrix, and more particularly tris based buffers reduced
stiffness
more than phosphate based buffers. Regarding ionic strength, peak stiffness
coincides
with maximum polymerization time at an ionic strength of about 0.3 M. As
phosphate concentration is increased, stiffness decreases, however the
concentration
of phosphate in a 1X PBS solution does not have a substantial effect on
stiffness. As
collagen content is increased the stiffness of the matrix is increased.
EXAMPLE 4
Three-Dimensional Imaging of Engineered ECM's by Confocal Reflection
Microscopy
Solutions of type I collagen or interstitial ECM components were
polymerized in a Lab-Tek chambered coverglass and imaged using a BioRad
Radiance*2100 MP Rainbow confocal/multiphoton microscope using a 60 x 1.4 NA
oil immersion lens. Optical settings were established and optimized for
matrices after
polymerization was complete. Samples were illuminated with 488 um laser light
and
the reflected light detected with a photomultiplier tube (PMT) using a blue
reflection
filter. A z step of 0.2 gm was used to optically section the samples. Because
the
resolution of the z axis is less than that of the x-y plane, the sampling
along the z axis
may be different from that of the x-y. Images were collected in the range of
10-25 gm
from the upper surface of the coverglass.
EXAMPLE 5
Quantification of Fibril Properties from Three Dimensional Images
Quantification of the fibril diameter distribution within engineered
extracellular matrices was conducted based upon two- and three- dimensional
image
sets obtained via electron and confocal microscopy techniques using methods
described within Brightman et al., Biopolymers 54:222-234, 2000. More
recently, a
* Trade-mark
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Matlab program with a graphical user interface was written for measurement of
fibril
diameters from these images. For three-dimensional confocal images, depth
attenuation was corrected by normalizing against a fitted logarithmic curve,
after
which images were binarized into white and black pixels using a threshold
value.
Three rectangles were outlined in the x-y plane across each fibril, with one
axis
aligned with the fibril. Average fibril diameter in each rectangle was
calculated as the
total white area divided by the rectangle's length. The average diameter of
each fibril
was taken to be the average of the three measurements, and the average
diameter in a
given matrix was calculated as an average of all measurements.
Length of fibril per volume was estimated by dividing the total white
volume of an image by the average cross-sectional area of fibrils in that
image. Due
to distortion in the z-plane, the fibril cross-sections in the image could not
be assumed
circular and calculated from diameter. Rather, the average cross-sectional
area was
found by expanding the rectangles described above into three-dimensional
boxes.
The cross-sectional area of a fibril in was found by dividing the total white
volume
contained in the box by the length of the box's axis aligned with the fibril.
A Matlab program has also been developed to determine fibril density
from two- and three- dimensional images. This method involves thresholding and
binarizing the image data to discriminate fibrils from the background. The
surface
area or volume representing fibrils is then quantified and normalized to the
surface
area or volume of the image.
EXAMPLE 6
Spectrophotometry of Extracellular Matrix Polymerization
The time-course of polymerization was monitored in a Lambda 35
UV-VIS spectrophotometer (Perkin-Elmer) equipped with a temperature-
controlled,
8-position cell changer as described previously by Brightman et al., 2000.
EXAMPLE 7
Rheometric Measurements of Extracellular Matrices
Mechanical properties of the matrices were measured using a TA
Instruments AR-2000 rheometer. Neutralized collagen or SIS was placed on the
peltier temperature-controlled lower plate at 6 C, and the 40-mm parallel-
plate
geometry was lowered to a 1-mm gap. The temperature was then raised to 37 C as
oscillation measurements were made every 30 seconds at 1 Hz and 5% strain.
After
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polymerization was complete, an oscillation frequency sweep was made at 5%
strain,
from 0.1 to 3 Hz. A shear creep test was then conducted with a shear stress of
1 Pa
for 1000 seconds.
EXAMPLE 8
Preparation of Reconstituted Bioactive Extracellular Matrices
Small intestinal submucosa was harvested and prepared from freshly
euthanized pigs as previously disclosed in U.S. Patent No. 4,956,178.
Intestinal
submucosa was powderized under liquid nitrogen and stored at -80 C prior to
use.
Digestion and solubilization of the material was performed by adding 5 grams
of
powdered tissue to each 100 ml of solution containing 0.1% pepsin in 0.01 N
hydrochloric acid and incubating with stirring for 72 hours at 4 C. Following
the
incubation period, the solubilized composition was centrifuged at 12,000 rpm
for 20
minutes at 4 C and the insoluble pellet was discarded. The supernatant was
dialyzed
extensively against 0.01 N HC1 at 4 C in dialysis tubing with a 3500 MWCO
(Spectrum Medical Industries). Polymerization of the solubilized extracellular
matrix
composition was achieved by dialysis against PBS, pH 7.4, at 4 C for about 48
hours.
The polymerized construct was then dialyzed against several changes of water
at
room temperature and was then lyophilized to dryness.
The polymerized construct had significant mechanical integrity and,
upon rehydration, had tissue-like consistency and properties. In one assay,
glycerol
was added prior to polymerization by dialysis and matrices with increased
mechanical
integrity and increased fibril length resulted.
EXAMPLE 9
Preparation of Extracellular Matrix Threads
Small intestinal submucosa was harvested and prepared from freshly
euthanized pigs as previously disclosed in U.S. Patent No. 4,956,178.
Intestinal
submucosa was powderized under liquid nitrogen and stored at -80 C prior to
use.
Digestion and solubilization of the material was performed by adding 5 grams
of
powdered tissue to each 100 ml of solution containing 0.1% w/v pepsin in 0.01
N
hydrochloric acid and incubating for 72 hours at 4 C. Following the incubation
period, the solubilized composition was centrifuged at 12,000 rpm for 20
minutes at
4 C and the insoluble pellet was discarded.
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The solubilized extracellular matrix composition (at 4 C) was placed in
a syringe with a needle and was slowly injected into a PBS solution at 40 C.
The
solubilized extracellular matrix composition immediately formed a filament
with the
diameter of the needle. If a blunt-tipped needle is used, straight filaments
can be
formed while coiled filaments can be formed with a bevel-tipped needle. Such
filaments can be used as resorbable sutures.
EXAMPLE 10
Lyophilization and Reconstitution of Solubilized Extracellular Matrix
Compositions
Frozen small intestinal submucosa powder that had been prepared by
cryogenic milling was centrifuged at 3000 x g for 15 minutes and the excess
fluid was
decanted. The powder (5% weight/volume) was digested and solubilized in 0.01 N
HC1 containing 0.1% weight/volume pepsin for approximately 72 hours at 4 C.
The
solubilized extracellular matrix composition was then centrifuged at 16,000 x
g for 30
minutes at 4 C to remove the insoluble material. Aliquots of the solubilized
extracellular matrix composition were created and hydrochloric acid (12.1 N)
was
added to create a range of concentrations from 0.01 to 0.5 N HC1.
Portions of the solubilized extracellular matrix composition were
dialyzed (MWCO 3500) extensively against water and 0.01 M acetic acid to
determine the effects of these media on the lyophilization product. Aliquots
of the
solubilized extracellular matrix composition in 0.01 M acetic acid were
created and
glacial acetic acid (17.4 M) was added to create a range of concentrations
from 0.01
to 0.5 M acetic acid. The solubilized extracellular matrix compositions were
frozen
using a dry ice/ethanol bath and lyophilized to dryness. The lyophilized
preparations
were observed, weighed, and dissolved at 5 mg/ml in either 0.01 N HC1 or
water. The
dissolution and polymerization properties were then evaluated. The results are
shown
in Tables 2-6.
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Table 3. Gross appearance of solubilized extracellular matrix compositions
following
lyophilization at various hydrochloric acid concentrations.
[HC1] (N) Appearance
0.01 Light, fluffy, homogenous, foam-like sheet; white to off-white in
color; pliable
0.05 Slightly wrinkled and contracted, some inhomogeneities in
appearance noted, slight brown tint, pliable to slightly friable in
consistency
0.10 Wrinkled, collapsed in appearance; inhomogeneities noted, some
regional "melting" noted; significant brown tint; friable
0.25 Wrinkled,
collapsed in appearance; increased inhomogeneities
noted, increased areas of regional "melting" noted; significant
brown tint; friable
0.50 Significant
collapse and shrinkage of specimen, dark brown
coloration throughout; dark brown in color; friable
Table 4. Dissolution properties of solubilized extracellular matrix
compositions
following lyophilization at various hydrochloric acid concentrations.
[HC1] (N) Reconstitution
Properties
Reconstitution H20 0.01 N HC1
Medium
0.01 Completely dissolved in 20-30 Completely dissolved in 20-30 minutes,
minutes, pH 4 pH 2
0.05 Majority dissolved in 2 hours; Majority dissolved in 40 minutes;
very
slight particulate noted, pH 3-4 slight particulate noted, pH 2
0.1 Incomplete dissolution Incomplete dissolution
0.25 Incomplete dissolution Incomplete dissolution
0.50 Incomplete dissolution Incomplete dissolution
Table 5. Polymerization properties of solubilized extracellular matrix
compositions
following lyophilization at various hydrochloric acid concentrations.
[HC1] (N) Polymerization
Properties
Reconstitution
Medium 1120 0.01 N HCI
0.01 Polymerized within 20-30 Polymerized within 10-20
_ minutes minutes
0.05 Weak polymerization noted at Polymerized within 20-30
45 minutes; significant lag time minutes
in polymerization
0.1 *No Polymerization *No Polymerization
0.25 _ *No Polymerization *No Polymerization
0.50 *No Polymerization *No Polymerization
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Table 6. Dissolution properties of solubilized extracellular matrix
compositions
following lyophilization at various acetic acid concentrations.
[Acetic Acid]
(M) Reconstitution Properties
Reconstitution in 1120 Reconstitution in 0.01 N HC1
0.01 Completely dissolved in 90 minutes, Completely dissolved in
90
pH 5 minutes, pH 1-2
0.05 Near complete dissolution after 90 Completely dissolved
in 90
minutes; small particulate remained, minutes, pH 1-2
pH 5
0.1 Completely dissolved in 90 minutes, Near complete
dissolution in 90
pH 5 minutes; small particulate, pH 1-
2
0.25 Completely dissolved in 90 minutes, Completely dissolved in
90
pH 5 minutes, pH 1-2
0.50 Near complete dissolution after 90 Completely dissolved
in 90
minutes; small particulate remained, minutes, pH 1-2
pH 5
Table 7. Polymerization properties of solubilized extracellular matrix
compositions
following lyophilization at various acetic acid concentrations.
[Acetic Acid]
(M) Polymerization Properties
Reconstitution
Medium 1120 0.01 N HC1
0.01 Polymerized within 5-10 minutes Polymerized within 5-10
minutes
0.05 Polymerized within 5-10 minutes Polymerized within 5-10
minutes
0.1 Polymerized within 5-10 minutes Polymerized within 5-10
minutes
0.25 Polymerized within 5-10 minutes Polymerized within 5-10
minutes
0.50 Polymerized within 5-10 minutes Polymerized within 5-10
minutes
These results show that lyophilization in HC1 and reconstitution of
solubilized extracellular matrix compositions in 0.01 N HC1 to 0.05 N HC1 or
in water
maintains the capacity of the components of the compositions to polymerize.
