The effect of devitalized trabecular bone on the formation of
osteochondral tissue-engineered constructs
Eric G. Lima
a
,d
, Pen-hsiu Grace Chao
b
, Gerard A. Ateshian
a
,e
, B. Sonny Bal
c
, James L. Cook
c
,
Gordana Vunjak-Novakovic
a
, Clark T. Hung
a
,*
aDepartment of Biomedical Engineering, Columbia University, 351 Engineering Terrace MC8904, 1210 Amsterdam Avenue, New York, NY 10027, United States bInstitute of Biomedical Engineering, National Taiwan University, Taipei, Taiwan
cComparative Orthopaedic Laboratory, University of Missouri, Columbia, MO, United States
dMaterials Characterization Laboratory, Cooper Union, 51 Astor Place, New York NY, 10003, United States
eDepartment of Mechanical Engineering, Columbia University, Department of Mechanical Engineering, 500 West 120th Street, Mail Code 4703, 220 S.W. Mudd
a r t i c l e
i n f o
Article history: Received 30 April 2008 Accepted 9 July 2008
Available online 20 August 2008
Keywords:
Cartilage tissue-engineering Biocompatibility
Bone
a b s t r a c t
In the current study, evidence is presented demonstrating that devitalized trabecular bone has an inhibitory effect on in vitro chondral tissue development when used as a base material for the tissue-engineering of osteochondral constructs for cartilage repair. Chondrocyte-seeded agarose hydrogel constructs were cultured alone or attached to an underlying bony base in a chemically defined medium formulation that has been shown to yield engineered cartilaginous tissue with native Young’s modulus (EY) and glycosaminoglycan (GAG) content. By day 42 in culture the incorporation of a bony base significantly reduced these properties (EY¼ 87 12 kPa, GAG ¼ 1.9 0.8%ww) compared to the gel-alone group (EY¼ 642 97 kPa, GAG ¼ 4.6 1.4%ww). Similarly, the mechanical and biochemical properties of chondrocyte-seeded agarose constructs were inhibited when co-cultured adjacent to bone (unattached), suggesting that soluble factors rather than direct cell–bone interactions mediate the chondro-inhibitory bone effects. Altering the method of bone preparation, including demineralization, or the timing of bone introduction in co-culture did not ameliorate the effects. In contrast, osteochondral constructs with native cartilage properties (EY¼ 730 65 kPa, GAG ¼ 5.2 0.9%ww) were achieved when a porous tantalum metal base material was adopted instead of bone. This work suggests that devitalized bone may not be a suitable substrate for long-term cultivation of osteochondral grafts.
Ó 2008 Elsevier Ltd. All rights reserved.
1. Introduction
Articular cartilage is a specialized connective tissue that bears
load and reduces friction across moving joints. It is composed of an
extracellular matrix that contains no nerves or blood vessels and
relatively few cells (5% volume). When damaged, articular cartilage
does not heal, but instead often degenerates further, leading to pain
and loss of function
[1]
. Due to the prevalence of osteoarthritis (OA)
and damage to articular cartilage, coupled with this poor intrinsic
healing response, there is a great demand for clinical intervention,
usually in the form of a highly invasive non-biological prosthetic
(such as total joint arthroplasty). Currently, the most common
biological alternative to arthroplasty entails the transplantation of
healthy osteochondral autografts (cartilage along with some of the
underlying bone) from a non-load bearing region to the affected
site
[1–5]
. Osteochondral grafts are better anchored than
chondral-only grafts and are less likely to be displaced by shearing forces
within the joint. While these autologous grafting procedures are
promising, they are limited both by the amount of tissue available
and donor-site morbidity associated with its harvest.
Tissue-engi-neering strategies, if successful, would alleviate these problems by
creating replacement tissues of the proper size and shape without
concurrent damage to other regions of the patient’s body.
Osteochondral constructs are designed to be press-fit into
pre-drilled cavities in the damaged joint, replacing the host cartilage
above while anchoring to the bone below. There are a great variety
of tissue-engineering approaches to form osteochondral constructs,
but typically they all entail the use of some form of tissue-scaffold.
The most important characteristics of the scaffold are its
mechanical properties, its porosity, and its biocompatibility.
Specifically it must be able to properly bear and transfer loads to the
surrounding tissue without being crushed. Since this environment
also includes shear forces, the interface strength between the
cartilage and bone regions should also be considered. The scaffolds
must also have the proper porosity to allow for cell infiltration and
nutrient transport and be biocompatible to mitigate immunogenic
*Corresponding author.E-mail address:cth6@columbia.edu(C.T. Hung).
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issues while allowing engineered tissue to develop functional
properties. The two regions of an osteochondral construct can be
formed from a single continuous scaffold
[6–8]
or from two
inde-pendent scaffolds that have been joined together
[9]
with each
region specialized to promote chondrogenic or osteogenic
development.
Although this development process would ideally take place
entirely within the defect site
[10–15]
, there is evidence that some
in vitro pre-culture may be necessary to have enough fortitude to
survive the harsh mix of inflammatory cytokines and concentrated
loads that are common in an injured joint
[16,17]
. The amount of
time dedicated to in vitro pre-culture must be considered in the
choice of engineered scaffolds, especially if the scaffold material is
designed to degrade over time. We have recently developed a tissue
culture protocol using chondrocyte-seeded agarose hydrogel and
the temporal addition of growth factors that yields engineered
tissue with native Young’s modulus (EY) and glycosaminoglycan
(GAG) content
[18]
. In Study 1 of the current set of experiments we
expand this system to include osteochondral constructs formed
using a devitalized trabecular bone as a base scaffold.
