U.S. patent application number 12/279172 was filed with the patent office on 2009-02-19 for plga/hydroxyapatite composite biomaterial and method of making the same.
This patent application is currently assigned to Nano Orthopedics, LLC. Invention is credited to Byung-Soo Kim.
Application Number | 20090048358 12/279172 |
Document ID | / |
Family ID | 38475500 |
Filed Date | 2009-02-19 |
United States Patent
Application |
20090048358 |
Kind Code |
A1 |
Kim; Byung-Soo |
February 19, 2009 |
PLGA/Hydroxyapatite Composite Biomaterial and Method of Making the
Same
Abstract
Tissue engineering is a growing field where new materials are
being developed for implantation into the body. One important area
involves bone graft materials to replace areas of bone lost to
trauma or disease. Traditionally, graft material may be harvested
from the bone of the individual receiving the graft material.
However, this requires an additional surgery and additional
recovery. Bone also may be taken from others, or even cadavers, but
this introduces biocompatibility problems as well as the risk of
disease transfer. Ideally, a biocompatible material is sought that
will act as a filler with appropriate mechanical strength,
encourage bone healing, and degrade to allow new bone ingrowth
without the risk of disease transfer. The present invention is a
new composite bone graft material made from biocompatible
poly(D,L-lactic-co-glycolic acid) (PLGA) and nano-sized
hydroxyapatite particles exposed on its surface using a gas foaming
particle leaching (GF/PL) method. A further embodiment of this
invention involves coating this PLGA/hydroxyapatite biomaterial
with an adherent, fast, uniform coating of a mineral such as
apatite. The PLGA polymer portion of the composite provides
sufficient mechanical strength to replace bone and is degradable
over time to allow new bone tissue ingrowth. The incorporated
hydroxyapatite particles increase the composite material's
osteogenic properties by providing sites for tissue attachment and
propagation. Finally, a uniform coating of mineral apatite on the
surface of this novel biomaterial composite further enhances its
osteogenic qualities.
Inventors: |
Kim; Byung-Soo; (Seoul,
KR) |
Correspondence
Address: |
LOTT & FRIEDLAND, P.A.
ONE EAST BROWARD BLVD., SUITE 1609
FORT LAUDERDALE
FL
33301
US
|
Assignee: |
Nano Orthopedics, LLC
Boca Raton
FL
|
Family ID: |
38475500 |
Appl. No.: |
12/279172 |
Filed: |
March 6, 2007 |
PCT Filed: |
March 6, 2007 |
PCT NO: |
PCT/US07/05693 |
371 Date: |
August 12, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60767137 |
Mar 6, 2006 |
|
|
|
Current U.S.
Class: |
521/76 ; 264/500;
523/115 |
Current CPC
Class: |
A61L 27/46 20130101;
A61L 27/44 20130101; A61L 2430/02 20130101; A61L 27/56 20130101;
A61L 2400/12 20130101; C08L 67/04 20130101; A61L 27/44
20130101 |
Class at
Publication: |
521/76 ; 523/115;
264/500 |
International
Class: |
A61F 2/02 20060101
A61F002/02; C08J 9/35 20060101 C08J009/35; B29C 43/00 20060101
B29C043/00 |
Claims
1. A bone graft biomaterial comprised of a biocompatible polymer
and hydroxyapatite composite wherein the hydroxyapatite particles
are less than 1.0 .mu.m in diameter.
2. The bone graft material of claim 1 wherein the scaffolds are
made via a solvent casting particulate leaching or gas foaming
particulate leaching process.
3. The bone graft biomaterial of claim 1 wherein the biomaterial
contains pores sized 200 .mu.m or less.
4. A bone graft biomaterial of PLGA/hydroxyapatite as in claim 1
further comprising an adherent, uniform coating of apatite on the
surface of the biomaterial.
5. A bone graft biomaterial of PLGA/hydroxyapatite as in claim 1
further comprising an adherent, uniform coating of apatite on the
surface of the biomaterial.
6. The bone graft material of claim 1 wherein the ratio of
biocompatible polymer to hydroxyapatite is from about 1:2 to about
2:1.
7. The bone graft material of claim 1 wherein the ratio of
biocompatible polymer to hydroxyapatite is about 1:1
8. The bone graft material of claim 1 wherein the ratio of
biocompatible polymer to NaCl ranges from 1:4 to 1:18.
9. The bone graft material of claim 1 wherein the ratio of
biocompatible polymer to NaCl is about 1:9.
10. The bone graft material of claim 3 wherein the pores are
created by washing out a salt.
11. The bone graft material of claim 3 wherein the pores are from
approximately 100 .mu.m to 200 .mu.m in diameter.
12. The bone graft material of claim 1 wherein the biocompatible
polymer is comprised of at least one of the following: poly
glycolic acid, poly lactic acid, poly lactic co glycolic acid.
13. The bone graft material of claim 1 wherein the biocompatible
polymer is poly lactic co glycolic acid.
14. A bone graft biomaterial comprised of a polylactic co glycolic
polymer and hydroxyapatite composite wherein the hydroxyapatite
particles are less than 1.0 .mu.m in diameter.
15. The bone graft material of claim 1 wherein the scaffolds are
made via a solvent casting particulate leaching or gas foaming
particulate leaching process.
16. The bone graft biomaterial of claim 1 wherein the biomaterial
contains pores sized 200 .mu.m or less.
17. A bone graft biomaterial as in claim 1 further comprising an
adherent, uniform coating of apatite on the surface of the
biomaterial.
18. A bone graft biomaterial as in claim 1 further comprising an
adherent, uniform coating of apatite on the surface of the
biomaterial.
19. The bone graft material of claim 1 wherein the ratio of
biocompatible polymer to hydroxyapatite is about 1:1.
20. The bone graft material of claim 3 wherein the pores are
created by washing out a salt.
21. The bone graft material of claim 3 wherein the pores are from
approximately 100 .mu.m to 200 .mu.m in diameter.
22. The bone graft material of claim 1 wherein the biocompatible
polymer is comprised of at least one of the following: poly
glycolic acid, poly lactic acid, poly lactic co glycolic acid.
23. The bone graft material of claim 1 wherein the biocompatible
polymer is poly lactic co glycolic acid.
24. A method of fabricating a bone graft material, comprising the
steps of: selecting biocompatible polymer and salt particles from
about 100-200 .mu.m in diameter, selecting hydroxyapatite particles
less than 1 .mu.m in diameter mixing the biocompatible polymer and
hydroxy apatite and NaCl particles loading the mixture of particles
into a mold, compressing the mixture with a very high pressure;
exposing the newly formed composite scaffold to high pressure gas
decreasing the gas pressure on the composite scaffold until the
pressure returns to ambient pressure, soaking the composite
scaffold in an aqueous solution to dissolve and leach out the salt
particles, rinsing and drying the new composite biomaterial. The
bone graft material of claim 1 wherein the ratio of biocompatible
polymer to hydroxyapatite is from about 1:2 to about 2:1.
25. The method of claim 24 wherein the ratio of biocompatible
polymer to hydroxyapatite is about 1:1
26. The method of claim 24 wherein the ratio of biocompatible
polymer to NaCl ranges from 1:4 to 1:18.
27. The method of claim 24 wherein the ratio of biocompatible
polymer to NaCl is about 1:9.
28. The method of claim 24 wherein the biocompatible polymer is
selected from at least one of the following: poly glycolic acid,
poly lactic acid, poly lactic co glycolic acid.
29. The method of claim 24, wherein the salt is NaCl.
30. The method of claim 24 wherein the compression pressure is from
1000 to 4000 psi.
31. The method of claim 24 wherein the compression pressure is 2000
psi.
32. The method of claim 24 wherein the high pressure gas is at a
pressure from 400 to 1600 psi
33. The method of claim 24 wherein the high pressure gas is at a
pressure of 800 psi.
34. The method of claim 24 wherein the method further comprises
soaking the biomaterial in a simulated body fluid (SBF) solution
comprising NaCl, NaHCO3, Na2SO4, KCl, K2HPO4, MgCl2.6H2O, and
CaCl2.2H2O in water.
Description
SUMMARY OF THE INVENTION
[0001] This invention is a novel biomaterial that is especially
useful in tissue engineering applications involving bone. It is a
composite of poly(D,L-lactic-co-glycolic acid) (PLGA) and
nano-sized hydroxyapatite, wherein the hydroxyapatite is highly
exposed on the biomaterial surface. A further embodiment of this
invention involves a ceramic, such as apatite, that is fastly,
highly, and uniformly coated on the biomaterial surface. This new
biomaterial is advantageous because it promotes bone cell
propagation and ingrowth better than current materials.
BACKGROUND
[0002] The ideal bone graft would replace bone defects, such as
those from disease or trauma, with a material that allows bone
cells to grow into the affected area, thus restoring the bone to
its original condition. Currently, autografts are the best material
for bone repair because they are biocompatible and there is little
risk of disease transfer. However, the downside of autografts is
that a separate operation must be performed to remove the person's
own bone. Allografts, which consist of bone from another
person/cadaver, are also available but carry the risk of immune
response and disease transfer that could lead to ultimate
failure.
