Note: Descriptions are shown in the official language in which they were submitted.
CA 02826771 2013-09-11
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BIOMIMETIC TISSUE GRAFT FOR LIGAMENT REPLACEMENT
CROSS REFERENCE TO RELATED APPLICATIONS
This application claims the priority benefit of U.S. Provisional Patent
Application
Serial No. 61/700,464 filed September 13, 2012.
TECHNICAL FIELD
The present disclosure is generally related to biomimetic ligaments suitable
for
grafting, and to apparatus and methods for the manufacture thereof.
BACKGROUND
The anterior cruciate ligament (ACL) is the most commonly injured ligament of
the
knee with over 200,000 patients in the United States diagnosed with ACL
injuries annually
(Pennisi E. (2002) Science 295: 1011). The ACL is critical to normal
kinematics and stability
in the knee. It stabilizes the knee joint and controls motion by connecting
the femur to the
tibia, effectively preventing abnormal types of motion. It prevents
dislocation and fracture of
bones in the joint by preventing excessive anterior translation of the femur
(Beynnon &
Fleming (1998) J. Biomech. 31: 519-525).
The ACL is a dense connective tissue comprised of four anatomically distinct
areas:
the ligament proper, non-mineralized fibrocartilage, mineralized
fibrocartilage, and
subchondral bone. The ligament proper is an avascular and aneurotic tissue
containing an
extracellular matrix (ECM) composed mainly of collagen type I and type III,
with a low
concentration fibroblast cells. As the ligament transcends to its bone
insertions, the ligament
proper gives way to a region of non-mineralized fibrocartilage. The ECM
becomes less
dense and regular, with a composition mainly of collagen type I and II.
Chondrocytes exist in
this matrix and at a much higher concentration than the concentration of
fibroblasts in the
ligament groper. As the fibrocartilage tissue becomes mineralized, the ECM
contains
hydroxyapatite (Ca10(PO4)6(OH)2) in addition to collagen. The collagen fibers
become more
regular and organized parallel to the ligament. Hypertrophic chondrocytes
exist in the
mineralized fibrocartilage tissue in similar concentrations to the non-
mineralized
fibrocartilage. The mineralized fibrocartilage increases in hydroxyapatite
concentration until
it transforms into subchondral bone. This linear variation in cellularity,
mineral content and
extracellular matrix composition results in a graduated change in stiffness
and allows for
effective load transfer from ligament to bone, minimizing stress
concentrations and
preventing failure (Petrigliano etal., (2006) Arthroscopy 22: 441-451). In
adults, the ACL
has an approximate length of 38mm and a diameter of 10mm.
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Biomechanical tests performed on the human knees have shown the femur-ACL-
tibia
complex to tolerate an ultimate load of 2160 ( 157) N and have a stiffness of
242 ( 28)
N/mm. These values correspond with reported ultimate load and a stiffness
values for the
ACL ligament proper from other sources. Tensile testing on bundles of the ACL
ligament
have shown it to have a Young's modulus of 516 ( 64) MPa (Zantop etal.,
(2006) Knee
Surg. Sports Traumatol. Arthrosc. 14: 982-992; Guilack et al. (2004)
Functional Tissue
Engineering. New York, NY. Springer). Tensile tests on samples of bovine
subchondral
bone have revealed an ultimate tensile strength of 3.5 ( 1.2) GPa. The same
tests revealed
a Young's modulus of 30 ( 7.5) MPa (Braidotti etal., (2000) J. Biomech. 33:
1153-1157).
Compression tests on the ACL ligament-to-bone interface revealed compressive
moduli of
0.32 ( 0.14) MPa and 0.68 ( 0.39) MPa for non-mineralized and mineralized
fibrocartilage,
respectively (Moffat et al., (2009) Clin. Sports Med. 28: 157-176).
Because it is avascular and has a low cell concentration, the natural healing
capabilities of the ACL proper are very limited. ACL injury can be treated
with physical
therapy by strengthening the muscles involved in knee motion, such as the
hamstring,
quadriceps, calf, hip, and ankle. While physical therapy can return full range
of motion, there
is an increased risk of further injury. Additionally, an ACL with reduced
function will result in
increased pressure on the meniscus, resulting in meniscus erosion and an
increased risk of
osteoarthritis. To avoid these risks and in the case of a severe ACL injury,
the torn ligament
can (and often must) be permanently replaced (Frank & Jackson (1997) J. Bone
Joint Surg.
Am. 79: 1556-1576).
Following injury, orthopedic reconstruction of the ACL is often performed to
regain
normal stability and kinematics. A major challenge in orthopedic
reconstruction surgery,
however, is the functional integration of soft tissue grafts with subchondral
bone. Because
integration between graft and bone is essential for musculoskeletal motion,
the fixation of
these grafts is crucial in the effective repair of injuries to ligaments and
tendons (Moffat et
al., (2009) Clin. Sports Med. 28: 157-176).
Currently, the implantation of autografts is the most common treatment, with
the
patellar tendon autograft being the most widely used. The process of
harvesting this graft
involves the surgical removal of the central one-third of the patellar tendon
along with a
section of bone from the patella as well as from the insertion point at the
tibia. This type of
graft is commonly called the "bone-patellar-bone" (BPB) graft. The BPB graft
is then fed
through a tunnel created by drilling though the tibia, drawn across the knee,
and into a tunnel
drilled though the femur. The bone section of the graft effectively provides
integration into
the tibial and femoral tunnels when set with interference screws (Laurencin &
Freeman
(2005) Biomaterials. 26: 7530-7536). This method is considered as the current
"gold
standard" despite several disadvantages, the most important being donor site
morbidity and
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danger of recurring instability (Fu etal., (1999) Am. J. Sports Med. 27: 821-
830; Fu etal.,
(2000) Am. J. Sports Med. 28: 124-130).
Another approach of ACL replacement is the use of allografts obtained from
cadavers, such as patellar tendon, hamstring tendon, and Achilles tendon
(Jackson et al.,
(1993) The Anterior Cruciate Ligament: Current and Future Concepts. Raven
Press, New
York, NY). The benefits of using allografts include the lack of a second
surgery to harvest
the graft, and the elimination of association of donor site morbidity. There
are, however,
significant disadvantages associated with the use of allografts, such as the
potential to
transmit disease, cause bacterial infection, elicit an unfavorable immunogenic
response from
the host, and the inability of the graft to be sterilized without altering the
mechanical
properties of the tissue (Snook etal., (1983) Clin. Orthop. Re/at. Res. 172:
11-13; Miller et
al., (1996) Review of Orthopaedics. W.B. Saunders Company, Philadelphia, PA).
As a result, an alternative ACL replacement graft is needed and has been
heavily
investigated. The early use of synthetic prostheses for ACL replacement made
of non-
resorbable materials such as polyethylene terephtalate (Leeds-Keio Ligament),
polypropylene (Kennedy Ligament Augmentation Device),
poly(tetrafluoroethylene) (GORE-
TEX.RTM), or carbon fibers, was abandoned due to severe inflammatory and
foreign body
reactions, poor mechanical properties, and very poor durability and abrasion
resistance
(Laurencin & Freeman (2005) Biomaterials. 26: 7530-7536; Duerselen etal.,
(1996)
Biomaterials 17: 977-982). To overcome these drawbacks, research has shifted
towards a
biocompatible model and is now focused on tissue-engineered solutions for ACL
reconstruction (Laurencin & Freeman (2005) Biomaterials 26: 7530-7536).
SUMMARY
The present disclosure encompasses embodiments of an artificial prosthetic
ligament, and methods of making thereof, advantageous for the replacement of a
torn or
injured ligament, and in particular of the anterior cruciate ligament (ACL).
The compositions
of the disclosure provide synthetic scaffolds that replicate to a significant
degree the various
regions of a native ACL and allow for the colonization of the scaffold by
cells that can
contribute to the matrix composition and to the mechanical strength of the
manufactured
ligament.
One aspect of the present disclosure, therefore, encompasses embodiments of a
biomimetic composition comprising: a biocompatible scaffold structure
comprising a sheet of
substantially parallel polymeric microfibers and a population of
hydroxyapatite nanoparticles
deposited on the sheet, wherein the population of hydroxyapatite nanoparticles
can be
distributed on the sheet in a pattern mimicking the mineralization of a native
ligament to
bone enthesis.
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In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can be biodegradable.
In embodiments of this aspect of the disclosure, the polymeric microfibers can
comprise poly(lactic acid).
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can further comprise at least one polypeptide deposited thereon.
In embodiments of this aspect of the disclosure, the at least one polypeptide
can be
selected from the group consisting of: an extracellular matrix polypeptide,
fibrin, fibrinogen, a
cell growth factor, and a cell differentiation inducer.
In embodiments of this aspect of the disclosure, the at least one polypeptide
deposited thereon can be fibrin.
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can comprise at least two polypeptides deposited thereon, and wherein one
polypeptide can
be fibrin deposited on the sheet of substantially parallel polymeric
microfibers and at least
one other polypeptide can be deposited on the fibrin.
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can further comprise a population of mesenchymal stem cells, or the progeny
thereof.
In embodiments of this aspect of the disclosure, the biomimetic composition
can
comprise: a biocompatible scaffold structure comprising a sheet of
substantially parallel
polymeric microfibers, a population of hydroxyapatite nanoparticles
distributed on the sheet
in a pattern mimicking the mineralization of a native ligament to bone
enthesis, fibrin
deposited on said sheet of polymeric microfibers, at least one polypeptide
deposited on the
fibrin, wherein the at least one other polypeptide can be selected to promote
the growth
and/or differentiation of a population of mesenchymal stem cells or progeny
thereof
colonizing the scaffold structure, and a population of mesenchymal stem cells
or progeny
thereof, wherein the biocompatible scaffold structure is configured for
replacing a native
ligament of a subject animal or human.
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can be configured for replacing a native anterior cruciate ligament.
Another aspect of the disclosure encompasses embodiments of a method of
forming
a biomimetic scaffold structure, the method comprising the steps of:
generating a sheet of
substantially parallel polymeric microfibers having polymeric nanofibers
deposited on the
surface thereof; distributing hydroxyapatite nanoparticles on the sheet in a
pattern mimicking
the mineralization of a native ligament to bone enthesis; and configuring said
sheet for
replacing a native ligament of a subject animal or human.
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In embodiments of this aspect of the disclosure, the method can further
comprise
contacting the sheet of substantially parallel polymeric microfibers with an
alkali; providing
exposed carboxyl groups; decreasing fiber diameter; and increasing surface
roughness.
In embodiments of this aspect of the disclosure, the method can further
comprise
providing fibrin on the surface of the sheet of substantially parallel
polymeric microfibers.
In embodiments of this aspect of the disclosure, the method can further
comprise the
step of colonizing the biomimetic scaffold with a population of mesenchymal
stem cells or
progeny.
In embodiments of this aspect of the disclosure, the method can further
comprise the
step of depositing a polypeptide on the fibrin on the surface of the sheet of
substantially
parallel polymeric microfibers, wherein the polypeptide is selected to promote
the growth
and/or differentiation of a population of mesenchymal stem cells or progeny
thereof
colonizing the biomimetic scaffold.
In embodiments of this aspect of the disclosure, the method of distributing
hydroxyapatite nanoparticles on the sheet in a pattern mimicking the
mineralization of a
native ligament can comprise microprinting the hydroxyapatite nanoparticles
onto the sheet
of substantially parallel polymeric microfibers or electrophoretically
depositing the
hydroxyapatite nanoparticles, thereby forming a density gradient of the
hydroxyapatite
nanoparticles mimicking the mineralization of a native ligament to bone
enthesis.
BRIEF DESCRIPTION OF THE DRAWINGS
Further aspects of the present disclosure will be more readily appreciated
upon
review of the detailed description of its various embodiments, described
below, when taken
in conjunction with the accompanying drawings.
Fig. 1 schematically shows the variation of four tissue areas on the proposed
prosthesis scaffold of the disclosure. The lower graph displays spatially
corresponding
mineral content to tissue areas.