The
results also show that lyophilization in acetic acid maintains the capacity of
the
components of the compositions to polymerize when the composition is
polymerized
in water or HC1. The solubility rate is lyophilization from 0.01 N HC1 >
lyophilization from 0.01 M acetic acid > lyophilization from water.
EXAMPLE 11
Preparation Of Solubilized Sis Composition
This procedure outlines a standard technique for the preparation of SIS
solution.
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1. Dissolution: of SIS powder in acetic acid with Pepsin
1.1. Preparation of acetic acid with pepsin
1.1.1. Prepare the desired volume of 0.5 M acetic acid (typically IL;
this requires 28.7 mL of 17.4 M glacial acetic acid).
1.1.2. Add the desired mass of pepsin to achieve a 0.1 % w/v solution
(typically 1 g, if I L of acetic acid is used).
1.1.3. Place the jar containing acetic acid and pepsin on a stir plate and
begin mixing.
1.2. Preparation of centrifuged SIS powder
1.2.1. Place SIS powder in 50 mL centrifuge tubes.
1.2.2. Centrifuge SIS powder at 3000 x g for 15 minutes.
1.2.3. Open centrifuge tubes, pour off and dispose of supernate.
1.2.4. Remove pellets from tubes. Measure out the desired mass to
achieve a 5% w/v solution (typically 50 g, if 1 L of acetic acid
was used). Previously prepared and frozen material may be
used, and excess centrifuged material may be frozen for later
use.
1.3. Add centrifuged SIS pellet material to acetic acid/pepsin solution. 1.4.
Cover and allow it to stir for 72 hours at 4 C.
2. Centrifugation of dissolved SIS
2.1. When removed from stirring, the SIS/pepsin solution should appear
viscous and somewhat uniform. Pour SIS/pepsin solution into
centrifuge jars. Balance jars as necessary.
2.2. This mixture should be centrifuged at 16,000 x g for 30 minutes at 4 C.
Refer to the operators manual or SOP for instructions on using the
centrifuge. If using the Beckman model J2-21, use the HO head at a
speed of 9500 rpm.
2.3. Remove jars of SIS from centrifuge. Pour the supemate into a clean jar.
Be careful not to disturb the pellet, and stop pouring if the SIS begins
to appear more white and creamy (this is pellet material).
3. Dialysis of SIS in water and hydrochloric acid
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3.1. Prepare dialysis tubing as follows:
3.1.1. Use dialysis tubing with MWCO 3500, diameter 29 lmll.
Handle dialysis tubing with gloves, and take care not to allow it
to contact foreign surfaces, as it may easily be damaged.
3.1.2. Cut dialysis tubing to the necessary length. (typically, 3 sections
of about 45 cm).
3.1.3. Wet tubing in millipore water, and leave tubing in the water until
each piece is needed.
3.1.4. Do the following with each length of tubing:
3.1.4.1. Place a clip near one end of the tubing.
3.1.4.2. Holding the tubing to avoid contact with foreign
surfaces, use a pipette to fill the tubing with SIS
solution. Each piece of tubing should receive roughly
the same volume of SIS (for example, if three lengths of
tubing are used, measure one third of the total volume
into each).
3.1.4.3. Place a clip on the open end of the dialysis tubing.
Avoid leaving slack. The tube should be full and taut.
3.1.4.4. Place the filled dialysis tubing in a container of 0.01 M
HCL with a stir bar.
3.1.4.5. Repeat the above steps to fill all lengths of tubing.
3.1. 5. Leave containers to stir at 4 C.
3.2. Details regarding changing the dialysis in 0.01 M HC1 are given below.
3.2.1. The 0.01 M HC1 in the dialysis containers must be changed
several times. This should be done as follows:
3.2.2. After changing the 0.01 M HCI, another change should not be
done for at least two hours.
3.2.3. Change the 0.01 M HC1 at least 10 times, over a period of at
least four days. This assumes a ratio of 200 mL SIS to 6 L of
0.01 M HC1. If a higher ratio is used, more changes may be
necessary.
3.2.4. When changing 0.01 M HC1, do not leave dialysis bags exposed
in the air or on the counter. Use tongs or forceps to move a
dialysis bag directly from one container to another. (It is okay
to have multiple diaiysis bags in one container.) Dump the first
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container in the sink, then refill it with millipore water. The
dialysis bags can now be placed in the newly filled container
while the other container or containers are changed.
4. Sterilization of SIS
4.1. Place dialysis bags of SIS in a solution of 0.18% Peracetic acid/4.8%
Ethanol. Leave to stir for two hours (more time may be necessary).
4.2. Translocate dialysis bags to 0.01 M HCI, and continue dialysis as before.
Continue for at least 2 days, changing HC1 at least 3 times daily.
4.3. When dialysis is complete, dialysis tubing filled with SIS should be
removed from the HC1.
4.4. Remove the clips. Cut open one end of the dialysis tubing and pour SIS
into a clean jar.
4.5. SIS should be refrigerated until use.
5. Lyophilization of SIS
5.1. Operating the Vertis Freezemobile
5.1.1. Make sure the condenser is free of any water. (The condenser is
the metal cylinder which opens on the front of the lyophilizer.)
Ensure that the black rubber collection tubing attached to the
bottom of the condenser is plugged. This can be accessed by
opening the grate on the front of the lyophilizer.
5.1.2. Close the door of the condenser, the top of the manifold, and all
sample ports. If the door of the condenser or the top of the
manifold are not forming a good seal apply a small amount of
vacuum grease to the rubber contact surfaces.
5.1,3. Turn on the "Refrigerate" switch. The indicator on the front of
the lyophilizer will show a light beside "Condenser" and
beneath "On." The light beneath "OK" will not illuminate until
the condenser is cooled. The condenser temperature is
indicated when the digital readout displays "Cl."
5 .1.4. When the "condenser" indicator light under "OK" is
illuminated, on the "Vacuum" switch. The indicator will show
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a light beside "Condenser" and beneath "On." The light
beneath "OK" will not turn on until the chamber is sufficiently
evacuated. The chamber pressure is indicated when the digital
readout displays "V 1."
5 .1.5. The rollers can be used for freezing a coat of material on the
inside surface of ajar. To use the rollers, first ensure that the
drain tube is plugged. (This can be accessed through the door
on the right side of the front of the lyophilizer.) Using 100%
Ethanol, fill the roller tank to a level several millimeters above
the top of the rollers. Under-filling will cause ineffective
cooling while over-filling will allow ethanol to leak into the
jars. The temperature of ethanol bath is indicated when the
digital readout displays "Ti." This bath is cooled when the
"Refrigerate" switch is turned on. The "Rollers" switch
controls the turning of the rollers, and may be switched off
when no jar is on the rollers.
5.2. Lyophilizing SIS
5.2.1. Lyophilization jars, glass lids, and rubber gaskets should be
cleaned with ethanol. Allow ethanol to evaporate completely
before use. Mid-size jars, lids, and gaskets (3-inch (7.62
cm) diameter) should be used to fit into the roller if
using the Virtis Freezemobile Jar lyophilization.
5.2.2. Pipette 75 mL of SIS solution into the lyophilization jar. Place
gasket and lid on jar.
5.2.3. Seal the jar by covering the openings with parafilm. Note the
small hole on the neck of the lid, which must be covered.
5.2.4. Place the jar of SIS on the lyophilizer rollers for a minimum of 2
hours.
5.2.5. Alternatively, the jar may be placed in a freezer until all material
is solid. In a -80 C freezer, this takes about 30 minutes.
5.2.6. Prepare a spigot on the lyophilizer by inserting a glass cock with
the tapered end out. The tapered end of the cock should be
coated with vacuum grease.
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5.2.7. Remove the jar of SIS from the rollers (or freezer). Place
springs on the hooks to hold the jar and lid together. Remove
the parafilm and place the neck of the lid of the jar over the
cock. Rotate the jar so that the holes in the lid and the cock do
not align. The spigot can be rotated so that the jar rests on the
top surface of the lyophilizer.
5.2.8. Turn the valve switch so that it points toward the jar of S1S.
5.2.9. More jars may be added to freeze-dry simultaneously, but add
jars one or two at a time. Wait until the vacuum pressure falls
to a reasonable range (e.g.,200 millitorr) to ensure that the last
jar is sealed before adding subsequent jars.
5.2.10. Leave the jars under vacuum for at least 24 hours.
5.2.11. After lyophilization is complete, turn the switch On the spigot
to point away from the jar. This will allow air into the jar.
5.2.12. Remove the jar from the cock.
5.2.13. Llyophilized material is not immediately used, it should be
stored in a dry environment. Use a large, sealable container with Dri-Rite or
another
desiccant, and place containers of lyophilized material therein.
6. Rehydration of lyophilized SIS
6.1. Place lyophilized SIS into a tube or jar.
6.2. Add the desired quantity of liquid (typically 0.01 N HC1) to the
container
of SIS.
6.3. Mixing may be accelerated by shaking, stirring, etc. Store container
under refrigeration until dissolution of SIS is complete.
STERILIZATION OF SOLUBILIZED SIS BY DIALYSIS AGAINST PERACETIC
ACM CONTAINING SOLUTION
1. Dialyze solubilized SIS against a large reservoir
containing
0.18 % peracetic acid/4.8% ethanol in water. Dialysis time may vary depending
upon
peracetic acid concentration, dialysis membrane molecular weight cut off,
temperature, etc.
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2. Transfer dialysis bags aseptically to reservoirs containing 0.01
N HC1. Dialyze extensively to reduce concentration of residual peracetic acid.
3. When dialysis is complete, dialysis tubing filled with
solubilized SIS should be removed from the dialysis tank aseptically.
4. Remove dialysis clips and pour or pipette solubilized SIS into a
sterile jar.
5. The disinfected solubilized SIS should be stored at 4 C
until
use.
STERILIZATION OF SIS BY DIRECT ADDITION OF PERACETIC ACID TO SIS
SOLUTION
1. Add 100% Ethanol and 32 wt % peracetic acid to
solubilized
SIS to create a solution with final concentration of 0.18% peracetic acid/4.8%
ethanol.
Stir well and leave for two hours.
2. Place solubilized SIS in aseptic dialysis bags. Dialyze against
sterile solution of 0.01 N HC1.
3. When dialysis is complete, dialysis tubing filled with
solubilized SIS should be removed from the dialysis tank aseptically.
4. Remove dialysis clips and pour or pipette solubilized SIS into a
sterile jar.
5. The disinfected solubilized SIS should be stored at 4 C until
use.
EXAMPLE 12
Engineered ECM Compositions Regulate Cell Behavior
The three-dimensional (3D) extracellular matrix (ECM) of tissues in
vivo represents a complex array of macromolecules that serves to provide
biochemical
and biophysical microenvironmental cues to resident cells. However, the exact
role
of any one biophysical feature or molecular component within the ECM in
regulating
cellular behavior has been difficult to elucidate due to the inherent
interdependence of
ECM compositional, structural, and mechanical properties. Recently, applicants
have
established that the 3D microstructural composition of fibrils within
engineered
ECMs created from purified type I collagen regulates cell-matrix adhesion,
matrix
remodeling, and proliferation properties of fibroblasts. It is further
anticipated that
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altering the ratios of collagen types I and III within engineered ECMs would
affect
the hierarchical assembly of fibrils, and therefore the ECM signaling
capacity.