[0003] In order to solve the problems associated with bone grafts,
many researchers have tried to develop artificial substances for
bone grafts. These artificial biomaterials need to possess several
qualities in order to be successful. First, the material must be
degradable to allow room for new bone to grow into the implant
site. Second, it must maintain mechanical strength similar to
native bone. Finally, the artificial biomaterial needs to be
osteoconductive; that is, it must allow bone cells to attach and
propagate on its surface.
[0004] Some of the materials that have shown promise as bone grafts
include calcium phosphate ceramics such as hydroxyapatite and
tricalcium phosphate. These particular ceramics are quite
biocompatible because they have characteristics similar to native
bone mineral. However, they are hard to shape and do not possess
the same mechanical properties as bone. Hydroxyapatite in
particular does not degrade quickly either, which inhibits new bone
from forming.
[0005] Another type of material that has sparked some interest is
degradable polymer. It is easy to shape and it degrades at a
predictable rate, thereby allowing new bone growth to replace it.
Some examples of degradable polymers are poly(glycolic acid),
poly(L-lactic acid), and poly(D,L-lactic-co-glycolic acid).
Although they are easily formed and have good mechanical strength,
degradable polymers alone are not ideal for bone grafts because
they are not very osteoconductive. New bone will not attach well or
grow well into this material.
[0006] It is possible to make a composite using a phosphate ceramic
in conjunction with a degradable polymer. Small particles of
ceramic can be included within the polymer scaffold material. These
particles will be partially exposed on the surface of the
biomaterial, thereby making the material more osteoconductive.
[0007] Most related methods for making a polymer/ceramic scaffold
biomaterial use organic solvents. This can be highly
disadvantageous because some residual solvent may remain in the
material. Almost all organic solvents are detrimental to cell and
tissue growth. Also, it has been noted that these processes may
actually leave behind a thin film of polymer that coats the ceramic
particles that are supposed to be exposed on the surface. This
unintentional thin film disrupts the osteoconductive nature of the
ceramic portion of these biomaterials.
[0008] The invention disclosed herein addresses these problems by
describing a polymer/ceramic biomaterial comprised of degradable
polymer and ceramic wherein the ceramic is highly exposed on the
surface of the biomaterial and the biomaterial is fabricated with
no use of organic solvents. Furthermore, an additional layer of a
mineral, such as apatite, can be coated on the surface of the
biomaterial in an adherent, fast, uniform fashion. Finally,
granules of the polymer/ceramic biomaterial with additional ceramic
coating can be fabricated.
[0009] All references cited within this application are expressly
incorporated by reference in their entirety.
BRIEF SUMMARY OF THE INVENTION
[0010] A preferred embodiment of the present invention is a
biomaterial comprised of poly(D,L-lactic-co-glycolic acid) (PLGA),
hydroxyapatite, and a possible coating of apatite. It is suitable
as an artificial bone graft material. The said biomaterial is
formed using a gas foaming particle leaching (GF/PL) method. GF/PL
introduces gas bubbles into the polymer matrix by saturating the
polymer with gas at high pressure, and then reducing the pressure
back to ambient conditions. In this case CO2 gas is used. Then,
salt particles are leached out of the matrix using distilled water.
Both gas foaming and particle leaching leave behind voids, which
form the pores of this biomaterial matrix.
[0011] The preferred embodiment is made by combining particles of
PLGA, hydroxyapatite, and sodium chloride in certain ratios and
then using the GF/PL method. The size and amount of each particle
will determine the general and interconnected porosity of the final
biomaterial. Initially the PLGA, hydroxyapatite, and sodium
chloride particles are sieved to obtain particles with a specific
size. Then these particles are combined in certain ratios and
loaded into a disk mold. The mixture is compressed at a high
pressure (around 2000 psi) for about 1 minute. It is believed that
pressures as low as 1000 psi and as high as 4000 psi will also
work. The resulting disk is then exposed to high pressure CO2 gas
(around 800 psi) for 48 hours. It is believed that pressures as low
as 400 PSI and as high as 1600 psi will also work. During this
time, the CO2 saturates the disk. After 48 hours, the CO2 gas
pressure is decreased to ambient pressure. This process leads to
nucleation and growth of CO2 pores within the polymer scaffold
portions of the biomaterial disk. The sodium chloride particles are
subsequently leached out of the material by immersing the disk in
distilled water for a lengthy period of time, thus leaving voids
formerly occupied by the sodium chloride particles. The final
material is highly porous with hydroxyapatite particles exposed on
the surface of its polymer network.
[0012] Furthermore, coating the surface of the PLGA/hydroxyapatite
scaffold with a bone-like apatite using a biomimetic process can
increase its osteogenic potential. This biomimetic process involves
soaking the biomaterial in a solution of simulated body fluid (SBF)
that has appropriate concentrations of ions dissolved in solution.
Certain ions will precipitate on the surface of the biomaterial and
form an apatite mineral coating.
[0013] In another embodiment, the PLGA/hydroxyapatite biomaterial
may be ground up and sieved to collect granules with a certain
size. These granules may then be soaked in the SBF and receive the
apatite coating that enhances its osteogenic properties.
[0014] A bone graft material according to the present teachings can
be provided in the form of a bone paste, a shaped solid, or a dry
pre-mix useful for forming such a paste or solid. The phrase "bone
paste" refers to a slurry or semi-solid composition of any
consistency that hardens to form a solid structure, and thus
includes, e.g., bone plasters, putties, adhesives, cements, bone
void fillers, and bone substitutes. As a result, the bone paste can
be any composition capable of being injected, molded, painted,
suffused, or placed into contact with a bone surface in vivo. The
"shaped solid" can take any form, including a pellet that can be
placed into a bone void or into contact with a bone surface in
vivo. The dry pre-mix can be provided in the form of a powdered
and/or granular material.
BRIEF DESCRIPTION OF THE DRAWINGS
[0015] FIG. 1. XRD patterns of apatite-coated (A) PLGA and (B)
PLGA/HA scaffolds that were incubated in 1.times. and 5.times.SBF
for five days.
[0016] FIG. 2. FT-IR spectra of apatite-coated (A) PLGA and (B)
PLGA/HA scaffolds that were incubated in 1.times. and 5.times.SBF
for five days.
[0017] FIG. 3. The mass change of apatite-coated PLGA (inverted
triangle) and PLGA/HA (circle) scaffolds incubated in 1.times.
(solid) and 5.times.SBF (open) for various time periods. The mass
change was expressed as the percent increase compared to the mass
of scaffolds incubated in a tris-buffer solution for identical
times. The initial mass was identical for both types of scaffolds
with *P<0.05 compared to the other groups.
[0018] FIGS. 4 (A)-(C). Scanning electron micrographs of (A,C)
surfaces and (B,D) cross-sections of the PLGA/HA composite
scaffolds fabricated by (A,B) the SC/PL method and (C,D) the GF/PL
method.
[0019] FIGS. 5(A)-(B) Pore size distributions of the: (A) SC/PL and
(B) GF/PL scaffolds
[0020] FIG. 6(A) A macroscopic image of three types of
scaffolds
[0021] FIG. 6(B) Microscopic images of stained GF/FL-no HA
scaffold,
[0022] FIG. 6(C) Microscopic images of stained SC/PL scaffold,
[0023] FIG. 6(D) Microscopic images of stained GF/PL scaffold.
[0024] FIG. 6 (E) XPS analyses of SC/PL scaffold
[0025] FIG. 6(F) XPS analyses of GF/PL scaffold.
[0026] FIG. 6(G) The atomic percentages of calcium exposed to the
scaffold surface in GF/PL and SC/PL scaffolds.
[0027] FIGS. 7 (A)-(C) The (A) growth rate, (B) alkaline
phosphatase (ALP) activity, and (C) calcium deposition of the
osteoblasts cultured on GF/PL (lined bars), SC/PL (blank bars) and
GF/PL-no HA (solid bars) scaffolds for eight weeks in vitro.
[0028] FIGS. 7 (A)-(B) (A) Cell-seeded GF/PL scaffold prior to
implantation. The scale is in centimeters. (B) A gross view of
cell/scaffold constructs retrieved at eight weeks after
implantation to the subcutaneous spaces of athymic
[0029] FIGS. 9(A)-(D) Histological evaluations of cell/polymer
constructs retrieved at five weeks after implantation to the
subcutaneous spaces of athymic mice. (A,C) H&E staining and
(B,D) Masson's trichrome staining.
[0030] FIGS. 10(A)-(F) Histological evaluations of cell/polymer
constructs retrieved at eight weeks after implantation to the
subcutaneous spaces of athymic mice. (A,C,E) H&E staining and
(B,D,F) Masson's trichrome staining.
[0031] FIG. 11 (A)-(B) (A) Bone formation area and (B) calcium
deposition in GF/PL (lined bars), SC/PL (blank bars) and GF/PL-no
HA (solid bars) scaffolds at eight weeks after implantation.
[0032] FIG. 12(A) Schematic illustration of the exposure of HA
nanoparticles to the surface of scaffolds fabricated by the GF/PL
and SC/PL methods.
[0033] FIG. 12(B) Schematic illustration of the fabrication of
PLGA/HA composite scaffolds with the GF/PL method.
[0034] FIGS. 13(A)-(B). Scanning electron micrographs of the
PLGA/HA composite scaffolds fabricated with (A) the SC/PL method
and (B) the GF/PL method.