Fig. 2 schematically illustrates an embodiment of an electrospinning apparatus
modified to deposit PLA nanofibers on a rotating drum: 1, Syringe fed by a
pump with a
charged lead from the voltage source to the needle; 2, polymer solution within
the syringe; 3,
ejected polymer forming micro-or nanofibers; 4, rotating mandrel; 5, grounding
point behind
the rotating mandrel; 6, a variable DC voltage supply.
Fig. 3 illustrates an embodiment of an electrospinner for electrospinning a
PLA
nanofiber mesh on a rotating collector drum.
Fig. 4 illustrates an embodiment of a method of electrospinning PLA nanofibers
to
form a 100-150pm mat. A strip over the aligned microfibers prevents
electrospun nanofibers
from binding to that area.
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Fig. 5 is a digital SEM image after EPD of nanoHA after 4 h on an aligned
electrospun PCL scaffold, with nanoHA displayed as texture on the fibers.
Fig. 6 schematically illustrates the EPD of nanoHA on the PLA nanofiber mat. A
nanoHA gradient is achieved by progressively raising the scaffold out of the
solution, as
denoted by the arrow.
Fig. 7 schematically illustrates the solution casting of PLA with nanoHA
dispersion
into a mold, then achieving nanoHA gradient via electrophoresis. Salt will be
dispersed after
completion of electrophoresis.
Fig. 8 schematically illustrates the solution casting of PLA with dispersed
nanoHA
and salt into a mold. Arrows indicate the removal of the partition after a
partial evaporation of
the solution, for bonding of the two casts.
Fig. 9 schematically illustrates a two-dimensional tri-culture model
separating three
regions of the scaffold with hydrogel dividers, 7.
Fig. 10 schematically illustrates a three-dimensional tissue engineered ACL
graft
after rolling up a 2-D scaffold.
Fig. 11 illustrates a cross-sectional representation of an electrospun PLA mat
indicating 900-1100pm areas selected for use.
Figs. 12A-121 illustrate the modification of 24-well tissue culture plates for
sample
fixation. Fig. '12A: sections of PE tubing for sample fixation. Fig. 12B:
Placement of lOmm
risers at the bottom of the well plate used as a frame. Fig. 12C: Placement of
the PE tubes
from "A" into the frame. Fig. 12D: Application of a biocompatible adhesive to
the flat,
polished end of the PE tubes. Fig. 12E: Placement of a glass slide with an
electrospun PLA
sample on four pre-glued PE tubes. Fig. 12F: Transfer of electrospun PLA
sample to the
glued ends of the four PE tubes. Fig. 12G: Separation of the PLA-capped tubes.
Fig. 12H:
PLA-capped tube before (left) and after (right) trimming of excess PLA. Fig.
121: Placement
of the PLA-capped tube face-down into a new, sterile 24-well culture plate.
Figs. 13A-13C are a series of digital SEM images of electrospun PLA fibers
using the
following uptake rates: Fig. 13A: Stationary; Fig. 13B: 3500 rpm; Fig. 13C:
7000 rpm.
Magnification = 1000x. Scale bars = 25pm.
Fig. 14 is a bar graph illustrating the frequency distribution of electrospun
PLA fiber
diameter at 0 rpm (stationary), 3500 rpm, and 7000 rpm uptake rates.
Fig. 15 is a graph illustrating the percent reduction in area of 7000 rpm PLA
nanofibers from various NaOH treatment times. Values were expressed as mean
S.E.M.
Percent reduction in area is dependent on NaOH treatment time and displays a
logarithmic
trend, r2=0.9816.
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,A
Fig. 16 is a graph illustrating fiber diameter of various uptake rates and
NaOH
treatment times on electrospun PLA. Values were expressed as mean S.E.M.
NaOH
treatments performed on 7000 rpm fibers.
Figs. 17A-17F are a series of digital SEM images of electrospun PLA fibers
subjected
to the following NaOH treatment times: Fig. 17A: 0 min (control); Fig. 17B: 1
min, arrow
indicates minor surface roughness; Fig. 17C: 5 min, arrows indicate visible
surface
roughness and pitting at the fiber interface; Fig. 17D: 10 min, arrow
indicates interface
pitting; Fig. 17E: 20 min, arrows indicate significant surface roughness and
pitting at the fiber
interface; Fig. 17F: 40 min, arrow indicates significant pitting and visibly
decreased fiber
diameter. Magnification = 50,000x. Scale bars = 1pm.
Fig. 18 is a graph illustrating the average percent swelling of the PLA fibers
in the
untreated and NaOH-treated conditions over 60 d. Values were expressed as mean
S.E.M.
Figs. 19A and 19B are a pair of digital SEM images of electrospun PLA fibers
subjected to fibrin immobilization in: Fig.19A: untreated condition, arrow
indicates a particle
of (presumably) fibrin found randomly throughout the sample; Fig. 19B: 20min
NaOH
treatment, arrows indicate connecting membranes of fibrin. Magnification =
50,000x. Scale
bars = 1pm.
Fig. 20 is a bar graph illustrating cellular proliferation over 14 d. There
was an overall
increase in proliferation at day 7, followed by a step-wise increase
manifesting at day 14.
Values expressed as mean S.E.M. ("p<0.01).
Fig. 21 is a graph illustrating gene expression profile for ALP over 14 d. At
day 14,
ALP expression was greatest in the NaOH-treated PLA + fibrin samples. Values
were
expressed as mean S.E.M. (**p<0.01).
Figs. 22A-22F are a series of digital SEM images of the hMSCs on the PLA
samples
at day 14 at 250x in BSE mode (left images) and 10,000x in SE mode (right
images). Figs
22A and 22B: Untreated PLA; Figs. 22C and 22D: 20 min NaOH-treated PLA; and
Figs. 22E
and 22F: 20 min NaOH-treated PLA + fibrin. Scale bar (left images) = 200pm,
(right images)
= 5pm.
Fig. 23 is a graph illustrating treatment-dependent variations of the stress-
strain
curves observed during mechanical testing of the PLA samples. (A) yield
stress, (B)
maximum stress, (C) strain at yield stress, (D) strain at failure, calculated
as the intersection
of a 45 line tangent to the stress-strain curve.
Fig. 24 is a graph illustrating yield stress of the aligned PLA fibers with
various
surface treatments, compared to natural ligament/graft tissue. Values were
expressed as
mean S.E.M. (PT=patellar tendon, "Ligament" includes ACL, PCL, and LCL).
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Fig. 25 is a graph illustrating maximum stress of the aligned PLA fibers with
various
surface treatments, compared to natural ligament/graft tissue. Values were
expressed as
mean S.E.M. (PT=patellar tendon, "Ligament" includes ACL, PCL, and LCL).
Fig. 26 is a graph illustrating the percent strain at failure of the aligned
PLA fibers
with various surface treatments, compared to natural ligament/graft tissue.
Values were
expressed as mean S.E.M.
Fig. 27 is a graph illustrating the elastic modulus of the aligned PLA fibers
at various
temperatures, compared to natural ligament/graft tissue. Values were expressed
as mean
S.E.M. (PT=patellar tendon, "Ligament" includes ACL, PCL, and LCL).
Fig. 28 is a graph illustrating an overlay of the stress-strain profiles from
a control, 20
min NaOH-treated sample, and 20 min NaOH-treated sample with fibrin at body
temperature. Dotted lines indicate the stress-strain profiles of the ACL for
comparison.
The drawings are described in greater detail in the description and examples
below.
The details of some exemplary embodiments of the methods and systems of the
present disclosure are set forth in the description below. Other features,
objects, and
advantages of the disclosure will be apparent to one of skill in the art upon
examination of
the following description, drawings, examples and claims. It is intended that
all such
additional systems, methods, features, and advantages be included within this
description,
be within the scope of the present disclosure, and be protected by the
accompanying claims.
DETAILED DESCRIPTION
Before the present disclosure is described in greater detail, it is to be
understood that
this disclosure is not limited to particular embodiments described, and as
such may, of
course, vary. It is also to be understood that the terminology used herein is
for the purpose
of describing particular embodiments only, and is not intended to be limiting,
since the scope
of the present disclosure will be limited only by the appended claims.
Where a range of values is provided, it is understood that each intervening
value, to
the tenth of the unit of the lower limit unless the context clearly dictates
otherwise, between
the upper and lower limit of that range and any other stated or intervening
value in that
stated range, is encompassed within the disclosure. The upper and lower limits
of these
smaller ranges may independently be included in the smaller ranges and are
also
encompassed within the disclosure, subject to any specifically excluded limit
in the stated
range. Where the stated range includes one or both of the limits, ranges
excluding either or
both of those included limits are also included in the disclosure.
Unless defined otherwise, all technical and scientific terms used herein have
the
same meaning as commonly understood by one of ordinary skill in the art to
which this
disclosure belongs. Although any methods and materials similar or equivalent
to those
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'
described herein can also be used in the practice or testing of the present
disclosure, the
preferred methods and materials are now described.
All publications and patents cited in this specification are herein
incorporated by
reference as if each individual publication or patent were specifically and
individually
indicated to be incorporated by reference and are incorporated herein by
reference to
disclose and describe the methods and/or materials in connection with which
the
publications are cited. The citation of any publication is for its disclosure
prior to the filing
date and should not be construed as an admission that the present disclosure
is not entitled
to antedate such publication by virtue of prior disclosure. Further, the dates
of publication
provided could be different from the actual publication dates that may need to
be
independently confirmed.
As will be apparent to those of skill in the art upon reading this disclosure,
each of the
individual embodiments described and illustrated herein has discrete
components and
features which may be readily separated from or combined with the features of
any of the
other several embodiments without departing from the scope or spirit of the
present
disclosure. Any recited method can be carried out in the order of events
recited or in any
other order that is logically possible.
Embodiments of the present disclosure will employ, unless otherwise indicated,
techniques of medicine, organic chemistry, biochemistry, molecular biology,
pharmacology,
and the like, which are within the skill of the art. Such techniques are
explained fully in the
literature.
It must be noted that, as used in the specification and the appended claims,
the
singular forms "a," "an," and "the" include plural referents unless the
context clearly dictates
otherwise. Thus, for example, reference to "a support" includes a plurality of
supports. In
this specification and in the claims that follow, reference will be made to a
number of terms
that shall be defined to have the following meanings unless a contrary
intention is apparent.
As used herein, the following terms have the meanings ascribed to them unless
specified otherwise. In this disclosure, "comprises," "comprising,"
"containing" and "having"
and the like can have the meaning ascribed to them in U.S. Patent law and can
mean"
includes," "including," and the like; "consisting essentially of" or "consists
essentially" or the
like, when applied to methods and compositions encompassed by the present
disclosure
refers to compositions like those disclosed herein, but which may contain
additional
structural groups, composition components or method steps (or analogs or
derivatives
thereof as discussed above). Such additional structural groups, composition
components or
method steps, etc., however, do not materially affect the basic and novel
characteristic(s) of
the compositions or methods, compared to those of the corresponding
compositions or
methods disclosed herein. "Consisting essentially of" or "consists
essentially" or the like,
9
CA 02826771 2013-09-11.
when applied to methods and compositions encompassed by the present disclosure
have
the meaning ascribed in U.S. Patent law and the term is open-ended, allowing
for the
presence of more than that which is recited so long as basic or novel
characteristics of that
which is recited is not changed by the presence of more than that which is
recited, but
excludes prior art embodiments.
Prior to describing the various embodiments, the following definitions are
provided
and should be used unless otherwise indicated.
Abbreviations
ACL, anterior cruciate ligament; ECM, extra cellular matrix; BPB, bone-
patellar-bone; PLA,
poly(lactic acid); PCL, poly(caprolactone); HA, hydroxyapatite; nanoHA,
hydroxyapatite
nanoparticles; PTFE, poly(tetrafluoroethylene); EPD, electrophoretic
deposition;mmP13,
matrix metalloprotease-13; IGF-1, insulin growth factor-1; CTGF, connective
tissue growth
factor; FGF, fibroblast growth factor; HDACs, histone deacetylases; TGF,
transforming
growth factor; BMP, bone morphogenic protein; H & E, hematoxylin and eosin;
BMPR, bone
morphogenetic protein receptor; BAP, bone-specific alkaline phosphatase; RT-
PCR, real
time polymerase chain reaction; UUMC, uniaxial unconfined micro-compression;
DIC, digital
image correlation; FE, finite element; hMSCs, human mesenchymal stem cells.