Among its many benefits, trabecular bone represents the gold
standard in terms of mechanical properties, and unlike degradable
substrates these properties do not change substantially over time in
culture, allowing for extensive in vitro cultivation. Bone also has an
interconnected porosity that is ideal for gelling chondrocyte-laden
agarose or other hydrogels. It is abundantly available for research
use and easily machined into a multitude of forms. Devitalized and
demineralized bone is already approved by the FDA for clinical use
as a scaffold to promote bone growth, as a source of osteoinductive
factors, and as allografts
[19]
. As such it is an alluring choice for an
osteochondral scaffold, however, the results of preliminary trials
[20]
using devitalized trabecular bone suggest there are
unantici-pated inhibitory effects on chondral development.
Study 1 extends these preliminary results, by examining the
effect of trabecular bone on chondral development when used both
directly in the formation of multi-phase osteochondral constructs
and indirectly when included in the culture medium unattached to
the developing gels.
An alternate approach to forming osteochondral constructs is to
culture the two phases separately and join them together after
certain engineering milestones have been met (i.e., mechanical or
chemical fortitude, etc.). Study 2 was designed to examine the
feasibility of this approach by delaying the introduction of bone to
the culture medium to allow for matrix development. Finally, in
Study 3 we examine porous tantalum metal, a highly promising
synthetic alternative to bone, as a tissue-engineering scaffold for
the formation of osteochondral constructs with functional
mechanical properties.
2. Materials and methods
2.1. Experimental design
Three studies were carried out in this manuscript (Fig. 1). In Study 1 the development of chondrocyte-seeded agarose hydrogel constructs (Study 1, Gel) and osteochondral constructs (Study 1, OC(bone)) were directly compared using the same tissue-engineering protocol. To assess the effects of soluble factors released by bone, Gel constructs were also co-cultured adjacent (but unattached) to devitalized bone (Study 1, Co-culture(bone)). To exclude the effect of soluble minerals Gel constructs were also cultured adjacent to demineralized bone (Study 1, Co-culture-(demin)). Finally, to test for the possibility that the inhibitory effect of bone is not related to soluble factors, but rather due to a decrease in the availability of growth factors through the absorption into bone, Gel constructs were cultured in medium without TGF-b3 (Study 1, No TGF).
In Study 2 the possibility of forming functional osteochondral constructs after separate cultivation of the chondral region was examined by delaying the intro-duction of bone to day 14 of culture (Study 2, Bone Introduced on day 14) or on day 28 (Study 2, Bone Introduced on day 28). Gel constructs without any Bone Introduced served as controls (Study 2, No Bone Introduced).
Study 1: Trabecular Bone is Examined as a Substrate for Osteochondral Tissue Engineering (But is Found to
Release Soluble Factors that Inhibit Chondral Development)
Gel
OC (bone)
Co-culture
(bone)
Co-culture
(demin bone)
No TGF-
β3
Study 2: The Introduction of Bone is Delayed to Allow for Matrix Development
No Bone Present in Medium
- No Bone Introduced
- Bone Introduced on Day 14
- Bone Introduced on Day 28
Bone Present in Medium
Study 3: Porous Tantalum Metal is Examined as an Alternate Substrate for Osteochondral Tissue Engineering
Day 0
Day 14
Day 28
Day 42
Gel
OC(bone)
OC(Tantalum)
Fig. 1. Schematic of experimental design. Study 1: Gel: single-phase chondrocyte-seeded agarose constructs serve as controls, OC(bone): multi-phase osteochondral constructs formed with a trabecular bone substrate, Co-culture(bone): single-phase gel cultured adjacent to bone to test for the effect of soluble factors, Co-culture(demin bone): Gel constructs cultured adjacent to demineralized bone to control for the effect of soluble minerals, No TGF-b3: Gel constructs cultured in medium without TGF-b3 to test for possible decrease in availability of growth factors through absorption into bone. Study 2: Bone is introduced to gel constructs at day 14 (Bone Introduced on day 14) or at day 28 (Bone Introduced on day 28) to examine the possibility of forming osteochondral constructs after gel is more developed. Study 3: OC(Tantalum): Multi-Phase osteochondral constructs formed with tantalum substrate, Gel and OC(bone) serve as controls. Each study was carried out separately and all groups were cultured for 42 days.
In Study 3 multi-phase osteochondral constructs are formed with a porous tantalum metal substrate (Study 3, OC(tantalum)) and compared to both Gel and OC(bone) constructs. Porous tantalum metal is a non-biological substrate currently being investigated as a component for total joint arthroplasty and has shown excellent early clinical results[21–26]. Each study was carried out separately and all groups were cultured for 42 days.
2.2. Cell isolation
Articular cartilage was harvested from bovine carpo-metacarpal (CMC) joints of freshly slaughtered 1–3 weeks old calves. Three to five joints were used for each study and cells were pooled from all joints. Cartilage chunks were digested in high-glucose Dulbecco’s Modified Eagle’s Medium (hgDMEM) supplemented with 5% fetal bovine serum and 390 U/ml collagenase type VI (Sigma Chemicals, St. Louis, MO) for 11 h at 37C with stirring. The resulting cell suspension was then filtered through a 70mm pore-size mesh and sedimented in a bench top centrifuge for 10 min at 1000g. Viable cells were counted using a hemacytometer and trypan blue.
2.3. Osteochondral substrate preparation
2.3.1. Devitalized bone
Cylindrical cores (w15 mm long) of trabecular bone were isolated from the subchondral region of bovine tibia using a diamond-tipped, hollow drill (Starlite, Rosemont, PA). Cores were rough cut to w6 mm in length and centered in a custom 4 mm thick stainless steel mold such that there were overhanging surfaces on both sides of the mold. These surfaces were then sanded flat with a hand-held device to ensure that the final bone cores had uniform dimensions (Ø 4 4 mm 50mm) with parallel faces cut normal to the axis of symmetry. Bone cores were then cleaned of marrow in one of the three ways: (1) with a water pick and subsequently sterilized in 70% ethanol for 4 h, (2) by washing in hypotonic buffer with detergent and DNAse and RNAse solutions, or (3) as provided by a commercial vendor through their FDA approved BioCleanse processes (RTI Biologics). To keep the quantity of bone
consistent between experiments, cleansed bone was sorted to within a 10% devia-tion in mass and volume. The experiments presented in Study 1 and Study 2 were repeated with each of these cleaning methods with no significant differences in results. The data presented here are averaged across all experiments. For the Co-culture(demin) group in Study 1 bone was demineralized in 6 N HCl for 12 h.