Engineered ECMs were created with altered ratios of collagen types III
and I ranging from 1:6 to 1:2. Application of confocal and scanning electron
microscopy showed that ECMs prepared with increasing amounts of type III
collagen
possessed an increasing number of small diameter fibrils. Furthermore, these
microstructural changes translated into alteration of matrix mechanical
properties.
Finally, results showed a biphasic response for fibroblast proliferation,
morphology,
and matrix remodeling.
EXAMPLE 13
Engineered ECM Compositions Regulate Stem Cell Differentiation
A multipotential mesenchymal stem cell line (D1) derived from mouse
bone-marrow stroma was obtained from American Type Culture Collection (ATCC).
D1 cells were propagated in Dulbecco's modified Eagle medium containing 4.5
g/L
glucose, 110 mg/L sodium pyruvate, 100 1J/m1 penicillin, 100 jag/m1
streptomycin,
and 10% fetal bovine serum (FIBS) within a humidified atmosphere of 5% carbon
dioxide at 37 C. Three-dimensional collagen ECMs were prepared by dissolving
native, acid-solubilized type I collagen from calf skin (Sigma Chemical Co,
St. Louis,
MO) in 0.01 N hydrochloric acid to achieve desired concentrations. As a final
purification step the isolated collagen obtained from Sigma Chemical was
dialyzed
against an acidic solution having a low ionic strength (0.01 N HC1) for 1-2
days, using
dialysis tubing or a membrane having a molecular weight cut-off selected from
a
range of about 3,500 to about 12,000 daltons. For sterile preparations of
collagen, the
purified collagen solution was layered onto a volume of chloroform. After
incubation
for 18 hours at 4 C, the collagen solution layer was carefully removed so as
not to
include the collagen-chloroform interface layer.
To produce 3D purified collagen matrices with microstructures of
varied collagen fibril dimensions (e.g., length, diameter, density), collagen
solutions
were polymerized under different conditions. Specifically, to create collagen
matrices consisting of collagen fibrils at increasing densities, collagen
solutions were
polymerized at final collagen concentrations of 1.0 to 3.0 mg/ml. The
polymerization
composition comprised a 10X phosphate buffered saline (PBS) with an ionic
strength
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of 0.14 N and a pH of 7.4. The specific formulation of the 10X phosphate
buffer is as
follows:
10X PBS, pH 7.4
1.37M NaC1
0.027M KC1
0.081M Na2HPO4
0.015M KH2PO4
5mM MgC12
1% w/v glucose
To create 10X PBS buffers of different pH, the ratio of Na2HPO4 and KH2PO4 is
varied. Ionic strength can be adjusted as an independent variable by varying
the
molarity of NaC1 only. To create the 3D matrix comprising cells suspended
within
3D matrix microenvironment the following components were mixed together:
1 ml solubilized collagen (e.g., type I collagen) in 0.01N HC1
150u1 10X PBS, pH 7.4
150 ul 0.1N NaOH
100 ul 13.57 mM CaC12
100 ul 0.01 N HCI
Final Volume 1.5 ml.
The composition is mixed well after each additional component is added. The
composition is then combined with a cell pellet of known cell number to create
desired cell density; mixed well; and allowed to polymerize. The resulting
polymerized 3D matrix has a final concentration of glucose and CaC12 of about
5.55mM glucose and about 0.9046mM CaCl2.
To create collagen matrices consisting of collagen fibrils that varied in
length and width, collagen solutions were polymerized at a pH selected from
the
range of 6.5-8.5. D1 cells were harvested in complete medium, collected by
centrifugation, and added as the last component before polymerization. Tissue
constructs were prepared at a relatively low cell density of 5 x 104 cells/ml.
Previous
studies by applicants have shown that this cell density is suitable for
maintaining cell
viability, minimizing cell-cell interaction, and allowing the study of the
dynamic
relationship between an individual cell and its surrounding ECM.
Polymerization of tissue constructs was conducted in 24-well culture
plates maintained in a humidified environment at 37 C. Immediately after
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polymerization (20 minutes or less), complete medium was added and the tissue
constructs were cultured for 48 hours at 37 C in a humidified environment
consisting
of 5% CO2 in air. After 48 hours, each of the constructs comprising D1 cells
seeded
within a specific ECM microstructure were cultured under 3 different
conditions:
1) complete medium no supplements
2) complete medium plus 10-7 M dexamethasone
3) complete medium plus 50 jig/m1 ascorbic acid
For comparison purposes, parallel experiments were conducted on D1
cells grown in a standard 2D format on tissue-culture plastic. Cell behavior
and
morphology were monitored throughout the duration of the experiment using
standard
brightfield microscopy. After 24 days in culture, tissue constructs were
histochemically stained with alcian blue, oil red 0, and alizirin red as
indicators of
chondrogenesis, adipogenesis, and osteogenesis.
Results:
The results of this experiment revealed the following:
1) multipotential stem cells seeded within engineered ECMs
proliferated at rates that were dependent upon microstructural composition of
the
engineered ECM and the media composition;
2) time-dependent patterns of cellular condensation and
aggregation exhibited by multipotential stem cells were dependent upon
microstructural composition of the engineered ECM and the media composition;
3) time-dependent differentiation of multipotential stem
cells
seeded within engineered ECMs was dependent upon microstructural composition
of
the engineered ECM and the media composition;
4) maintenance of precursor or multipotential cells in an
undifferentiated state in vitro was dependent upon microstructural composition
of the
engineered ECMs and the media composition;
5) patterns of cellular proliferation/differentiation for cells grown
within 3D were different from those observed for cells grown in 2D on tissue
culture
plastic; and
6) Decreasing the cell density of viable multipotential stem cells
within engineered ECMs led to clonal growth of a large population of cells
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representing a single cell lineage. Estimates of optimal cell densities for
clonal
growth range from about 10 cells/ml to about 103 cells/ml and depend upon the
specific seeding efficiencies. For most cells, cell survival is known to
decrease with
seeding density.
EXAMPLE 14
Engineered ECM Compositions Regulate Unipotential Stem Cell Differentiation
A unipotential stem (precursor) cell line (L1) derived from mouse and
representing pre-adipocytes was obtained from American Type Culture Collection
(ATCC). Li cells were propagated in Dulbecco's modified Eagle medium
containing
4.5 g/L glucose, 110 mg/L sodium pyruvate, 100 U/ml penicillin, 100 ug/m1
streptomycin, and 10% fetal bovine serum (FBS) within a humidified atmosphere
of
5% carbon dioxide at 37 C. To enhance cell viability, cells representing
passage
numbers greater than 5 were maintained in complete media in which the
penicillin
and streptomycin were reduced to 25 U/ml and 25 ig/ml, respectively.
Preparation of tissue constructs representing Li cells seeded within 3D
engineered ECMs of different microstructural compositions was carried out as
described in Example 13. Immediately after polymerization (20 minutes or
less),
complete medium was added and the tissue constructs were cultured for 48 hours
at
37 C in a humidified environment consisting of 5% CO2 in air. After 48 hours,
each
of the constructs comprising Li cells seeded within a specific ECM
microstructure
were cultured under 3 different conditions:
1) complete medium no supplements; medium changed every 2
days thereafter;
2) complete medium no supplements and post differentiation
medium treatment every 2 days thereafter; and
3) differentiation medium and post differentiation medium
treatment every 2 days thereafter.
The differentiation medium consists of DMEM supplemented with
10% FBS, 25 U/ml penicillin, 25 jig/m1 streptomycin, 115 lag/MI methyl-
isobutyl
xanthine, 101.1g/m1 insulin, and 5 x 10-7M dexamethasone. The post
differentiation
medium consisted of DMEM supplemented with 10% FBS, 25 U/ml penicillin, 25
g/ml streptomycin, and 10 p.g/m1 insulin. For comparison purposes, parallel
experiments were conducted on Li cells grown in a standard 2D foiinat on
tissue-
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culture plastic. Cell behavior and morphology were monitored throughout the
duration of the experiment using standard brightfield microscopy.
Results:
The results of this experiment revealed the following:
1) unipotential stem (precursor) cells seeded within engineered
ECMs proliferated at rates that were dependent upon microstructural
composition of
the engineered ECM and the media composition;
2) time-dependent patterns of cellular condensation and
aggregation exhibited by unipotential stem cells were dependent upon
microstructural
composition of the engineered ECM and the media composition;
3) time-dependent differentiation of unipotential stem cells into
mature adipocytes seeded within engineered ECMs was dependent upon
microstructural composition of the engineered ECM and the media composition;
4) maintenance of precursor cells in an undifferentiated state in
vitro was dependent upon microstructural composition of the engineered ECMs
and
the media composition; and
5) patterns of cellular proliferation/differentiation for
cells grown
within 3D were different from those observed for cells grown in 2D on tissue
culture
plastic.
EXAMPLE 15
Effect of Fibril Microstructure and Mechanical Properties of 3D ECM on
Cultured
Stem Cells
Multi-potential stem cells derived from the bone marrow of mice (Dls;
ATCC) were suspended at 5 x 104 cells/ml within purified type I collagen
solutions
(Sigma Chemical Co.) at varying collagen concentrations ranging from 1.5-3.6
mg/ml
using the procedures described in Example 13. Tissue constructs consisting of
D1
cells entrapped within a 3D ECM were formed by inducing self-assembly
(polymerization) at pH 7.4, 137 mM NaC1, and 37 C. For this specific example,
an
increase in collagen concentration as a self-assembly parameter, was used to
generate
a 3D ECM microenvironment in which the density of the resultant fibrils and
stiffness
(linear or elastic modulus) of the matrix were systematically increased. The
3D
constructs and resident cells were maintained in one of three different media
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formulations (Table 8) at 37 C in a humidified environment consisting of 5%
CO2 in
oxygen for periods of time up to 4 weeks. Basal medium consisted of Dulbecco's
modified Eagle's medium supplemented with 4 mM L-glutamine, 4.5g/L glucose,
1.5
g/L sodium bicarbonate, 1 mM sodium pyruvate, 10% fetal bovine serum, 100 U/ml
penicillin, and 100 ,g/m1 streptomycin. For comparison purposes, D1 cells
also were
cultured in a parallel fashion in the standard 2D format on the surface of
tissue culture
plastic.
Table 8: Medium formulations used to culture D1 cells
Medium Designation Medium Formulation
A Basal medium supplemented with 1 j_iM dexamethasone,
0.5 mM isobutylmethylxanthine, 1 tig/m1 insulin
Basal medium supplemented with 0.1 ILLM dexamethasone,
8 ,g/m1 ascorbic acid, 5 mM13-glycerophosphate
Basal medium with no additives
After various periods of time, the proliferative and differentiation
status of the cells were determined qualitatively or quantitatively.
Qualitative
evaluation of cell number and morphology was conducted several times a week
using
light microscopy. Real-time RT-PCR was used to quantify and compare the
expression levels of CFBA1 (runx2), LPL (lipoprotein lipase), and procollagen
II as
indicators of osteogenesis (bone foiniation), adipogenesis (fat formation),
and
chondrogenesis (cartilage formation), respectively. Histochemical stains,
including
alkaline phosphatase and oil red 0, were applied to whole mount or
cryosectioned
samples to detect osteogenic and adipogenic activity, respectively. In some
cases
immunohistochemical staining was used to corroborate results.
