U.S. patent application number 09/770620 was filed with the patent office on 2001-08-23 for tissue implant.
Invention is credited to Bader, Daniel L., Lee, David A., Stephens, Myra D..
Application Number | 20010016772 09/770620 |
Document ID | / |
Family ID | 10808861 |
Filed Date | 2001-08-23 |
United States Patent
Application |
20010016772 |
Kind Code |
A1 |
Lee, David A. ; et
al. |
August 23, 2001 |
Tissue implant
Abstract
A biochemical implant is described which comprises at least two
matrix components. The first matrix component is composed of
collagen with a porous macrostructure with the ability to withstand
tensile or shear forces. The second matrix component is a hydrated
alginate gel which substantially fills the porous macrostructure of
the first component and exerts a swelling pressure. The implant
additionally comprises a population of cells comprising
chondrocytes, fibrochondrocytes, fibroblasts or osteoblasts, or
precursors thereof.
Inventors: |
Lee, David A.; (St. Albans,
GB) ; Bader, Daniel L.; (London, GB) ;
Stephens, Myra D.; (London, GB) |
Correspondence
Address: |
STERNE, KESSLER, GOLDSTEIN & FOX PLLC
1100 NEW YORK AVENUE, N.W., SUITE 600
WASHINGTON
DC
20005-3934
US
|
Family ID: |
10808861 |
Appl. No.: |
09/770620 |
Filed: |
January 29, 2001 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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09770620 |
Jan 29, 2001 |
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09188165 |
Nov 9, 1998 |
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09188165 |
Nov 9, 1998 |
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PCT/GB98/00673 |
Mar 6, 1998 |
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Current U.S.
Class: |
623/14.12 ;
623/17.16; 623/23.75 |
Current CPC
Class: |
A61L 27/48 20130101;
A61L 27/3856 20130101; A61L 27/3843 20130101; A61L 2430/06
20130101; A61L 27/3817 20130101; A61L 27/56 20130101; A61L 27/3852
20130101; A61L 27/58 20130101; A61L 27/3804 20130101; A61L 27/48
20130101; A61L 27/56 20130101; A61L 27/58 20130101; A61F 2/30
20130101; C08L 89/06 20130101; C08L 5/08 20130101; C12N 5/00
20130101; A61L 27/48 20130101; C08L 5/08 20130101 |
Class at
Publication: |
623/14.12 ;
623/17.16; 623/23.75 |
International
Class: |
A61F 002/38; A61F
002/44 |
Foreign Application Data
Date |
Code |
Application Number |
Mar 7, 1997 |
GB |
9704749.2 |
Claims
1. A biomechanical implant comprising at least two matrix
components, the first matrix component having a porous
macrostructure with the ability to withstand tensile or shear
forces, the second matrix component being a hydrated gel which
substantially fills the porous macrostructure of the first
component and exerts a swelling pressure, the implant additionally
comprising a population of animal cells having a differentiated
biomechanical phenotype or having an undifferentiated phenotype
with the potential to differentiate into such cells.
2. A biomechanical implant as claimed in claim 1, in which the
implant is a full or partial replacement for a tissue where there
is a mechanical interaction in both the loaded and unloaded states
sufficient to set up an equilibrium between major matrix
components.
3. A biomechanical implant as claimed in claim 2, in which the
tissue comprises cartilage, temperomandibular joint meniscus, knee
meniscus or an intervertebral disc.
4. A biomechanical implant as claimed in any one of claims 1 to 3,
in which the first matrix component is composed of a biodegradable
or bioresorbable polymer.
5. A biomechanical implant as claimed in claim 4, in which the
biodegradable or bioresorbable polymer comprises collagen, gelatin,
poly-L-lactic acid, poly-glycolic acid, polycaprolactone,
polyhydroxybutarate, polyanhydrides or a copolymer thereof
6. A biomechanical implant as claimed in claim 4 or claim 5, in
which the biodegradable or bioresorbable polymer is modified by
crosslinking agents.
7. A biomechanical implant as claimed in any one of claims 4 to 6,
in which the first matrix component further comprises a coating of
hydroxyapatite, fibronectin, collagen, vitronectin, hyaluronan,
proteoglycans, glycosaminoglycans or RGD (Arg-Gly-Asp).
8. A biomechanical implant as claimed in any one of claims 4 to 7,
in which the coating is restricted to specified portions of the
implant.
9. A biomechanical implant as claimed in any one of claims 4 to 8,
in which the first matrix component has a groove or microporosities
in the range 0.1-50 .mu.m, suitably 1-25 .mu.m, preferably 3-10
.mu.m.
10. A biomechanical implant as claimed in any one of claims 4 to 9,
in which the polymeric components forming the first matrix
component have an average pore size ranging from 10-1000 .mu.m,
suitably 30-250 .mu.m, preferably 50-150 .mu.m with a total
porosity ranging from 50-98%, suitably 60-95%, preferably
70-90%.
11. A biomechanical implant as claimed in any one of claims 1 to 10
which is in the form of woven or non-woven fibres, sponges, foams,
or a series of plates or sheets joined by interconnecting strands,
fibres or adjoining plates.
12. A biomechanical implant as claimed in any one of claims 1 to
11, in which the hydrated gel which forms the second matrix
component is a hydrated polymeric gel.
13. A biomechanical implant as claimed in any one of claims 1 to
12, in which the hydrated polymeric gel is biodegradable or
bioresorbable.
14. A biomechanical implant as claimed in claim 12 or claim 13, in
which the hydrated polymeric gel comprises alginate, agarose,
carrageenans, glycosaminoglycans, proteoglycans, polyethylene oxide
or collagen monomers.
15. A biomechanical implant as claimed in any one of claims 12 to
14, in which the hydrated gel is heterogeneous in nature.
16. A biomechanical implant as claimed in any one of claims 12 to
15, in which the hydrated gel is cross-linked.
17. A biomechanical implant as claimed in any one of claims 12 to
16, in which the hydrated gel further comprises collagen,
hydroxyapatite, growth factors, cytokines, cell attachment factors,
chemotactic factors, angiogenic factors or enzymes.
18. A biomechanical implant as claimed in any one of claims 1 to
17, in which the animal cells in the implant are autologous,
allogeneic or xenogeneic with respect to the host.
19. A biomechanical implant as claimed in claim 18, in which the
cells comprise chondrocytes, fibrochondrocytes, fibroblasts or
osteoblasts, or sub-populations thereof, or precursors thereof.
20. A biomechanical implant as claimed in claim 18 or claim 19, in
which the cells are derived from a tissue biopsy.