[0035] FIGS. 14(A)-(C) The compressive modulus (A) and tensile
modulus (C) of the SC/PL and GF/PL scaffolds. (B) Typical tensile
stress strain curve of the SC/PL and GF/PL scaffolds.
[0036] FIGS. 15(A)-(E) Macroscopic images of the three types of
scaffolds FIGS. 16(A)-(C) Photographs showing wettability of the
three types of scaffolds.
[0037] FIG. 17 (A) Photograph showing defect created in rat
cranium.
[0038] FIG. 17(B) Photograph showing the implanted scaffold
[0039] FIG. 17 (C) Gross view of the cranium containing the GF/PL
HA scaffold.
[0040] FIG. 17(D) Gross view of the cranium containing the GF/PL no
HA scaffold
[0041] FIGS. 18(A)-(F) Histological evaluations of specimens
retrieved at 8 weeks after implantat showing Hematoxylinandeosin
(H&E) staining (A,D) GF/PL-no HA scaffold, (B,E) SC/PL
scaffold, and (C,F) GF/PL scaffold at 8 weeks. (A, B, and C) Defect
edge and (D, E, and F) midsection of the scaffolds.
[0042] FIGS. 19(A)-(F) Histological evaluations of specimens
retrieved at 8 weeks after implantation. Masson's trichrome
staining. (A,D) GF/PL-no HA scaffold, (B,E) SC/PL scaffold, and
(C,F) GFIPL scaffold at 8 weeks. (A,B, and C) Defect edge and (D,
E, and F) mid-section of the scaffolds (blank arrow, original bone;
solid arrow, new bone).]
[0043] FIG. 20 Graph showing higher bone formation in GF/PL
scaffolds compared to SC/PL and OF/PL no HA scaffolds.
[0044] FIGS. 21(A)-(C) Images of micro CT analysis of specimens
retrieved 8 weeks after implantation. A-GF/PL no HA; B SC/PL and C
GF/PL scaffold.
DESCRIPTION OF THE INVENTION
[0045] The present invention is a novel biomaterial with special
characteristics that allow it to perform well as a bone graft
material. It is comprised of a degradable
poly(D,Llactic-co-glycolic acid) polymer scaffolding with
incorporated, nano-sized hydroxyapatite particles made by a gas
foaming and particle leaching (GF/PL) method. A further embodiment
of this invention describes the same biomaterial with an adherent,
highly uniform apatite coating.
[0046] The method of constructing a PLGA polymer scaffold using
GF/PL is described thoroughly in the journal article titled, "Open
pore biodegradable matrices formed with gas foaming" (Harris L D,
Kim B S, and Mooney D J; J Biomed Mater Res, 42, 396-402, 1998).
This entire article is hereby incorporated by reference. The
research reported in this article found that the porosity and pore
size of the PLGA scaffold can be controlled by the salt/PLGA ratio
and respective particle sizes. Also, the pores of the matrix are
interconnected and highly uniform. In this manner, a useful
scaffold can be created without the use of organic solvents or high
temperatures.
[0047] Although the method for constructing a polymer scaffold
using GF/PL is already known, the addition of nano-sized
hydroxyapatite particles to this specific polymer scaffolding has
not been taught in the prior art. An article has recently been
published by the inventor of this application. It describes the
addition of hydroxyapatite particles to the PLGA scaffold. It is
called, Poly(lactide-co-glycolide)/hydroxyapatite composite
scaffolds for bone tissue engineering (Kim S S, Park M S, Jeon O,
Choi C Y, and Kim B S; Biomaterial, 27, 1399-1409, available online
Oct. 5, 2005). This article is also hereby incorporated by
reference. It is important to note that this article is not by
another. Students of Dr. Byung-Soo Kim, the sole inventor of the
present invention, conducted the research for the article, but the
ideas for the invention are uniquely those of Dr. Byung-Soo Kim
[0048] The research describes how porous PLGA/HA composite
scaffolds were fabricated by the modification of the previously
described GF/PL method (Harris L D, Kim B S, and Mooney D I; J
Biomed Mater Res, 42, 396-402, 1998). The most significant
modification of the previous method is the non-obvious addition of
nano-sized hydroxyapatite particles that end up being highly
exposed on the polymer surface. It is not obvious to one of
ordinary skill in the art to add nano-sized hydroxyapatite
particles to the polymer matrix and to expose the hydroxyapatite
particles on the polymer surface.
[0049] PLGA/HA composites were prepared with 75:25 PLGA particles
(diameter=100-200 Pm, molecular weight=100,000 Da, Birmingham
Polymers, Birmingham, Ala.), HA nanoparticles
(diameter=approximately 100 nm, Berkeley Advanced Biomaterials
Inc., Berkeley, Calif.), and sodium chloride particles
(diameter=100-200 Pm, Sigma, St. Louis, Mo.). The PLGA pellets were
ground using a Tekmar grinder (Bel-Art Products, Pequannock, N.J.)
and sieved to obtain particles ranging from 100 to 200 Pm. The salt
particles were sieved to yield a range of sizes from 100 to 200 Pm.
The polymer particles were mixed with the HA and NaCl particles.
The PLGA/HA/NaCl mass ratio was 1:1:9. The mixture was loaded into
a disk mold (diameter=1.35 cm; Aldrich Chemical Co., Milwaukee,
Wis.) and compressed at 2000 psi for 1 min using a Carver
Laboratory Press (Fred S. Carver, Inc., Menominee Falls, Wis.) to
yield solid disks with a thickness of about 1.7 mm. The samples
were then exposed to high pressure CO2 gas (800 psi) for 48 hours
to saturate the polymer with the gas. Then, decreasing the gas
pressure to ambient pressure created a thermodynamic instability.
This led to the nucleation and growth of CO2 pores within the
polymer scaffolds. The sodium chloride particles were subsequently
removed from the scaffolds by leaching the scaffolds in distilled
water for 48 hours.
[0050] While the above materials are preferred, it is believed that
the present invention will work with polymers of diameters from
50-400 Pm and with HA particles having diameters of 50-200 Pm and
salt particles having diameters from 50-300 Pm. It is believed that
the ratios of polymer to hydroxy apatite to NaCL can vary by as
much as 50% without deviating from the spirit of this
invention.--
[0051] The process for creating these PLGA/HA composite
biomaterials can be summarized by the steps of: (1) grinding PLGA
to small particles, (2) sieving the PLGA and sodium chloride
particles to yield particles with a 100-200 Pm diameter, (3) mixing
the particles PLGA/HA/NaCl in a mass ratio of 1:1:9, (4) loading
the mixture of particles into a disk mold, (5) compressing the
mixture with a very high pressure for a certain amount of time, (6)
exposing the newly formed disk to high pressure CO2 gas long enough
to saturate the disk, (7) decreasing the pressure on the disk until
it returns to ambient pressure, (8) soaking the disk in distilled
water to dissolve and leach out the sodium chloride particles.
[0052] A further embodiment of the invention involves forming a
uniform mineral coating of apatite on the surface of the
PLGA/hydroxyapatite biomaterial. This apatite layer enhances the
osteogenic potential of the biomaterial scaffold.
[0053] The apatite layer is created by incubating the scaffolds in
an ion rich simulated body fluid (SBF) solution. The solution is
prepared by dissolving reagent grade NaCl, NaHCO3, Na2SO4, KCl,
K2HPO4, MgCl2.6H2O, and CaCl2.2H2O in distilled deionized water.
1.times.SBF has the same ion concentrations as blood plasma while
5.times.SBF has ion concentrations five times greater than blood
plasma. The pH is adjusted to 6.4 with
tris(hydroxymethyl)aminomethane.
[0054] The described PLGA/hydroxyapatite biomaterial can be coated
with apatite relatively quickly because the exposed hydroxyapatite
particles act as nucleation sites for the growth of the mineral
apatite layer in SBF solution. Although the method for coating of
polymeric biomaterial with apatite by incubating the biomaterial in
SBF solution is already known, accelerated coating by incubating
polymeric biomaterial with nano-hydroxyapatites exposed on the
biomaterial surface has not been taught in the prior art. A study
was conducted to compare the formation of an apatite layer on both
PLGA and PLGA/nanohydroxyapatite scaffolds created by the GF/PL
method. Each was incubated in 1.times. and 5.times.SBF for up to
five days. A series of brief evacuation-repressurization cycles
were performed to force the solution into the pores of the
scaffold. The SBF was refreshed every day. After various incubation
times the samples were removed, rinsed, and dried in vacuum before
being characterized.
[0055] The PLGA and PLGA/nano-hydroxyapatite specimens were
characterized. The morphologies of the scaffolds were examined by
scanning electron microscopy (SEM; JSM6330F, JEOL, Tokyo, Japan)
after platinum coating. X-ray diffraction (XMD) spectra were
obtained using an X-ray diffractometer (D/MAX-2500, Rigaku Co.,
Tokyo, Japan) with a mixed incidence of 1.degree. and a 2.theta.
scanning rate of 2.5.degree./min in the range of 10-60.degree.. Cu
K.alpha. radiation, with a voltage of 40 kV and a current of 100
mA, was used for the diffraction. The XRD results are shown in FIG.
1. It is apparent that the PLGA/hydroxyapatite results show the
higher intensity peaks expected from greater apatite formation on
the PLGA/hydroxyapatite composite biomaterial.