Definitions
In describing and claiming the disclosed subject matter, the following
terminology will
be used in accordance with the definitions set forth below.
The term "biomimetic" as used herein refers to a material or structure
designed to
resemble and/or function in a manner similar to a structure, organ, or tissue
found in a native
state in an animal or human. In the embodiments of the disclosure, the
biomimetic
compositions herein disclosed are suitable for the replacement structurally
and functionally
of a ligament such as, but not limited to, an ACL.
The term "biocompatible" as used herein refers to a material that does not
elicit any
undesirable local or systemic effects in vivo.
The term "scaffold" as used herein refers to a material that can provide a
supporting
structure on which animal cells may attach and proliferate. The scaffold may
be configured
to resemble the shape and size of a native animal or human structure or
physical feature
that it is desired to replace.
The term "substantially parallel" as use herein refers to the arrangement of
microfibers wherein at least 80%, preferably at least 85%, more preferably at
least 90%,
more preferably at least 95%, and most preferably at least 99% of the
microfibers are
parallel to one another.
The term "extracellular matrix polypeptide" as used herein refers to a
polypeptide
found in the extracellular matrix of an animal or human tissue including, but
not limited to, a
CA 02826771 2013-09-11
fibrous protein, a glycosaminoglycan, heparan sulfate, chondroitin sulfate,
keratan sulfate,
collagen, elastin, fibronectin, laminin, and the like.
The term cell growth factor as used herein refers to a naturally occurring
substance
capable of stimulating cellular growth, proliferation and cellular
differentiation. Usually it is a
protein or a steroid hormone. Growth factors are important for regulating a
variety of cellular
processes. Growth factors typically act as signaling molecules between cells.
Examples are
cytokines and hormones that bind to specific receptors on the surface of their
target cells.
They often promote cell differentiation and maturation, which varies between
growth factors.
For example, bone morphogenic proteins stimulate bone cell differentiation,
while fibroblast
growth factors and vascular endothelial growth factors stimulate blood vessel
differentiation
(angiogenesis). While growth factor implies a positive effect on cell
division, cytokine is a
neutral term with respect to whether a molecule affects proliferation. Some
cytokines can be
growth factors, such as G-CSF and GM-CSF. Individual growth factor proteins
tend to occur
as members of larger families of structurally and evolutionarily related
proteins such as, but
not limited to, Adrenomedullin (AM), Angiopoietin (Ang-2), Autocrine motility
factor, Bone
morphogenetic proteins (BMP-2, BMP-4, BMP-6), Brain-derived neurotrophic
factor (BDNF),
Epidermal growth factor (EGF), Erythropoietin (EPO), Fibroblast growth factor
(FGF-2 (FGF-
13), FGF-4), Glial cell line-derived neurotrophic factor (GDNF), Granulocyte
colony-
stimulating factor (G-CSF), Granulocyte macrophage colony-stimulating factor
(GM-CSF),
Growth differentiation factor-9 (GDF-5, GDF-7, GDF-8, GDF9, GDF-11),
Hepatocyte growth
factor (HGF), Hepatoma-derived growth factor (HDGF), Insulin-like growth
factor (IGF-1,
IGF-2), Migration-stimulating factor, Myostatin (GDF-8), Platelet-derived
growth factor
(PDGF), Thrombopoietin (TPO), Transforming growth factor-alpha (TGF-a),
Transforming
growth factor-beta (TGF-6), Tumor necrosis factor-alpha (TNF-a), and Vascular
endothelial
growth factor (VEGF).
Description
The human anterior cruciate ligament (ACL) is ruptured over 200,000 times per
year
(or an incidence of 1 in 3000) in the United States, resulting in over $1
billion of medical
costs annually. The current gold standard for surgical repair is the patellar
tendon autograft,
but this treatment is far from optimal due to lengthy recovery time, the
potential for
developing arthritis, associated donor site morbidity, and degenerative joint
disease. These
limitations have prompted the development of a tissue-engineered solution.
However, many
attempts at creating an ACL replacement graft have failed due to issues with
inhomogeneous cellular infiltration and lack of surface penetration though the
scaffold, the
lack of multiple cellular phenotypes mimetic of native ACL tissue, poor
mechanical
properties, and poor biocompatibility.
11
CA 02826771 2013-09-11
The implantable devices of the disclosure are engineered biomimetic ligament
grafts
suitable for use in ligament replacement, including, but not limited to, that
of the anterior
cruciate ligament (ACL). The replacement implants can consist of a
biocompatible
degradable polymeric scaffold seeded with mesenchymal stem cells (or a
combination of
different phenotypes). The scaffold can be effectively populated with the
appropriate cells
mimetic of the native tissue by the use of selected materials such as, but not
limited to,
poly(lactic acid), fibrin, and nanohydroxyapatite (nanoHA) particles, area-
dependent
compositional modifications, surface topography, biochemical manipulations,
and selective
growth environments in vitro. The mechanical properties of the scaffold
support the
development and linear progression of four anatomically distinct tissue areas:
subchondral
bone, mineralized fibrocartilage, non-mineralized fibrocartilage, and the
ligament proper.
Following cellular confluence on the surface of the scaffold, it can be rolled
up, transitioning
the two-dimensional scaffold to a three-dimensional graft, prior to
implantation. Optionally,
the scaffold can be placed in a bioreactor to allow the cells within the
scaffold to further bind
to each other and the scaffold material. The finished constructs resemble
anatomical and
mechanical properties of such as the native human ACL. The device can then be
surgically
implanted in a similar manner to, for example, a patellar tendon graft during
ACL
reconstruction surgery, eliminating the need for secondary surgery to obtain
an autograft, or
the use of an allog raft.
The disclosure provides embodiments of a two-dimensional scaffold biomimetic
for
replacement of a ligament such as, but not limited to, the ACL. In the mimetic
ligaments of
the disclosure, cellular confluence on the scaffold can be obtained before
rolling the scaffold
up into a three-dimensional structure, and thereby ensuring true cellular
penetration
throughout the entirety of the graft. The materials chosen are fully
biocompatible and
degradable; as the cells spread out and multiply within the scaffold, the
synthetic scaffold
material is broken down and replaced by cells and cell-derived material.
Similarly, the
surface topography of the scaffold encourages direction cell growth in
specific regions,
providing mechanical properties similar to the native tissue.
Tissue engineering of a ligament such as an ACL requires cells that are
capable of
producing a ligament-like extracellular matrix and a degradable, thee-
dimensional scaffold
that supports tissue development with close mechanical properties to the
native structure.
Furthermore, the degradable scaffold must exhibit biocompatibility and high
interconnecting
porosity to promote cell attachment, ingrowth, and differentiation.
Accordingly, the
biomimetic structures of the disclosure comprise a biocompatible and
degradable polymeric
scaffold seeded with mesenchymal stem cells (or a combination of different
cell phenotypes)
that can effectively populate the scaffold though the use of ligament-specific
materials, area-
dependent compositional modifications, surface topography, biochemical
manipulations, and
12
CA 02826771 2013-09-11
,
selective growth environments in vitro. Embodiments of the scaffold can be
engineered to
provide four anatomically distinct tissue areas: subchondral bone, mineralized
fibrocartilage,
non-mineralized fibrocartilage, and the ligament proper, to create a cell-
populated,
biomimetic composition of appropriate size, shape, and mechanical function for
the ligament
to be replaced.
A desirable property of any device or prosthetic implanted in the body is the
selection
of materials that are biocompatible, i.e. biologically inert materials that
can be integrated into
the body and do not elicit an adverse immunogenic response. For an engineered
ligament
graft such as an ACL graft, the mechanical properties of the materials
desirably should
resemble those of the native ligament to effectively provide similar
biomechanical function.
Furthermore, it is most advantageous that post-surgical graft-to-bone enthesis
be
established, so that the graft material is able to integrate into bone to
facilitate adequate
anatomical function.
A variety of biocompatible and degradable materials of natural and synthetic
origin
have been studied, including collagen type I (Dunn et al., (1995) J. Biomed.
Mater. Res. 29:
1363-1371; Caruso & Dunn (2005) J. Biomed. Mater. Res. A. 73: 388-397),
hyaluronic acid
(Cristino et al., (2005) J. Biomed. Mater. Res. A. 73: 275-283), silk (Altman
et al., (2002)
Biomaterials 23: 4131-4141; Chen etal., (2003) J. Biomed. Mater. Res. A. 67:
559-570), and
resorbable polymers such as poly(lactic acid) (PLA) and
poly(caprolactone)(PCL)
(Hasegawa etal., (1999) Clin. Orthop. Re/at. Res. 358: 235-243; Lu etal.,
(2005)
Biomaterials 26: 4805-4816; Thomas et al., (2006) J. Nanosci. Nanotechnol. 6:
487-493).
Collagen and hyaluronic acid degrade too rapidly to establish adequate cell
ingrowth and
lack sufficient desirable mechanical properties to be used as a scaffold for
such as an ACL
graft. Silk, PLA, and PCL, however, have more suitable characteristics as a
scaffold
material, with slower degradation rates and desired mechanical properties
(Laurencin &
Freeman (2005) Biomaterials 26: 7530-7536; Altman et al. (2002) Biomaterials,
23; 4131-
4141; Duerselen et al. (2001) J. Biomed. Mater. Res. 58: 666-6720).
A variety of polymers can be modified to obtain functional properties and
design
flexibility desirous in a scaffold. Similarly, biodegradability can be
achieved by tailoring some
of these polymers (Murugan & Ramakrishna (2007) Tissue Eng. 13: 1845-1866). As
such,
embodiments of biomimetic ligaments of the disclosure may advantageously
comprise such
as poly(lactic acid), poly(caprolactone), or a combination thereof. The
biomimetic ligaments
further comprise dispersed nano-hydroxyapatite particles in a poly(lactic
acid) matrix of the
disclosure.
Poly(lactic acid) (PLA) as used herein refers to an aliphatic polyester
derived from
renewable resources, such as corn starch or sugarcane. It is a biodegradable
thermoplastic,
and the degradation product, lactic acid, is metabolically innocuous, making
it an
13
CA 02826771 2013-09-11
. a
õ
advantageous material for medical applications. As such, it is one of the few
biodegradable
polymers approved for human clinical use.
The degradation of PLA involves random hydrolysis of its ester bonds to form
lactic
acid that enters the tricarboxylic acid cycle to be excreted as water and
carbon dioxide. The
degradation rate can vary by altering factors such as structural
configuration, morphology,
stresses, crystallinity, molecular weight, copolymer ratio, amount of residual
monomer,
porosity and site of implantation, and the like, by methods well known to
those in the art.
Another suitable polymer for use in the scaffolds of the disclosure is
poly(caprolactone) (PCL) derived by chemical synthesis from petroleum. It is a
semi-
crystalline, resorbable, aliphatic polyester that biodegrades by hydrolysis of
ester linkages
and eventual intracellular phagocytosis. PCL degrades at a lower rate than PLA
and is
useful in long term, implantable drug delivery systems.
The biomimetics of the disclosure further incorporate hydroxyapatite (HA),
which is
the main inorganic mineral present in animal teeth and bones. It is
biocompatible,
degradation-resistant and highly osteotropic as it is one of the few
biomaterials able to
establish a substantial continuity between itself and bone. Nanohydroxyapatite
(NanoHA)
particles have been incorporated into polymer tissue constructs using various
tissue
fabrication techniques such as electrospinning and solution or melt casting,
resulting in a
significant increase in mechanical properties and improved cell binding and
proliferation,
resulting in improved biointegration. In nanoscale synthetic form with a
similar chemical
composition to that of bone, nanoHA can be bound in vivo to a great variety of
molecules
such as enzymes and proteins. Accordingly, growth factors can be bound to
nanoHA
particles incorporated into the biomimetic ligament grafts of the disclosure
to tailor cell
integration and differentiation.
The biomimetic ligaments of the disclosure provide scaffold structures that
resemble
and facilitate the development of the four tissue types found in ligaments
such as the ACL
and ACL-to-bone interface, as illustrated in Fig. 1. The morphology of the
scaffolds of the
disclosure, accordingly, spatially correspond to these tissue types over four
distinct areas:
the ligament proper (area 1), non-mineralized fibrocartilage (area 2),
mineralized
fibrocartilage (area 3), and subchondral bone (area 4). This material can be
formed by such
as a modified electrospinning device as schematically shown in Fig. 2.