21. A method of repairing damaged connective tissue, in which the
method comprises the step of wholly or partially replacing the
damaged tissue with a biomechanical implant as defined in any of
claims 1 to 20.
22. The use of at least two matrix components, the first matrix
component having a porous macrostructure with the ability to
withstand tensile or shear forces, the second matrix component
being a hydrated gel which substantially fills the porous
macrostructure of the first component and has the ability to
withstand compressive loading, and a population of animal cells
having a differentiated biomechanical phenotype or having an
undifferentiated phenotype with the potential to differentiate into
such cells, in the preparation of a biomechanical implant for
repair of damaged connective tissue.
23. The use as claimed in claim 22, in which the tissue to be
repaired is cartilage, temperomandibular joint menisci, knee
menisci or intervertebral discs.
24. A biomechanical implant comprising at least two matrix
components the first matrix component having a porous
macrostructure with the ability to withstand tensile or shear
forces, the second matrix component being a hydrated gel which
substantially fills the porous macrostructure of the first
component and has the ability to withstand compressive loading, the
implant additionally comprising a population of animal cells having
a differentiated biomechanical phenotype or having an
undifferentiated phenotype with the potential to differentiate into
such cells, as a combined preparation for simultaneous, separate or
sequential use in the preparation of a biomechanical implant for
repair of damaged connective tissue.
25. A method for the manufacture of a biomechanical implant as
defined in any one of claims 1 to 20, in which the method comprises
the steps of: (a) infiltrating a solution of the uncrosslinked
precursors of the hydrated gel into the first matrix component, (b)
crosslinking the gel in situ to form the second matrix component,
(c) incorporating the cells into the matrix, and (d) applying a
mechanical loading.
26. A method as claimed in claim 25, in which the uncrosslinked
precursors of the hydrated gel are infiltrated into the first
matrix component by diffusion, vacuum infiltration or
centrifugation.
27. A method as claimed in claim 25 or claim 26, in which the cells
are incorporated into the matrix (i) by direct attachment to the
first matrix component, (ii) by addition of the cells to the
solution of the second matrix component prior to infiltration into
the first matrix component, or (iii) by attachment of the cells to
the first matrix component and immobilisation of cells within the
second matrix component.
Description
[0001] The present invention relates to a tissue implant for use in
damaged load-bearing cartilaginous tissue, such as the meniscus and
articular cartilage.
[0002] Many non-calcified skeletal connective tissues such as
articular cartilage, menisci, ligaments, tendons and intervertebral
disc have a mechanical function and, as such, are subjected to
dynamic mechanical loading during physiological activity. For
example, joint forces are considerable during normal physiological
activities and commonly exceed 2000 N or three times body weight
and occur repetitively approximately 1 million times per year. Thus
articular cartilage will be commonly subjected to contact stresses
exceeding 5 MPa.
[0003] The cartilaginous tissues are composed of cells embedded
within an extensive extracellular matrix. The functional behaviour
of these tissues is best understood when it is considered as
biphasic consisting of a fluid phase of interstitial water
containing dissolved inorganic salts and a solid phase containing
the collagen-proteoglycan organic solid matrix. The two phases
together permit fluid flow through the permeable solid phase. It is
the physicochemical interaction of the various components of the
extracellular matrix which is responsible for the mechanical
properties of the healthy tissues. For example in articular
cartilage, there is a physico-chemical equilibrium between the
osmotic swelling pressure (P.sub.swelling) of the proteoglycan gel
which is balanced by the hydrostatic pressures (P.sub.elastic) due
to the tensile stresses generated within the collagen fibre
network. This balance exists even in unloaded articular cartilage.
It is altered when the tissue is loaded in compression by an
applied hydrostatic pressure (P.sub.applied) resulting in a net
pressure differential (.DELTA.p) and fluid flow away from the
compressed tissue. The appropriate equation is given by
.DELTA.p=P.sub.applied+(P.sub.elastic-P.sub.swelling)
[0004] This fluid flow will result in an increased proteoglycan
concentration within the tissue and a change in the relative
magnitudes of the stresses in the two solid components of articular
cartilage. If the compressive load remains constant the rate of
fluid flow decreases with time and eventually reduces to zero at a
new state of equilibrium. This timne-dependent creep behaviour is
characteristic of all viscoelastic soft tissues.
[0005] The cells, although occupying less than 10% of the tissue
volume, are necessary for the synthesis and maintenance of matrix
levels and are, therefore, crucial to the structural integrity and
function of the tissue. It is known that the cells, whether
chondrocytes in cartilage, fibrochondrocytes in menisci, or
tenocytes and fibroblasts in tendon and ligament respectively, are
able to alter their metabolic activity in response to applied
loads. Both the level of strain applied and the dynamic frequency
are known to be important in determining this response. These
processes are believed to be major factors in determining cellular
activity in these tissues.
[0006] The mechanisms by which cells detect and respond to
mechanical load are termed mechanotransduction pathways and are
complex and poorly understood. mechanotransduction events may be
resolved into extracellular components including cell deformation,
hydrostatic pressures and streaming potentials, followed by
intracellular signalling events such as intracellular calcium
fluxes, cAMP production and cytoskeletal alterations which finally
lead to altered effector cell response. However, understanding such
mechanisms has been complicated by the need to investigate the
processes at a number of different levels, including the cellular
level.
[0007] Injuries to soft tissues are extremely common in hospital
clinics. Indeed, soft tissue replacements amount to an estimated
35% of the world market for all medical devices (Materials
Technology Foresight in Biomaterials, Institute of Materials,
London (1995).
[0008] In the case of articular cartilage and knee menisci,
traumatic damage is common in young active people. Natural repair
is often poor due to limited vascularity. If untreated, damage
commonly leads to progressive degenerative changes, such as
osteoarthritis, in the injured tissue and associated structures.
Current solutions, include the use of artificial joints but these
implants have a relatively short lifespan of 12-15 years with
subsequent replacements lasting for shorter periods. This is a
special problem for patients who have received an implant early in
life.
[0009] There have been many options proposed for the repair of soft
tissues. These generally involve synthetic materials, biological
materials or a combination of the two. The former solutions have
the advantages of providing a structure which is immunologically
acceptable and with the mechanical integrity required of load
bearing structures. However their instability in the body leads to
relatively poor long term performance. Biological solutions
traditionally involve autografts, allografts or xenografts,
depending on their source of tissues. Each of these options has
proved to be far from ideal with, for example, autografts leading
to donor site morbidity and allografts and xenografts leading to
graft rejection.
[0010] Other common solutions involve augmentation devices
incorporating both synthetic structures and biological grafts.