2.3.2. Tantalum metal
Porous tantalum metal (Zimmer) was prepared using wire cut electron discharge machining (to maintain the scaffold porosity[27]) to final construct dimensions of Ø 4 4 mm.
2.4. Casting of chondral or osteochondral constructs
One volume of chondrocyte suspension (at 60 106cells/ml) was then mixed with an equal volume of 4% low-melt agarose (Type VII, Sigma) at 37C to yield a final cell concentration of 30 106cells/ml in 2% agarose. For Gel specimens the chondrocyte/agarose suspension was cast into slabs and cored using a sterile disposable punch (Miltex, York, PA) to final dimensions (Ø 4 mm 2.3 mm).
For osteochondral specimens 60ml of molten cell-laden agarose suspension was poured into the cylindrical wells of a custom mold. Osteochondral substrates were pressed into the gel from above to the desired depth (adjusted through a temporary retaining ring). With this technique a multi-layered construct was formed with the following dimensions: a 2 mm agarose-only top region, a 2 mm agarose þ substrate interface region, and a 2 mm substrate-only region.
2.5. Growth medium
The growth medium was changed every other day and consisted of hgDMEM with 1 PSF (100 units/ml Penicillin, 100mg/ml Streptomycin, 0.25mg/ml Fungi-zone), 0.1mMdexamethasone, 50mg/ml ascorbate 2-phosphate, 40mg/mlL-proline, 100mg/ml sodium pyruvate, and 1 ITS þ premix (insulin, human transferrin, and selenous acid, Becton Dickinson, Franklin Lakes, NJ). Chemically defined medium
500
600
700
800
Day 0
Day 14
Day 28
Day 42
§*
0
100
200
300
400
Gel
OC (bone)
No TGF
E
Y(KPa)
Day 0
Day 14
Day 28
Day 42
Day 0
Day 14
Day 28
Day 42
*
*
*
*
*
*
§*
+
‡
+
*+
+
Co-culture
(bone)
Co-culture
(Demin bone)
Gel
OC (bone)
Co-culture
No TGF
(bone)
Co-culture
(Demin bone)
Gel
OC (bone)
Co-culture
No TGF
(bone)
Co-culture
(Demin bone)
5
6
7
§*
1
2
3
4
5
GAG (%ww)
*
4
0
1
2
3
*
*
0
Collagen (%ww)
Fig. 2. Results of Study 1. Gel: single-phase chondrocyte-seeded agarose constructs (n ¼ 14), OC(bone): multi-phase osteochondral constructs formed with a trabecular bone substrate (n ¼ 14), Co-culture(bone): single-phase gel cultured adjacent to bone (n ¼ 15), Co-culture(demin bone): gel constructs cultured adjacent to demineralized bone (n ¼ 4), No TGF-b3: gel constructs cultured in medium without TGF-b3 (n ¼ 4). *p < 0.05 vs. previous time point. xp < 0.05 vs. all other groups at same time point. zp < 0.05 vs. OC(bone), Co-culture(demin), or Co-culture(bone). þp < 0.05 vs. No TGF.
was further supplemented with 10 ng/ml of transforming growth factor beta 3 (TGF-b3, R&D Systems, Minneapolis, MN) for the first 14 days of culture.
2.6. Material testing
The equilibrium Young’s modulus (EY) is commonly used as a measure of the behavior of cartilage that has been allowed to reach equilibrium after a known load or displacement has been applied. Constructs were tested for EYin unconfined compression using a custom computer-controlled testing system[28]. An initial 0.02 N tare load was applied, followed by a compression to 10% strain (of the chondral region) at a strain rate of 0.05%/s. EYwas calculated from the equilibrium stress at 10% strain. Previous studies have shown EYto remain invariant across strain magnitudes ranging from 0 to 20%[18].
To determine the shear strength at the interface, the gel region of osteochondral constructs was cut in half, and mounted in a custom mold as to allow a platen to come into contact with the newly created flat surface (Fig. 6A). A linear displacement (10mm/s) was then applied and the load measured. The shear strength at the interface was calculated in three ways, as is commonly expressed in the literature
[29,30]. Peak load was determined as the highest force before failure (Fig. 6B, asterisks). Shear stiffness was determined by curve fitting the linear region on the force/displacement curve (Fig. 6B, bold line), Energy to Failure was determined by integrating the area under the force/displacement curve to the peak load and normalizing by interface area (Fig. 6B, shaded area).
2.7. Biochemical content
The biochemical content of each sample was assessed by first measuring sample wet weight, lyophilizing for 24 h, and then measuring the sample dry weight. Once dry, the samples were digested in proteinase-K overnight at 56C, as described previously[31]. Aliquots of digest were analyzed for glycosaminoglycan (GAG) content using the 1,9-dimethylmethylene blue dye-binding assay[32]. A further aliquot was acid hydrolyzed in 12 N HCl at 110C for 16 h, dried over NaOH, and resuspended in assay buffer[31]. Ortho-hydroxyproline (OHP) content was then determined via a colorimetric assay by reaction with chloramine T and dimethyla-minobenzaldehyde[33]. OHP content was converted to total collagen content using the conversion of 1:10 ratio of OHP:collagen[34]. Each biochemical constituent (GAG and collagen) was normalized to tissue wet weight.