[0056] Fourier transformed infrared spectroscopy (FT-IR) spectra
were obtained using a FT-IR spectrometer (Avatar 360, Nicolet
Instrument Corp., Madison, Wis., USA) with a resolution of 8 cmi1.
For FT-IR analysis, the scaffolds were cut into fine particles,
milled with potassium bromide, and pressed into transparent thin
discs. The scaffold mass increase during apatite formation in SBF
was measured using an analytical balance accurate to 10-4 g
(EPG214C, Ohaus Corp., Pine Brook, N.J., USA), and the data is
shown in FIG. 2.
[0057] Finally, the scaffolds were air-dried and then vacuum dried.
The mass increase from apatite formation was expressed as a percent
increase compared to the scaffold mass when incubated in a
tris-buffer at the same pH value, at the same temperature, and for
the same time intervals. The data in FIG. 3 shows that the
PLGA/hydroxyapatite scaffolds (circles) gained the greatest mass
due to apatite formation. Furthermore, it is also evident that SBF
solutions with higher ion concentrations lead to greater apatite
deposition.
[0058] SEM micrographs of the apatite coated PLGA samples revealed
that the apatite layer was not uniform. There were areas of bare
PLGA, which shows that there was poor apatite deposition. However,
the PLGA/hydroxyapatite scaffolds showed a more desirable, uniform
apatite layer.
[0059] More significantly, the apatite-coated PLGA/hydroxyapatite
scaffolds exhibited noticeably improved cell growth and
mineralization, when seeded with osteoblast cells, compared with
apatite-coated PLGA scaffolds in vitro. This result supports the
hypothesis that the uniform apatite layer is favorable for
osteogenic properties.
[0060] The biomimetic apatite coating process is enhanced by
introducing nano-sized hydroxyapatite nucleation sites and by using
concentrated SBF solution. This coating is advantageous because it
conveys better osteogenic properties to the PLGA/hydroxyapatite
biomaterial.
[0061] A further embodiment of this invention involves the coating
of PLGA/nanohydroxyapatite particles (rather than scaffolds) with a
biomimetic, adherent, and uniform apatite coating. The particles
may be the product of a reaction process or be ground down from
bulk PLGA/nano-hydroxyapatite composite to a size of 30-2000 Pm.
The particles will be sieved to isolate particles with a more
narrow size distribution depending on the desired application.
These particles will then be soaked in SBF solution to coat them
with a uniform layer of biomimetic apatite.
[0062] Most of the previous methods for fabricating
polymer/bioceramic composite scaffolds, such as the solvent casting
and particulate leaching (SC/PL) method or the phase separation
method, use organic solvents. However, residual solvents in the
scaffolds may be harmful to transplanted cells or host tissues.
Furthermore, the polymer coating on the ceramics created by polymer
solutions may hinder the exposure of the ceramics to the scaffold
surfaces (FIG. 4A), which could decrease the chance that osteogenic
cells make contact with the bioactive ceramics.
[0063] The preferred embodiment of the present invention relies on
gas forming and particulate leaching (GF/PL) methods to fabricate
PLGA/HA composite scaffolds for bone tissue engineering. This
method efficiently exposes the bioceramic on the scaffold surfaces
and avoids the use of organic solvents. To reduce the amount of HA
(which degrades extremely slowly in vivo) required, and to increase
the HA exposure to the scaffold surface, HA particles approximately
100 nm in size rather than micro-sized particles, are used to
fabricate the composite scaffolds. The HA exposure at the scaffold
surface in GF/PL scaffolds was compared to that in SC/PL scaffolds
see FIGS. 4 A-C.
EXAMPLE 1
[0064] Porous PLGA/HA composite scaffolds were fabricated by the
modification of a previously described GF/PL method of 24. Harris L
D, Kim B S, Mooney D J. Open pore biodegradable matrices formed
with gas foaming. J Biomed Mater Res 1998; 42: 396-402. PLGA/HA
composites were prepared with 75:25 PLGA particles
(diameter=100-200 mm, molecular weight=100,000 Da, Birmingham
Polymers, Birmingham, Ala.), HA nanoparticles
(diameter=approximately 100 nm, Berkeley Advanced Biomaterials
Inc., Berkeley, Calif.), and sodium chloride particles
(diameter=100-200 mm, Sigma, St. Louis, Mo.). The PLGA pellets were
ground using a Tekmar grinder (Bel-Art Products, Pequannock, N.J.)
and sieved to obtain particles ranging from 100 to 200 mm. The salt
particles were sieved to yield a range of sizes from 100 to 200 mm.
The polymer particles were mixed with the HA and NaCl particles.
The PLGA/HA/NaCl mass ratio was 1:1:9. The mixture was loaded into
a disk mold (diameter=1.35 cm; Aldrich Chemical Co., Milwaukee,
Wis.) and compressed at 2000 psi for 1 min using a Carver
Laboratory Press (Fred S. Carver, Inc., Menominee Falls, Wis.) to
yield solid disks with a thickness of 1.7 mm. The samples were then
exposed to high pressure CO2 gas (800 psi) for 48 h to saturate the
polymer with the gas. Then, decreasing the gas pressure to ambient
pressure created a thermodynamic instability. This led to the
nucleation and growth of CO2 pores within the polymer scaffolds.
The NaCl particles were subsequently removed from the scaffolds by
leaching the scaffolds in distilled water for 48 h. PLGA scaffolds
without HA were also fabricated by the GF/PL method and used as a
control (GF/PL-no HA).
[0065] Porous PLGA/HA scaffolds were also fabricated by the
modification of a previously described SC/PL methods of Wei G, Ma P
X. Structure and properties of nano-hydroxyapatite/polymer
composite scaffolds for bone tissue engineering. Biomaterials 2004;
25:4749-57; Cho S W, Kim I K, Lim S H, Kim D I, Kang S W, Kim S H,
et al. Smooth muscle-like tissues engineered with bone marrow
stromal cells. Biomaterials 2004; 25:2979-86; Cho S W, Kim S S,
Rhie J W, Cho H M, Choi CY, Kim B S. Engineering of volume-stable
adipose tissues. Biomaterials 2005; 26:3577-85; Kim B S, Jeong S I,
Cho S W, Nikolovski J, Mooney D J, Lee S H, et al. Tissue
engineering of smooth muscle under a mechanically dynamic
condition. J Microbiol Biotech 2003; 13:841-5, and were used as
another control. In this process, PLGA was dissolved in methylene
chloride (J. T. Baker, Phillipsburg, N.J.) at a 10% (w/v)
concentration, and HA and NaCl were added to the PLGA solution at
the same sizes and ratios as for the GF/PL scaffolds. This mixture
was loaded into Teflon cylinders (diameter=21.5 mm, height=25 mm;
Cole-Parmer Instrument Company, Vernon Hills, Ill.). Following
solvent evaporation, the polymer disks with entrapped salt
particles were removed from the molds. The salt was removed by
immersing disks in distilled water for 48 h.
[0066] The porosity of fabricated scaffolds was measured using
mercury intrusion porosimetry (Autopore IV 9500, Micromeritics
Instrument Corporation, Norcross, Ga.). A contact angle of 1301 for
mercury on the scaffold was used for this analysis. The pore
structures of the scaffolds were examined using a scanning electron
microscope (SEM, JEOL, Tokyo, Japan). Compression and tensile tests
were performed with an Instron mechanical tester (Instron 4201,
Instrons, Canton, M A). The scaffold samples were cut into
1.times.1 cm2 for compression testing. For tensile testing, the
samples (1.times.1 cm2) were attached to cardboard using epoxy
glue. The sample was centered in a 7 mm slot in the center of the
cardboard and then glued to standardize the gauge length.
Compression and tensile tests were performed with a constant strain
rate of 1 mm/min. The moduli were determined from the slopes in the
initial elastic portion of the stress-strain diagram. To examine
the distribution and extent of surface exposure of HA in the
scaffolds, the HA exposed to the scaffold surface was visualized
with a hydrophilic dye (trypan blue, Sigma) staining. The residual
dye was removed by sonication in 100% ethanol. Afterwards, the
surface of the PLGA/HA scaffolds was examined with a microscope
(Camseope, Samtech, Seoul, Korea). To examine the chemical
composition of the scaffold surface, we carried out X-ray
photoelectron spectroscopic (XPS; Sigma Probe, ThermoVG Scientific,
West Sussex, UK) analyses, evaluating the O 1s, C 1s, Ca 2p, and P
2p peaks. The residual pressure in the spectrometer was
1.1.times.10.sup.-8 Pa, and a Mg anode (1.25 keV) powered at 250 W
was used as an X-ray source. The constant pass energy was 23 eV.
All XPS data were acquired at a nominal photoelectron takeoff angle
of 551. The area of the XPS peaks was determined after background
subtraction, and the atomic percentage was determined by
normalizing the peak area of each element by the total peak areas
of all elements.