These devices depend on tissue ingrowth and regeneration induced by
the successful transfer of stress from the synthetic material to
the natural tissues. This stress transfer process would need to
change with time after implantation as the tissue regenerates.
Clinical reports are not generally convincing due to, for example,
the inadequate initial performance of the synthetic component of
the augmentation device.
[0011] Commonly used procedures for treatment of articular
cartilage lesions include the Pridie technique which involves
drilling or abrasion of the joint surface to release repair cells
which form a fibrocartilaginous repair tissue. Other procedures
involve the use of carbon fibre rods or mats associated with
drilling and the use of allografts and xenografts. These techniques
lead to the formation of soft fibrocartilaginous tissue which has
limited long-term stability. Options for the treatment of damaged
menisci currently include, surgery to remove the damaged portion of
meniscus (20 years ago, before the meniscus was better understood,
damaged menisci were removed completely), implantation of a plastic
meniscus or using fibrin-glue to glue back the torn portion of
meniscus. However, in plastic meniscal implants the use of
hydrogels is often inadequate and there are problems with
shear/stress forces on the implant. The use of fibrin glue is also
unsatisfactory because the meniscus still contains a point of
weakness and is a non-homogeneous anisotropic structure.
[0012] The relative failure of many surgical, synthetic and graft
solutions has led to the growing interest in the development of
cell-seeded repair systems for solving a number of clinical
problems related to connective tissues such as articular cartilage,
menisci and ligaments. These systems have also been called
tissue-engineered repair systems. Typically autologous or allogenic
cells are isolated from a tissue biopsy removed from a site remote
from the injury. The cells are expanded in cell culture and seeded
in a suitable 3D resorbable scaffold material, which when implanted
into the defective or damaged site elicit a biological repair.
[0013] There are some reports describing in vitro and in vivo
evaluation of cell-engineered systems for repair of load-bearing
cartilaginous tissues. Examples of such systems have been described
by Brittberg et al (New England Journal of Medicine 331 889-95
(1994)), WO 89/00413, WO 90/09769, WO 91116867, WO 90/12603, WO
95/31157 and Paige et al (Plastic and Reconstructive Surgery 97(1)
168-180 (1996)).
[0014] Brittberg et al describe a procedure in which autologous
chondrocytes, expanded in culture, are transferred into the defect.
This procedure does not involve a scaffold material and thus a
periosteal graft is required, with associated donor site morbidity,
to retain the cells within the defect. The implanted cells are
denuded of matrix and therefore have no mechanical integrity on
implantation. Also the technique is only suitable for lesions which
do not include the sub-chondral bone.
[0015] Stone (WO 89/00413, WO 90/09769, WO 91/16867) describe
prostheses for the intervertebral disc, meniscus or other similar
tissue. Each prosthesis is said to include a dry, porous, volume
matrix of biocompatible and bioresorbable fibres. These are
described as being interspersed with glycosaminoglycan molecules,
which may provide attachment sites for cross-links to the fibres.
The shape of such devices can either be manufactured to replicate
the outer surface contour of the biological structure which it is
designed to replace, or can be manufactured to a larger-shape and
trimmed down to size. For long term in vivo performance, these
prostheses require the infiltration of functional fibrochondrocytes
to provide a scaffold for the regenerating tissue structure.
However, this prosthesis is not cell seeded and thus does not
depend upon the use of seeded cells into the scaffold prior to
implantation. Additionally, the nature of the mechanical
interaction has not been specified.
[0016] The choice of biomaterials has to date been largely
empirical based on biocompatibility and the maintenance of cell
morphology and function. Additionally, the structural design of
devices has been generally ignored. Materials such as collagen in
various forms, poly-1-lactic acid and alginates have been employed.
Most of the post implantation analysis has involved histological
and biochemical analysis of repair tissue, with only a few reports
assessing its mechanical integrity. The mechanical integrity of the
device pre-implantation has been largely ignored. For example,
Vacanti (WO 90/12603) and Kim et al (Plastic and Reconstructive
Surgery 94(2) 233-237 (1994)) describe a method in which
chondrocytes are seeded onto a bioresorbable polymer scaffold for
transfer into a cartilaginous defect. Whilst the constructs may
have a defined shape and size, the mechanical properties do not
approach that of articular cartilage. The compressive modulus of
articular cartilage is approximately 5-10 MPa. Problems associated
with ensuring an initial even cell density and maintenance of
chondrocytes phenotype are not addressed. The application of a
coating to the scaffold was proposed but only to induce cell
attachment. The method relies upon the synthesis of sufficient
cartilaginous matrix to fill voids within the scaffold before any
significant mechanical integrity is achieved.
[0017] Paige et al (Plastic and Reconstructive Surgery 97(1)
168-180 (1996)) propose the use of cells embedded in alginate
alone. However, these devices are approximately one hundred times
less stiff than cartilage raising questions about the mechanical
functionality of the device.
[0018] WO 95/31157 relates to an anatomically specific
bioresorbable device for healing of voids in soft tissues, such as
in articular cartilage and the associated subchondral bone. The
device is described as being an internal three-dimensional
structure in fibrous form, termed the macrostructure, with voids
which are partially or totally filled with a polymer gel, forming a
microstructure. The gel provides a carrier material for selected
chondrocytes and in conjunction with chemical mediators into the
cartilage region of the device to enhance tissue regeneration. The
gel phase is designed to be resorbed after 72 hours in vivo. This
prosthesis, therefore, only depends upon the seeded microstructure
to deliver the cells and to enable the short term transfer of cells
on to the fibrous scaffold. The design does not depend upon the
total filling of the voids by the microstructure. There is no
consideration or mention of the physicochemical interaction of the
macrostructure and the microstructure or the resulting mechanical
integrity of the prosthesis at any time following implantation.
[0019] None of the aforementioned methods will produce a construct
with a mechanical function which resembles that possessed by the
tissues to be repaired. In addition none of the aforementioned
methods address the transfer of mechanical load to cells within the
device leading to mechanotransduction-induced desired cell
response. It has now been discovered that an approach which
considers both the mechanical and biological aspects of implants
for repair of damaged load-bearing connective tissue can provide a
superior device for implantation.
[0020] According to a first aspect of the present invention there
is provided a biomechanical implant comprising at least two matrix
components, the first matrix component having a porous
macrostructure with the ability to withstand tensile or shear
forces, the second matrix component being a hydrated gel which
substantially fills the porous macrostructure of the first
component and exerts a swelling pressure. The implant additionally
comprises a population of animal cells having a differentiated
phenotype or having an undifferentiated phenotype with the
potential to differentiate into such cells.