Cells grown in basal culture medium with no additives (medium
formulation C) on standard tissue culture plastic (A) and within 3D ECM
micro environments of controlled fibril density and stiffness showed distinct
growth
patterns and morphologies. Results showed as the fibril density and stiffness
of the
3D ECM microenvironment increased, the proliferative capacity of the cells
decreased. The dependence of D1 proliferation on the stiffness of the 3D ECM
microenvironment was noted for all media formulations studied. More
specifically,
D1 cells grown on plastic or within the low stiffness 3D ECM microenvironment
showed an increased number of spindle-shaped cells. Within 2 weeks the cells
on
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plastic reached confluence and formed a sheet of cuboidal shaped cells. On the
other
hand, spindle-shaped cells were evident within the low stiffness ECM even
after 4
weeks of culture. These cells appeared to remain undifferentiated and
populated the
ECM uniformly. Growth patterns indicative of isolated clonogenic events were
higher in frequency within ECMs of increased stiffness.
The observed differences in the growth patterns and morphologies
adopted by cells grown in the 2D and 3D microenviromnents suggested that the
multi-
potential cells were being directed down distinct differentiation patterns.
Limited
directed differentiation appeared to occur for cells grown on plastic or
within the low
stiffness ECMs (1.5 mg/ml). Interestingly, D1 cells grown within 3D ECMs of
high
stiffness (3.4 mg/ml) formed regional aggregates of cells indicative of
osteogenesis
and/or skeletal myogenesis. Osteogenesis but not myogenesis events were also
observed with engineered ECMs of moderate stiffness (3 mg/ml). The biochemical
composition of the media also could be varied to enhance the differentiation
of cells
down a specific pathway or to maintain cells in a relatively undifferentiated
state.
Specifically, cells grown in medium formulation A demonstrated a high
frequency of
adipogenesis on plastic and within 3D ECMs of low fibril density and
stiffiless (1.5
mg/ml). As the fibril density and stiffness of the 3D ECM microenvironment
increased, adipogenesis events decreased and osteogenesis increased. Medium
formulation B appeared to support differentiation of D1 cells into fat
(adipogenesis)
and (bone) osteogenesis on plastic. Limited areas of osteogenesis and
adipogenesis
were noted amongst a large number of spindle-shaped cells for D1 cells grown
within
ECMs of low stiffness under these same medium conditions. As the stiffness of
the
3D ECM increased, cells more uniformly developed regional areas of
osteogenesis
and myogenesis-like events. A 2D projection of one confocal image revealed
cells
organized or fused to form a multi-cellular structure reminiscent of a
myotube. These
events were limited to 3D ECM microenvironments of high stiffness (3.4 mg/ml
and
greater). While these myotube-like events were noted in all three medium
formulations, they appeared to occur more frequently in medium formulations B
and
C. The cells of the myotube-like structure were stained immunohistochemically
for
F-actin to demonstrate the fusion of and connectivity of the actin
cytoskeleton
between individual cells.
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Real-time RT-PCR confirmed that biophysical features of the 3D ECM
microenvironment (e.g., fibril density and ECM stiffness) could be modulated
to
regulate stem cell growth and differentiation. Fig. 8 shows the differences in
gene
expression patterns for D1 cells grown for two weeks on tissue culture plastic
(Plastic) and within 3D engineered ECMs prepared at low (1.5 mg/ml), moderate
(3.0
mg/ml), and high fibril density and stiffness (3.6 mg/ml). Again, cells
subjected to
each of these 2D and 3D culture formats were maintained in one of three
different
media foimulations (Table 8).
The tissue specific genes CBFA1 (runx2), LPL (lipoprotein lipase),
and Pro Col II (procollagen II) were selected as indicators of osteogenesis,
adipogenesis, and chondrogenesis, respectively. Results showed that cells
grown for
2 weeks on 2D plastic in the basal medium (no additives) remain relatively
undifferentiated, more specifically, limited expression of the osteogenic,
adipogenic,
and chondrogenic indicators. On the other hand, D1 cells show an increase in
LPL
(adipogenesis) when cultured on plastic in the presence of Medium A or Medium
B.
The expression of LPL correlates well with the observed fat cell morphology
developed within the cultures. Interestingly, the gene expression patterns
developed
by cells grown within a 3D ECM microenvironment were dramatically different
from
those observed for cells grown on plastic. Specifically, the expression of
CBFA1,
indicative of osteogenesis, could be enhanced by growing the cells within 3D
ECMs
of increased stiffness or Medium B. Again, the increased expression of CFBA1
correlated well with cell morphologies and histochemical staining.
Interestingly,
chondrogenesis events as indicated by high procollagen II expression appeared
to be
enhanced within DI cells cultured in 3D ECMs of high stiffness.
The starting cell density was also a critical determinant of the stem cell
fate within the 3D culture formats studied. Clonal growth and cell
differentiation
events were favored by increasing the ECM stiffness and/or by decreasing the
starting
cell density within a given 3D ECM format, Adipogenesis was favored by
decreasing
the ECM stiffness and/or by increasing the cell density. Interestingly,
adipogenesis
was observed within high stiffness 3D ECMs only when the cell seeding density
approached 1 x 106 cells/ml and above.
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EXAMPLE 16
Cell Culture
Low passage neonatal human dermal fibroblasts (NHDFs) were
obtained from Cambrex Bioproducts (Walkersville, MD). NHDFs were propagated
in fibroblast basal medium supplemented with human recombinant fibroblast
growth
factor, insulin, gentamicin, amphotericin-B, and fetal bovine serum (FBS)
according
to manufacturer's recommendations. Cells were grown and maintained in a
humidified atmosphere of 5% CO2 at 37 C. Cells representing a limited passage
number of 20 or less were used for all experiments.
EXAMPLE 17
Preparation of 3D Engineered ECMs and 3D Tissue Constructs
Purified type I and type III collagens, that were solubilized from
bovine deimis and human placenta, respectively, were obtained from Sigma
Chemical
Company (St. Louis, MO). Three-dimensional engineered ECMs were prepared at a
constant collagen type I concentration of 1.5 mg/ml and type III collagen
concentrations of either 0, 0.25, 0.50, and 0.75 mg/ml (Table 9) using the
general
procedures described in Example 13. The polymerization buffer consisted of 10X
phosphate buffered saline (PBS) with an ionic strength of 0.14 M and a pH of
7.4.
All 3D engineered ECMs and tissue constructs were polymerized in vitro within
a
humidified environment at 37 C. To determine the cellular signaling capacity
of each
3D ECM microenvironment, 3D tissue constructs were formed by first harvesting
NHDFs in complete media and then adding the cells as the last component to the
collagen solutions prior to polymerization. Tissue constructs were prepared at
a
relatively low cell density of 5x104 cells/ml in order to minimize cell-cell
interactions.
Immediately after polymerization (20 minutes or less), complete medium was
added
and the tissue constructs were maintained at 37 C in a humidified environment
consisting of 5% CO2 in air.
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Table 9. Summary of formulations for 3D engineered ECMs prepared with varied
ratios of collagen types I and III.
Type I Collagen Type III Collagen Total Collagen Collagen Type VIII
Type III Collagen
(mg/nil) (mg/ml) (mg/ml) Ratio (% of Total)
1.5 0 1.50 0 0
1.5 0.25 1.75 6:1 14.3%
1.5 0.50 2.00 3:1 25.0%
1.5 0.75 2.25 2:1 33.3%
A summary of the results of the data generated by the experiment of
Example 13-17 is provided in Fig. 9.
EXAMPLE 18
Preparation of Two-Dimensional (2D) ECM Surface Coatings
To prepare 2D surfaces coated with the different ECM compositions,
solutions containing collagen type 1(1.5 mg/ml) and varying concentrations of
collagen type III (0, 0.25, 0.5, and 0.75 mg/ml) were aliquoted (300 ill/well)
into
tissue culture plates (24-well) and air-dried within a laminar flow hood for
approximately 18 hours. Well plates containing 2D ECM surface coatings were
equilibrated with PBS, pH 7.4, prior to seeding the NHDFs at a density of
2.5x104
cells/well. Complete medium was added and the NHDFs on the surface of the 2D
ECM coatings were maintained at 37 C in a humidified envirorunent consisting
of 5%
CO2 in air.
EXAMPLE 17
Qualitative and Quantitative Analysis of 3D ECM Microstructural Composition
Two quantitative parameters describing the 3D ECM microstructural
composition, fibril area fraction (a 2D approximation of 3D fibril density)
and fibril
diameter, were determined based upon confocal reflection and scanning electron
microscopy (SEM) images. Prior to microstructural analysis, engineered 3D ECM
constructs were polymerized within four-well Lab-Tek coverglass chambers
(Nalge
Nunc International, Rochester, NY) and placed within a humidified environment
at
37 C where they were maintained for approximately 15 hours. For measurements
of
fibril area fraction, the confocal microscope was used to obtain high
resolution, 3D,
reflection images of the component collagen fibrils within each ECM (Brightman
at
al., Biopolymers 54: 222-234, 2000; Voytik-Harbin et al., Methods Cell Biol
63: 583-
597, 2001). Three images (at least 10 m in thickness) were taken at random
locations within each of 2 specimens representing a given 3D ECM composition.
The
confocal image stacks were then read into Matlab (The Mathworks, Natick, MA),
and
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2D projections, representing 21 z-sections, of each image were created and a
threshold chosen for binarization. Using a built-in function in Matlab, the
area
occupied by collagen fibrils (white pixels) was calculated, converted to pM2
based
upon the pixel sizes, and normalized to the total image area.
Fibril diameter measurements were made by applying Imaris 4.0
(Bitplane Inc., Saint Paul, MN) to both confocal reflection and SEM images of
engineered ECM constructs. For SEM imaging, engineered ECM constructs were
fixed in 3% glutaraldehyde in 0.1M cacodylate at pH 7.4, dehydrated with
ethanol,
and critical point dried. Samples were sputter-coated with gold/palladium
prior to
imaging. Duplicate samples were imaged in a JEOL (Peabody, MA) JSM-840 SEM
using 5 kV accelerating voltage and a magnification of 3,000X. Digital images
were
captured using 1280x960 resolution and 160 second dwell time. From each image
obtained from duplicate samples, forty fibrils were chosen at random (10
fibrils per
quadrant). Five lines were drawn perpendicular to the long axis of each fibril
using
the measurement tool in Imaris (Brightman at al., Biopolymers 54: 222-234,
2000).
The average number of pixels representing the fibril diameter was then
converted into
Jim based upon the known pixel size.
EXAMPLE 18
Measurement of Tensile Mechanical Properties of 3D Engineered ECMs
Specimens for mechanical testing were prepared by polymerizing each
soluble ECM formulation in a "dog bone" shaped mold as described previously
(Roeder et al., J Biomech Eng, 124: 214-222, 2002). In brief, the mold
consisted of a
glass slide and a piece of flexible silicone gasket. The gauge section of the
mold
measured 10 mm long, 4 mm wide, and approximately 1.5 mm thick. Neutralized
ECM solution was added to each mold and allowed to polymerize at 37 C in a
humidified environment where they were maintained for 18-20 hours prior to
tensile
loading. Polypropylene mesh was embedded in the ends of each 3D ECM construct
to facilitate clamping for mechanical loading. Low-magnification, 4D images
(x, y, z,
and time) of each ECM construct during uniaxial tensile loading were acquired
using
an integrated mechanical loading-stereomicroscope set-up. This set-up involved
interfacing a modified (Roeder et al., J Biomech Eng, 124: 214-222, 2002.)