[0067] Osteoblasts were isolated from the calvaria of neonatal
(less than one day old) Sprague-Dawley rats (SLC, Tokyo, Japan) by
an enzymatic digestive process. The calvaria were isolated, and all
connective tissues were carefully removed. The parietal bones were
minced into pieces measuring about 1.times.1 mm.sup.2 using sterile
surgical scissors. Osteoblasts were isolated by an enzyme solution
containing 1.37 mg/ml collagenase type I (Sigma) and 0.5 mg/ml
trypsin (Sigma). Following 30 min of incubation, the released cells
were discarded to prevent contamination with other cell types. The
minced bones were redigested with the enzyme solution for 30 min,
and the supernatant was transferred to the culture medium,
Dulbecco's Modified Eagles Medium (DMEM, Gibco BRL, Gaithersburg,
Md.) containing 10% (v/v) fetal bovine serum (Gibco BRL), 1% (v/v)
penicillin-streptomycin (Gibco BRL), 10 mM b-glycerophosphate
(Sigma), 50 mg/ml L-ascorbic acid (Sigma), and 100 nM dexamethasone
(Sigma). This process was repeated three times, and then finally
the collected solution was centrifuged for 10 min at 1500 rpm.
Cells were plated into tissue culture flasks and cultured in a
humidified incubator at 37.degree. C. with 5% (v/v) CO.sub.2.
[0068] The fabricated scaffolds were sterilized by ethylene oxide
gas and pre-wetted in the culture medium for 12 h. Aliquots of 50
ml of the cell suspension (4.0.times.10.sup.7 cells/ml,
2.0.times.10.sup.6 cells/scaffold) were seeded onto the tops of the
pre-wetted scaffolds. The scaffolds were left undisturbed in an
incubator for 3 h to allow the cells to attach to the scaffolds. An
additional 1 and 10 ml of culture medium were added to each
scaffold at 6 and 8 h, respectively. The cell/scaffold constructs
were cultured in a humidified incubator at 37.degree. C. with 5%
(v/v) CO.sub.2 for eight weeks. The medium was changed everyday.
Analytical assays were performed at 7, 14, 28, and 56 days.
[0069] To determine the seeding efficiency and cell growth on the
scaffolds, cell numbers were determined by quantitative DNA assays
(n=3). DNA was isolated using a Wizard Genomic DNA Purification kit
(Promega, Madison, Wis.). For DNA isolation, the cell/scaffold
constructs were washed twice with phosphate-buffered saline. The
specimens were placed in a 1.5-ml tube and crushed with a
homogenizer (PowerGen 125, Fisher Scientific, Germany). DNA was
isolated according to the kit protocol, and DNA content was
measured with an ultraviolet absorbance spectrophotometer (JASCO
V-530, Tokyo, Japan) at 260 nm. The cell numbers were calculated
from a DNA standard curve of identical cells.
[0070] The alkaline phosphatase (ALP) production of osteoblasts
cultured on scaffolds was measured spectroscopically (n=3) using
the methods of Ekholm M, Hietanen J, Tulamo R M, Muhonen J,
Lindqvist C, Kellomaki M, et al. Tissue reactions of subcutaneously
implanted mixture of epsilon-caprolactone-lactide copolymer and
tricalcium phosphate. An electron microscopic evaluation in sheep.
J Mater Sci Mater Med 2003; 14:913-8. The osteoblast/scaffold
constructs were washed with PBS, homogenized with 1 ml Tris buffer
(1 M, pH 8.0, Sigma), and sonicated for 4 min on ice. Aliquots of
20 ml were incubated with 1 ml of a p-nitrophenyl phosphate
solution (16 mM, Sigma) at 30 lC for up to 5 min. The production of
p-nitrophenol in the presence of ALP was measured by monitoring
light absorbance at 405 nm.
[0071] The amount of calcium deposited in the cell-scaffold
constructs was measured using a previously reported method (n=3) of
Jaiswal N, Haynesworth S E, Caplan A I, Bruder S P. Osteogenic
differentiation of purified, culture-expanded human mesenchymal
stem cells in vitro. J Cell Biochem 1997; 64:295-312. After the
cell-scaffold constructs were rinsed twice with PBS and homogenized
with 0.6 N HCl, calcium was extracted by shaking for 4 h at
4.degree. C. The lysate was then centrifuged at 1000 g for 5 min,
and the supernatant was used to determine calcium content. To
measure the amount of calcium produced by the seeded osteoblasts,
the calcium content of the PLGA/HA scaffold itself was also
measured, and the calcium content of the scaffold itself was
subtracted from the total calcium content of the lysate. The
calcium concentration in the cell lysates was quantified
spectrophotometrically with cresolphthalein complexone (Sigma).
Three minutes after the addition of reagents, the absorbance of the
samples was read at 575 nm using a microplate reader (Multiskan
Spectrum, Thermo Electron Co., Vantaa, Finland). The calcium
concentration was calculated from a standard curve generated from a
serial dilution of a calcium standard solution (Sigma).
[0072] The surface and cross-sectional morphologies of the
scaffolds and cell-scaffold constructs were examined using a SEM.
The samples were washed twice with PBS, prefixed in 1% (v/v)
buffered glutaraldehyde for 1 h, and fixed in 0.1% (v/v) buffered
formaldehyde for 24 h. The fixed samples were dehydrated in
ascending grades of ethanol, dried, and mounted on aluminum stubs
using double-sided carbon tape. The specimens were coated with gold
using a Sputter Coater (Cressington 108, Cressington Scientific
Instruments, Cranberry, Pa.) and examined with SEM at an
acceleration voltage of 10 kV.
[0073] In addition to the culture of cell-scaffold constructs in
vitro, cell scaffold constructs were implanted into the
subcutaneous space of athymic mice (BALB/c-nu, 7 weeks old, female,
SLC, Tokyo, Japan). After the mice were anesthetized with an
intramuscular administration of ketamine hydrochloride (50 mg/kg,
Yuhan Co., Seoul, Korea) and xylazine hydrochloride (5 mg/kg, Bayer
Korea Ltd., Seoul, Korea), small incisions were made on the dorsal
skins of six mice. Four pouches per animal were made by blunt
dissection in subcutaneous sites, and cell-seeded scaffolds were
immediately implanted into the pouches (n=4). Subsequently, the
skin was closed with 5-0 Vicryl sutures (Ethicon, Lenneke Marelaan,
Belgium). The mice were housed singly after surgery and received
humane care in compliance with the Hanyang University Guidelines
for the care and use of laboratory animals. The implants were
retrieved for analysis at five and eight weeks after
implantation.
[0074] Cell-scaffold constructs were retrieved from athymic mice at
five and eight weeks after implantation (FIG. 8B), fixed in 10%
(v/v) buffered formaldehyde, dehydrated in ascending grades of
ethanol, and embedded in paraffin. The tissue blocks were sectioned
at a 4-mm thicknesses and stained with hematoxylin and eosin
(H&E) and Masson's trichrome. The Masson's trichrome-stained
mid-portion sections were examined with a microscope for
histomorphometry. The percentage of bone occupying space within the
constructs was measured using an image analysis system (KS400,
Zeiss, Munich, Germany) coupled to a light microscope. The bone
formation area was expressed as the percentage of bone area in the
available pore space (bone area/pore area.times.100%).
[0075] Quantitative data were expressed as the mean standard
deviation. Statistical comparisons were carried out using analysis
of variance (ANOVA, SAS Institute Inc., Cary, N.C.). A value of
p<0:05 was considered to be statistically significant.
[0076] Gas foaming and the subsequent salt leaching of scaffolds
containing a high percentage (90%) of NaCl particles (diameter
range 100-200 mm) led to the formation of highly porous, open pore
structures with no evidence of an external, nonporous skin layer
(FIGS. 4C and D). The pore structure observed in the cross-sections
of the GF/PL scaffolds was similar to that of the scaffolds
fabricated by the SC/PL method (FIGS. 4A and B). The SC/PL method
produced scaffolds with pore sizes of approximately 100-200 mm
(FIG. 5A). In contrast, the GF/PL process resulted in scaffolds
with two levels of porosity: interconnected macropores (100-200 mm)
were created by the leaching of the NaCl particles, and smaller,
closed pores (10-45 mm) were created by the nucleation and growth
of gas pores within the polymer particles (FIG. 5B). The average
porosities of the GF/PL and SC/PL scaffolds were 91.+-.3% and
85.+-.3% respectively.
[0077] The mechanical properties of the scaffolds were assessed
using compressive and tensile mechanical tests. The GF/PL scaffolds
exhibited enhanced mechanical properties as compared to the SC/PL
scaffolds. The average compressive moduli were 2.3.+-.0.4 and
4.5.+-.0.3 MPa (p<0:05) for the SC/PL and GF/PL scaffolds,
respectively. The average tensile moduli were 2.0.+-.0.1 and
26.9.+-.0.2 MPa (p<0:05) for the SC/PL and GF/PL scaffolds,
respectively. These data represent a 99% increase in the
compression modulus and a 1331% increase in tensile modulus,
demonstrating the positive effects of the GF/PL fabrication process
in enhancing the mechanical properties of the scaffolds.
[0078] To determine whether the scaffold fabrication process
affects the extent of HA exposure at the scaffold surface, the
exposed HA was stained with a hydrophilic dye. HA was stained more
abundantly in the GF/PL scaffolds than in the SC/PL scaffolds
(FIGS. 6A, C and D). The surface composition of the PLGA/HA
composite scaffolds was also analyzed with XPS. The amount of Ca in
the GF/PL scaffold surface was greater than in the SC/PL scaffold
surface (FIGS. 6E and F). The atomic ratio of the Ca exposed on the
scaffold surface was 156% higher on the GF/PL scaffold surface
compared with the SC/PL scaffold surface (FIG. 6G).