[0021] The biomechanical implants according to the present
invention are advantageous in that they provide a composite
construct with mechanical properties under compressive loading
which are greater than the sum of the compressive properties of the
two components individually. In this manner, the composite
construct may contribute to the restoration of the biomechanical
function of the damaged tissue immediately post-implantation.
[0022] The implant may be a full or partial replacement for tissues
where there is a mechanical interaction in both the loaded and
unloaded states sufficient to set up an equilibrium between major
matrix components, such as those with a mechanical function, where
compression is the predominant type of loading. Suitable tissues
for repair can include but are not limited to tissues such as
cartilage, temperomandibular joint menisci, knee menisci and
intervertebral discs.
[0023] The first matrix component may suitably be composed of a
biodegradable or bioresorbable polymer such as collagen, gelatin,
poly-L-lactic acid, poly-glycolic acid, polycaprolactone,
polyhydroxybutarate, polyanhydrides and copolymers of the above.
However, the specific use of non-resorbable polymeric components,
or of non-polymeric resorbable components such as soluble
bioglasses is not precluded. The first matrix component may
preferably be composed of collagen.
[0024] The mechanical properties, antigenicity and degradation rate
of the polymers comprising the first matrix component can be
modified by the use of crosslinking agents such as glutaraldehyde,
formaldehyde, tissue transglutaminase and acid aldehyde or by heat
treatment. The polymers forming the first matrix component can be
modified by alteration to the surface properties of the polymer.
This can be achieved by coating with components known to influence
cell attachment and differentiation such as hydroxyapatite,
fibronectin, collagen, vitronectin, hyaluronan, proteoglycans,
glycosaminoglycans or RGD (Arg-Gly-Asp) tripeptide cell attachment
sequence. Coating may be restricted to specified regions of the
first matrix component. Alternatively, attachment can be influenced
by specific surface ultrastructural features such as grooves or
microporosities within the range 0.1-50 .mu.m, suitably 1-25 .mu.m,
preferably 3-10 .mu.m.
[0025] The polymeric components forming the first matrix component
preferably may possess a porous macrostructure consisting of
interconnecting channels. The average size of the pores may range
from 10-1000 .mu.m, suitably 30-250 .mu.m, preferably 50-150 .mu.m
with a total porosity ranging from 50-98%, suitably 60-95%,
preferably 70-90%. The macrostructure can therefore be in the form
of woven or non-woven fibres, sponges or foams, or a series of
plates or sheets joined by interconnecting strands, fibres or
adjoining plates. The macrostructure of the first matrix component
can be non-homogeneous in terms of porosity, pore size and/or
predominant polymeric fibre/plate/sheet orientation.
[0026] In biomechanical implants to replace articular cartilage,
the dimensions of the first matrix component may be approximately
0.05-6 mm, suitably 0.075-4 mm, preferably 0.1-3 mm in thickness,
corresponding to the thickness of articular cartilage. The other
dimensions of the construct suitably may correspond to the
dimensions of the defect to be repaired. The macrostructure
conveniently possesses a surface zone with an orientation
predominantly parallel to the surface. This zone preferably
corresponds to the natural superficial zone of the cartilage,
thereby representing approximately 10-20% of the total thickness of
the construct in the range 0.005-1.2 mm, suitably 0.0075-0.8 mm,
preferably 0.1-0.6 mm. The surface zone also preferably may possess
a relatively low porosity, within the range 50-98%, suitably
60-95%, preferably 65-80% with average pore size between 10-1000
.mu.m, suitably 30-250 .mu.m, preferably 50-100 .mu.m. The
predominant macrostructural organisation within the remainder of
the construct depth may be suitably perpendicular to the surface
with a porosity between 50-98%, suitably 60-95%, preferably 75-90%
and average pore size in the range 10-1000 .mu.m, suitably 30-250
.mu.m, preferably 50-150 .mu.m. The lowermost zone, representing
0.05-1 mm in thickness, may be coated with a bioactive material
such as hydroxyapatite to enhance integration with the underlying
sub-chondral bone.
[0027] In the preferred embodiment for the replacement of the knee
menisci, the overall dimensions of the first matrix component
preferably correspond to the dimensions of the meniscal fragment to
be replaced. The predominant orientation of the polymeric
macrostructure of the first matrix component may be parallel to the
surface of the construct in the outermost 100-300 .mu.m of the
construct corresponding to the superficial zone of the natural
meniscus. The porosity within this region can be between 50-98%,
suitably 60-95%, preferably 65-80% with pore size in the range
10-1000 .mu.m, suitably 30-250 .mu.m, preferably 50-100 .mu.m. The
central core of the polymeric macrostructure of the first matrix
component may be predominantly radial in orientation with minor
linking structures oriented at 90.degree. to the predominant
orientation. The porosity in this region may be between 50-98%,
suitably 60-95%, preferably 75-90% with pore size in the range
10-500 .mu.m, suitably 30-250 .mu.m, preferably 50-150 .mu.m.
[0028] The hydrated gel which forms the second matrix component may
be a hydrated polymeric gel such as the hydrated gels of alginate,
agarose, carrageenans, glycosaminoglycans, proteoglycans,
polyethylene oxide or collagen. In many situations the hydrated gel
may be alginate. Preferably, the hydrated gel is biodegradable or
bioresorbable. Suitably, the second matrix component can withstand
compressive loading.
[0029] Alginate is a polysaccharide, derived from brown seaweeds,
composed of D-mannuronic and L-guluronic acid monosaccharide
subunits. Whilst the sodium salt of alginate forms viscous
solutions, alginate can form hydrated gels in the presence of
divalent cations such as calcium due to cross linking through the
negatively charged carboxyl groups residing on the L-guluronic acid
residues. The viscosity of the uncrosslinked solutions and thereby
the mechanical strength of crosslinked gels can be influenced by
altering the average chain length of the alginate or by altering
the proportion of D-mannuronic acid and L-guluronic acid residues
within the polysaccharide. These factors may also influence the
rate of resorption of the alginate.
[0030] The concentration of the hydrated gel forming the second
matrix component may be heterogeneous within the macrostructure
formed by the first matrix component of the implant. This can
provide functionally graded mechanical properties throughout the
depth of the implant.
[0031] The hydrated gel forming the second matrix component can be
modified by the addition of factors, which are immobilised within
the gels after cross-inking or which are contained within the
hydrated fraction of the gel and are free to diffuse. These factors
include collagen, hydroxyapatite, growth factors, cytokines, cell
attachment factors, chemotactic factors, angiogenic factors and
factors which may influence the attachment of the construct to the
existing tissue such as hydroxyapatite, tissue transglutaminase and
matrix degrading enzymes. The degree to which these factors are
immobilised is dependent on the size and charge of the factor and
the concentration and thereby pore size of the alginate gel. A
particular example of this modification is the incorporation of
collagen within the alginate gel in order to increase the overall
collagen content of the composite device. This may be of particular
importance in the preparation of devices for meniscal repair.