Minimat
2000 miniature materials tester (Rheometric Scientific, Inc., Piscataway, NJ)
with a
Stemi 2000-C Stereomicroscope (Carl Zeiss MicroImaging; Thomwood, NY)
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mounted with a DFC480 high-resolution color digital camera (Leica
Microsystems,
Cambridge, IN. Strategically placed right-angle prisms (Edmund Industrial
Optics,
Barrington, NH) were used to monitor changes in specimen thickness (z-
direction)
throughout the loading process. The image field was positioned to include the
clamp
that was attached to the load cell in order to provide a "fixed" frame of
reference
throughout the loading process. Each ECM construct was loaded uniaxially at an
extension rate of 1 mm/min (corresponding to a strain rate of i= "-=,' 0.04
/min) until
failure. Images were collected at a rate of 0.1 frames/sec to provide
sequential images
at 0.64% strain intervals. Changes in the width (x-direction) and thickness (z-
direction) dimensions of the specimen's gauge section were measured directly
from
low-magnification digital camera images representing the width and thickness
of the
specimen and used to calculate cross-sectional area. The mechanical behavior
of each
specimen, including engineering stress (a,), true stress (o-t), and applied
strain (sap) were
calculated from load-displacement recordings provided by the Mini-mat. Applied
strain was calculated by simplifying the Lagrangian strain definition
(Malvern,
Introduction to the Mechanics of a Continuous Medium. Upper Saddle River, NJ:
Prentice-Hall, 1969) for a simple stretch X (new length divided by original
length) as
indicated below
sap = (22 (1)
Engineering stress was calculated as
o-e ¨
A, (2)
where, F was the force recorded by the Minimat and A, was the initial cross-
sectional
area (width x thickness) within the center of the specimen (Callister et al.,
Materials
Science and Engineering: An Introduction. 3rd edition. New York, NY: John
Wiley &
Sons, 1994). For calculation of true stress, the actual cross-sectional area
of each
specimen at a specific load was imaged, quantified, and substituted for A, in
the
engineering stress equation above. From the resulting stress-strain
relationships
ultimate strength (maximum stress achieved during tensile loading), failure
strain
(strain at which specimen fails), and linear or elastic modulus (stiffness;
slope of
linear region of stress-strain curve) were determined.
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EXAMPLE 19
Multi-Dimensional Confocal Imaging of Cell-ECM Interactions
All multi-dimensional imaging was performed on a Bio-Rad Radiance
2100 MP Rainbow (Bio-Rad, Hemel Hempstead, England) multi-photon/confocal
system adapted to a TE2000 (Nikon, Tokyo, Japan) inverted microscope with a
heated stage set at 37 C (ALA Scientific Instruments, Westbury, NY). A custom-
designed environmental chamber was adapted to the microscope to provide tissue
constructs with a sterile environment of 5% CO2 in humidified air (Pizzo et
al., J Appl
Physiol 98: 1909-1921, 2005). For each of the engineered ECMs studied, at
least 5
individual cells were repeatedly monitored during the first 5 hours following
construct
polymerization. During the collection of time-lapse images, the confocal
microscope
was used in a reflection (back-scattered light) mode to obtain image stacks of
an
individual cell and the component collagen fibrils of its surrounding ECM as
described previously (Brightman et at., Biopolymers 54: 222-234, 2000; Voytik-
Harbin et al., Methods Cell Biol 63: 583-597, 2001). Images were collected at
30-
minute intervals and a z-step of 0.5 j.tm to minimize exposure of the tissue
constructs
to radiation from the confocal microscope laser.
EXAMPLE 20
3D Cell Morphometric Analysis
Three-dimensional confocal images used for qualitative and
quantitative analyses of NBDF morphology were collected using the confocal
microscope in a combination reflection-epifiuorescence mode (Voytik-Harbin et
al.,
Methods Cell Biol 63: 583-597, 2001; Voytik-Harbin et al., .Microsc Microatial
9: 74--
85, 2003). Immediately following the 6-hr time-lapse imaging, tissue
constructs were
stained with the vital dye, Cell Tracker Green (Molecular Probes, Eugene, OR),
to
facilitate discrimination of the cell from the surrounding collagen ECM. The
processed image stack was used to determine fundamental morphological
parameters
including number of cytoplasmic projections, cell volume, 3D cell surface
area,
length, width, and height as described previously (Pizzo et at., .1 Appl
Physiol 98:
1909-1921, 2005). Since each cell had a relatively unique orientation within
the 3D
matrix, these morphological parameters were defined based on a cellular
coordinate
system. Morphological evaluation was conducted on a total of 10 to 23 cells
for each
of the 3D ECM compositions studied.
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EXAMPLE 21
Determination of 3D Average Local Principal Strains and Points of Maximum
Local
Principal Strain
Consecutive time-lapse confocal reflection images representing the
time-dependent deformation to the collagen fibril microstructure induced by an
individual resident cell provided the basis for quantification of local ECM
remodeling
in terms of 3D displacements and strains. Strains were quantified using an
incremental digital volume correlation ODVC) algorithm developed previously by
our
laboratory (Roeder et al., J Bionlech Eng 126: 699-708, 2004). To determine 3D
average local principal strains within the surrounding ECM induced by an
individual
cell, first the 15 x 15 x 3 grid of displacements (174.2 x 174.2 x 24 m3
total volume)
from the IDVC algorithm were converted into 3D strains with x, y, and z
directions
based on the confocal coordinate system. The strains in the entire volume were
then
averaged in each of the three confocal directions to give a 3 x 3 symmetric
matrix of
average strains, Eavg, such that
xx xy 8 xz
Eavg = xy yy yz (3)
C yzzz
where gii are average strains in the confocal coordinate system directions.
This
average strain matrix in Equation (3) was then solved using eigenvector
analysis
(Strang et al., Linear Algebra and Its Applications. 3rd edition. San Diego,
CA:
Academic Press, 1988) to determine 3 average principal strains (EL E2, E3) and
associated directions such that,
E, 0 0
Eavg = [V] = [V] = 0 E2 0 (4)
0 0 E3 _
where [V] is a 3 x 3 matrix such that the column vectors (VI, V2, V3) are the
directions
of the principal strains given by
{v]= [r' V2 V3] (5)
Therefore, the deformation induced by each cell had a unique set of average
principal
strains and directions in 3D.
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Another analysis was performed to determine on a finer scale where
the maximum local principal strains within the 3D ECM occurred in relationship
to
the cell. This analysis involved determination of local principal strains El,
E2, and B3,
each with unique principal direction, at each of the 15 x 15 x 3 grid points.
The
maximum compressive El, E2, and E3 were then identified within the image
volume.
The location of each maximum compressive principal strain was known in terms
of its
IDVC grid location and also in m. The distance from these three-maximum
principal strain locations to the center of the cell body in 3D could then be
determined
using simple vector relationships. The locations of the maximum compressive
principal strains did not necessarily occur at the same grid locations for
each cell.
EXAMPLE 22
Labeling and Visualization of Actin Cytoskeleton within 3D Engineered Tissue
Constructs
Tissue constructs formed by seeding NHDFs within specific 3D ECM
formulations during polymerization were prepared in four-well Lab-Tek
coverglass
chambers (Nalge Nunc International, Rochester, NY) for visualization of the F-
actin
cytoskeleton. At specified timepoints, constructs were fixed and permeabilized
with a
solution containing 0.1% Triton 100X and 3% paraformaldehyde, post-fixed in 3%
paraformaldehyde, and treated with 1% bovine serum albumin to minimize non-
specific binding. The constructs were then stained overnight at 4 C with Alexa
Fluor
488 Phalloidin (Molecular Probes, Eugene, OR) and rinsed. Three-dimensional
images of the F-actin distribution within an individual cell as well as its
surrounding
ECM were collected simultaneously using confocal microscopy in a combined
epifluorescence and reflection mode. When necessary, images were deconvolved
using AutoDeblur*(Autoquant Imaging, Inc., Watervliet, New York).
EXAMPLE 23
Qualitative and Quantitative Determination of Cell Proliferation
Quantification of NHDF proliferation and its dependency on the 3D
ECM microenvironment involved preparing 3D tissue constructs within 24-well
tissue-culture plates using an alarnarBlue-based proliferation assay as
described
previously (Pizzo et at, J Appl Physiol 98: 1909-1921, 2005; Voytilc-Harbin et
al., In
Vitro Cell Dev Biol All171134: 239-246, 1998). For comparison purposes, the
proliferative capacity of NHDF was also determined for an equivalent number of
cells
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seeded directly onto the plastic surface of a well-plate as well as 2D plastic
surfaces
coated with different ECM compositions consisting of type I collagen in the
presence
of varying amounts of type III collagen. At time points representing 24 and 48
hours
after construct polymerization and/or cell seeding, each well and tissue
construct was
examined microscopically to observe the viability, number, and morphology of
the
cells. The medium from each well then was replaced with fresh medium
containing
the metabolic indicator dye alamarBlue (10% v/v; BioSource International,
Inc.,
Camarillo, CA). Twenty-four hours later, dye reduction was monitored
spectrofluorometrically using a FluoroCount*Microplate Fluorometer (Packard
Instruments, Meriden, CT) with excitation and emission wavelengths of 560 Dm
and
590 urn, respectively. Background fluorescence measurements were determined
from
wells containing only dye reagent in culture medium. Maximum levels of
relative
fluorescence were determined from alamarBlue solutions that were autoclaved to
induce complete dye reduction. The mean and the standard deviation values for
all
fluorescence measurements were calculated and subseqUently normalized with
respect
to the background and maximum fluorescence readings. All experiments were
performed in triplicate and repeated at least three times. When relevant,
statistical
analyses were performed using Matlab and included an analysis of variance
(ANOVA). The Tukey-Kramer method for multiple comparisons (p<0.05) was then
applied. The two-tailed t-test (a 0.05) was applied for pairwise comparisons.
EXAMPLE 24
Three-Dimensional Microstructural Composition of Engineered ECMs Depends upon
Collagen Type I/III Ratio
This study utilized the application of confocal microscopy in a
reflection mode and SEM facilitated microstructural analysis from both 2D and
3D
perspectives as well as at two different limits of resolution. SEM provided
high-
resolution (approximately 10 mu) 2D images of ECM microarchitecture after
specimens had been critical point dried. On the other hand, confocal
reflection
microscopy allowed visualization of the 3D microstractural organization of
component collagen fibrils within engineered ECMs in their fully hydrated or
native
state; however the resolution obtained with confocal imaging is approximately
200
DM, twenty times less than that obtained with SEM.