[0079] Both types of the PLGA/HA composite scaffolds allowed for
the adhesion and proliferation (FIG. 7A) of the seeded rat
calvarial osteoblasts over the 56-day in vitro culture period. The
initial cell seeding density of 2.00.times.10.sup.6 cells/scaffold
resulted in 1.33.times.10.sup.6 cells/scaffold remaining attached
to the GF/PL scaffold after one day in culture, giving an adhesion
percentage of 66.5%. For the SC/PL scaffold, the cell adhesion
efficiency was 62.0%. Osteoblasts grew more rapidly in the GF/PL
scaffolds than in the SC/PL scaffolds (FIG. 7A). The average cell
density of the GF/PL scaffolds was 2.48.times.10.sup.6
cells/scaffold after four weeks in culture, while that of the SC/PL
scaffolds was 2.19.times.10.sup.6 cells/scaffold, corresponding to
86.5% and 69.7% increases in cell density for the GF/PL and SC/PL
scaffolds, respectively (FIG. 8).
[0080] The ALP activity of the osteoblasts cultured on both types
of PLGA/HA composite scaffolds increased during the four-week
culture period and decreased at eight weeks (FIG. 7B). In contrast,
the ALP activity of the osteoblasts grown on the PLGA scaffolds
without HA was low and did not show significant changes during the
culture period. The osteoblasts on the GF/PL scaffolds showed
significantly higher (p<0:05) levels of ALP activity compared to
the osteoblasts on the SC/PL scaffolds during the first four weeks
of culturing, but showed no significant differences at eight
weeks.
[0081] The calcium deposition by cultured osteoblasts was
significantly higher (p<0:05) on the GF/PL scaffolds than on the
SC/PL scaffolds during the 8-week culture period (FIG. 7C). The
deposition on both types of the PLGA/HA scaffolds gradually
increased during the culture period. On the PLGA scaffolds without
HA, calcium deposition was significantly lower than on both types
of the PLGA/HA scaffolds. The calcium deposition on the PLGA
scaffolds remained constant at low levels for the first four weeks,
and increased slightly at eight weeks.
[0082] The implantation of both types of the osteoblast-seeded
PLGA/HA composite scaffolds resulted in new bone formation in vivo
in ectopic sites at five and eight weeks after implantation. Five
weeks after implantation, a small amount of woven bone was detected
in both the SC/PL (FIGS. 9A and B) and the GF/PL (FIGS. 9C and D)
scaffolds. Eight weeks after implantation, osteogenesis had
progressed, and more bone with lamellar structures appeared (FIG.
10C-F). Histomorphometric analyses of the mid-portion sections of
the regenerated tissues showed enhanced bone formation in the GF/PL
scaffolds, compared with the SC/PL scaffolds and the PLGA scaffolds
with no HA, at five and eight weeks after implantation (FIG. 11A).
In contrast, the cell-seeded PLGA scaffolds with no HA had produced
nearly no new bone in vivo for eight weeks. Most of the pores of
the PLGA scaffolds with no HA were filled with loose fibrous
connective tissues without evidence of bone formation at five and
eight weeks after implantation (FIGS. 10A and B).
[0083] The calcium deposition in the regenerated tissues was much
more extensive in the GF/PL scaffold group than in the SC/PL
scaffold group at five and eight weeks after implantation (FIG.
11B), although the calcium deposition in both groups increased with
the implantation period. The calcium deposition in the PLGA
scaffold group with no HA was far less than that in both
HA-containing scaffold groups.
[0084] The PLGA/HA scaffolds fabricated by the GF/PL method
exhibited a higher exposure of HA at the scaffold surface and much
better bone formation in vitro and in vivo than those fabricated by
the conventional SC/PL method. As compared with other methods for
fabricating biodegradable polymer/ceramic composite scaffolds, the
GF/PL method has a number of advantages.
[0085] First, the GF/PL process avoids the use of organic solvents.
Residual organic solvents remaining in scaffolds may damage
transplanted cells and surrounding tissues. Furthermore, exposure
to organic solvents may inactivate biologically active factors.
Therefore, the GF/PL process may cause less denaturation of the
growth factors incorporated within the scaffolds.
[0086] Second, the GF/PL method can efficiently expose bioceramics
at the surface of the polymer/bioceramic composite scaffolds.
Staining with a hydrophilic dye and XPS analysis showed that the
GF/PL method exposed a significantly higher extent of HA at the
scaffold surface than did the conventional SC/PL method in this
study (FIG. 6). The SC/PL method causes the polymer coating on the
bioceramics by polymer solutions, which hinders the exposure of
bioceramics on the scaffold surfaces, while the 5 weeks 8 weeks
GF/PL method, which does not use a polymer solution, efficiently
exposes the bioceramics on the scaffold surface. Therefore, a GF/PL
scaffold can increase the chances of osteogenic cells to make
contact with the bioactive ceramics, which enhances osteoblast
differentiation and growth.
[0087] Third, the GF/PL scaffolds exhibit enhanced mechanical
properties as compared to the SC/PL scaffolds. The GF/PL scaffolds
had significantly higher compressive and tensile moduli than the
SC/PL scaffolds. This might be due to a closer packing of the
polymer chains under the high pressure in the GF/PL process and to
be tensile alignment of the polymer chains by the polymer
elongation that occurs during the foaming. In addition, the
residual solvent in the SC/PL scaffolds may function as a
plasticizer and make the polymer more ductile. Although the GF/PL
composite scaffolds showed greatly enhanced mechanical properties,
as compared with the SC/PL composite scaffolds, the measured
compressive moduli of the prepared scaffolds is rather low compared
to that of human bone. This might be due to the highly porous
structure of the fabricated scaffolds and the poor mechanical
properties
[0088] The use of GF/PL scaffolds resulted in enhanced osteogenic
potentials in both in vitro and in vivo experiments. Since the
SC/PL and GF/PL scaffolds have similar physical properties such as
porosity, pore size, and interconnectivity, the difference in
osteogenic ability between the two scaffold types might be due to
their different surface chemistries. Enhanced bone formation in
vitro and in vivo on the GF/PL scaffolds may result from the direct
contact of seeded cells with the HA particles exposed on the
scaffold surface, which stimulate the cell proliferation and
osteogenic differentiation. In contrast, the HA particles would be
coated with the polymer, which would hinder interaction with cells
in the SC/PL scaffolds. Since the SC/PL scaffold has a large
portion of HA particles buried in PLGA polymer, the degradation of
PLGA will expose HA particles on its surface. However, the PLGA
degradation requires a long time and there will be no acceleration
of bone formation by HA during this period.
[0089] The ALP activity on the PLGA scaffold without HA did not
show any significant changes during the culture period in vitro,
but the calcium concentration increased at 56-days in this group.
This disparity could be due to the fact that ALP is an early marker
for osteogenic differentiation and usually peaks early, while
mineralization occurs continuously over the in vitro culture
period.
[0090] In this study, we used nano-sized HA particles to fabricate
PLGA/HA composite scaffolds. Since the highly crystalline HA
degrades over long periods of time in vivo, the incompletely
degrading residual HA may hinder or slow the complete bone healing.
Therefore) to reduce the total amount of HA while enhancing the HA
distribution on the scaffold surface, we used nano-sized HA
particles, which have a high surface area, to fabricate the
composite scaffolds, instead of micro-sized HA particles.
Furthermore, the nano-sized HA particles showed improved
bioactivity and osteointegration when implanted in the bone defect
sites, as compared with the micro-sized HA particles. It has been
also reported that protein adsorption and cell adhesion can be
enhanced by using nano-sized HA particles instead of micro-sized HA
particles.
EXAMPLE 2
[0091] Increasing interest has currently been focused on
polymer/ceramic composite materials as bone substitutes because
these materials have advantages over ceramic scaffolds and polymer
scaffolds for bone tissue engineering. Calcium phosphate-based
ceramics, such as hydroxyapatite (HA) and tricalcium phosphate,
have been used as bone substitutes, but these materials have poor
mechanical performance. Most synthetic polymer biomaterials have
low surface wettability due to their composition of noncharged
elements. Such hydrophobic surfaces are unfavorable to osteogenic
cells as they show a lower proliferative and a higher apoptotic
rate on hydrophobic surfaces than on hydrophilic surfaces. In
addition, these polymeric biomaterials have a bioinert surface that
lacks bioactive functions for bone formation, therefore evoking
minimal tissue responses. An essential requirement for bone grafts
is the ability to create a bond with the living host bone through
the formation of a biologically active bonelike apatite layer on
the surface of the bone grafts. Bioinert bone substitutes often
become encapsulated with fibrous tissues, thereby resulting in
disturbed bone formation. Therefore, the addition of calcium
phosphate ceramics to biodegradable polymers, such as poly(glycolic
acid), poly(L-lactic acid), and poly(D,L-lactic-co-glycolic acid)
(PLGA), would allow for better surface wettability as well as
enhanced osteoconductivity.
[0092] Most of the previously available methods for the fabrication
of polymer/ceramic composite scaffolds, such as the solvent casting
and particulate leaching (SC/PL) method and the phase separation
method, use organic solvents. However, residual solvents in the
scaffolds may be harmful to transplanted cells or host tissues.