[0032] As discussed above, preferably both the first and second
matrix components are resorbable and over a period of time can be
resorbed and cleared in vivo by hydrolysis, enzymic degradation,
depolymerisation and diffusion to be replaced by natural matrix
components synthesised by the cells of the implant. In this manner,
the original construct can be replaced by an entirely natural
repair tissue with functional integrity which, due to the presence
of cells and the appropriate mechanical and biochemical signals,
can reach a stable equilibrium in terms of cell number,
architecture and matrix turnover.
[0033] The animal cells in the implant can be autologous,
allogeneic or xenogeneic with respect to the host, preferably the
cells are autologous. Conveniently, the cells can be chondrocytes,
fibrochondrocytes, fibroblasts, osteoblasts, or sub-populations
thereof, which have a differentiated phenotype. Alternatively,
precursors of the aforementioned cell types may be used which have
the potential to differentiate into such cells.
[0034] The cells can be isolated from a tissue biopsy or from bone
marrow, using methods common to those skilled in the art. If
insufficient cell numbers are available at isolation the cells can
be allowed to proliferate in tissue culture prior to seeding into
the construct. During the proliferation phase the cells may be
cultured as a monolayer on a tissue culture treated substrate and
maintained in tissue culture medium such as Dulbeccos Modified
Eagle's Medium supplemented with, for example, between 1 and 20%
fetal calf serum or autologous human serum. Alternatively the cells
may be cultured in serum free medium supplemented with mitogens on
tissue culture plastic modified by the immobilisation of specific
attachment factors. In another alternative approach, the
differentiated cells may be seeded at a specified seeding density
within alginate beads and cultured in tissue culture medium
supplemented with serum or mitogenic growth factors. The cells may
be isolated by dissolving the beads in a sodium citrate saline
solution followed by collagenase digestion. These cells may be
passaged subsequently by seeding and culture in alginate beads at
the same seeding density as previously described. The cells may be
cultured within a suitable rotating bioreactor.
[0035] The biomechanical implants of the present invention
represent an improvement over the known devices of the prior art
for at least the following reasons.
[0036] The interaction between the first and second matrix
components can provide enhanced mechanical properties which are
greater than the sum of the properties of the two components
individually. In this manner, the implant mimics the physiochemical
interaction within natural cartilaginous tissues.
[0037] The presence of immobilised cells within the second matrix
component ensures a uniform cell density throughout the construct
over an extended period of time. Cell density throughout the first
matrix component can be varied by altering its porosity which
permits heterogeneities in cell density to be achieved.
[0038] Contact inhibition of proliferation typically occurs for
cells cultured as a monolayer and cells seeded onto a single factor
scaffold such as that described in WO 90/12603 therefore reach a
limiting cell number which does not fill the voids within the
scaffold. Whereas the present invention allows for cells to be
maintained in a 3-dimensional state which does not induce contact
inhibition of proliferation to occur until the entire 3-dimensional
volume is filled. In this way, a greater cell density is achieved
or a lower initial seeding density may be used.
[0039] Cell morphology is known to influence cell phenotype and the
type, quantity and organisation of extracellular matrix produced by
the cells. The present invention permits cells to be maintained in
either a rounded or elongated morphology, thereby allowing control
of cell phenotype to be achieved. The meniscus contains two
sub-populations of cells, one which is elongated and fibroblastic
in phenotype, the other rounded and chondrocytic in phenotype. The
phenotypes are interchangeable by the use of appropriate cell
culture conditions. The present invention permits the restoration
of the two sub-populations by seeding cells onto the first matrix
component and within the second matrix component.
[0040] The biomechanical implant of the present invention also
permits the transfer of mechanical load applied to the construct to
the cells seeded within it in a defined manner. The inherent
mechanical integrity of the construct allows mechanical load to be
applied either in vitro or in vivo. Different cellular strain
regimes and mechanotransduction pathways can be induced by seeding
the cells onto the first matrix component or within the second
matrix component.
[0041] Furthermore, it is known that mechanical strain influences
cell behaviour in many load-bearing connective tissue such as
articular cartilage, meniscus, tendon, ligament, intervertebral
disc, bone and skin. In many cases this phenomenon can be crucial
to the maintenance of tissue structure and function. The exact
mechanism by which cells sense and respond to strain, known as
mechanotransduction, is unclear but is believed to include cell
deformation, alterations in fluid pressure, streaming potentials,
fluid flow, alterations in nutrient concentration and transport,
alteration in pH and variations in the osmotic environment. The
preferred embodiments of the present invention may induce
alterations in cell behaviour in response to mechanical loading
through many of the aforementioned mechanotransduction pathways.
Mechanical loading may be applied to the cells within the construct
prior to implantation using specially designed technologies or
applied to the cells within the construct post implantation through
defined exercise regimes or through externally applied regimes such
as continuous passive motion regimes. Mechanical loading regimes,
applied in a static or dynamic manner may take a variety of forms
including, although by no means exclusively, uniaxial compression
or tension or hydrostatic pressure. Static and dynamic peak strain
amplitude may be in the range between 0.5-30%, suitably 1-25%,
preferably 5-20% whilst dynamic frequencies should range from
0.001-10 Hz, suitably 0.01-5 Hz, preferably 0.1-3 Hz.
[0042] The are also several advantages in the application of
mechanical strain which may be transferred to the cells within the
construct prior to implantation.
[0043] Firstly, the use of optimal strain windows of biosynthetic
response can quantitatively increase metabolic parameters known to
influence the success of tissue engineered repair devices such as
cell proliferation, proteoglycan synthesis and collagen
synthesis.
[0044] Secondly, qualitative reorganisation and/or orientation of
newly synthesised matrix components in response to strain may be
induced which can provide anisotropy and, potentially, improved
gross mechanical properties. Orientation of cells along strain
fields may induce long-term production of oriented matrix. Oriented
cell proliferation may also restore the cell architecture, such as
columns which are a feature of articular cartilage and are
important in the structural/functional relationship within the
tissue. The cells in the implant may suitably possess a
pre-implantation strain history and can therefore be adapted to
respond to mechanical strain post-implantation.
[0045] According to a second aspect of the present invention there
is provided a method of repairing damaged connective tissue, in
which the method comprises the step of wholly or partially
replacing the damaged tissue with a biomechanical implant as
defined above.