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Both confocal and SEM images showed that ECMs prepared with
increased amounts of type III collagen possessed an increased number or
density of
collagen fibrils. The fibril area fraction (Fig. 2A) was quantified from
confocal
reflection images and showed a nearly linear increase with type III collagen
over the
range studied. Engineered ECMs prepared from type I collagen alone had a
fibril area
fraction of 12.0 1.4% compared to 21.5 2.6% for those fowled in the presence
of the
highest concentration (0.75 mg/ml) of type III collagen. In addition to this
effect on
fibril density, increased levels of type III collagen resulted in a downward
shift in the
fibril diameter distribution (Table 10) and (Fig. 2B). The mean fibril
diameter as
determined from SEM images for ECMs prepared with type I collagen alone was
115.2 23.2 pm. The mean fibril diameter showed a significant (p<0.05)
decrease to
94.8 23.0 pm and 87.0 17.0 tun for ECMs in which collagen III was added at
levels of 0.25 mg/ml and 0.75 mg/ml, respectively. In general, fibril diameter
measurements made from confocal reflection images corroborated SEM results;
however, fibril diameter values obtained from confocal images were greater
than
those obtained using SEM (Table 10) since confocal imaging was conducted on
unprocessed, fully hydrated specimens. It should be noted that fibril diameter
measurements made using confocal reflection imaging were considered somewhat
less accurate and less precise since fibril diameters were near the limit of
resolution
for this imaging technique.
Table 10. Collagen fibril diameter measurement data for 3D engineered ECMs
prepared from type I collagen in the absence and presence of type III collagen
as
determined from scanning electron (SEM) and confocal reflection (CRM) images.
Fibril Type I Collagen (1.5 mg/ml)
Diameter Type I Collagen (1.5 mg/ml) Type III Collagen (0.75
mg/ml)
(nm) SEM CRM SEM CRM
Mean SD 115.16 412.63 76.35 87.04 17.00
384.60 71.96
23 A 8
Median 112 408 86 378
Range 78-194 200-664 56-176 236-628
EXAMPLE 25
Mechanical Properties of Engineered ECMs Depend upon Collagen Type I/III Ratio
Previously, we showed that engineered ECMs prepared at increasing
concentrations of type I collagen featured an increase in fibril density but
no
significant change in fibril diameter (Roeder et al., J Biomech Eng, 124: 214-
222,
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2002). Furthennore, this change in ECM microstructure, specifically an
increase in
collagen fibril density, was found to be positively correlated with ECM
tensile
strength and stiffness (linear or elastic modulus (Roeder et al., J Biomech
Eng, 124:
214-222, 2002)). Traditionally, mechanical properties for collagen-based
matrices
have been calculated based upon engineering stress (Osborne et al., Med Biol
Eng
Comput 36: 129-134, 1998; Ozerdem et al., J Biomech Eng 117: 397-401, 1995;
Roeder et al., J Biomech Eng, 124: 214-222, 2002), which assumes no change in
, specimen cross-sectional area during mechanical loading. However, since it
is known
that our engineered ECMs exhibit Poisson's ratios on the order of 2 to 4
(Roeder et
al., J Biomech Eng, 124: 214-222, 2002; Voytik-Harbin et al., Microsc
Microanal 9:
74-85, 2003), true stress was calculated to account for the significant
reduction in
cross-sectional area experienced by the scaffolds during testing. Since our
experimental set-up facilitated the continuous monitoring of changes in
specimen
cross-sectional area during tensile loading, true stress calculated parameters
were
considered to most accurately reflect mechanical behavior of the ECMs.
ECMs engineered from type I collagen in the presence of type III
collagen over the range of 0 to 0.75 mg/ml (type III collagen content of 0 to
33.3%)
showed biphasic responses in terms of true stress calculated parameters
ultimate
strength and stiffness. The mean ultimate strength obtained for ECMs prepared
from
1.5 mg/ml type I collagen alone was 136.7 49.9 kPa. Addition of collagen III
resulted in significant reductions in ultimate strength, with 66.7 4.2 kPa
(p=0.0016)
and 75.1 22.7 kPa (13=0.0085) values being measured for ECMs prepared at
collagen
III levels of 0.25 mg/ml and 0.75 mg/ml, respectively. The linear or elastic
modulus
(stiffness), as determined from the linear region of the stress-strain curve,
also showed
reductions of 32% (1)=0.0002) at the 0.25 mg/ml collagen III level and 18%
(p=0.189)
at the 0.75 mg/ml collagen III level compared to those where no collagen III
was
added. A decline in failure strain with increasing type III collagen content
was noted.
Specifically, failure strain values decreased significantly from 62.2 12.2%
when no
collagen III was added to 53.3 1.4% (p=0.048) and 43.0 5.9% (p=0.002) when the
type III collagen content was 0.25 mg/ml and 0.75 mg/ml, respectively.
Finally,
increasing the type I collagen content from 1.5 to 3 mg/ml increased ECM
ultimate
strength and stiffness, continuing previous findings (Roeder et al., J
Bioniech Eng,
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124: 214-222, 2002). ECMs prepared at 3 mg/ml type I collagen had ultimate
strength and stiffness values that were 2.2 and 3.5 times, respectively, those
obtained
for ECM prepared at 1.5 mg/ml type I collagen.
EXAMPLE 26
3D Cell Morphology Depends upon Collagen Type I/III Ratio
The ability of cells to sense and respond to changes in the 3D ECM
microenvironment that resulted from the addition of type III collagen
initially was
assessed by determining and comparing 3D cell morphology and cell-induced ECM
remodeling (defomiation and reorganization of component collagen fibrils).
Three-
dimensional morphometric analyses for cells seeded within the different ECM
microenvironments were conducted at 6 and 12 hours following tissue construct
,
formation. ECM remodeling by individual cells was repeatedly monitored during
a 5
to 6 hour time window shortly after construct formation.
Notable differences in 3D cell morphometric parameters were detected
at both 6 and 12 hours as the cells probed and adapted to their extracellular
microenvironment.
One of the more prominent differences noted at 6 hours was that cells
seeded within engineered ECMs prepared at the lowest type III collagen content
(0.25
mg/nil) took on a rounded morphology with multiple short projections. This
cell
morphology contrasted that observed within ECMs prepared with 0.5 mg/ml and
0.75
mg/ml type III collagen. At these higher collagen III levels a large
percentage of cells
took on a more spindle-shaped cell body with fewer but prominent lengthy
projections. Cells seeded within ECMs prepared from type I collagen alone took
on a
more spindle, bipolar shape and possessed the fewest (on average 2 to 4) but
longest
projections at the 6-hour time point.
By 12 hours, the morphological differences that resulted from the
various ECM microenvironments were subtler, largely owing to varying levels of
ECM remodeling induced by cells at this time. At 12 hours, cells seeded within
all
ECM formulations appeared relatively spindle, bipolar-shaped. However,
qualitative
and morphometric analyses indicated that as the type III collagen content
increased
within the ECM, cells showed a statistically significant decrease in length (p
<0.05),
a subtle increase in width, and a decline in the length-to-width ratio as
indicated (Fig.
3A-C). Based upon both qualitative and quantitative 3D morphology data, it
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appeared that cells grown in ECMs containing type III collagen took on a more
contracted cell state. Despite the observed changes in cell shape, no
significant
differences were noted in 3D cell surface area (Fig. 3D) or cell volume at
either time
point. Consistent with previous studies (Pizzo et al., J Appl Physiol 98: 1909-
1921,
2005), a larger proportion of cells grown within ECMs of higher total collagen
content possessed an increased number of cytoplasmic projections at both 6-
and 12-
hour time points; however, this effect on projection number was less obvious
when
the collagen content was altered by adding type III collagen rather than type
I
collagen. Collectively these results demonstrate that cells adapt their shape,
including
the number and length of their projections, in response to ECMs that vary in
collagen
type ratio. Furthermore, the morphological differences between cells
appeared to
be related to stiffiless properties of the ECM.
EXAMPLE 27
Collagen Type I/III Ratio of 3D Microenvironment Affects Contractile State of
the
Cell and ECM Remodeling
The collagen type I/III ratio also affected the ability of individual cells
to deform and reorganize the component collagen fibrils of their surrounding
ECM.
Repeated monitoring of interactions between a cell and its surrounding
collagen
fibrils within a live tissue construct by confocal reflection microscopy
provided a
means of visualizing and quantifying this response over a 5 to 6 hour time
window.
An IDVC algorithm (Roeder et al., J Biomech Eng 126: 699-708, 2004) was
applied
to consecutive confocal image stacks and used to determine 3D displacements
and
strains as they occurred locally to a given cell and adjacent collagen
fibrils. Data
generated from this algorithm provided the basis for 1) quantification of
volumetric
strain induced by a single cell within a tissue construct; 2) a detailed
analysis of
average local principal strains for each imaged volume; and 3) determination
of
magnitudes and locations for points within the image volume where maximum
principal strains, El, E2, and E3, occurred. This data was then compiled and
used to
compare the mechanical status of a large number of individual cells grown
within the
different ECM formulations.
Results showed that cells grown in ECMs containing type III collagen
were less able to contract and remodel the surrounding matrix as the type III
collagen
content increased or type 11111 ratio decreased. Qualitative perspectives and
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corresponding volumetric strain data obtained for representative cells grown
within
type I collagen ECMs prepared with low (0.25 mg/ml) and high (0.75 mg/ml) type
III
collagen concentrations are shown (Fig. 4). Comparison of average local
principal
strains induced by cells grown within the different ECM formulations indicated
that
cells grown at low type III collagen levels (0,25 mg/ml) induced higher strain
(approximately 3 to 3.5 greater) in each of the three principal directions
compared to
those grown at high type III collagen levels (0.75 mg/ml) and these
differences were
significant for E2 and E3 (p<0.05; Fig. 4). However, it is important to note
that type
III collagen containing ECMs with total collagen contents of 1.75 mg/ml to
2.25
mg/ml were characterized by 3D average local principal strain levels that were
about
2 to 3 times greater than ECMs prepared of type I collagen alone and a total
collagen
content of 1 mg/ml. Analysis of the locations and magnitudes for points. of
maximum
principal strain in the 1-, 2-, and 3-direction revealed that cells grown
within
engineered ECMs of type III collagen content of 0.25 mg/ml induced strain
values
that were approximately twice that exerted by cells grown within ECMs
containing
0.75 mg/m1 type III collagen (Fig. 5). Furthermore, in general, points of
maximum
principal strain for all three directions occurred at distances further from
the center of
the cell (Fig. 6) when grown in ECMs at the low versus high type III collagen
content.
Specifically, maximum principal strains were observed at distances of 40-50
!Am from
the center of the cell for ECMs containing 0.25 mg/ml type III collagen.
However,
cells within ECMs prepared with a type III collagen content of 0.75 mg/ml
generated
maximum principal strains at distances of only 15-25 1mi from the center of
the cell.
Although the addition of type III collagen enabled cells to induce large
principal
strains within their ECMs, the distance at which maximum principal strain
occurred
was considerably less than that found for ECMs prepared at low levels of type
I
collagen alone (Fig. 6). More specifically, ECMs prepared at 1 mg/ml type I
collagen
yielded, on average, points of maximum principal strain for 1-, 2-, and 3-
directions at
distances of 48 pm, 45 m, and 52 pm from the center of the cell,
respectively. It
was also noted that the locations of the maximum principal strain were often
associated with and occurred along major cell projections, especially for
cells grown
at the high collagen type I/III ratios. Furthermore, fibril deformation
patterns were
dependent upon the collagen type I/III ratio. Remodeling of ECMs containing
type
III collagen was characterized by fibril condensation around the cell
periphery. On
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the other hand, ECMs prepared from type I collagen alone showed regional areas
of
fibril alignment. The difference in ECM remodeling, as indicated by both
qualitative
fibril deformation and quantified strains suggested differences in mechanical
properties between fibrils foimed from homotypic type I and heterotypic type
I/III
fibrils.