Electrospinning formation of a nanofiber PLA mat can be followed by the
incorporation or deposit of nanoHA onto the mat via such as electrophoretic
deposition
(EPD). EPD is a two-step process that allows rapid deposition rates with a
high degree of
control over deposition thickness. First, charged colloidal-sized
hydroxyapatite nanoparticles
(nanoHA) of, but not limited to, between about 0.2pm to about 40pm in
suspension migrate
towards a counter-charged electrode at which deposition occurs. Secondly,
these particles
14
CA 02826771 2013-09-11
= I 3
I
are deposited (discharged and flocculated) on to the substrate, as described
by Wei et al.,
(2005) J. Mater. Sci. Mater. Med. 16: 319-324, incorporated herein by
reference in its
entirety, and as shown in Fig. 5.
A nanoHA gradient may be generated in a scaffold according to the disclosure
by
progressively raising the nanofiber mat out of the solution during the EPD
process. The
deposition thickness depends on the time allowed for the EPD process, and the
deposition
gradient will depend on the rate of raising the nanofiber mat out of the
solution, as illustrated
in Fig. 6.
Another aspect of the disclosure provides embodiments of a cast PLA film. A
first
manufacturing step involves casting to produce a thin porous PLA film with a
dispersion of
nanoHA particles over a distally increasing gradient relative to the
microfibers. This can be
accomplished by preparing a solution of PLA and nanoHA and then casting the
solution into
a section of a mold partitioned by a thin, inert film with desired dimensions
such as, but not
limited to, about 15mm height and about 100 to about 150 pm in thickness.
While still in the
solution state, a nanoHA gradient can be formed via electrophoresis. Then an
even
dispersion of granular salt (approximately 100 pm) can be incorporated into
the solution,
followed by overlaying a microfiber mat of the disclosure onto the 1mm
casting, as shown in
Fig. 7. A second step then involves casting a solution of PLA, granular salt,
and evenly
dispersed nanoHA (with a greater concentration than the previous PLA-nanoHA
solution)
into the remaining mold space, as shown in Fig. 8.
The solvent can be allowed to partially evaporate until a viscous solution
remains,
and then the partition may be removed, allowing the two separate casts to
bond. The
remaining solvent can be evaporated, and the resulting film removed from the
mold and
washed out to remove the salt and residual solvent. By selectively using salt
granules with
close dimensions to the thickness of the cast film, porosity can be obtained
through the film.
Due to its hydrophobic nature and smooth morphology, the surface of PLA alone
is
not an ideal substrate for cellular integration and the subsequent synthesis
of new tissue.
The use of NaOH to introduce carboxyl groups and increase surface roughness
can assist in
the effective binding of cells, but the combination of the surface carboxyl
groups with an
intermediary protein (fibrin) has a potent effect on cellular attachment and
proliferation.
Fibrin is a natural polymer in the human body that is critical for hemostasis
and
wound healing. Fibrin can be created ex vivo though the rapid enzymatic
polymerization of
fibrinogen with thrombin from either allogeneic or autologous sources. Due to
a natural
binding affinity, the immobilization of many growth factors as well as
improved cell seeding
efficiency and uniformity of cell distribution is possible (Ahmed et al.,
(2008) Tissue Eng. Part
B Rev. 2008; Miller et al., (2009) Comb. Chem. High Throughput Screen 12: 604-
618; Peng
et al., (2004) Blood 103: 2114-2120; Schense & Hubbell (1999) Bioconjug. Chem.
10: 75-
CA 02826771 2013-09-11
81). Furthermore, the biocompatibility and ease of processing from autologous
sources
eliminates immunological concerns (Cummings et al., (2004) Biomaterials 25:
3699-3706).
A combination of atomic force microscopy (AFM) and fluorescence microscopy was
used to measure the strain at failure (extensibility) of fibrin (Liu et al.,
(2006) Science 313:
634). The strain at failure was high, with 332% and 226% elongation in the
cross-linked and
uncross-linked states, respectively, indicating that these properties of
fibrin can affect the
overall mechanical resilience of PLA when immobilized on the surface.
Accordingly, the
effects of stepwise surface degradation and morphological modifications on the
mechanical
properties of highly-aligned electrospun PLA nanofibers were investigated.
Surface
morphology and fiber diameter were altered though induced hydrolytic
degradation via
NaOH treatments over various time periods and then mechanical properties were
evaluated
under a tensile load and compared to natural ligament and ligament-replacement
graft
materials from literature. Fibrin was immobilized on the surface of a NaOH
treated sample,
which was then re-evaluated under tensile load at room temperature (21 C) and
body
temperature (37 C) to assess the effects of the fibrin and the combined
effects of elevated
temperature and fibrin on the mechanical properties.
The mechanical studies of unidirectionally electrospun PLA nanofiber mats of
the
disclosure following various stepwise surface treatments were investigated.
Induced
hydrolytic degradation of the fibers in 0.25M NaOH for various time periods
followed by
immobilization of fibrin on the hydrolyzed fiber surfaces were shown to
significantly affect
yield stress, maximum stress, and strain at failure. The combination of 20 min
of NaOH
hydrolysis followed by the surface immobilization of fibrin effectively
altered the mechanical
properties to values within the range of human ligament and current ligament-
replacement
graft materials. Furthermore, mechanical tests performed at 37 C on a fibrin-
coated NaOH-
treated sample further indicated a stress-strain profile closer to human ACL
tissue than the
other treated/untreated samples. The elastic modulus and strain at yield
stress were not
significantly affected by the various surface treatments, and the values
reported at 37 C
were considerably higher (approximately 2.5x) than the upper range of human
ligament and
ligament-replacement graft materials. These properties are inherent of the
molecular weight
of the polymer used and thus can be altered during synthesis. The most
advantageous
combination of fiber orientation/alignment, induced hydrolytic degradation,
and
immobilization of fibrin can result in modification of the mechanical
properties of an
electrospun PLA tissue scaffold, so as to be mechanically mimetic of human
ligament and
ligament-replacement graft tissue.
Embodiments of biomimetic ligaments of the disclosure can advantageously
promote
the differentiation of mesenchymal stem cells (MSCs). Since the ligaments of
the disclosure
are composed of four differing tissue areas, each area requires a unique
biochemical
16
CA 02826771 2013-09-11
. . 1
I
approach that considers three factors: growth factors, retroviral gene
infection, and adhesive
ligands. These three factors can be considered where appropriate for the four
tissue areas
that comprise a ligament such as an ACL.
Area 1: The Ligament Proper
The ligament proper advantageously has a high degree of fibroblast
differentiation
and deposition of collagen type I and III. These factors may be incorporated
though the
addition of connective tissue growth factor (CTGF). CTGF has effectively been
shown to
induce MSC differentiation into fibroblasts. Research indicates that CTGF-
stimulated MSCs
lose surface mesenchymal epitopes, express broad fibroblastic characteristics,
and increase
synthesis of collagen type I and Ill (Lee etal., (2010) J. Clin. Invest. 120:
3340-3349; Wang
etal., (2006) Circulation 114: 1200-1205). As an added benefit, the CTGF/MSC
derived
fibroblasts exhibit diminished abilities to differentiate into non-
fibroblastic lineages including
osteoblasts, chondrocytes and adipocytes (Lee etal., (2010) J. Clin. Invest.
120: 3340-
3349).
To further stimulate the high levels of type I collagen deposition required in
the
ligament, the use of insulin growth factor-1(IGF-1), a fibroblast-interacting
growth factor that
selectively induces type I collagen deposition, can be incorporated in
addition to CTGF
(Miller etal., (2009) Comb. Chem. High Throughput Screen 12: 604-618; Jonsson
etal.,
(1993) Bioscience Reports 13: 297-302).
Area 2: Non-mineralized Fibrocartilage
In area 2, the MSCs must differentiate into hypertrophic chondrocytes.
Transforming
growth factor beta-1(TGF-131) can be used to achieve this differentiation
path. TGF-131 has
been shown in studies to induce chondrogenic differentiation of MSCs, with
production of
cartilage specific proteoglycans and cartilage type II (Park etal., (Sept.
2010) Biomaterials;
Augello & De Bari (2010) Hum. Gene Ther. 21: 1226-1238; Buxton etal., (2010)
Tissue Eng.
Part A). The TGF-13 family of growth factors has successfully been used and is
a key
requirement in in vitro chondrogenic assays (Augello & De Bad (2010) Hum. Gene
Ther. 21:
1226-1238; Danisovic etal., (2009) Gen. PhysioL Biophys. 28: 56-62; Salinas
etal., (2007)
Tissue Eng. 13: 1025-1034; Lim et a/., (2010) J. Materials Sc.: Materials in
Medicine 21:
2593-2600). The deposition of type I collagen can be stimulated in the same
manner as
previously described for area 1 through the use of IGF-1.
Area 3: Mineralized Fibrocartilage
Mineralized fibrocartilage requires type X collagen deposition and the
differentiation
of MSCs into hypertrophic chondrocytes. The use of bone morphogenic proteins
(BMPs)
can be used to facilitate this response. Within the TGF superfamily, BMPs
offer potential
options to regulate tissue development for area 3, as they have been
identified to induce
osteogenesis as well as promote both chondrogenesis and chondrocyte
hypertrophy (Chen
17
CA 02826771 2013-09-11
=
et at., (2004) Growth Factors 22: 233-2341; Miller et al., (2009) Comb. Chem.
High
Throughput Screen 12: 604-618; Volk etal., (1998) J. Bone Miner. Res. 13: 1521-
1529;
Shen etal., (2010) J. Ce// Biochem. 109: 406-416; Takemoto etal., (2010) J.
Orthop.
Trauma 24: 564-566; Zhou etal., (Aug. 2010) Int Orthop.). The use of BMP-7 has
been
shown to induce both chondrogenesis or osteogenesis, express collagen type X,
and
decrease the expression of collagen type I (Zhou etal., (Aug. 2010) Int
Orthop.). To inhibit
the osteogenic differentiation potential of BMP-7 and promote chondrogenic
differentiation of
the MSCs, it can be advantageous to incorporate TGF- 01. Studies show that in
the
presence of TGF- f31, BMP-7 promotes chondrogenic rather than osteogenic
differentiation
(Shen etal., (2010) J. Cell Biochem. 109: 406-416; Miyamoto et at., (2007) J.
Orthopaedic
Sci. 12: 555-561). The presence of hydroxyapatite nanoparticles on the
scaffold can
provide the mineralization of area 3 mimetic of a native bone-ligament
interface.
Area 4: Subchondral Bone
Area 4 requires the formation of subchondral bone; therefore, cell
differentiation into
osteoblasts is required. Furthermore, it requires type X collagen deposition
and inhibition of
type I collagen. However, many growth factors for bone induce collagen
deposition of all
types. Because it is desired to selectively induce bone only in this area, a
specific growth
factor or group of growth factors is required. As with area 3, the BMP family
of growth
factors can be advantageously used. Two BMP growth factors can serve the
purposes of
area 4 optimally: BMP-2, which selectively induces osteogenesis, and BMP-7, to
promote
osteogenesis and formation of collagen type X while inhibiting collagen type I
(Chen etal.,
(2004) Growth Factors 22: 233-241; Volk et a/., (1998) J Bone Miner. Res. 13:
1521-1529;
Takemoto et al., (2010) J. Orthop. Trauma 24: 564-566). Additionally,
incorporating
hydroxyapatite nanoparticles on the scaffold can further induce accelerated
osteogenesis
in combination with BMP-7 and BMP-2 (Li etal., (Sept 2010) J Biomed. Mater.
Res. A).
Additional Modifications
If necessary, vascular endothelial growth factor (VEGF) can be incorporated
into the
entirety of the scaffold to promote angiogenesis; this can be determined
following in vitro
studies. Similarly, the biomimetic peptide, RGD can be attached to the
scaffold, with the
goal of enhancing cell/material associations. These mimetic peptides can
facilitate cell
adhesion by engaging cell surface integrin receptors (Hennessy et al., (2009)
Biomaterials
30: 1898-1909; Anderson et al., (2009) Biomacromolecules 10: 2935-2944).