2.8. Histological analysis
Gel samples were fixed in acid formalin ethanol[35], paraffin embedded, sectioned (8mm thick), and stained to view proteoglycan or total collagen or type II collagen distribution as described previously[36]. For osteochondral constructs histological specimens were prepared through plastic embedding and stained at the
Department of Surgical Sciences, University of Wisconsin. Live/dead assays were carried out using manufacture’s protocol (Molecular Probes).
2.9. Statistics
Statistics were performed with the Statistica (Statsoft, Tulsa, OK) software package. Each data point represents the mean and standard deviation. Groups were examined for significant differences by analysis of variance (a¼ 0.05), with EY, GAG, or OHP as the dependent variable using Tukey’s Honest Significant Difference Test (HSD).
3. Results
3.1. Study 1: comparing the development of chondral and
osteochondral constructs in the presence of devitalized trabecular
bone
In Study 1, both osteochondral (OC(bone)) and chondral (Gel)
constructs
developed
significantly
better
mechanical
and
biochemical properties over time (
Fig. 2
). When the two groups
were compared against each other, however, the osteochondral
group consistently underperformed. The day 42 values, for example,
were EY
¼ 87 12 kPa and GAG ¼ 1.9 0.8%ww for the
osteochon-dral group compared to EY
¼ 642 97 and GAG ¼ 4.6 1.4%ww for
the Gel group. Collagen values were not significantly different
between the two groups. DNA quantification (not shown) indicated
a 30% increase in cell number over the culture period with no
significant differences between the two groups.
Live/dead staining revealed the presence of vital cells (98% live
and 2% dead) in all three regions of the osteochondral constructs,
including the bone-only region where no cells were initially seeded
(not shown). These cells most likely migrated from the periphery of
the agarose region where they appeared to have attached to the
underlying bony substrate and adopted an elongated morphology.
Immunohistological staining indicated the continued deposition of
type II collagen in all three regions (
Fig. 3
), suggesting that the
chondrocyte phenotype was maintained, even with the change in
morphology.
Fig. 3. Type II collagen deposition (green) in upper gel region (A), interface region (B) and by chondrocytes migrating to the lower bone region (C). Red indicates bone. Chondrocytes (grey) maintain phenotype in all three regions. Scale bar ¼ 1 mm. Magnification ¼ 40.
The presence of separate bone plugs in the co-culture
experi-ments (Co-culture(bone)) resulted in significantly lower E
Yand GAG
by day 42 than the Gel groups and no statistical differences from the
osteochondral group (
Fig. 2
). The demineralization of the bone
Co-culture(demin) did not ameliorate these effects, yielding no
statis-tical differences in EY, GAG, or collagen from the osteochondral
groups and yielding statistically lower EY
and GAG than the Gel
group (
Fig. 2
).
Constructs cultured without TGF-
b
3 (No TGF) resulted in
significantly lower EY
and GAG when compared to Gel groups, but
higher E
Ythan the osteochondral group (
Fig. 2
). Histological
staining indicated a well-distributed extracellular network in all
groups (not shown). Von Kossa staining was negative for calcium
accumulation in agarose constructs and, therefore, did not indicate
an osteogenic phenotype (not shown).
3.2. Study 2: delaying the introduction of bone to allow for
extracellular development
The addition of bone to the culture medium resulted in lower EY
values regardless of whether the bone was added later in culture
(
Fig. 4
). The introduction of bone on day 14 resulted in an E
Ythat
was 15% of the Gel group by day 28 and 25% of Gel by day 42.
Likewise introduction of bone on day 28 resulted in 58% of the EY
of
Gel group by day 42. GAG values were similarly lower between the
two bone groups and the Gel group, with the exception of the day 42
Bone Introduced on day 28 group vs. Gel. There were likewise no
significant differences in collagen values between any of the groups.
3.3. Study 3: generating osteochondral constructs using a porous
tantalum base substrate
By day 42, osteochondral groups formed with bone substrate
(OC(Bone)) developed significantly lower E
Yand GAG than Gel
groups, as observed in the previous two studies (
Fig. 5
).
Osteo-chondral constructs formed using a porous tantalum base, on the
other hand, were not adversely affected by the scaffold and
developed an EY
of 730 65 kPa; a value within the range of native
cartilage
(500–1500 kPa)
[16]
.
Gross
morphology
indicated
a robust, cartilage-like chondral layer by day 42 in OC(tantalum)
specimens. The chondral region in OC(bone) specimens appeared in
some cases to have developed a gradient of extracellular
deposi-tion, becoming whiter and denser farther from the bony substrate.
Unfortunately, there was some shrinkage in the histological
prep-aration of the specimens of Study 3, nevertheless staining clearly
indicated rich accumulation of proteoglycans in both the Gel and
the OC(tantalum) groups, with less intense staining in OC(bone)
groups. Shear testing (
Fig. 6
) showed that the integration strength
of the chondral region to the tantalum base was greater than 200%
that of the OC(bone) groups whether it was expressed as peak load,
stiffness, or Energy to Failure (only peak load shown). By
compar-ison OC(tantalum) groups developed 28% of the Energy to Failure
observed in native osteochondral specimens (
Fig. 6
). All shear
failure occurred along the interface between the gel region and
bony substrate region of the osteochondral constructs.
4. Discussion
Taken together these results demonstrate that devitalized
trabecular bone has an inhibitory effect on in vitro chondral tissue
development when used as a base material for the
tissue-engi-neering of osteochondral constructs for cartilage repair. There is no
established level of mechanical competency that must be achieved
to assure construct survival within the joint. Orthopedic surgeons
have developed empirical guidelines for the transplantation of
allogenic or autologous cartilage grafts with properties similar to
the cartilage surrounding the defect, but there are no guidelines
regarding the implantation of engineered cartilage tissues that
have material properties less than native cartilage. Success using
these tissues will likely depend on parameters such as the construct
stiffness, defect size, and anatomic location, which together will
determine the stress and strain on the implant. It is important to
develop tissue-engineering protocols that result in stiffer, more
functional tissues, as this will likely broaden their range of clinical
applicability, possibly to include the tissue-engineering of total
joint replacements (biological arthroplasty)
[7,37,38]
. Additionally,
implantation of functional tissues with more mature elaborated
matrix may better resist inflammatory cytokine-induced tissue
degradation
[16]
.