Based collectively on the observed effects of type III collagen on ECM
microstructure/mechanical properties as well as differences in the 3D cell
morphology
and cell-induced ECM remodeling within these matrices, it was hypothesized
that
varying the collagen type I/III ratio within the ECM microenvironment
modulated the
contractile state of resident cells. To test this hypothesis, cells were
seeded within the
3D ECM microenvironments and the organization of cytoskeletal actin was
visualized
using confocal microscopy 6 hours post polymerization. Results showed
prominent
actin stress fiber formation for cells within ECMs containing collagen III.
Well-
organized actin bundles (stress fibers) were even observed within ECMs
containing
the highest collagen III concentration, 0.75 mg/ml, despite the high total
collagen
content and fibril density. Cells containing a few scattered actin filaments
were
observed in ECMs prepared from type I collagen alone, but only at low collagen
concentrations of 1.5 mg/ml and below. Cells with diffuse actin staining
patterns
were noted within ECMs prepared at collagen I levels greater than 1.5 mg/ml.
Diffuse actin staining patterns were observed for cells grown in engineered
ECMs
representing type I collagen ECMs prepared at concentrations of 1 mg/ml and 3
mg/ml. A few organized actin bundles were noted in engineered ECMs created
from
1 mg/ml type I collagen and a large number of organized actin bundles running
parallel along the major cytoplasmic projections or long axis of cells grown
within
engineered type I collagen ECMs formed in the presence of 0.75 mg/ml type III
collagen. Collectively, results showed that stress fiber formation, which is
indicative
of the contractile state of the cell, was positively related to cell-induced
local ECM
remodeling and strain and inversely related to ECM stiffness.
EXAMPLE 28
Collagen Type I/III Ratio within 3D Engineered ECMs but not 2D ECM Surface
Coatings Modulates Cellular Proliferation
To deteimine the effect of the collagen type I/III ratio on the
fundamental proliferative behavior of cells, NHDFs were seeded within the
different
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ECM formulations. For comparison purposes, parallel studies were conducted in
which fibroblasts were seeded onto tissue culture plastic. The number of
living cells
present at 24 and 48 hours following cell seeding was quantified indirectly
using the
metabolic indicator dye alamarBlue and confirmed qualitatively. Consistent
with
previous studies (Pizzo et al., J Appl Physiol 98: 1909-1921, 2005), cells
grown
within a 3D ECM microenvironment proliferated at decreased rates compared to
those
grown in a 2D format on tissue culture plastic (Fig. 7A). Fibroblast
proliferation was
enhanced in ECMs with increased type III collagen content (Fig. 7A). Since the
type
I collagen content was kept constant, increasing the amount of type III
collagen also
increased the overall collagen content. Although the total number of cells
within all
ECM formulations increased between 24 and 48 hours, the total number of
fibroblasts
was greatest for ECMs prepared with the highest type III collagen
concentration for
both time points. When type III collagen was added at levels below 0.25 mg/ml,
in
the range of 0.02 mg/ml to 0.10 mg/ml, the proliferative capacity of the
resident cells
was lower than that obtained for 1.5 mg/ml type I collagen alone.
Since the addition of type III collagen affected not only
microstructural-mechanical properties but also the macromolecular composition
of
the engineered ECMs, it was uncertain if changes in NHDF proliferation were a
result
of differences in biophysical or biochemical signals (cues) inherent in the 3D
ECM
microenvironments. To isolate the biochemical and biophysical variables,
traditional
experimental methods involving creation of 2D ECM surface coatings consisting
of
varied collagen I/III ratios to evaluate cell-ECM interactions were applied.
NHDF
were seeded onto the ECM-coated surfaces and proliferation monitored. No
significant difference was observed in cell proliferation due to type III
collagen
content at either the 24- or 48-hour time points (Fig. 7B). All coatings
showed a
significant increase (p<0.05) in cell number between the 24- and 48-hour time
points.
And at the 48-hour time point, cells seeded on plastic showed significantly
greater
proliferation than those seeded on any of the ECM coated surfaces (p<0.05).
EXAMPLE 29
Comparison of Structure-Function of Engineered ECM Formulations
Various engineered ECM formulations were compared to analyze
three-dimensional microstructure-mechanical properties, including fibril area
fraction,
fibril diameter, and stiffness of the engineered ECM (Table 11). The various
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engineered ECM formulations were also compared in regards to ECM contraction,
morphology, and cell proliferation (Table 11).
Table 11. Comparison of Structure-Function of Engineered ECM Formulations
1.5 mg/ml Type I
Engineered ECM 1.5 mg/ml 3 mg/ml
1.0 mg/ml Type I + 0.75 mg/ml
Formulation Type I Type I
Type III
3D ECM Microstructure-Mechanical Properties
Fibril Area
++ +++ +++
Fraction (Density)
Fibril Diameter ++ ++ ++
Stiffness -H- +++
Cellular Response:
ECM Contraction/Morphology/Proliferation
ECM Contraction +++ ++ ++/+++
Distance +++ +++ ++
Number of
++ +++
projections
Length ofMedium- Medium-
Medium-Long Short
_ projections Long Long
Long-
Morphology Long-Spindle Stellate Short-Spindle
Spindle
Cytoskeletal Stress-
Stress-fibers Diffuse Stress-fibers
Actin fibers
Proliferation ++ ++ +++
EXAMPLE 30
Recent studies have demonstrated that human adipose-derived stem
cells (ASC) derived from adult human adipose tissue secrete bioactive levels
of
multiple angiogenic and antiapoptotic growth factors including granulocyte-
macrophage colony stimulating factor (GM-CSF), VEGF, hepatocyte growth factor
(HGF), bFGF, and transfolining growth factor-I3 (TGF-f3), and are able to
enhance
blood flow and minimize death of ischemic muscle tissue [Rehman et al., 2004,
Circulation 109: 1292-8]. These results are important because they indicate
that
autologous delivery of ASC, which are readily available from liposuction under
local
anesthesia, may be a novel and uniquely feasible therapeutic option to enhance
angiogenesis and tissue rescue in ischemia. However, quantitative analysis of
cell
delivery has documented that the majority of peripheral blood mononuclear
cells or
ASC injected via intramyocardial, intracoronary, and interstitial retrograde
coronary
venous (IRV) in an ischemic swine model are not retained in the heart
immediately
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following delivery and that the processes of delivery were highly
inconsistent. In
addition, examination of ASC surviving 1 week following intramuscular
injection
showed reduction of cell numbers to 25% of the injected cells over this
period,
suggesting limited cell survival [Rehman et al., 2004, Circulation 109: 1292-
8]; this is
further corroborated in the myocardial system by survival of approximately 20%
or
less of initially retained mesenchymal stem cells over 4 weeks post-injection.
The survival, proliferation, and differentiation properties of human
APC and EPC cells implanted within three dimensional matrices will be
investigated
using both standard cell culture media or by suspension in any of the
formulations of
"ready-to-assemble" components of self-assembling 3D matrix microenvironments,
in
which the microstructure, composition, and mechanical properties are
quantified and
systematically varied. The delivery efficiency and subsequent engraftment
(cell
survival and differentiation) of human ASC or endothelial progenitor cells
derived
from human cord blood (EPC) implanted within an animal model of hindlimb
muscle
ischemia will also be investigated. More particularly, cells will be delivered
with or
without injectable 3D matrix microenvironments in which the "instructive" or
signaling properties are controlled and systematically varied.
Methods:
A series of in vitro experiments will be conducted to determine the
effect of specific biophysical features of a cell's 3D ECM microenvironment on
the
fundamental behavior of human adipose-derived stem cells (ASC) and highly
proliferative endothelial progenitor cells derived from human cord blood
(EPC). ASC
will be harvested from human adipose tissue as described previously [Rehman et
al.,
2004, Circulation 109: 1292-81. Cultures of endothelial progenitor cells will
be
obtained from umbilical veins using established procedures [Ingram et al.,
2004,
Blood 104:2752-2760]. 3D ECM microenvironments in which specific biophysical
features including fibril density, length, and width and stiffness are
systematically
varied will be created from purified collagen as described previously [Pizzo
et al.,
2005, J. App. Phys. 98:1909-1921; Roeder et al., 2002 J. Biomech Eng. 124: 214-
22213,17]. In addition, 3D microenvironments in which composition is
systematically varied by including ECM molecules such as type III collagen,
hyaluronic acid, VEGF, bFGF will also be investigated. These molecules were
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chosen based upon their known role in neovascularization and cardiac muscle
development. In all cases, cells will be added as the last component of the
solubilized
collagen matrix and the suspension will be injected over 30 seconds through a
25Ga
needle (paralleling intramuscular injection for in vivo systems) into a well
plate and
polymerized at 37 C. Immediately following polymerization (less than 30
minutes),
complete medium will be added to all constructs.
For these studies, cell seeding densities ranging between 1x105 to
lx107 cells/ml will be evaluated. Fundamental cell behaviors including
survival,
morphology, proliferation, and differentiation will be determined using
techniques
established previously [Pizzo et al., 2005, J. App. Phys. 98:1909-1921]. In
some
cases, cells will be prelabeled with CellTracker dyes or transfected with GFP
and
analyzed in 3- or 4-dimensions using confocal microscopy in a combination
reflection-fluorescence mode. Outcomes will be compared to those from control
"deliveries" in which cells are injected into media within culture plates,
parallel to the
situation for cell injection into a tissue environment in the absence of a
solubilized,
self-assembling matrix.
In addition to the in vitro culturing of cells within the 3D
microenvirornnents, the 3D cell containing matrices will be implanted via
injection
into either noinial or ischemic muscle in vivo, using the hindlimb model of
muscle
ischemia that the March lab has established and published in the preliminary
findings
concerning adipose stem cells [Rehman et al., 2004, Circulation 109: 1292-8].
Briefly, nude mice are employed so that cells of human origin can be studied
in the
absence of xenogeneic barriers. The ilio-femoral artery is surgically ligated
and
excised as described previously, in the left hindlimb only. The right hindlimb
thus
serves as a non-ischemic control. The musculature of the distal legs (e.g.,
tibialis
anterior) then can be used as a well-demarcated delivery site for 1000
injections into
normal (right) and ischemic (left) muscle, that are perfothied under direct
visualization. Injections of precisely defined numbers of ASC or EPC will be
conducted 1 day following the surgical induction of ischemia in mice, with
groups of
5 animals for each condition to be evaluated. The conditions will include
control
injections in saline (the previous standard) or in soluble self-assembling
matrices.
The cells will be labeled with GFP to permit enumeration by subsequent flow
cytometry following muscle dissociation, as well as microscopic evaluation of
the
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anatomy of engraftment and differentiation in selected mice. Mice injected
will be
sacrificed at either 3 hours post-injection, to quantify the number of cells
retained
acutely following delivery; and at 2 weeks post-injection, to determine
precisely the
cell survival over time following the injection. Cells will be counted by flow
cytometry with the addition of fluorescent particles to penult precise
volumetric
enumeration. A total of 60 mice will be used in this study (e.g., 2 cell types
x 3
ECMs x 5 animals/group x 2 timepoints). The key endpoints will be quantitation
of
cell retention, and subsequent survival and engraftment into muscle or
vasculature in
the noinial and ischemic muscles.