Furthermore, polymer coating on the ceramic particles by polymer
solutions may hinder the exposure of the ceramics to the scaffold
surfaces [FIG. 12(A)], which decreases the chance of contact
between the osteogenic cells and the bioactive ceramics. We thought
the gas foaming and particulate leaching (GF/PL) method would
efficiently expose bioactive ceramics on the scaffold surfaces and
that these efficiently exposed ceramics would thus enhance the
osteoconductivity and wettability of the scaffolds. In the present
study, we tested this hypothesis by transplanting scaffolds to rat
skull critical size defects and examining the bone formation. The
HA exposure to the scaffold surface was compared between GF/PL
scaffolds and SC/PL scaffolds. Bone regeneration using GF/PL
scaffolds was evaluated in vivo and compared with that using SC/PL
scaffolds.
[0093] Porous PLGAIHA composite scaffolds were fabricated by the
modification of a previously described GF/PL method [FIG. 10(B)] of
Harris L D, Kim B S, Mooney D J. Open pore biodegradable matrices
formed with as foaming, J Biomed Mater Res 1998; 42: 396-402;
PLGA/HA composites were prepared with 75:25 PLGA particles
(diameter=100-200 mm, molecular weight=100,000 Da; Birmingham
Polymers, Birmingham, Ala.), HA nanoparticles
(diameter=approximately 100 nm; Berkeley Advanced Biomaterials,
Berkeley, Calif.), and sodium chloride particles (diameter=100-200
mm; Sigma, St. Louis, Mo.). The mixed PLGA/HA/NaCl mass ratio was
1:1:9. The mixture was loaded into a disk mold (diameter=13.5 mm;
Aldrich Chemical, Milwaukee, Wis.) and compressed at 2000 psi for 1
min using a Carver Laboratory Press (Fred S. Carver, Menominee
Falls, Wis.) to yield solid disks with a thickness of 1 mm. The
samples were exposed to high pressure CO2 gas (800 psi) for 48 h to
saturate the polymer with the gas. Next, a thermodynamic
instability was created by decreasing the gas pressure to ambient
pressure. The NaCl particles were subsequently removed from the
scaffolds by leaching the scaffolds in distilled water for 48 h.
PLGA scaffolds without HA were also fabricated by the GF/PL method
(GF/PL-no HA).
[0094] Porous PLGA/HA scaffolds were also fabricated by the
modification of a previously described SC/PL method set forth in Lu
H H, El-Amin S F, Scott K D, Laurencin C T, Three-dimensional,
bioactive, biodegradable polymer-bioactive glass composite
scaffolds with improved mechanical properties support collagen
synthesis and mineralization of human osteoblast-like cells in
vitro, J Biomed Mater Res A 2003; 64:465-474; Cho S W, Kim I K, Lim
S H, Kim D I, Kang S W, Kim S H, et al. Smooth muscle-like tissues
engineered with bone marrow stromal cells, Biomaterials 2004;
25:2979-86; and Kim B S, Jeong S I, Cho S W, Nikolovski J, Mooney D
J, Lee S H, et al., Tissue engineering of smooth muscle under a
mechanically dynamic condition, J Microbiol Biotech 2003; 13:841-5
and used as another control. In this process, PLGA was dissolved in
methylene chloride (J. T. Baker, Phillipsburg, N.J.) at a 10% (w/v)
concentration, and HA and NaCl were added to the PLGA solution at
the same size and ratio as those of the GF/PL scaffolds. This
mixture was loaded into Teflon cylinders (diameter=21.5 mm,
height=25 mm; Cole-Parmer Instrument Company, Vernon Hills, Ill.).
Following solvent evaporation, polymer disks with entrapped salt
particles were removed from the molds. The salt was removed by
immersing disks in distilled water for 48 h. The dimension of
fabricated SC/PL scaffolds was same with that of GF/PL
scaffolds.
[0095] The morphologies of the scaffolds were examined using a
scanning electron microscope (SEM; JSM-6330F, JEOL, Tokyo, Japan).
Samples were dehydrated in ascending grades of ethanol, dried, and
mounted on an aluminum stub using a double-sided carbon tape. The
specimens were coated with platinum using a Sputter Coater
(Cressington 108, Cressington Scientific Instruments, Cranberry,
Pa.) and examined with SEM at an acceleration voltage of 10 kV. The
porosity of fabricated scaffolds was measured using mercury
intrusion porosimetry (Autopore IV 9500, Micromeritics Instrument
Corporation, Norcross, GA). A contact angle of 1308 for mercury on
the scaffold was used for this analysis. Compression and tensile
tests were performed with an Instron mechanical tester (Instron
4201, Instron.RTM., Canton, M A). The scaffold samples were cut
into 1.times.1 cm.sup.2 for compression testing. For tensile
testing, the samples (1.times.1 cm.sup.2) were attached to
cardboard using epoxy glue. The sample was centered in a 7 mm slot
in the centre of the card board and then glued to standardize the
gauge length. Compression and tensile tests were performed with a
constant strain rate of 1 mm/min. The moduli were determined from
the slopes in the initial elastic portion of the stress-strain
diagram. To examine the distribution and the extent of surface
exposure of HA in the scaffolds, HA exposed to the scaffold surface
was visualized with a hydrophilic dye (trypan blue) staining or von
Kossa's silver staining. For trypan blue (Sigma) staining, the
residual dye was removed by sonication in 100% ethanol. For the von
Kossa staining, scaffolds were immersed in 2% (w/v) silver nitrate
(Sigma) solution and placed directly in front of a bright lamp for
30 min. Scaffolds were then rinsed in distilled water. The surface
of the PLGA/HA scaffolds was then examined with a microscope
(Camscope, Samtech, Seoul, South Korea). The scaffolds were then
tested for their wettability.
[0096] For this, trypan blue solution was dropped on top of the
scaffold and the time required for complete absorption of the
solution into the scaffold was measured.
[0097] Sprague-Dawley rats (8-week-old males, n=24; SLC, Tokyo,
Japan) were anesthetized with an intraperitoneal injection (5 mg/kg
body wt.) of a 4:1 solution of ketamine hydrochloride (Ketara,
Yuhan, Seoul, South Korea) and xylazine (Rompun, Bayer Korea,
Seoul, South Korea). After shaving the scalp hair, a longitudinal
incision was made in the midline of cranium from the nasal bone to
the posterior nuchal line, and the periosteum was elevated to
expose the surface of the parietal bones. Using a surgical trephine
bur (Ace Surgical Supply, Brockton, Mass.) and a low-speed
micromotor, a circular and transosseous defect (8 mm in diameter)
was produced in the parietal bone. The drilling site was irrigated
with saline and the bleeding point was clectrocauterized. The
defect was filled with the fabricated scaffolds. The periosteum and
skin were then closed in layers with resorbable 4-0 Vicryl.RTM.
(Ethicon, Edinburgh, UK) sutures. The rats were housed singly after
surgery and received humane care in compliance with Seoul National
University Guidelines for the care and use of laboratory animals.
The implants were retrieved 8 weeks after implantation for
analyses.
[0098] Following euthanasia, the craniums including implanted
scaffolds were retrieved at 8 weeks after surgery. Samples were
fixed immediately in a 10% (v/v) neutral buffered formalin
solution. One specimen from each condition was scanned using a
desktop X-ray 3D micro computed tomography (CT; Skyscan 1072.RTM.,
SkyScan bvba, Aartselaar, Belgium) to analyze bone formation. A
microfocus X-ray tube with a focal spot of 8 mm was used as a
source and a 1024.times.1024 12-bit digital CCD X-ray detector were
used. The coronal view was imaged at 8 mm slices. Each coronal CT
image was reconstructed by V-works.TM. (CyberMed, Seoul, South
Korea) to visualize the three-dimensional volume image of new bone
and to examine the microarchitecture of the regenerated bone
tissue.
[0099] After collection of the micro CT scans, all samples (n=24)
were decalcified in EDTA (pH 6.0) for 7 days and embedded in
paraffin. The tissue blocks were sectioned at a 4 mm thickness and
stained with hematoxylin and eosin (H & E) and Masson's
trichrome. The Masson's trichrome-stained mid-portion sections were
examined with a microscope for histomorphometry. The percentage of
bone-occupying space within the constructs was measured using an
image analysis system (KS400, Zeiss, Munich, Germany) coupled to a
light microscope. The bone formation area was expressed as percent
bone area in available pore space (bone area/pore area=100%).
[0100] Quantitative data were expressed as the mean 6 standard
deviation. Statistical comparisons were carried out using analysis
of variance (ANOVA, SAS Institute, Cary, N.C.). A value of
p<0.05 was considered to be statistically significant.
[0101] Gas foaming and subsequent salt leaching of scaffolds
containing a high percentage of NaCl particles led to the formation
of highly porous structures with no evidence of an external, non
porous skin layer [FIG. 13(B)]. The GF/PL scaffolds exhibited
highly porous and open-pore structures. The pore structure ob
served in the cross-sections of the GF/PL scaffolds was similar to
that of the SC/PL scaffolds [FIG. 13(A)]. The average porosities of
the GF/PL and SC/PL scaffolds were (91.+-.3) % and (86.+-.3) %,
respectively.