Our previous work has optimized culture conditions for growing
engineered cartilage with material properties similar to native
cartilage in less than 8 weeks
[18]
. For these chondral-only
engi-neered tissues to be clinically relevant, however, they must be
integrated with an underlying bony substrate that provides graft
fixation upon implantation in a manner similar to native
osteo-chondral grafts. As several weeks of culture are necessary for
development of functional tissue properties, the scaffolds used in
the construct design must foster tissue elaboration in long-term
culture. In this context, devitalized trabecular bone initially seemed
600
800
No Bone Introduced
Bone Introduced on Day 14
*
200
400
600
E
Y(kPa)
Bone Introduced on Day 28
+§
*
+
*
0
0
14
28
42
0
14
28
42
0
14
28
42
+
8
10
No Bone Introduced
Bone Introduced on Day 14
2
4
6
GAG (%ww)
Bone Introduced on Day 28
+§
*
*
*
+
0
5
6
No Bone Introduced
Bone Introduced on Day 14
1
2
3
4
Collagen (%ww)
Bone Introduced on Day 28
*
*
0
Fig. 4. Study 2: Co-culture of Gel with bone introduced to a subset of controls at different times in culture (either on day 14 or on day 28). *p < 0.05 vs. previous time point. þp < 0.05 vs. No Bone Introduced. xp < 0.05 Bone Introduced on day 14 vs. Bone Introduced on day 28. n ¼ 10–12 per group.
to be a promising choice for the engineering of osteochondral
constructs
[37,39–41]
. We had anticipated few problems in its use
as vital and devitalized bone are used clinically as a scaffold for
bone repair
[42–44]
. Our findings showing the apparent inhibitory
effect of bone on chondral development (Study 1) were, therefore,
unexpected.
When faced with the results of Study 1, we initially
hypothe-sized that the observed detrimental effects were a result of the
methods in which we had cleaned and sterilized the freshly
har-vested bony substrates. This was not the case, however, as we tried
two additional cleaning methods, including the use of industrially
prepared bone tissue. None of these techniques led to significant
improvements in the properties measured.
We had also speculated that the differences in the mechanical
properties were not inherent to the tissues themselves, but an
artifact of having porous substrate present during mechanical
testing. This could be possible if, for example, compressive tests led
to permanent deformation of the hydrogel into the porous bony
substrate, or if small variations in the parallelism of the bony
surfaces propagated into larger errors in the mechanical testing
results. To account for this we tested constructs both on and off
their bony substrates, but removing the bone prior to testing did
not affect the results.
Thus the most likely remaining mechanisms in which bone was
adversely affecting tissue development were as follows: (1)
minerals and/or chemical factors associated with bone were
leech-ing into the growth medium, (2) nutrient transport was limited by
the presence of bone, or (3) TGF-
b
3 was being sequestered away in
the bone matrix and was, therefore, unavailable to cells.
The results of our co-culture experiment in Study 2 clearly
indicated that excessive mineral content, or limitations on
nutri-ents, or sequestering away of TGF-
b
3, were not the main factors
suppressing construct development. Cartilage constructs
co-cultured with devitalized bone matrix showed a decrease of
*
*
Day 14
Day 42
600
800
1000
day 0
day 14
day 28
day 42
+ §
B42
B14
Bone
0
200
400
E
Y(kPa)
*
*
*
*
T14
T42
Tantalum
Gel
OC (Tantalum)
OC (Bone)
Gel
OC (Tantalum)
OC (Bone)
Gel
OC (Tantalum)
OC (Bone)
8
10
day 0
day 14
day 28
+ §
*
Day 42
Gel
2
4
6
GAG (%ww)
day 42
*
Gel
Day 42
Tantalum
5
day 14
0
*
*
Day 42
Bone
2
3
4
day 28
day 42
0
1
Collagen (%ww)
Fig. 5. Left: Osteochondral constructs with porous tantalum metal as a substrate (OC(tantalum)) compared to a trabecular bone substrate (OC(bone)) or no substrate(Gel). *p < 0.05 vs. previous time point. þp < 0.05 vs. Gel. xp < 0.05C(bone) vs. OC(tantalum). n ¼ 12 per group. Right: Images of developing osteochondral constructs showing more robust growth for tantalum substrate over bone, OC(bone) on day 14 (B14) and day 42 (B42), OC(tantalum) on day 14 (T14) and day 42 (T42). Histological staining for GAG (alcian blue staining, dark purple) on day 42 for Gel, OC(tantalum), and OC(bone). Deeper intensity of purple indicates the presence of more GAG. Some shrinkage occurred in processing. Scale bar ¼ 1 mm.
mechanical properties and GAG levels. Similarly, constructs
cultured in medium that had been preconditioned with bone (data
not shown) also yielded significantly lower E
Ythan the Gel group.
Collectively, our findings suggest that soluble chemical mediators
were inhibiting the observed chondral tissue development. We
speculate that osteoinductive factors released by bone may
contribute to suppression of the chondrogenic phenotype
[19]
.
Bone matrix is known to contain intrinsic cytokines and growth
factors
[45]
that have a wide and largely unknown range of effects
on cell development. The type and concentration of these factors
vary even between the bones of the same animal
[45,46]
. Ongoing
studies are aimed at elucidating the specific chemical factor(s).