[0046] This aspect of the present invention also contemplates the
use of at least two matrix components the first matrix component
having a porous macrostructure with the ability to withstand
tensile or shear forces, the second matrix component being a
hydrated gel which substantially fills the porous macrostructure of
the first component and has the ability to withstand compressive
loading, and a population of animal cells having a differentiated
phenotype or having an undifferentiated phenotype with the
potential to differentiate into such cells, in the preparation of a
biomechanical implant for repair of damaged connective tissue.
[0047] The implant may be most applicable to tissues where there is
a mechanical interaction in the loaded and unloaded state
sufficient to set up an equilibrium between major matrix
components, such as those with a mechanical function, where
compression is the predominant type of loading. Suitable tissues
for repair can include but are not limited to tissues such as
cartilage, temperomandibular joint menisci, knee menisci and
intervertebral discs.
[0048] According to a third aspect of the present invention, there
is provided a biomechanical implant comprising at least two matrix
components the first matrix component having a porous
macrostructure with the ability to withstand tensile or shear
forces, the second matrix component being a hydrated gel which
substantially fills the porous macrostructure of the first
component and has the ability to withstand compressive loading, the
implant additionally comprising a population of animal cells having
a differentiated phenotype or having an undifferentiated phenotype
with the potential to differentiate into such cells, as a combined
preparation for simultaneous, separate or sequential use in the
preparation of a biomechanical implant for repair of damaged
connective tissue.
[0049] According to a fourth aspect of the present invention, there
is provided a method for the manufacture of a biomechanical implant
a s defined above, in which the method comprises the steps of:
[0050] (a) infiltrating a solution of the uncrosslinked precursor
of the hydrated gel into the first matrix component,
[0051] (b) crosslinking the gel in situ to form the second matrix
component,
[0052] (c) incorporating the cells into the matrix, and
[0053] (d) applying mechanical loading.
[0054] Where the hydrated gel is alginate, a solution of sodium
alginate at concentrations ranging from 0.5-10% (w/v), suitably,
1-8% (w/v), preferably 2-6% (w/v) may be prepared in water, salt
solution or tissue culture medium. The solution can be infiltrated
into the porosity of the first matrix component by diffusion,
vacuum infiltration or centrifugation in a manner to ensure that
all the pores are filled. Sodium alginate solutions swell,
markedly, when separated from an isotonic saline solution by a
permeable barrier. Typically, a 50% increase in weight occurs
during the first 4 hours of water uptake. In this manner sodium
alginate entrapped within the first matrix component can be
immersed in saline and will exert a physicochemical swelling
pressure which is resisted by the first matrix component thereby
setting up tensile forces within the first matrix component in the
unloaded state. This relies upon the ability of the first matrix
component to resist swelling and not exhibit marked swelling in its
own right. This ability is determined by the structural composition
of the first matrix component.
[0055] The composite construct can then be immersed in a solution
containing a salt of a divalent cation to induce crosslinkage of
the alginate in situ to form a hydrated gel. The most preferred
crosslinking solution contains calcium chloride at concentrations
ranging from 10-500 mM, suitably 25-250 mM, preferably 50-150 mM.
In this manner the alginate gel can be immobilised within the
porosity of the first matrix component.
[0056] The cells can be incorporated into the construct in a number
of ways, which include direct attachment to the first matrix
component, addition of the cells to the solution of the second
matrix component prior to infiltration into the first matrix
component or attachment of the cells to the first matrix component
and immobilisation of cells within the second matrix component to
produce two sub-populations of cells within the construct with
different morphologies.
[0057] Direct attachment of the cells to the first matrix component
can be achieved by the addition of a suspension of cells in medium
to a suitably prepared piece of the first matrix component.
Infiltration of the cell suspension into the porosity of the first
matrix component can draw cells into it which subsequently attach
and spread on the polymeric surface of the first matrix component.
The cells may assume a flattened or elongated morphology.
Subsequent culture in suitable tissue culture medium may cause
proliferation and migration of the cells which can lead to the
formation of a uniform covering of cells over the entire surface of
the first matrix component. At any stage after the cells have
attached to the first matrix component, infiltration and
crosslinking of the second matrix component can be performed.
[0058] Addition of the cells to the solution of the second matrix
component prior to infiltration into the first matrix component
achieves incorporation of the cells into the matrix as the
infiltration draws the cells into the porous structure of the first
matrix component in association with the second matrix component.
The cells can become immobilised within the hydrated gel after
crosslinking of the second matrix component. Thereby the cells may
retain a rounded morphology and remain unattached directly to the
first matrix component.
[0059] Attachment of a proportion of cells to the first matrix
component and immobilisation of cells within the second matrix
component can produce two sub-populations of cells within the
construct with different morphologies. Alternatively, expanded
sub-populations of cells, such as cells isolated from the
superficial and deep zones of articular cartilage, may be seeded in
either or both matrix components within the distinct zones of the
implant. The construct can be implanted directly after cell seeding
or the construct containing cells can be maintained in tissue
culture to allow the synthesis of extracellular matrix components
by the cells and proliferation of the cells prior to implantation.
Defined loading regimes can be applied to the cells within the
construct during the culture phase using specially designed
technologies as described in the examples.
[0060] Preferred features and characteristics of the second and
subsequent aspects are as for the first aspect mutatis
mutandis.
[0061] The invention will now be described by way of example with
reference to the accompanying Examples and drawings which are
provided for the purposes of illustration and neither of which are
to be construed as being limiting on the present invention. In the
following description, reference is made to a number of drawings,
in which:
[0062] FIG. 1 shows a photomicrograph of a haematoxylin and eosin
stained section prepared from a meniscal fibrochondrocyte seeded
collagen gel fixed immediately after casting. Magnification
.times.250.
[0063] FIG. 2 shows a photornicrograph of a haematoxylin and eosin
stained section prepared from a collagen sponge construct seeded
with bovine chondrocytes and cultured for 14 days. Magnification
.times.250.
[0064] FIG. 3 shows a photomicrograph of a haematoxylin and eosin
stained section prepared from a bovine chondrocyte seeded alginate
construct which had been cultured for 7 days. Magnification
.times.250.
[0065] FIG. 4 shows a photomicrograph of a haematoxylin and eosin
stained section prepared from a bovine chondrocyte seeded agarose
construct which had been cultured for 3 days. Magnification
.times.250.
[0066] FIG. 5 shows a photomicrograph of a haematoxylin and eosin
stained section prepared from a collagen gel/alginate composite
construct seeded with meniscal fibrochondrocytes which had been
cultured for 3 days. Magnification .times.500.