EXAMPLE 31
Effect of Hyaluronic Acid Content in 3D Matrices on Cell Behavior
MATERIALS AND METHODS
Cell culture.
Low passage neonatal human dennal fibroblasts (NHDFs), growth
media, and passing solutions were obtained from Cambrex Bioproducts
(Walkersville,
MD). NHDF were propagated in fibroblast basal medium supplemented with human
recombinant fibroblast growth factor, insulin, gentamicin, amphotericin B, and
FBS
according to manufacturer's recommendation. Cells were maintained in a
humidified
atmosphere of 5% CO2 at 37 C and cell passage numbers representing 15 or less
were
used for all experiments.
Engineered 3D tissue constructs.
To investigate the effect of hyaluronic acid (HA) on ECM assembly
and signaling, type I collagen matrices with varied HA concentrations were
prepared.
Native (acid solubilized) type I collagen prepared from calf skin (Sigma) and
hyaluronic acid prepared from bovine vitreous humor (Sigma) were each
dissolved in
0.01 N hydrochloric acid (HC1) to achieve desired concentrations. Dissolved
collagen
was sterilized by exposure to chloroform overnight at 4 C. Three-dimensional
engineered ECMs were prepared similar to those described in Example 13 at a
constant collagen type I concentration (2 mg/ml) and hyaluronic acid
concentrations
of between 0 and 1.0 ing/ml. The polymerization buffer consisted of 10X
phosphate
buffered saline (PBS) with an ionic strength of 0.14 M and a pH of 7.4. All 3D
engineered ECMs and tissue constructs were polymerized in vitro within a
humidified
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environment at 37 C. To determine the cellular signaling capacity of each 3D
microenvironment, 3D tissue constructs were formed by first harvesting NHDFs
in
complete media and then adding the cells (5 x 104 cells/ml) as the last
component to
the collagen solutions prior to polymerization. Immediately following
polymerization
complete media was added and the constructs were maintained in a humidified
atmosphere of 5% CO2 in air at 37 C.
Qualitative and Quantitative Analysis of 3D ECM Microstructure
Two quantitative parameters describing the 3D fibril microstructural
composition of the ECM, fibril area fraction (a 2D approximation of 3D fibril
density)
and fibril diameter, were determined based upon confocal reflection and
scanning
electron microscopy (SEM) images. Prior to microstructural analysis,
engineered 3D
ECM constructs were polymerized within four-well Lab-Tek coverglass chambers
(Nalge Nunc International, Rochester, NY) and placed within a humidified
environment at 37 C where they were maintained for approximately 15 hours. For
measurements of fibril area fraction, the confocal microscope was used to
obtain high
resolution, 3D, reflection images of the component collagen fibrils within
each ECM.
Three images (at least 10 pm in thickness) were taken at random locations
within
specimens representing a given 3D ECM composition. The confocal image stacks
were then read into Matlab (The Mathworks, Natick, MA), and 2D projections of
each image were created and a threshold chosen for binarization. Using a built-
in
function in Matlab, the area occupied by collagen fibrils (white pixels) was
calculated, converted to tnn2 based upon the pixel sizes, and normalized to
the total
image area.
Fibril diameter measurements were made by applying Imaris 4.0
(Bitplane Inc., Saint Paul, MN) to both confocal reflection and SEM images of
engineered ECM constructs. For SEM imaging, engineered ECM constructs were
fixed in 3% glutaraldehyde in 0.1M cacodylate at pH 7.4, dehydrated with
ethanol,
and critical point dried. Samples were sputter-coated with gold/palladium
prior to
imaging. Samples were imaged in at least duplicate with a JEOL (Peabody, MA)
JSM-840 SEM. From each image obtained, twenty fibrils were chosen at random (5
fibrils per quadrant). Five lines were drawn perpendicular to the long axis of
each
fibril using the measurement tool in Imaris. The average number of pixels
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representing the fibril diameter was then converted into m based upon the
known
pixel size.
Dynamic Mechanical Testing of 3D Engineered ECMs
Mechanical properties of the engineered ECMs were measured using a
TA Instruments (New Castle, DE) AR-2000 rheometer. Soluble ECM preparations
were adjusted to specific polymerization conditions and placed on the peltier
temperature-controlled lower plate at 22 C, and the 40-mm parallel-plate
geometry
was lowered to a 1-mm gap. The temperature was then raised to 37 C to initiate
polymerization. The peltier heated plate required about 1 minute to stabilize
at 37 C.
Measurements of storage modulus G' and loss modulus G" of the polymerizing
material under controlled-strain oscillatory shear were made every 30 seconds
under
oscillation at 1 Hz and 0.1% strain for a proscribed time. This strain was
sufficiently
small to ensure that it did not affect the kinetics of polymerization. Two
hours and
thirty minutes after polymerization, a shear creep test was conducted with a
shear
stress of 1 Pa for 120 seconds, Creep data was interpreted with a standard
four-
element Voigt spring dashpot model. Next a frequency sweep of controlled-
strain
oscillatory shear was made at 0.1% strain, from 0.01 to 20 Hz. Following the
frequency sweep, a continuous shear stress ramp from 0.1 to 10.0 Pa over 2
minutes
was applied. Finally, the specimen was subjected to unconfined compression at
a rate
of 10 m/sec.
Qualitative and Quantitative Determination of Cell Proliferation
Quantification of NHDF proliferation and its dependency on the 3D
ECM microenvironment involved preparing 3D tissue constructs within 24-well
tissue-culture plates. For comparison purposes, the proliferative capacity of
NHDF
was also determined for an equivalent number of cells seeded directly onto the
surface
of tissue culture plastic. At timepoints representing 24 and 48 hours after
construct
polymerization and/or cell seeding, each well and tissue construct was
examined
microscopically to observe the viability, number, and morphology of the cells.
The
medium from each well then was replaced with fresh medium containing the
metabolic indicator dye alamarBlue (10% v/v; BioSource International, Inc.,
Camarillo, CA). Approximately 18 hours later, dye reduction was monitored
spectrofluorotnetrically using a FluoroCount Microplate Fluorometer (Packard
Instruments, Meriden, CT) with excitation and emission wavelengths of 560 nm
and
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590 nm, respectively. Background fluorescence measurements were determined
from
wells containing only dye reagent in culture medium. Maximum levels of
relative
fluorescence were determined from alamarBlue solutions that were autoclaved to
induce complete dye reduction. The mean and the standard deviation values for
all
fluorescence measurements were calculated and subsequently normalized with
respect
to the background and maximum fluorescence readings.
Time-Lapse Imaging of Cell-ECM Interactions
Tissue constructs representing NHDFs seeded at 5 x 104 cells/ml
within 3D engineered ECMs with defined microstructural and biochemical
compositions were evaluated using time-lapse confocal microscopy. Beginning 1
hour after polymerization, 2 to 3 cells were repeatedly monitored using the
confocal
microscope in a reflection (back-scattered light) mode to obtain image stacks
of the
individual cell and its surrounding matrix as described previously (Voytik-
Harbin, et
al., Microscopy and Microanalysis, 9:74-85, 2003). Images were collected at 30-
minute intervals and a z-step of 0.5 mm to minimize exposure of the tissue
constructs
to radiation from the argon laser.
Determination of Volumetric Strain
Consecutive confocal reflection images representing temporal
deformation induced by a resident cell on its surrounding ECM microstructure
provided the basis for the quantification of local displacements and strains
in 3D.
Within each image, subvolumes of 32 x 32 x 20 pixels in the x, y, and z
directions,
respectively, were established. Each subvolume represented a group of voxels
centered around a given point at which displacement values were sought. Each
image
subvolume provided a unique 3D voxel intensity pattern that allowed
correlation
pattern matching between consecutive images using an incremental digital
volume
correlation (IDVC) algorithm developed previously by our laboratory (Roeder et
al.,
J. Biomech. Eng. vol. 124, pp. 214-222 (2002)). The IDVC algorithm provided
strain-state data, including principal strains and their associated
directions, for all grid
point locations. Grid points were established in 512 ' 512¨pixel images that
were 32
pixels apart in both x- and y-directions, with 24-pixel spacing in the z-
direction.
Principal strains determined for the length (E L) , width (Ew), and height
(EH) directions
were used to calculate volumetric strain (Er) based upon the following
formula:
Er = EH Ew + EL + (Ew = EH)-1-(EL = EH)-1-(EL = Ew)d-(EL = Ew = EH).
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Determination of 3D Cell Morphology
Prior to imaging at either 6 or 12 hours after construct polymerization,
tissue constructs were stained with the vital dye Cell Tracker Green
(Molecular
Probes, Eugene, OR) to facilitate discrimination of the cell from the
surrounding
collagen ECM. Confocal image stacks were then collected in a combined
reflection-
epifluorescence mode for determination of cell morphology and fibril
microstructural
organization.
Results
Fibril diameter distribution was measured, as determined from
scanning electron microscopy images, for engineered matrices prepared from
type I
collagen in the presence of varied amounts of hyaluronic acid. Over the range
of
hyaluronic acid concentrations tested, no significant difference was observed
in mean
fibril diameter. Mean fibril diameter measurements were 80.8 18.3 [Am, 72.2
13.0
pm, and 72.1 11.8 lam ( standard deviation) for engineered matrices
prepared from
2 mg/ml type I collagen containing 0, 0.5 mg/ml, and 1.0 mg/ml hyaluronic
acid,
respectively. Interestingly, it did appear that the variation (standard
deviation) of
fibril diameter measurement decreased with increasing hyaluronic acid content.
No
observable or quantitative differences in fibril area fraction measurements
were
determined for the engineered matrices prepared with and without hyaluronic
acid.
While hyaluronic acid did not dramatically effect the fibril
microstructure of engineered matrices, the polymerization rate was found to
decrease
with increasing hyaluronic acid content as indicated by a decreased slope of
the G'
versus time plot. Furthermore, as the hyaluronic acid content increased, the
engineered matrices showed an increase in compliance and an increase in their
compressive stiffness, respectively. These results demonstrate that the 3D
fibril
microstructure as well as the viscous fluid component provide critical
determinants of
the overall mechanical properties of the engineered ECMs.
Studies comparing the cell response to 3D ECM microenvironments
prepared with various hyaluronic acid content have shown no significant
difference in
the proliferation properties of neonatal human dermal fibroblasts. However,
analysis
of cell morphology and mtrix contraction (remodeling) by cells indicate that
hyaluronic acid alters the mechanics of cell-ECM interactions. Analyses of the
magnitude and spatial distribution of local, 3D strain induced by a resident
cell
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within an engineered matrix microenvironment revealed that the addition of
hyaluronic acid reduces the ability of fibroblasts to effectively contract and
induce
alignment of surrounding collagen fibrils. In other words, the extent of
fibril
deformation and realignment (remodeling) by cells is decreased and more
uniformly
distributed around the cell in the presence of increased concentrations of
hyaluronic
acid.
Results of these studies show that while the addition of hyaluronic acid
does not dramatically affect the 3D fibril microstructure of the resultant
engineered
matrices, it does affect the mechanical properties, likely by changing the
properties of
the viscous fluid component. Furthermore, systematic variation of the viscous
fluid
component as a specific design criteria for 3D engineered matrices does affect
the
mechanisms by which resident cells mechanically manipulate (contract) or
remodel
their ECM microenvironment.