As previously stated, many growth factors can be immobilized onto fibrin via a
natural binding affinity without the need for covalent cross-linking or
chemical modification
(Miller etal., (2009) Comb. Chem. High Throughput Screen 12: 604-618; Peng
etal.,
(2004) Blood 103: 2114-2120). It has been shown that growth factors containing
heparin-
binding domains bind to fibrin as well. This is verified by research
indicating fibroblast
18
CA 02826771 2013-09-11
growth factor-2 (FGF-2) having high affinity binding domains for fibrin and
heparin (Peng
etal., (2004) Blood 103: 2114-2120). Similarly, members of the TGF superfamily
such as
TGF481, BMP-2 also demonstrated adequate binding with fibrin and heparin. IGF-
1 can
be engineered with a heparin binding domain that has been shown to effectively
bind to
fibrin (Campbell et al., (1999) J. Biol. Chem. 274: 30215-3021).
CTGF can be bound to fibrin though intermediary binding with fibronectin
(Yoshida
& Munakata (2007) Biochim. Biophys. Acta 1770: 672-680). It is therefore
postulated that
multiple combinations of growth factors can be selectively bound to a fibrin-
coated scaffold
over specific spatial domains.
Transitioning the 2-D Scaffold into a 3-D Graft
Following the appropriate surface modifications, hMSCs can be cultured onto
the
scaffold and allowed to differentiate and thereby produce a variable ECM
spatially
corresponding to the aforementioned bioactive areas. The scaffold can then be
rolled up to
the proper diameter creating a 3-D ligament graft with hierarchical ECM and
phenotype
variation (including bone fixations at either end), corresponding to natural
ligament and a
ligament-replacement graft.
One aspect of the present disclosure, therefore, encompasses embodiments of a
biomimetic composition comprising: a biocompatible scaffold structure
comprising a sheet of
substantially parallel polymeric microfibers and a population of
hydroxyapatite nanoparticles
deposited on the sheet, wherein the population of hydroxyapatite nanoparticles
can be
distributed on the sheet in a pattern mimicking the mineralization of a native
ligament to
bone enthesis.
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can be biodegradable.
In embodiments of this aspect of the disclosure, the polymeric microfibers can
comprise poly(lactic acid).
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can further comprise at least one polypeptide deposited thereon.
In embodiments of this aspect of the disclosure, the at least one polypeptide
can be
selected from the group consisting of: an extracellular matrix polypeptide,
fibrin, fibrinogen, a
cell growth factor, and a cell differentiation inducer.
In embodiments of this aspect of the disclosure, the at least one polypeptide
deposited thereon can be fibrin.
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can comprise at least two polypeptides deposited thereon, and wherein one
polypeptide can
be fibrin deposited on the sheet of substantially parallel polymeric
microfibers and at least
one other polypeptide can be deposited on the fibrin.
19
CA 02826771 2013-09-11
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can further comprise a population of mesenchymal stem cells, or the progeny
thereof.
In embodiments of this aspect of the disclosure, the biomimetic composition
can
comprise: a biocompatible scaffold structure comprising a sheet of
substantially parallel
polymeric microfibers, a population of hydroxyapatite nanoparticles
distributed on the sheet
in a pattern mimicking the mineralization of a native ligament to bone
enthesis, fibrin
deposited on said sheet of polymeric microfibers, at least one polypeptide
deposited on the
fibrin, wherein the at least one other polypeptide can be selected to promote
the growth
and/or differentiation of a population of mesenchymal stem cells or progeny
thereof
colonizing the scaffold structure, and a population of mesenchymal stem cells
or progeny
thereof, wherein the biocompatible scaffold structure is configured for
replacing a native
ligament of a subject animal or human.
In embodiments of this aspect of the disclosure, the biocompatible scaffold
structure
can be configured for replacing a native anterior cruciate ligament.
Another aspect of the disclosure encompasses embodiments of a method of
forming
a biomimetic scaffold structure, the method comprising the steps of:
generating a sheet of
substantially parallel polymeric microfibers having polymeric nanofibers
deposited on the
surface thereof; distributing hydroxyapatite nanoparticles on the sheet in a
pattern mimicking
the mineralization of a native ligament to bone enthesis; and configuring said
sheet for
replacing a native ligament of a subject animal or human.
In embodiments of this aspect of the disclosure, the method can further
comprise
contacting the sheet of substantially parallel polymeric microfibers with an
alkali; providing
exposed carboxyl groups; decreasing fiber diameter; and increasing surface
roughness.
In embodiments of this aspect of the disclosure, the method can further
comprise
providing fibrin on the surface of the sheet of substantially parallel
polymeric microfibers.
In embodiments of this aspect of the disclosure, the method can further
comprise the
step of colonizing the biomimetic scaffold with a population of mesenchymal
stem cells or
progeny.
In embodiments of this aspect of the disclosure, the method can further
comprise the
step of depositing a polypeptide on the fibrin on the surface of the sheet of
substantially
parallel polymeric microfibers, wherein the polypeptide is selected to promote
the growth
and/or differentiation of a population of mesenchymal stem cells or progeny
thereof
colonizing the biomimetic scaffold.
In embodiments of this aspect of the disclosure, the method of distributing
hydroxyapatite nanoparticles on the sheet in a pattern mimicking the
mineralization of a
native ligament can comprise microprinting the hydroxyapatite nanoparticles
onto the sheet
of substantially parallel polymeric microfibers or electrophoretically
depositing the
CA 02826771 2013-09-11
,
hydroxyapatite nanoparticles, thereby forming a density gradient of the
hydroxyapatite
nanoparticles mimicking the mineralization of a native ligament to bone
enthesis.
The following examples are put forth so as to provide those of ordinary skill
in the art
with a complete disclosure and description of how to perform the methods and
use the
compositions and compounds disclosed and claimed herein. Efforts have been
made to
ensure accuracy with respect to numbers (e.g., amounts, temperature, etc.),
but some errors
and deviations should be accounted for. Unless indicated otherwise, parts are
parts by
weight, temperature is in C, and pressure is at or near atmospheric. Standard
temperature
and pressure are defined as 20 C and 1 atmosphere.
It should be noted that ratios, concentrations, amounts, and other numerical
data
may be expressed herein in a range format. It is to be understood that such a
range format
is used for convenience and brevity, and thus, should be interpreted in a
flexible manner to
include not only the numerical values explicitly recited as the limits of the
range, but also to
include all the individual numerical values or sub-ranges encompassed within
that range as if
each numerical value and sub-range is explicitly recited. To illustrate, a
concentration range
of "about 0.1% to about 5%" should be interpreted to include not only the
explicitly recited
concentration of about 0.1 wt% to about 5 wt%, but also include individual
concentrations
(e.g., 1%, 2%, 3%, and 4%) and the sub-ranges (e.g., 0.5%, 1.1%, 2.2%, 3.3%,
and 4.4%)
within the indicated range. The term "about" can include 1%, 2%, 3%, 4%,
5%, 6%,
7%, 8%, 9%, or 10%, or more of the numerical value(s) being modified.
EXAMPLES
Example 1
Production of an Aligned Electrospun PLA Nano fiber Mat: A 20% w/v solution of
PLA with an
intrinsic viscosity of 1.6 dL/g (Medisorb 100L Poly(L-Lactide), Lakeshore
Biomaterials,
Birmingham, AL) using 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP) as a solvent
was prepared.
Unidirectional electrospinning equipment was used; the PLA/HFIP solution was
transferred
to a 10m1 syringe affixed to a 20G x 20cm septum-penetrating needle connected
to a voltage
source, and held constant at 15kV. The flow rate was established by a syringe
pump set at
1m1/h for 2 h.
Fibers were collected at a 15cm distance on a 25cm x 15cm sheet of aluminum
foil
attached via cellophane tape to a cylindrical aluminum mandrel 67cm in
diameter and
15.5cm in length, rotating at 7000 rpm (24.6 m/s) with a stationary grounding
point 3.5cm
behind the mandrel, as illustrated in Figs. 2 and 5. Placing the grounding
point behind the
mandrel as opposed to grounding the mandrel yielded fiber uptake over a
narrower and
more consistent area (approximately 9cm width perpendicular to the rotation
axis compared
to 15cm when using a grounded mandrel).
21
CA 02826771 2013-09-11
After the electrospinning was complete, the resulting PLA fiber mat was not
handled
and left in the electrospinning chamber for 24 h to allow excess solvent to
evaporate, thus
avoiding curling or warping. The foil with the PLA fiber mat was then removed
from the
mandrel and placed in a vacuum desiccator environment for 72 h to evaporate
residual HFIP
and water. Due to the random deposition of fibers during the electrospinning
process, the
thickness of the fiber mat was inconsistent over the approximate 9cm wide
fiber uptake area.
Specifically, the deposition would be greater towards the center of the
samples, with
decreasing thickness towards the outer edges. Multiple thickness measurements
were
taken across the transverse direction of the fibers from each sample (without
the foil) using a
thermomechanical analyzer (TMA); only the areas with a thickness range of
between about
900pm to about 1100pm were used throughout the experimentation (Fig. 11). This
typically
resulted in two useable sections per PLA mat, of approximately 20cm x 2cm.
Example 2
NaOH Treatment of Electrospun PLA: Electrospun PLA samples were placed into a
glass
beaker containing an aqueous solution of 0.25M NaOH at room temperature (21
C). The
samples were immersed for various time periods ranging from 1-40 min, most
typically about
min, depending on the specific aim. Five samples were processed at a time and
monitored to ensure complete immersion at all times and prevent clinging to
each other and
to the beaker wall. After their respective treatment times, the samples were
removed from
20 the NaOH solution and immediately immersed in a 500m1 beaker of
deionized water to dilute
any residual NaOH solution remaining on the samples. The samples were then
rinsed 3
times with deionized water,
placed on a low-lint paper cloth and dried in a desiccator for 72 h.
Example 3
Immobilization of Fibrin on the PLA: PLA samples were covered with a thin
layer of fibrin to
assess various parameters of each specific aim. Modifying methods to produce a
homogenous surface layer of fibrin on a variety substrates (Campbell et al.,
(2005)
Biomaterials 26: 6762-6770; Campbell etal., (1999) J. Biol. Chem. 274 30215-
30221),
incorporated herein by reference in their entireties, 50 ml of an aqueous
solution containing
1.1 U/ml of bovine thrombin (Sigma Aldrich Co., St. Louis, MO) and 5mm of
calcium chloride
was prepared and warmed to 37 C. Similarly, a 50m1 aqueous solution
containing 4mg/m1
of bovine fibrinogen (Sigma Aldrich Co., St. Louis, MO) and 50mm of HEPES-
buffered saline
was prepared and warmed to 37 C. When combined, the thrombin/fibrinogen
solution was
fully polymerized into fibrin within 15 min. To effectively immobilize fibrin
on the surface of
the PLA samples, the thrombin and fibrinogen solutions were combined,
immediately
followed by immersion of the PLA. After a few seconds, the PLA samples were
removed
and placed in an incubator at 37 C for 15 min, allowing the thrombin and
fibrinogen to self-
22
CA 02826771 2013-09-11
. , .
assemble into fibrin on the exposed surface of the PLA mat. The slides were
then triple
rinsed with PBS to remove unbound excess fibrin, and placed in a desiccator
overnight to
dry.
Example 4
Preparation of Samples for Cell Studies: The PLA fiber mat was cut into 15mm x
75mm
samples with the longer side parallel to the alignment of the fibers, and the
foil was removed.
The samples were then subjected to the various processing parameters necessary
for each
specific aim. The samples were then placed on individual 25mm x 75mm glass
slides, held
in place by their static charge.
Example 5
Sample Fixation to the Tissue Culture Plates: Due to their thin size
(approximately 1000pm),
flexibility, and high static attraction to almost every surface, handling the
PLA samples
without damaging them was difficult. Additionally, once immersed in an aqueous
cell culture
environment, the samples would float to the top of the surface, fold onto
themselves along
the fiber alignment direction, or cling to the sides of the cell culture well.