[0067] FIG. 6 shows a photomicrograph of a haematoxylin and eosin
stained section prepared from a collagen sponge/alginate composite
construct seeded with chondrocytes which had been cultured for 1
day. Magnification .times.250.
[0068] FIG. 7 shows a photornicrograph of a haematoxylin and eosin
stained section prepared from a collagen sponge/agarose composite
construct seeded with chondrocytes which had been cultured for 7
days. Magnification .times.250.
[0069] FIG. 8a shows a graph representing glycosaminoglycan
synthesis by chondrocytes seeded in agarose constructs and
subjected to 15% gross compressive strain at various frequencies
for 48 hr. The values have been normalised to unstrained control
levels (100%).
[0070] FIG. 8b shows a graph representing [.sup.3H]-thymidine
incorporation by chondrocytes seeded in agarose constructs and
subjected to 15% gross compressive strain at various frequencies
for 48 hr. The values have been normalised to unstrained control
levels (100%).
[0071] FIG. 9 shows a graph representing [.sup.3H]-thymidine
incorporation by chondrocytes isolated from the superficial and
deep zones of cartilage and seeded in agarose constructs and
subjected to 15% gross compressive strain at various frequencies
for 48 hours. The values have been normalised to unstrained control
levels (100%).
EXAMPLE (1)
Cell Isolation and Expansion in Culture
[0072] Chondrocytes were isolated from either bovine cartilage or
human cartilage obtained after amputation due to osteosarcoma.
Meniscal cells were isolated from human knee menisci obtained after
amputation due to osteosarcoma. The cartilage or meniscal tissue
was cut into cubes of no more than 2 mm.sup.3 and sequentially
digested with 1% (w/v) pronase (BDH Ltd. Poole, England) in
Dulbecco's Modified Eagles Medium, 2% HEPES buffer, 1% Glutamine
and 1% pencillin/streptomycin (referred hereafter as DMEM)
containing 20% FCS for 1 h at 37.degree. C. followed by collagenase
Type 1A (Sigma) in DMEM +20% FCS for 2 h at 37.degree. C. The
enzyme concentration (i.u./ml) is optimised for each new batch of
collagenase. Debris was removed from the cell containing medium by
filtration through a nylon sieve (Falcon 70 .mu.m pore size) and
the cells separated by centrifugation at 1000 g for 5 minutes. The
cell pellet was washed .times.2 with DMEM+20% FCS and resuspended
in an appropriate volume of medium. Cell density and viability was
determined by the Trypan Blue exclusion test.
[0073] Where necessary, the cell population will be expanded to the
desired number by repeat passage. Cells will be cultured as a
monolayer in tissue culture flasks in DMEM+10% FCS at 37.degree.
C./5%CO.sub.2 with medium changes every two days. At confluence the
cells will be lifted with 0.2% trypsin/EDTA and divided into three
flasks for further culture.
EXAMPLE (2)
Seeding of Fibrochondrocytes into Collagen Gels
[0074] Human fibrochondrocyte seeded collagen gels were prepared by
mixing 12 ml of double strength Dulbeccos Modified Eagle Medium
(DMEM, Gibco, Paisley, UK), 3 ml of a cell suspension
(1.times.10.sup.6 cell/ml in DMEM), 3 ml foetal calf serum (FCS,
Gibco, Paisley, UK), 3 ml 0.1 M NaOH and 9 ml collagen solution (3
mg/ml collagen type I in 0.5M acetic acid, ICN, Paisley, UK). All
components were maintained at 4.degree. C. during mixing. Aliquots
(1.5 ml) of the solution were added to wells and gelled at
37.degree. C./5%CO.sub.2 for 30 minutes. The gels were maintained
at 37.degree. C./5%CO.sub.2 in DMEM+ 10% FCS+150 .mu.g/ml ascorbate
with medium changes every 2 days. Cell-induced contraction of the
gel occurred during the first 3 days. Histology revealed the
presence of cells with typical fibroblastic morphology evenly
distributed within a disorganised fibrous collagen matrix, as shown
in FIG. 1.
EXAMPLE (3)
Seeding of Chondrocytes onto Collagen Sponges
[0075] Bovine chondrocytes were seeded onto collagen sponges by
absorption. Cylinders of absorbable sponge 3 mm thick and 6 mm in
diameter (Helistat.RTM. Colla-Tec, Inc. USA) were cut. Medium (100
.mu.l) containing 1.5.times.10.sup.7 cells/ml was either added
dropwise to the surface of the dry matrix or the sponge was
immersed in cell containing medium. The sponges containing cells
were cultured in DMEM+20% FCS for up to 24 days with medium changes
every two days. Histology using haematoxylin and eosin staining
revealed the presence of clusters of chondrocytes within the sponge
matrix. The majority of cells exhibited rounded morphology although
a proportion were flattened or fibroblastic in appearance as shown
in FIG. 2. The density was not homogeneous with the greatest
proportion of cells at the periphery of the sponge. Evidence of
synthesis of proteoglycan rich matrix around the cells was
noted.
EXAMPLE (4)
Seeding of Chondrocytes in Alginate
[0076] An equal volume of 4% (w/v) alginate (Kelco International
Ltd. Waterfield, Tadworth, Surrey, UK) in Earle's buffered salt
solution (EBSS) was mixed with 1.6.times.10.sup.7 cells/ml in DMEM
containing 20% FCS by gentle agitation. Cross-linking of the
alginate suspension was effected by dialysis against a solution of
100 mM CaCl.sub.2 in DMEM/10% FCS. The crosslinked gel constructs
were cultured in DMEM+20% FCS+ 150 .mu.g/ml ascorbate for up to 24
days with medium changes every 2 days. After 6 days in culture a
proportion of the chondrocytes within the alginate gel had
undergone proliferation to form small clusters as shown in FIG. 3.
The cells retained their original rounded morphology and
synthesised a `halo` of cartilaginous matrix.
EXAMPLE (5)
Seeding of Chondrocytes into Agarose
[0077] During preparation of a cell suspension containing
2.times.10.sup.7 cells/ml, a 6% suspension of agarose (type VII,
Sigma, Poole, England) in distilled water was melted by autoclaving
and subsequently cooled to 37.degree. C. The chondrocyte suspension
was added to an equal volume of agarose to give a final
concentration of 1.times.10.sup.7 cells/ml in 3% agarose. The
agarose suspension was plated in a tissue culture dish and gelled
at 4.degree. C. for 20 minutes. Cylindrical constructs, 5 mm
diameter.times.5 mm height were formed and cultured in DMEM+20%
FCS+150 .mu.g/ml ascorbate for up to 24 days. Histology using
haematoxylin and Safranin-O staining revealed the presence of an
even distribution of cells with a rounded morphology as shown in
FIG. 4. After 3 days a significant `halo` of cartilaginous matrix
was present around the cells and cell clusters formed by
proliferation were noted after 6 days.