In light of our observations of tissue maturation-independent
deleterious effects of long-term culture with bone, one approach
would be to cultivate the gel phase separately until maturation and
then integrating this functional engineered chondral tissue to the
bone with some fixation method such as sutures, glues or solders
[9,47,48]
immediately prior to implantation. This approach has the
advantage of allowing culture conditions to be optimized for each
phase of the osteochondral construct, but precludes any
cell-mediated development of an interface between the two halves
prior to implantation. However, the results of the current study do
not support this approach as the addition of bone at later time
points in culture resulted in poorer mechanical and biochemical
properties when compared to bone-free controls.
The generation of osteochondral constructs was most successful
when bone was substituted with a non-biological alternative. Using
porous tantalum metal we were able to achieve native Young’s
modulus values and GAG and collagen content similar to
chondral-only constructs. The integration strength (between the layers) of
tantalum/agarose scaffolds was on par with, or exceeded, values
reported in the literature
[29,49,50]
, but remained below native levels
[30,51]
. Improvements on the integration strength are likely to be
achieved with improvements of the collagen content in the tissue,
although currently there is no established level of integration strength
required for tissue-engineered constructs to be successful in vivo.
In the context of creating functional engineered, osteochondral
grafts with chondral regions that possess mechanical properties
similar to native tissue and that are integrated to the underlying
bony base via cell-elaborated matrix, porous tantalum metal
appears to be a suitable bone substitute. Tantalum metal was
chosen as a non-biological substrate control because of its current
clinical application in orthopedic implants and its promise as an
osteochondral scaffold for tissue-engineering
[23–25]
. An
addi-tional benefit of a tantalum base may be realized upon
implanta-tion since it may reduce direct contact with the underlying marrow
and blood supply and, therefore, lessen the inflammatory effects
associated with allogenic cells
[19,52,53]
. Tantalum metal has also
been shown to be osteo- and chondroinductive
[21–26]
and may,
therefore, promote integration between the two graft halves in
culture as well as development of the subchondral plate after
implantation. Agarose hydrogel has been used extensively in
cartilage biology and in tissue-engineering efforts for maintaining
long-term chondrocyte suspension cultures
[54–57]
. It is currently
being explored as a scaffold component of a next generation
autologous cartilage implantation (ACI) procedure called Cartipatch
[58]
where it is currently being tested in phase III clinical trials
[59]
.
In this context, our findings support the use of a biohybrid
osteo-chondral graft design comprised of a cell-seeded hydrogel and
underlying tantalum metal substrate for promoting functional
development of tissue-engineered osteochondral constructs in
long-term culture. If an appropriate human cell source is identified
(e.g., allogenic or autologous chondrocytes, stem cells, etc.), we
anticipate the possibility for immediate clinical application of the
agarose-tantalum biohybrid system.
5. Conclusions
In the current study, devitalized trabecular bone had an
inhib-itory effect on in vitro chondral tissue development when used as
a base material for the tissue-engineering of osteochondral
constructs for cartilage repair. Changes in the method of bone
preparation or the timing of its introduction did not ameliorate the
effects. In contrast, osteochondral constructs with native cartilage
properties were achieved when a porous tantalum metal base
material was adopted instead of bone. This work suggests that
devitalized bone may not be a suitable substrate for long-term
cultivation of osteochondral grafts.
Acknowledgments
This research was funded by the following grants: National
Institutes for Health grants: R01 AR46568, R21 AR53530,
P41-EB002520, and R01 DE16525.
2
Tantalum
1
Bone
*
Integration Strength Measurement
Day 42 Tantalum Group
Day 0 Native Cart
Peak load (N)
1.4 0.6
6.0 0.8
2.9 ± 12
21.4 ± 13.4
0
Day 0
Day 14
Day 42
Peak Shear Load (N)
*
Stiffness (N/mm)
Energy to Failure(J/m
2)
109 ± 49
386 ± 123
A
B
C
Fig. 6. Shear strength at the interface of osteochondral constructs. (A) Methodology for the preparation of osteochondral constructs for shear testing. (B) Analysis of data: peak load determined at highest force before failure (asterisk). Shear stiffness determined by curve fit to linear region (Linear line), Energy to Failure determined by integrating area under curve to peak load and normalizing by interface area (shaded area). (C) Peak load between bone and tantalum groups. *p < 0.05 for OC(bone) vs. OC(tantalum). Table: comparison between tantalum group at day 42 and native cartilage.
Musculoskeletal Transplant Foundation grant: CU07-194.
Dr. Sonny Bal has received financial support from Zimmer
unrelated to the current project.
References
[1] Hangody L, Modis L. Surgical treatment options for weight bearing articular surface defect. Orv Hetil 2006;147(46):2203–12.
[2] Hangody L, Feczko P, Bartha L, Bodo G, Kish G. Mosaicplasty for the treatment of articular defects of the knee and ankle. Clin Orthop Relat Res 2001;(391 Suppl.):S328–36.
[3] Hangody L, Fules P. Autologous osteochondral mosaicplasty for the treatment of full-thickness defects of weight-bearing joints: ten years of experimental and clinical experience. J Bone Joint Surg Am 2003;85-A(Suppl. 2):25–32. [4] Marcacci M, Kon E, Zaffagnini S, Iacono F, Neri MP, Vascellari A, et al. Multiple
osteochondral arthroscopic grafting (mosaicplasty) for cartilage defects of the knee: prospective study results at 2-year follow-up. Arthroscopy 2005;21(4): 462–70.
[5] Williams 3rd RJ, Ranawat AS, Potter HG, Carter T, Warren RF. Fresh stored allografts for the treatment of osteochondral defects of the knee. J Bone Joint Surg Am 2007;89(4):718–26.
[6] Jiang CC, Chiang H, Liao CJ, Lin YJ, Kuo TF, Shieh CS, et al. Repair of porcine articular cartilage defect with a biphasic osteochondral composite. J Orthop Res 2007;25(10):1277–90.