However, for
analysis of cell proliferation and alkaline phosphatase (ALP) activity, PLA
samples were
required that completely covered the bottoms of the culture plate wells; any
human
mesenchymal stem cell (hMSC) attachment to the bottom of the tissue culture-
treated wells
and subsequent proliferation/differentiation would impede the accuracy of the
quantitative
results.
To facilitate a consistent means of obtaining a flat, continuous, and
undamaged
sample surface at the bottom of the culture well, the following sample
fixation technique was
implemented: Polyethylene (PE) tubing (LLDPE Value-Tube, Advanced Technology
Products, Milford Center, OH) with 15.875mm 0.D and 12.7mm I.D. was cut into
18mm
sections, as shown in Fig. 12A. The dimensions resulted in the PE sections
fitting snugly in
the wells of a 24-well cell culture plate (Beckton Dickinson and Co, Franklin
Lakes, NJ). One
end of the tubes was ground flat and smooth using progressively finer grit
sandpaper (240,
320, 400, and 600, respectively). The sections were washed with water,
sonicated in a bath
of acetone to remove any surface oils and ink product identification markings,
followed by a
1 h soak in ethanol for sterilization.
To create a frame to hold the samples during preparation, a 24-well plate was
modified by inserting lOmm sections of the PE tubing at the base of four of
the wells on one
side of the plate to act as risers (see Fig. 12B). Four PE sections were
inserted into the
modified 24-well plate at a time (polished side up) and a biocompatible
adhesive (Mastisol,
Ferndale Laboratories, Inc., Ferndale, MI) was brushed onto the polished
surface (Figs. 12C
and 12D). 15mm x 75mm sections of the various PLA treatments were placed on a
25mm x
75mm glass microscope slide (held by static charge), aligned with the four PE
tubes, and
23
CA 02826771 2013-09-11
. , .
pressed firmly and evenly on the tubes. This resulted in the effective
transfer of the PLA
section to the row of PE tubes (Figs. 12E and 12F). The PLA was then cut with
scissors
between the PE tubes (Fig. 12G), and the separated tubes were removed from the
well plate
frame, individually "stamped" against a rubber surface to ensure complete
adhesion of the
PLA to the PE tube, and placed in a vacuum-desiccator to allow the adhesive to
fully dry
overnight. After the adhesive was dry, the PLA was trimmed around the PE
tubes, the tubes
were placed in sterile 24-well culture plates (Figs. 12H and 121). The
modification resulted in
flat PLA mats affixed to the bottom of the each of the wells with 1.26cm2 of
useable surface
area. All of the prepared 24-well plates were placed under a UV-light hood for
2 h, followed
by rinsing the wells with sterilized deionized water 12 hours prior to
seeding.
Example 6
Preparation of Samples for Imaging: To characterize the morphology of the
hMSCs in
response to the various PLA treatment/processing conditions, separate samples
were
prepared for SEM imaging following cell culture. Unlike the samples prepared
for analysis of
proliferation and ALP assessment, it was not essential that the PLA cover the
entirety of the
well bottom since only the surface of the PLA was observed for qualitative
characterization
of cell morphology. Samples of each of the PLA treatment conditions were cut
into 1cm2
sections, and affixed to a 15mm diameter round glass coverslip (Ted Pella
Inc., Redding,
CA) and allowed to dry in a vacuum desiccator overnight. The samples were then
inserted
face-up into sterile 12-well tissue culture plates and placed under a UV hood
for 2 hours for
sterilization, followed by rinsing the wells with sterilized deionized water
12 hours prior to
seeding.
Example 7
Human Mesenchymal Stem Cell Culture: hMSCs were passaged, retaining only cells
from
passage numbers 4-6. The cells were cultured in mesenchymal stem cell basal
medium
(MSCBM) (Lonza, Inc. Walkersville, MD) supplemented with MSCBM SINGLEQUOTS.RTM
(Lonza, Inc. Walkersville, MD). Upon reaching confluence, cells were removed
from the
culture surface and deactivated by adding an equal volume of Dulbecco's
modified eagle
medium (DMEM) supplemented with 10% fetal bovine serum (FBS), 1% amphotericin
B, 1%
penicillin,1% streptomycin, and 1% L-glutamine. The hMSCs were then
centrifuged at 1,000
rpm for 5 min and re-suspended at a concentration of 15,000 cells per 677pL of
DMEM. The
solution was measured into 677pL aliquots and transferred to the 24-well cell
culture plates
with the PE tubes prepared for the study. The 12-well cell culture plates were
also prepared
for sample imaging.
Cell cultures were held constant under standard culture conditions (37 C, 95%
relative humidity, 5% CO2). Every 3-4 days, old media was replaced with 677pL
of fresh
media. For the 24-well plate samples, the cells were cultured, trypsinized and
harvested at
24
CA 02826771 2013-09-11
day 1, 7, and 14, then cryogenically stored in Eppendorf tubes at -80 C until
analysis of
cellularity was performed and alkaline phosphatase (ALP) activity assessed.
For the 12-well
plate imaging samples, the samples were fixed using 2.5% glutaraldehyde, and
2%
paraformaldehyde in a sodium cacodylate buffer (0.2M, pH 7.4) with deionized
water.
Following initial fixation, the samples were rinsed several times with PBS for
a
minimum of 15 min, followed by post-fixation with 1% sodium tetroxide in 0.1M
phosphate
buffer for 1 h. After re-rinsing with PBS several times for 15 min, the
samples were
dehydrated using a series of graded ethanol: 70% for 15 min, 95% for 15 min,
and 3
changes of 100% for 10 min each. The samples were then subjected to chemical
drying
using 2 parts 100% ethanol and 1 part hexamethyldisilazane (HDMS) for 15 min,
1 part
100% ethanol and 2 parts HDMS for 15 min, then 2 changes of 100% HDMS for 15
min
each. All residual solution was aspirated from the samples, and then the
samples were
allowed to air-dry under a fume hood overnight. The samples were then affixed
to a mount,
sputter-coated, and placed in the SEM. Images were acquired at an accelerating
voltage of
5kV.
Example 8
Analysis of Proliferation: Proliferation was analyzed using hMSCs harvested
from the 24-well
plates at d 1, 7, and 14. Picogreen assay (Molecular Probes, Eugene, OR) was
used to
identify the double stranded DNA content according to manufacturer
specifications. 100pL
cell extracts were placed in 96-well plates for analysis. Picogreen dye was
added into the
sample preparations then incubated in the dark for 15 min. Using a fluorescent
microplate
reader (Synergy HT, BIO-TEK Instruments, Winooski, VT) filtered at 485/528
(excitation/emission) was used to measure the double-strand DNA content.
Example 9
Alkaline Phosphatase Activity: Quantitative ALP activity was assessed on the
hMSCS
harvested from the 24-well plates at 14 d. Using a fluorimetric SensoLyte FDP
Alkaline
Phosphatase Assay Kit (Anaspec, San Jose, CA), 50pL aliquots of each sample
were
assayed for ALP content on a fluorescent microplate reader (using the same
parameters as
described in Example 8, above) then compared to a standard correlating known
ALP content
to fluorescence levels. A picogreen analysis was performed to normalize ALP
expression
via DNA content by determining the specific ALP activity present within each
sample.
Example 10
Statistical Analysis: The results presented herein are representative data
sets with
experiments performed using six samples (n=6) for each condition at 1, 7, and
14 day time
points. Values were expressed as standard error of mean (S.E.M.). Using SPSS
software
(SPSS, Chicago, IL), one-way analysis of variance (ANOVA) was performed to
quantify any
significant differences between conditions at each time point. Tukey multiple
comparison
CA 02826771 2013-09-11
1.
tests were conducted to further determine significant differences between
pairs. A value of
p<0.05 was considered significant for all tests.
Example 11
Production of Electrospun PLA Mats with Altered Mandrel Velocity: For the
purpose of
evaluating mandrel rotation velocity on fiber diameter, alignment, and
crystallinity, two PLA
fiber mats were produced using the same parameters with the exception of
altered mandrel
rotation speed of 0 rpm (stationary) and 3500 rpm (12.3m/s), respectively.
These two PLA
mats did not undergo any of the subsequent NaOH treatments.
Example 12
Characterization of Fiber Alignment, Diameter, and Surface Morphology: SEM was
used to
analyze fiber alignment, diameter, and morphology of the PLA samples.
Additionally,
qualitative assessment of the surface morphology and binding efficacy of
fibrin on PLA in the
untreated vs. NaOH treated condition was evaluated using SEM. Sections of the
7000 rpm,
3500 rpm, stationary (0 rpm), and various NaOH-treated and fibrin-coated
samples were
affixed to a mount, sputter-coated with gold palladium and observed under SEM
(Quanta
650FEG, FEI Co., Hillsboro, OR) at an accelerating voltage of 5kV. The SEM
micrographs
were analyzed to measure fiber diameter utilizing image-analyzing software.
Example 13
Assessment of Crystallinity by Differential Scanning Calorimefry: DSC analyses
were
performed on PLA samples of the following morphology: as received (pellet),
thin film (cast
from the electrospinning solution), unaligned electrospun fibers (0 rpm),
aligned electrospun
fibers (7000 rpm), and 20min NaOH treated aligned electrospun fibers (7000
rpm). The
degree of crystallinity for each sample was calculated using the following
formula:
A) Crystallinity = (AHexp/AH100) x 100
where AHexp and AH100 are the melting enthalpy values for the experimental
sample and
a fully crystalline sample, respectively.
Example 14
Swelling Study: To evaluate physical changes that may have occurred under
physiological
conditions over time, the PLA in the untreated and NaOH treated conditions
were placed in a
PBS solution to evaluate swelling (PBS uptake) over various time periods. A
PLA fiber mat
was cut into multiple 1cm2 samples and the foil was removed. Half of the
samples were
subjected to the 20 min NaOH treatment protocol. Samples of both conditions
(untreated/
NaOH treated) were placed into separate 25m1 glass vials containing PBS
solution (pH 7.4),
sealed, and incubated at 37 C for 1, 7, 14, 30, and 60 d. At the end of each
interval, the
samples were removed from the PBS, dabbed on a lint-free paper towel to remove
surface
PBS, and immediately weighed to obtain wet mass. The samples were then placed
in a
desiccator for 72 h to dry thoroughly. The samples were removed from the
desiccator and
26
CA 02826771 2013-09-11
1
. .
immediately weighed to identify dry mass. Percent swelling was calculated
using the
following equation:
% Swelling = (Mw-Md)/Md x 100
where, Mw is the wet mass after the various incubation periods, and Md is the
mass after the
sample has been dried. All samples were run in triplicate (n = 3) for
statistical validity. This
method to measure swelling was modified from K.S (2008) J. Oral Tissue
Engineering 6: 77-
87, incorporated herein by reference in its entirety for measuring PLA
degradation over time.
Example 15
Sectioning and Measurement of the Electrospun PLA Mats: To obtain consistent
samples for
accurate mechanical testing, the PLA fiber mats were cut into 5mm x 45mm
sections, with
the 45mm length parallel to the alignment of the fibers. Following any surface
treatment
performed, the thickness of each sample was measured in the different
locations along the
sample width by the TMA, and averaged to calculate cross-sectional area for
tensile testing.
Mechanical Testing of the PLA Nanofiber Mats: A micromechanical testing system
(Minimat
Model 2000, TA Instruments, New Castle, DE) was used on the various samples to
establish
mechanical behavior under uniaxial tension. The distance between the clamps
was set at
25mm. Sections of adhesive-backed 320 grit sandpaper were affixed to the
contact surfaces
of the test clamps to provide better contact and minimize slipping of the bulk
fibers under
load. The samples were placed in the clamps, and centered with calipers (
0.25mm) to
ensure uniaxial loading along the fiber alignment direction prior to
tightening of the clamps.
A heated chamber was used for some of the samples to establish mechanical
properties at
temperatures up to 37 C (normal human body temperature). Testing was
performed in
tension mode with a load cell of 200N and at a strain rate of 5mm/min until
failure. The
mechanical properties of the samples were then defined based on these uniaxial
data
obtained.