EXAMPLE (6)
Preparation of Two Component Constructs by Diffusion Techniques
[0078] Fibrochondrocyte seeded collagen gels were cast as described
in example 2 and cultured for 24 hr. The gels were placed in a
suspension of 2% (w/v) alginate in DMEM+10% FCS:EBSS (1:1). The
alginate was crosslinked by dialysis against a solution of 100 mM
CaCl.sub.2 in DMEM/10% FCS. The constructs were cultured in
DMEM+10% FCS +150 .mu.g/ml ascorbate for up to 24 days with medium
changes every 2 days. Histology revealed the presence of evenly
distributed cells with a fibroblastic morphology throughout the
composite as shown in FIG. 5. Alginate content was inhomogeneous
with greater concentration at the periphery of the construct.
[0079] Collagen sponges with or without prior culture were
infiltrated with 2.0% (w/v) alginate containing 1.5.times.10.sup.7
chondrocytes/ml in DMEM+20% FCS:EBSS (1:1) or 2% (w/v) agarose type
VII containing 4.times.10.sup.6 cells/ml in DMEM+20% FCS:EBSS (1:1)
by dropwise addition to the surface of the sponge. The alginate was
crosslinked by bathing the construct in a solution of 100 mM
CaCl.sub.2 in DMEM/10% FCS. Agarose containing constructs were
gelled by incubation at 4.degree. C. for 30 minutes. The constructs
were cultured in DMEM+20% FCS+150 .mu.g/ml ascorbate for up to 24
days with medium changes every 2 days. Cells, associated with the
agarose were found evenly distributed within the porous
microstructure of the collagen sponge and exhibited rounded
morphology as shown in FIG. 6. There was evidence of cartilaginous
matrix production around the cells.
EXAMPLE (7)
Preparation of Two Component Constructs by Vacuum Infiltration
[0080] Collagen gels or sponges 6 mm diameter with or without prior
culture of cells were placed in a 96 well sterile vacuum manifold
plate (Multiscreen.RTM. Millipore UK), 100 .mu.l of 2% (w/v)
alginate in DMEM+10% FCS:EBSS (1:1) containing 1.5.times.10.sup.7
cells/ml was layered on top of the collagen gel or sponge. A gentle
vacuum was applied to incorporate the alginate and cells into the
collagen gel or sponge. Crosslinking of the alginate was effected
by bating the construct in a solution of 100 mM CaCl.sub.2 in
DMEM/10% FCS. An even distribution of cells was seen throughout the
porous microstructure of the collagen sponge. The cells were
associated with the alginate which was also evenly distributed and
the cells possessed a rounded morphology as shown in FIG. 7.
EXAMPLE (8)
Preparation of Two Component Constructs by Low Speed
Centrifugation
[0081] Collagen gels or sponges with or without prior culture of
cells were placed in wells of a tissue culture plate. 2% alginate
in Earle's buffered salt solution (EBSS) containing 4.times.
10.sup.6 cells/ml in DMEM and 10% FCS was gently layered on top of
the collagen gel or sponge. The tissue culture plate was
centrifuged at 100 g for 5 minutes to incorporate the alginate and
cell suspension within the collagen gel or sponge. Crosslinking of
the alginate was effected by bathing the construct in a solution of
100 mM CaCl.sub.2 in DMEM/10% FCS.
EXAMPLE (9)
Mechanical Testing of Two Component Constructs and Their
Components
[0082] The mechanical characteristics of the collagen gels and
composites were evaluated. Uniformly sized cores of each material
to be tested were mounted between two large plattens and tested to
failure in compression at a crosshead speed of 1 mm/min using a
universal testing machine (Instron Model 1122). The tangent
modulus, an estimate of the material stiffness, was calculated from
the linear region using the following equation: 1 modulus ( MPa ) =
Ft dt A
[0083] Where F=compressive force; t=Original thickness of material;
dt=change in thickness; A=sample cross-sectional area. Another
parameter, the equilibrium modulus, is calculated from a load
relaxation experiment conducted at a prescribed strain of 20%.
[0084] Composite constructs consisting of 2% Keltone HV alginate
infiltrated by vacuum into collagen sponges had a significantly
greater tangent modulus than either collagen or alginate alone,
indicating physicochemical interaction between the components as
shown in Table 1.
1TABLE 1 Mechanical evaluation of single and two component collagen
and alginate constructs. Tangent Equilibrium Construct Modulus
Modulus Collagen 85 kPa 32 kPa sponge/alginate construct Collagen
sponge <1 kPa <1 kPa construct Alginate 25 kPa 9 kPa
construct.
EXAMPLE (10)
Mechanical Stimulation of Cells in Hydrated Gels
[0085] A specially designed cell straining apparatus (Dartec Ltd,
Stourbridge, UK) described in the literature was used to apply
compressive strain to chondrocytes seeded in agarose as described
in example 6 (Lee, D. A., and Bader, D. L., Journal of Orthopaedic
Research 15 (2) 181-188 (1997). The constructs were subjected to a
compressive strain amplitude of 15% at three dynamic frequencies
(0.3 Hz, 1 Hz, 3 Hz). In addition static compressive strain was
applied. Controls were unstrained. The compressive strain was
applied for 48 hr and GAG synthesis, cell proliferation and total
protein synthesis in response to strain were assessed. GAG
synthesis was stimulated at 1 Hz as shown in FIG. 8a, whilst cell
proliferation was stimulated by all of the dynamic strain regimes
as shown in FIG. 8b.
EXAMPLE (11)
Mechanical Stimulation of Chondrocyte Sub-populations in Hydrated
Gels
[0086] Chondrocytes were isolated, separately, from the superficial
and deep zones of bovine cartilage and seeded within 3 % agarose
constructs. The constructs were subjected to a compressive strain
amplitude of 15% at three dynamic frequencies (0.3 Hz, 1 Hz, 3 Hz).
In addition, static compressive strain was applied. Controls were
unstrained. The compressive strain was applied for 48 hours and
cell proliferation in response to strain was assessed. Cell
proliferation was stimulated in superficial zone cells by all of
the dynamic strain regimes as shown in FIG. 9. By contrast, deep
zone cells were unaffected by both static and dynamic strain
regimes as shown in FIG. 9.
* * * * *