[7] Schek RM, Taboas JM, Hollister SJ, Krebsbach PH. Tissue engineering osteo-chondral implants for temporomandibular joint repair. Orthod Craniofac Res 2005;8(4):313–9.
[8] Tuli R, Nandi S, Li WJ, Tuli S, Huang X, Manner PA, et al. Human mesenchymal progenitor cell-based tissue engineering of a single-unit osteochondral construct. Tissue Eng 2004;10(7–8):1169–79.
[9] Schaefer D, Martin I, Shastri P, Padera RF, Langer R, Freed LE, et al. In vitro generation of osteochondral composites. Biomaterials 2000;21(24): 2599–606.
[10] Orr TE, Patel AM, Wong B, Hatzigiannis GP, Minas T, Spector M. Attachment of periosteal grafts to articular cartilage with fibrin sealant. J Biomed Mater Res 1999;44(3):308–13.
[11] Dennis JE, Caplan AI. Porous ceramic vehicles for rat-marrow-derived (Rattus norvegicus) osteogenic cell delivery: effects of pre-treatment with fibronectin or laminin. J Oral Implantol 1993;19(2):106–15 [discussion 36–7].
[12] Kawamura S, Wakitani S, Kimura T, Maeda A, Caplan AI, Shino K, et al. Articular cartilage repair. Rabbit experiments with a collagen gel-biomatrix and chondrocytes cultured in it. Acta Orthop Scand 1998;69(1):56–62. [13] Wakitani S, Goto T, Young RG, Mansour JM, Goldberg VM, Caplan AI. Repair of
large full-thickness articular cartilage defects with allograft articular chon-drocytes embedded in a collagen gel. Tissue Eng 1998;4(4):429–44. [14] Caplan AI, Elyaderani M, Mochizuki Y, Wakitani S, Goldberg VM. Principles of
cartilage repair and regeneration. Clin Orthop 1997;342:254–69.
[15] Gao J, Dennis JE, Solchaga LA, Goldberg VM, Caplan AI. Repair of osteochondral defect with tissue-engineered two-phase composite material of injectable calcium phosphate and hyaluronan sponge. Tissue Eng 2002;8(5):827–37. [16] Lima EG, Tan AR, Tai T, Bian L, Stoker AM, Ateshian GA, et al. Differences in
interleukin-1 response between engineered and native cartilage. Tissue Eng 2008.
[17] Duda GN, Maldonado ZM, Klein P, Heller MO, Burns J, Bail H. On the influence of mechanical conditions in osteochondral defect healing. J Biomech 2005;38(4):843–51.
[18] Lima EG, Bian L, Ng KW, Mauck RL, Byers BA, Tuan RS, et al. The beneficial effect of delayed compressive loading on tissue-engineered cartilage constructs cultured with tgf-beta3. Osteoarthr Cartil 2007;15(9):1025–33. [19] Urist MR. Bone: formation by autoinduction. Science 1965;150(698):893–9. [20] Lima EG, Mauck RL, Gasinu S, Ateshian GA, Hung CT. Functional tissue
engi-neering of free-swelling and dynamically loaded osteochondral constructs. San Francisco, California: Orthopaedic Research Society; 2004.
[21] Wigfield C, Robertson J, Gill S, Nelson R. Clinical experience with porous tantalum cervical interbody implants in a prospective randomized controlled trial. Br J Neurosurg 2003;17(5):418–25.
[22] Mardones RM, Reinholz GG, Fitzsimmons JS, Zobitz ME, An KN, Lewallen DG, et al. Development of a biologic prosthetic composite for cartilage repair. Tissue Eng 2005;11(9–10):1368–78.
[23] Bobyn JD, Toh KK, Hacking SA, Tanzer M, Krygier JJ. Tissue response to porous tantalum acetabular cups: a canine model. J Arthroplasty 1999;14(3):347–54. [24] Levine B, Sporer S, Della Valle CJ, Jacobs JJ, Paprosky W. Porous tantalum in reconstructive surgery of the knee: a review. J Knee Surg 2007;20(3):185–94. [25] Christie MJ. Clinical applications of trabecular metal. Am J Orthop
2002;31(4):219–20.
[26] Hacking SA, Bobyn JD, Toh K, Tanzer M, Krygier JJ. Fibrous tissue ingrowth and attachment to porous tantalum. J Biomed Mater Res 2000;52(4):631–8. [27] Deglurkar M, Davy DT, Stewart M, Goldberg VM, Welter JF. Evaluation of
machining methods for trabecular metal implants in a rabbit intramedullary osseointegration model. J Biomed Mater Res B Appl Biomater 2007;80(2): 528–40.
[28] Soltz MA, Ateshian GA. Experimental verification and theoretical prediction of cartilage interstitial fluid pressurization at an impermeable contact interface in confined compression. J Biomech 1998;31(10):927–34.
[29] Allan KS, Pilliar RM, Wang J, Grynpas MD, Kandel RA. Formation of biphasic constructs containing cartilage with a calcified zone interface. Tissue Eng 2007;13(1):167–77.
[30] Broom ND, Oloyede A, Flachsmann R, Hows M. Dynamic fracture character-istics of the osteochondral junction undergoing shear deformation. Med Eng Phys 1996;18(5):396–404.
[31] Kelly TA, Ng KW, Wang CC, Ateshian GA, Hung CT. Spatial and temporal development of chondrocyte-seeded agarose constructs in free-swelling and dynamically loaded cultures. J Biomech 2005.
[32] Farndale RW, Sayers CA, Barrett AJ. A direct spectrophotometric microassay for sulfated glycosaminoglycans in cartilage cultures. Connect Tissue Res 1982;9(4):247–8.
[33] Stegemann H, Stalder K. Determination of hydroxyproline. Clin Chim Acta 1967;18(2):267–73.