Example 16
Optimization of a hydroxyapatite nanoparticle (nanohydroxyapatite. nanoHA)
Bioink: An
aqueous solution containing 5% w/v of hydroxyapatite nanoparticles (Nanocerox
Inc., Ann
Arbor, MI), 11% v/v propylene glycol (C3H802) (used as a dispersant and
humectant), and
85% v/v deionized water was mixed, and then sonicated to 2 h to break up
hydroxyapatite
nanoparticles agglomerates. The average size of the hydroxyapatite
nanoparticles in the
solution was characterized using dynamic light scattering. An additional batch
of the
hydroxyapatite nanoparticles bioink (not used for any ensuing cell studies)
was produced
with the addition of 1% w/v of a fluorescent dye (calcein) for the purpose of
characterizing
printed patterns under fluorescent microscopy.
The hydroxyapatite nanoparticles bioinks were loaded into a cartridge, and
placed
into bioinkjet printer (JetLabll, Microfab Technologies Inc, Plano, TX) with a
50pm diameter
27
CA 02826771 2013-09-11
nozzle. This type of piezoelectric printer permits drop-on demand control of
the
hydroxyapatite nanoparticles bioink deposition, with multiple jetting
parameters that can be
customized to fine tune droplet volume, velocity, and firing frequency.
Pattern parameters
such as droplet spacing and overlap could be controlled, and multiple arrays
with variable
dimensions could be written in script form and uploaded to the device for
printing. Two CCD
cameras provided feedback on the jetting parameters: droplets firing out of
the nozzle and
drop deposition on the target substrate, respectively.
Example 17
Design of a Printed Gradient Pattern: The ligament-to-subchondral bone
transition in the
ACL spans a width of 1400pm based on mechanical heterogeneity. Proximal to the
ligament
end of that interface, the fibrocartilage covers an approximate 300pm distance
with roughly
150pm of mineralization, so the total distance of mineralized tissue
(mineralized
fibrocartilage and subchondral bone) across the interface is approximately
1250pm. Using
that dimension as a reference, multiple pattern-programming scripts were
written to
determine the most effective means of replicating a gradient of mineralization
as seen in the
ACL.
By adjusting such parameters as droplet spacing and droplet overlay, a variety
of
gradients with a height of 1250pm were printed on a silicon wafer and imaged
with the CCD
camera on the bioprinter. The printed patterns were compared to histological
images of the
ACL-bone interface (enthesis), resulting in the selection of the gradient with
the most similar
anatomical profile.
Example 18
Inkjet Printing of Hydroxy apatite Nanoparticles on the PLA Nano fibers:
Electrospun PLA
fiber mats were cut into 15mm x 75mm samples with the longer side parallel to
the alignment
of the fibers. To optimize the mechanical and morphological properties, the
samples were
subjected to a 20 min NaOH treatment, and then placed in a desiccator for 72
h. The
samples were then placed on individual 25mm x 75mm glass slides, held in place
by their
static charge. The glass slides with the 15mm x 75mm NaOH treated PLA samples
were
loaded onto the stage of the bioinkjet printer. Four identical 1.25mm x lOmm
gradient
patterns were printed onto the aligned PLA fibers, with the gradient variation
parallel to the
fibers to replicate the anatomical orientation of an ACL microstructure.
Example 19
Characterization of the Printed Hydroxyapatite Nano particles: SEM coupled
with EDS was
used to characterize the distribution of hydroxyapatite nanoparticles on the
PLA fibers via
printing. A section of the PLA mat with the printed hydroxyapatite
nanoparticles gradient
was affixed to a mount, sputter-coated with gold-palladium and placed in the
SEM. The
sample was examined and imaged using an accelerating voltage of between 5-
15kV,
28
CA 02826771 2013-09-11
, =
depending on magnification (due to beam damage on the PLA fibers). Spectra of
mineral
distribution were obtained using energy dispersive X-ray analysis (EDS) (INCA
Energy 200
EDS, Oxford Instruments, Oxfordshire, UK) at 200x magnification and an
accelerating
voltage of 20kV to determine Ca and P content. The spectra were collected over
200pm
square regions over various areas of the 1250pm printed hydroxyapatite
nanoparticles
gradient pattern, as well as areas beyond the gradient without any visible
hydroxyapatite
nanoparticles droplets to serve as a control.
Example 20
Stability Evaluation of the Printed Hydroxyapatite Nanoparticles Gradient: Due
to various
aqueous processing and experimental conditions (immobilization of fibrin, cell
culture, etc.),
the hydroxyapatite nanoparticles immobilized on the surface of the NaOH-
treated PLA fibers
were evaluated to assess their stability in an aqueous environment. PLA
samples were
printed with the hydroxyapatite nanoparticles-calcein bioink to produce the
gradient. The
samples were cut in half along the direction of the aligned fibers at the
center of the gradient.
One of the halves of the samples were immersed in separate 25mL vials
containing a
solution of phosphate buffered saline (pH 7.4), and incubated at 37 C for 72
h, changing the
PBS daily. Following the 72 h incubation, the samples were taken out and
placed in a
desiccator for 72 h. The hydroxyapatite nanoparticles-calcein samples were
affixed to a
microscope slide and imaged under fluorescence microscopy to qualitatively
compare any
changes in the gradient morphology.
Example 21
Immobilization and Characterization of Fibrin on the PLA: A portion of the
hydroxyapatite
nanoparticles printed samples were coated with fibrin to evaluate any
encapsulation effects
on the printed particles and gauge hMSCs response as a result of the
combination the fibrin
and hydroxyapatite nanoparticles. SEM micrographs were obtained to
qualitatively assess
any modification of surface morphology and encapsulation of the hydroxyapatite
nanoparticles in fibrin on the surface of the PLA fibers.
Example 22
Fiber Alignment, Diameter, and Morphology: PLA nanofibers were produced
utilizing
electrospinning. Fibers were collected on three uptake conditions: stationary
(0 rpm), 3500
rpm mandrel rate, and 7000 rpm mandrel rate. SEM images exhibited nanofibrous
morphology for all conditions with varying degrees of inter-fiber spacing,
diameter, and
alignment (Figs. 13A-13C). The stationary sample had relatively straight
fibers that were
randomly oriented with an average diameter of 2608 353nm (Fig. 13A). The
3500 rpm
sample showed a more ordered alignment and orientation perpendicular to the
axis of
mandrel rotation, reduced inter-fiber spacing compared to the stationary
sample, and an
average fiber diameter of 1396 312 nm (Fig. 13B). As the uptake rate
increased to 7000
29
CA 02826771 2013-09-11
,
rpm, the fibers achieved a highly-aligned morphology perpendicular to the axis
of mandrel
rotation. The inter-fiber spacing was significantly reduced in comparison to
the stationary
and 3500 rpm samples. Similarly, the average fiber diameter was reduced to 760
96 nm
with greater regularity (Fig. 130). These findings indicate that increased
uptake rates for
electrospun PLA resulted in increased fiber order and alignment, and the rates
were
inversely proportion to fiber diameter and spacing.
Average fiber diameters ranged from about 760 nm to about 2608 nm depending on
the uptake condition. The range of fiber diameters within a specific uptake
condition was
also uptake-rate dependent. Fig. 14 shows the frequency distribution of fiber
diameters for
the electrospun PLA at the different uptake conditions. Increased uptake rates
resulted in
decreasing distributions of the fiber diameters. This was shown with the
stationary, 3500
rpm, and 7000 rpm rates and their standard error of mean (S.E.M.) (n = 50)
distributions of
50, 44, and 14, respectively.
To facilitate a fiber diameter mimetic of the 50-500nm diameter collagen
bundles, the
PLA electrospun at the 7000 rpm rate was subjected to 0.25M NaOH hydrolysis
treatments
of 1,5, 10, 20, and 40 min time intervals. Table 1 shows the average diameter
and
reduction in cross-sectional fiber area as a result of the NaOH treatments.
Table 1: NaOH treatment effects on 7000 rpm PLA nanofibers.
NaOH Treatment Average diameter (nm) Reduction in fiber area
(%)
Time (min) S.E.M* S.E.M*
0 (control) 760 13.6
1 700 13.6 13.6
3.4
5 601 11.1 36.4
2.3
10 564 12.3 43.5
2.3
457 12.5 62.5 2.1
40 399 13.1 70.9
2.0
*n=50
20 The 1, 5 and 10 min treatments all had fiber diameters above the upper
limit of the
collagen bundles. Treatments at 20 and 40 min produced fibers that fall within
the upper
range of the collagen bundles with average diameters of 457 89nm and 399
92nm,
respectively. However, the 40 min fibers were found to be very brittle and
would separate
easily during handling, while the 20 min samples maintained a structural
integrity more
similar to the untreated samples.
The final cross-sectional area of the fibers was shown to be dependent of NaOH
treatment time, with the percent reduction in area following a logarithmic
trend as treatment
time was increased, as shown in Fig. 15. The combined effects of increasing
the
electrospun PLA uptake rate followed by increasing intervals of hydrolytic
degradation
treatments on fiber diameter are illustrated in Fig. 16.
CA 02826771 2013-09-11
The various NaOH treatments also had an effect on the surface morphology of
the
fibers, as shown in Figs 17A-17F. The control sample (Fig. 17A) had a smooth
and
continuous surface morphology. There were some small cracks (less than 100nm)
present
at certain points on the fibers, believed to be caused by
expansion/contraction of the
polymer and subsequent cracking of the gold-palladium coating during sample
preparation,
or by thermal damage from the electron beam during SEM imaging. The fibers
displayed no
pitting or altered shape/thickness along the length of the fibers. The 1 min
sample displayed
relatively smooth fiber morphology with a very minor amount of pitting and
surface
roughness, indicated by the arrow in Fig. 17B.
There were lightly colored spots on the fibers that were found throughout all
of the 1
min samples, believed to be points of increased surface roughness. When
observed at the
visible edge of a fiber, there appeared to be a dip in surface topography.
Attempts at
viewing the spots at a higher magnification were not possible due to thermal
damage caused
by the electron beam on the polymer when magnification was increased.
The 5 and 10 min samples in Figs. 17C and 17D, respectively, exhibited more
pronounced degradation compared to the control and 1 min samples. A higher
degree of
surface roughness could be observed, and there was visible pitting at the
interface between
fibers, indicated by the arrows. The 20 and 40 min samples displayed a degree
of surface
roughness and interface pitting, as indicated by the arrows in Figs. 17E and
17F,
respectively. Additionally, a reduced overall fiber diameter was visibly
apparent when
compared to the control sample. These data presented indicate that obtaining
PLA
nanofibers via unidirectional electrospinning with a mandrel rotational
velocity of about 24.6
m/s, followed by a 20 min treatment in 0.25M NaOH solution will result in a
polymer matrix
with fiber alignment and diameter mimetic of native collagen bundles found in
human
ligament tissue.
Example 22
Analysis of Crystallinity: Using an intrinsic viscosity value of 1.6 clUg, the
molecular weight of
the PLLA used in this study was calculated to be 56,273 using the Mark-Houwink
equation in
the following form:
[ri] = 5.45 x 104.M,"3
where 11 and M 73 are the intrinsic viscosity and average molecular weight,
respectively. To
calculate crystallinity, the value of 75.57 J/g was used, corresponding to the
DSC
thermogram of a fully crystalline sample of PLLA with a similar molecular
weight as the PLLA
used herein. The results of the DSC analysis are shown in Table 2.
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CA 02826771 2013-09-11
Table 2. Differential scanning calorimetry (DSC) data of various PLA
morphologies
Sample As Solution Electrospun Electrospun Electrospun
Received Cast Thin Fibers Fibers Fibers
(Pellet) Film (Random) (Aligned)
(Aligned, 20min
NaOH treated)
Melting 60.58 34.06 41.54 54.14 54.62
Enthalpy
(J/g)
Melting 178.72 174.93 174.74 174.17 174.55
Temperature
( C)
Crystallization 80.16 45.07 54.97 71.64 72.28
(0/0)
The various morphologies had little effect on the melting temperatures.
However, the
melting enthalpy was significantly reduced with the solution cast compared to
the as
received sample. Melting enthalpy and subsequent crystallization increased
when the PLA
solution was electrospun into randomly oriented fibers. Crystallinity was
further increased
when the fibers were collected on the 7000 rpm rotating mandrel. The reduced
diameters of
the aligned fibers indicate a significant degree of fiber stretching, and
subsequently, strain-
induced crystallization. The 20 min NaOH treated aligned fibers displayed a
slight increase
in crystallinity compared to the untreated aligned fibers, likely due to
polymer chain scission
caused by the NaOH hydrolysis and the subsequent reduction in molecular weight
on the
fiber surface.