[0102] The mechanical properties of the scaffolds were assessed
using compressive and tensile mechanical tests. The GF/PL scaffolds
exhibited enhanced mechanical properties as compared to the SC/PL
scaffolds. The average compression modulus was 2.3.+-.0.4 and
4.5.+-.0.3 MPa (p<0.05) for the SC/PL and GF/PL scaffolds, F3
respectively [FIG. 14(A)]. The average tensile modulus was 2.0 6
0.1 and 26.9 6 0.2 MPa (p<0.05) for the SC/PL and GF/PL
scaffolds, respectively [FIG. 14(B,C)]. These data represent a 99%
increase in compression modulus and a 1331% increase in tensile
modulus.
[0103] To determine whether the scaffold fabrication process
affected the extent of HA exposure to the scaffold surface, the
exposed HA was stained with von Kossa's silver nitrate [FIG. 15(A)]
and a hydrophilic trypan blue dye [FIG. 4(B)]. HA was stained more
abundantly in the GF/PL scaffolds than in the SC/PL scaffolds with
both staining methods (FIG. 15). In contrast, GF/PL scaffolds
without HA (GF/PL-no HA scaffolds) showed no positive staining with
either staining method [FIG. 15(A-C)].
[0104] To evaluate whether the hydrophilicity of the PLGA scaffold
and PLGA/HA composite scaffolds could be improved by the addition
of HA and by the application of different fabrication processes,
respectively, the wettability of GF/PL, SC/PL, and GF/PL-no HA
scaffolds was measured. When a trypan blue dye solution was dropped
to the scaffolds, the solution was completely absorbed into the
GF/PL scaffold within 5 s [FIG. 16(C)]. However, it was not
absorbed at all into F5 the GF/PL-no HA scaffold even after 60 min
because of the hydrophobic character of the scaffold [FIG. 16(A)].
The SC/PL scaffolds absorbed the dye solution slowly within 15 min
[FIG. 16(B)]. The faster wetting of the GF/PL scaffold compared to
that of the SC/PL scaffold correlated with the amount of HA exposed
onto the scaffold surface.
[0105] The implantation of both types of the PLGA/HA composite
scaffolds into critical size defects in rat skulls resulted in
enhanced bone formation in vivo compared with the PLGA scaffold
(FIG. 17). Eight weeks after implantation, new bone with lamellar
structures and osteoid formation was appreciated in the SC/PL
[FIGS. 18(B,E) and FIGS. 19(B,E)] and GF/PL [FIGS. 18(C,F) and
19(C,F)] scaffolds at F8 the defect edges and midsites of the
grafts. The bone formation area in the GF/PL scaffold was higher
than that in the SC/PL scaffold and GF/PL-no HA scaffold at 8 weeks
after implantation (FIG. 20). The PLGA polymer with no HA produced
very little new bone in vivo in the 8 weeks following implantation
[FIGS. 18(A,D) and 19(A,D)]. Most of the pores of the PLGA
scaffolds were filled with loose fibrous connective tissues without
evidence of bone formation [FIGS. 18(D) and 19(D)]. Over time, the
PLGA seemed to continuously degrade without adverse reactions, and
had not completely resorbed during the 8 weeks following
implantation.
[0106] Micro CT evaluation allowed the mineralized tissues to
distinguish from the remaining soft tissues present inside the
defects (FIG. 21). The reconstructed three-dimensional images
showed the formation of new bone inside both types of the PLGA/HA
composite scaffolds. Within the scaffolds, dispersed
irregular-shaped mineralized tissues were found throughout the
implant. Mineralized tissue areas were significantly larger in the
GF/PL scaffolds than in the SC/PL and GF/PL-no HA scaffolds.
[0107] As compared to the conventional methods for the fabrication
of biodegradable polymer/ceramic composite scaffolds, the GF/PL
method has several advantages. First, the GF/PL method avoids the
use of organic solvents. Residual organic solvents remaining in the
scaffold may damage transplanted cells and surrounding tissues.
Furthermore, exposure to organic solvents may inactivate
biologically active factors. Therefore, the GF/PL method may cause
less denaturation of growth factors incorporated within the
scaffold. Second, the GF/PL scaffold exhibits a higher exposure of
HA to the scaffold surface than the SC/PL scaffold. The SC/PL
method causes the coating of the HA by the polymer solution, which
hinders the exposure of HA to the scaffold surfaces, while the
GF/PL method efficiently exposes HA to the scaffold surface due to
the lack of a requirement of a polymer solution. Therefore, the
GF/PL scaffold can increase the chances of contact between the
osteogenic cells and the bioactive ceramics that are exposed on the
scaffold surface.
[0108] In addition, the enhanced exposure of hydrophilic HA
nanoparticles to the scaffold surface affected the hydrophilicity
of the scaffold. The hydrophobic surfaces of most synthetic polymer
biomaterials are unfavorable to osteogenic cells, because they show
a lower proliferative rate and a higher apoptotic rate on
hydrophobic surface than on hydrophilic surfaces. There fore the
addition of HA to the PLGA scaffold may enhance the scaffolds
hydrophilicity as well as its osteoconductivity. The scaffold
fabrication process also affected the hydrophilicity of the PLGA/HA
scaffolds. The faster wetting of the GF/PL scaffold compared with
the SC/PL scaffold was likely due to higher amount of HA exposed to
the surface of the GF/PL scaffold.
[0109] The GF/PL scaffold exhibited enhanced bone regeneration in
vivo when compared to the SC/PL scaffold. The PLGA in the PLGA/HA
scaffolds is bioinert for bone formation. However, the HA exposed
to the scaffold surface stimulates bone formation. Enhanced bone
formation on the GF/PL scaffolds may result from the direct contact
of migrating osteogenic cells from the surrounding tissues with the
HA particles exposed to the scaffold surface, which then stimulates
cell proliferation and osteogenic differentiation. In contrast, the
SC/PL scaffold had HA particles coated with the polymer, which
hindered interaction of HA with the osteogenic cells and thus
hindered the bone formation process.
[0110] In this study, we used nanosized HA particles to fabricate
PLGA/HA composite scaffolds. Since the highly crystalline HA
degrades over a long period in vivo, incompletely degraded residual
HA may hinder or slow the complete bone healing. Therefore, to
reduce the total amount of HA while enhancing the HA exposure to
the scaffold surface, nanosized HA particles, which have a high
surface area, were used to fabricate the composite scaffolds
instead of microsized HA particles. Furthermore, nanosized HA
particles showed improved bioactivity and osteointegration when
implanted to the bone defect site compared to the microsized HA
particles. Protein adsorption and cell adhesion have also been
reported to be enhanced by the use of nanosized HA particles
compared to microsized HA particles.
[0111] The PLGA/HA composite scaffolds of the present invention
show enhanced hydrophilicity and osteoconductivity compared with
the SC/PL scaffolds. This enhancement was most likely due to a
higher extent of exposure of HA particles to the scaffold surface.
The biodegradable polymer/bioceramic composite scaffolds fabricated
by the GF/PL method could enhance bone regeneration efficacy for
the treatment of bone defects compared with conventional composite
scaffolds.
[0112] One of skill in the art will be readily aware that while
polyglycolic acid polymers are most preferred due to their
osteoconductive and osteoinductive properties, other polymers can
be used to achieve similar results to the present invention. Such
other polymers include but are not limited to a bioabsorbable, or
biodegradable, synthetic polymer such as a polyanhydride,
polyorthoester, polylactic acid, and copolymers or blends thereof.
Non-degradable materials can also be used to form the matrix.
Examples of suitable materials include ethylene vinyl acetate,
derivatives of polyvinyl alcohol, teflon, and nylon. The preferred
non-degradable materials are a polyvinyl alcohol sponge, or
alkylation, and acylation derivatives thereof, including esters.
Collagen can be used, but is not as controllable and is not
preferred. These materials are all commercially available.
Non-biodegradable polymer materials can be used, depending on the
ultimate disposition of the growing cells, including
polymethacrylate and silicon polymers.
[0113] Those of skill in the art are familiar with the use of bone
graft materials. The present invention can be used in the same
manner as any other bone graft material. In a preferred delivery
system, the bone graft material is mixed with polyethylene glycol
to form a paste. The material can be premixed and sold in a syringe
for easy application or can be mixed at the point of use and
delivered via any convenient means. When provided in a dry state,
any suitable biocompatible fluid can be used to wet the material
and create a paste for administration to the patient. Examples of
such biocompatible fluid wetting agents include, but are not
limited to: dextrose, glucose, maltose or sodium chloride
solutions, blood, serum, platelet concentrate, bone marrow
aspirate, and synovial fluid. A biological fluid can be used in the
form obtained from the biological source, or it can be processed by
application of one or more desired useful techniques, examples of
which include, separation techniques, such as filtration (macro-,
micro-, or ultra-filtration); purification techniques, such as
dialysis; concentration techniques; and sterilization
techniques.
[0114] One of skill in the art will recognize that other biological
components including but not limited to proteins, growth factors,
cells, stem cells, osteoblasts or such other components that will
promote bone growth or maturation of the bone graft.
[0115] The description of the teachings is merely exemplary in
nature and, thus, variations that do not depart from the gist of
the teachings are intended to be within the scope of the teachings.
Such variations are not to be regarded as a departure from the
spirit and scope of the teachings.
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