[34] Vunjak-Novakovic G, Martin I, Obradovic B, Treppo S, Grodzinsky AJ, Langer R, et al. Bioreactor cultivation conditions modulate the composition and mechan-ical properties of tissue-engineered cartilage. J Orthop Res 1999;17(1):130–8. [35] Lin W, Shuster S, Maibach HI, Stern R. Patterns of hyaluronan staining are
modified by fixation techniques. J Histochem Cytochem 1997;45(8):1157–63. [36] Kelly TA, Fisher MB, Oswald ES, Tai T, Mauck RL, Ateshian GA, et al. Low-serum media and dynamic deformational loading in tissue engineering of articular cartilage. Ann Biomed Eng 2008.
[37] Hung CT, Lima EG, Mauck RL, Takai E, LeRoux MA, Lu HH, et al. Anatomically shaped osteochondral constructs for articular cartilage repair. J Biomech 2003;36(12):1853–64.
[38] Ateshian GA, Hung CT. Patellofemoral joint biomechanics and tissue engi-neering. Clin Orthop Relat Res 2005;436:81–90.
[39] Lima EG, Mauck RL, Han SH, Park S, Ng KW, Ateshian GA, et al. Functional tissue engineering of chondral and osteochondral constructs. Biorheology 2004;41(3–4):577–90.
[40] Hung CT, Mauck RL, Wang CC, Lima EG, Ateshian GA. A paradigm for func-tional tissue engineering of articular cartilage via applied physiologic defor-mational loading. Ann Biomed Eng 2004;32(1):35–49.
[41] Tognana E, Chen F, Padera RF, Leddy HA, Christensen SE, Guilak F, et al. Adjacent tissues (cartilage, bone) affect the functional integration of engi-neered calf cartilage in vitro. Osteoarthr Cartil 2005;13(2):129–38. [42] Mulliken JB, Glowacki J, Kaban LB, Folkman J, Murray JE. Use of demineralized
allogeneic bone implants for the correction of maxillocraniofacial deformities. Ann Surg 1981;194(3):366–72.
[43] Ellegaard B, Karring T, Loe H. The fate of vital and devitalized bone grafts in the healing of interradicular lesions. J Periodontal Res 1975;10(2):88–97. [44] Dallari D, Fini M, Stagni C, Torricelli P, Nicoli Aldini N, Giavaresi G, et al. In vivo
study on the healing of bone defects treated with bone marrow stromal cells, platelet rich plasma and freeze-dried bone allografts, alone and in combina-tion. J Orthop Res 2005.
[45] Gao J, Knaack D, Goldberg VM, Caplan AI. Osteochondral defect repair by demineralized cortical bone matrix. Clin Orthop Relat Res 2004;(427 Suppl.):S62–6.
[46] Urist MR, Jurist Jr JM, Dubuc FL, Strates BS. Quantitation of new bone formation in intramuscular implants of bone matrix in rabbits. Clin Orthop Relat Res 1970;68:279–93.
[47] Schaefer D, Martin I, Jundt G, Seidel J, Heberer M, Grodzinsky A, et al. Tissue-engineered composites for the repair of large osteochondral defects. Arthritis Rheum 2002;46(9):2524–34.
[48] Sherwood JK, Riley SL, Palazzolo R, Brown SC, Monkhouse DC, Coates M, et al. A three-dimensional osteochondral composite scaffold for articular cartilage repair. Biomaterials 2002;23(24):4739–51.
[49] Tam HK, Srivastava A, Colwell Jr CW, D’Lima DD. In vitro model of full-thick-ness cartilage defect healing. J Orthop Res 2007;25(9):1136–44.
[50] Nettles DL, Vail TP, Morgan MT, Grinstaff MW, Setton LA. Photocrosslinkable hyaluronan as a scaffold for articular cartilage repair. Ann Biomed Eng 2004;32(3):391–7.
[51] Kumar P, Oka M, Nakamura T, Yamamuro T, Delecrin J. Mechanical strength of osteochondral junction. Nippon Seikeigeka Gakkai Zasshi 1991;65(11):1070–7. [52] Sun L, Hu Y, Ning Z. Immunological comparison of differently treated allografts of bone. Zhonghua Wai Ke Za Zhi. [Chinese Journal of Surgery] 1996; 34(8):460–3.
[53] Sun L, Hu Y, Ning Z, Liang Z. The correlation between immune rejection and osteoinduction of allogeneic bone grafting. Chin Med J 1998;111(9):818–22. [54] Buschmann MD, Gluzband YA, Grodzinsky AJ, Kimura JH, Hunziker EB.
Chondrocytes in agarose culture synthesize a mechanically functional extra-cellular matrix. J Orthop Res 1992;10(6):745–58.
[55] Benya PD, Shaffer JD. Dedifferentiated chondrocytes reexpress the differenti-ated collagen phenotype when cultured in agarose gels. Cell 1982;30(1):215–24. [56] Mauck RL, Soltz MA, Wang CC, Wong DD, Chao PH, Valhmu WB, et al. Func-tional tissue engineering of articular cartilage through dynamic loading of chondrocyte-seeded agarose gels. J Biomech Eng 2000;122(3):252–60. [57] Mauck RL, Wang CC, Oswald ES, Ateshian GA, Hung CT. The role of cell seeding
density and nutrient supply for articular cartilage tissue engineering with deformational loading. Osteoarthr Cartil 2003;11(12):879–90 [OARS, Osteo-arthritis Research Society].
[58] Selmi TA, Verdonk P, Chambat P, Dubrana F, Potel JF, Barnouin L, et al. Autologous chondrocyte implantation in a novel alginate–agarose hydrogel: outcome at two years. J Bone Joint Surg Br 2008;90(5):597–604.
[59] Trials N. Multicenter randomized trial comparing cartipatchÔ versus mosaicoplasty. 2008.