Example 23
Swelling: Fig. 18 displays the percent swelling of the PLA fibers in the
untreated and NaOH
treated conditions over a 60-day period. Samples in the untreated condition
displayed
minimal swelling (0.07% to 1.1%), likely due to the high hydrophobicity of
PLA. Conversely,
the NaOH treated samples displayed modest swelling (4.04% to 9.2%) due to
their
increased surface roughness and surface carboxyl groups. It is presumed that
this
effectively reduces hydrophobicity and increases PBS uptake.
Example 24
Immobilization of Fibrin: Samples of the 7000 rpm PLA in the untreated and
20min NaOH
treatment condition were subjected to surface immobilization of fibrin. As
shown in Figs.
19A and 19B, there was a difference in the surface morphology of the untreated
PLA (Fig.
19A) and the NaOH treated sample (Fig. 19B). The surfaces of the untreated
fibers were
smooth, continuous, and closely resembled the control sample shown in Fig.
17A, with the
exception of particles randomly scattered along the surface, presumably
comprised of fibrin
(indicated by an arrow in Fig. 19A).
Conversely, the NaOH treated sample displays a different morphology from the
untreated sample with the added fibrin, as well as the sample with the same
NaOH
32
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. .
,
treatment in Fig. 17E. The pitting and rough surface topography seen in the
samples with
the same 20min NaOH treatment was eliminated and was generally smooth with an
apparently confluent superficial layer of fibrin. Another indicator of
confluence was the
connecting membranes at the interface between fibers, indicated by the arrows
in Fig. 19B.
It is considered that the highly hydrophobic nature of the untreated PLA may
inhibit effective
adhesion of fibrin, but when hydrolytically degraded by NaOH, the subsequent
addition of
the carboxyl groups and increase in surface roughness results in an increased
binding
affinity of fibrin.
Example 25
Cell Studies: To determine their proliferative rate over time, hMSCs were
cultured on three
conditions of the PLA fiber mats (untreated, 20 min NaOH, and 20 min NaOH +
fibrin) and
their long term proliferation was assessed over 14 d, as shown in Fig. 20. By
day 14,
cellular proliferation had progressed in a fashion highlighting a decrease in
proliferative rate
based on modifications to the PLA fiber mats. Proliferation proceeded most
rapidly on the
untreated PLA fiber mats, approaching a cellular DNA content of 113.13 3.66
ng/well. Cells
cultured on 20 min NaOH-treated PLA mats displayed slightly less proliferation
over time.
The addition of fibrin onto the 20 min NaOH-treated PLA mat further slowed
proliferation,
showing little increase in cellularity between d 1 and 14. Considering that
proliferation is
known to plateau at the onset of differentiation, it is possible that NaOH
treatment modified
the PLA surface properties in a manner that facilitated osteogenic
differentiation and that the
addition of fibrin does so further.
To assess mineralization, hMSCs were cultured on the modified PLA fiber mats
for
up to 14 d, following which physical ALP activity was determined utilizing a
quantitative
biochemical assay (Fig. 21). These values were normalized by DNA content as a
means to
assess the activity of each individual cell as opposed to total ALP activity.
By day 14, the 20
min NaOH + fibrin condition exhibited statistically greater ALP activity
relative to the other
conditions.
SEM images of the hMSCs on the various PLA samples at day 14 are displayed in
Figs. 22A-22E. Fig. 22A displays the untreated PLA sample at 250x
magnification; an
almost confluent layer of cells can be seen on the sample with individual
cells displaying an
elongated and spindle-like morphology, aligned in the general direction of the
PLA fibers. A
higher magnification (50,000x) image of the same sample is shown in Fig. 22B.
The PLA
fibers are visible with the cells exhibiting finger-like projections extending
towards the fibers.
Fig. 220 displays the 20min NaOH treated sample with a visible cell density
slightly less
than seen on the untreated sample and a higher degree of cell alignment, but
with similar
cell morphology. The higher magnification image corresponding to the 20min
NaOH sample
is shown in Fig. 22D. There is not a significant amount of discernible
difference at high
33
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. .
magnification between the untreated sample and the 20min NaOH sample. Fig. 22E
shows
the fibrin coated NaOH-treated sample. There is a reduction in cell density on
the surface
compared to the other samples at that magnification. The morphology of some of
the cells
appears to be slightly less elongated or spindle-like, adopting a more
rectangular or cuboidal
shape.
The high magnification image shows small interconnecting strings of fibrin
between
the PLA fibers and more pronounced cellular processes integrating with the
fibers and fibrin.
The various images in Figs. 22A-22E correlate to the quantitative data
obtained from the
analysis of proliferation and the ALP study; the untreated sample had a
visibly higher
quantity of cells on the surface with a diminishing amount for the NaOH-
treated and the
NaOH + fibrin-treated samples, respectively. Additionally, a subtle change in
morphology
(from elongated/spindle-like to shorter/more cuboidal) is visible in the
fibrin coated sample
when compared to the others, indicating the early onset of osteogenic
differentiation. The
cells spread and oriented themselves along the direction of the fibers. This
is advantageous
in an engineered ligament. An aligned orientation of cells and collagen fibers
occurs in native
ligament tissue. Subsequent differentiation, and the ECM production by the
hMSCs on the
scaffolds of the disclosure can follow the same morphology due to the presence
of the
aligned PLA fibers.
Based on the results, osteogenic differentiation appears to have been enhanced
by
the 20 min NaOH treatment and even further by the addition of fibrin as
indicated by
decreased proliferation and enhanced ALP activity. This behavior being
exhibited on the
NaOH treatment conditions is likely based on an increase in surface roughness
since
surface topography has been shown to modify cellular behavior, as described
earlier. The
addition of fibrin led to even further enhanced osteogenic activity.
During NaOH hydrolysis, carboxyl groups were added to the exposed surface of
the
PLA fibers and surface roughness was visibly increased. The affinity of the
fibrin and
improved cellular response to the hydrolyzed fibers may have been a function
of surface
roughness, surface carboxyl groups, or both.
Example 26
Uniaxial tensile tests of the aligned electrospun PLA fibers were performed
using ten
samples (5 dry, 5 wet) per surface treatment condition including a control
(unmodified
aligned electrospun sample). The data revealed considerable variety in the
mechanical
properties based on the surface treatments employed on the samples. The wet
samples did
not deviate beyond the range of the data obtained from the dry samples of each
given
condition. One-way ANOVA performed on the wet/dry samples for each treatment
condition
yielded p-values ranging from 0.12 - 0.28, indicating relatively equal sample
means.
34
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=
Therefore, the data obtained for the wet and dry samples were combined.
Numerical results
are reported as mean standard error of mean (S.E.M.).
Fig. 23 displays a representation of the two conditions obtained throughout
the
mechanical study. The yield stress (A) is the point at which the linear
elastic region yields
and begins to exhibit non-linear deformation. The maximum stress is the
highest value of
stress obtained throughout the test, and can differ from the yield stress due
to strain-induced
crystallization of the polymer fibers (B). The strain at yield stress is the
percent elongation at
the end of the linear elastic region (C), and the strain at failure is the
percent elongation
when the fibers begin to rupture, quantitatively assessed in this study by the
intersection of a
downward 45 line tangent to the stress/strain curve at the onset of failure
(D). This method
was implemented because a consistent means of establishing the point of total
failure was
difficult, due to the stair-step shape of the stress-strain curve during
failure as fibers were
breaking irregularly rather than in unison.
The yield and maximum stresses for the samples are illustrated in Figs. 24 and
25,
respectively. The 1 min samples indicated a slight loss of yield strength
(50.1 MPa 0.78),
while 5 and 10 min samples displayed slightly higher yield strength than the
control sample,
with values of 54.1 MPa 0.29, 53.0 MPa 0.79, and 52.5 MPa 0.59,
respectively.
The variety of 20 min samples (20 min, 20 min + fibrin, and 20 min + fibrin at
37 C)
exhibited a modest decline in yield stress (47.4 MPa 1.1, 48.6 MPa 0.66, and
40.6 MPa
1.3, respectively) and displayed properties most similar to the ACL and
ligament tissue
data, particularly the 20 min + fibrin at 37 C sample. The 30 min and 40 min
samples had a
sharp decline in yield stress with a wider distribution of values (29.2 MPa
3.9 and 21.5 MPa
4.5, respectively).
The maximum stress values were higher than the yield stress for the 1, 5, 10,
and
20min + fibrin samples (55.4MPa 1.2, 65.2MPa 0.48, 57.8MPa 1.2, and 58.0MPa
0.89,
respectively). For the 1, 5, and 10 min samples, strain induced
crystallization is likely
responsible for the increase in strength caused by the hydrolytic degradation
by NaOH and
subsequent lowering of the molecular weight on the PLA fiber surface,
corresponding to an
increase in degree of crystallization during strain (Kulkarni et al., (2007)
Surface and
Interface Analysis 39: 740-746). In terms of simple crystallization kinetics,
a shorter polymer
chain (lower molecular weight) can align into ordered lamellae easier than a
longer chain.
Fig. 26 displays the percent elongation at failure, with the 1 min, 5 min and
10 min
NaOH-treated sample values higher than the control sample (36.4% 1.6, 40.5%
1.8,
36.7% 1.7, and 29.2% 1.4, respectively), which corresponds to the
aforementioned strain
induced crystallization observed for the same samples. The 20 min, 30 min and
40 min
samples (23.3% 3.3, 8.4% 0.60, and 6.5% 1.4, respectively) exhibited
elongation at
failure less than the control sample, with the 30 min and 40 min samples
failing shortly after
CA 02826771 2013-09-11
the onset of plastic deformation with less than 10% elongation at failure. The
highest values
of strain at failure were observed for the 20 min + fibrin (49.6% 2.2) and
20 min + fibrin at
37 C (38.0% 4.5) samples, both within range of natural ACL tissue.
A combination of factors can result in a preferred response, where increased
strength
caused by decreased molecular weight (due to hydrolysis) is not affected by
reduced
strength caused by reduced fiber diameter (also due to hydrolysis). The 5 min
and 20 min +
fibrin samples both display these properties and are both within the range of
the natural ACL
and graft tissue for yield/maximum stress. The 5 min sample had higher
stresses (yield and
maximum) and a greater percent elongation at failure than all of the other
NaOH-treated
samples. The 20 min + fibrin sample also displays a maximum stress higher than
the yield
stress, believed to be caused by a combination of strain-induced
crystallization kinetics and
the large extensibility of fibrin. However, the percent elongation at failure
for the 20 min +
fibrin samples is the highest out of all the samples tested, and it is within
the range of natural
ACL tissue.
The addition of fibrin on the surface of the fibers fortified the reduced
diameter
caused by the 20 min NaOH-treated PLA, effectively negating the effects of
lowered yield
and maximum stresses and the strain at failure, when compared to the 20 min
sample
without fibrin.
The elastic modulus for the PLA was consistent (2542 MPa 23) for all of the
samples tested (n=80) at room temperature; however, the values were about 4-
fold higher
than the values of the patellar tendon (643 MPa 53) and between 7- to 40-
fold higher than
the various ligament data (65-345 MPa), as seen in Fig. 27. Samples were also
tested at 29
C and 37 C to see the effects of temperature on the modulus, yielding 1884 MPa
39 and
1636 MPa 70, respectively. The 37 C (body temperature) sample was
approximately 2.5-
fold higher than the patellar tendon; however, it indicated a 36% reduction in
modulus
compared to samples tested at room temperature.
To visually understand the overall effects of the various surface treatments
and
temperature on the mechanical behavior of the PLA fibers, Fig. 28 displays the
stress-strain
profiles of a control, 20 min NaOH-treated sample, and 20 min NaOH-treated
sample with
fibrin at body temperature. The stepwise process of electrospinning a highly
aligned PLA
nanofiber mat, followed by a 20 min NaOH hydrolysis treatment, then
immobilizing fibrin on
the hydrolyzed surface of the fibers results in a mechanical profile more
similar to human
ligament tissue than unmodified PLA.
36
