1.Hani a. Awad -Chondrogenic Differentiation of Adipose-Derived 2003

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Biomaterials 25 (2004) 3211–3222

Chondrogenic differentiation of adipose-derived adult stem cells in

agarose, alginate, and gelatin scaffolds

Hani A. Awada, M. Quinn Wickhama, Holly A. Leddya, Jeffrey M. Gimbleb,Farshid Guilaka,*

aDepartment of Surgery, Division of Orthopaedic Surgery, Duke University Medical Center, Durham, NC 27710, USAbPennington Biomedical Research Center, Louisiana State University, Baton Rouge, LA, USA

Received 9 September 2003; accepted 29 September 2003

Abstract

The differentiation and growth of adult stem cells within engineered tissue constructs are hypothesized to be influenced by cell-

biomaterial interactions. In this study, we compared the chondrogenic differentiation of human adipose-derived adult stem

(hADAS) cells seeded in alginate and agarose hydrogels, and porous gelatin scaffolds (Surgifoam), as well as the functional

properties of tissue engineered cartilage constructs. Chondrogenic media containing transforming growth factor beta 1 significantly

increased the rates of protein and proteoglycan synthesis as well as the content of DNA, sulfated glycosaminoglycans, and

hydroxyproline of engineered constructs as compared to control conditions. Furthermore, chondrogenic culture conditions resulted

in 86%, and 160% increases ( po0:05) in the equilibrium compressive and shear moduli of the gelatin scaffolds, although they didnot affect the mechanical properties of the hydrogels over 28 days in culture. Cells encapsulated in the hydrogels exhibited a

spherical cellular morphology, while cells in the gelatin scaffolds showed a more polygonal shape; however, this difference did not

appear to hinder the chondrogenic differentiation of the cells. Furthermore, the equilibrium compressive and shear moduli of the

gelatin scaffolds were comparable to agarose by day 28. Our results also indicated that increases in the shear moduli were

significantly associated with increases in S-GAG content (R2 ¼ 0:36; po0:05) and with the interaction between S-GAG and

hydroxyproline (R2 ¼ 0:34; po0:05). The findings of this study suggest that various biomaterials support the chondrogenicdifferentiation of hADAS cells, and that manipulating the composition of these tissue engineered constructs may have significant

effects on their mechanical properties.

r 2003 Elsevier Ltd. All rights reserved.

Keywords: Alginate; Agarose; Hydrogel; Collagen; Gelatin; Cartilage tissue engineering; Stem cell; Differentiation

1. Introduction

Tissue engineering is a promising therapeutic ap-

proach that combines cells, biomaterials, and environ-

mental factors to induce differentiation signals into

surgically transplantable formats and promote tissue

repair and/or functional restoration [1–5]. Despite many

advances, tissue engineers still face significant challenges

in repairing or replacing tissues that serve predomi-

nantly biomechanical functions such as articular carti-

lage. One obstacle has been the development of a

competent cartilage scaffold that: (a) has mechanical

properties and capability to withstand the large contact

stresses and strains of an articulating joint, (b) allows

functional tissue growth, and (c) provides for appro-

priate cell–matrix interactions to stimulate tissue growth

[6,7]. An evolving discipline termed ‘‘functional tissue

engineering’’ (FTE) seeks to address the functional

challenges of tissues such as cartilage by attempting to

define sets of criteria that must be satisfied in order to

overcome challenges associated with developing a

successful tissue-engineered graft [6].

Challenges related to the cellular component of an

engineered tissue include cell sourcing, expansion, and

differentiation as well as regulatory and production

issues, such as sterility, safety, storage, shipping, quality

control, and scale-up [8]. The use of human adipose

derived adult stem (hADAS) cells represents a feasible

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*Corresponding author. Tel.: +1-919-684-2521; fax: +1-919-681-

8490.

E-mail address: [email protected] (F. Guilak).

0142-9612/$- see front matterr 2003 Elsevier Ltd. All rights reserved.

doi:10.1016/j.biomaterials.2003.10.045


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approach to many of these issues [9]. Adipose tissue is

routinely available in liter quantities from liposuction

surgeries and yields an average of 400,000 cells per ml of

tissue after expansion, providing the numbers of cells

necessary for many tissue engineering applications [9].

The hADAS cells are pluripotent, expressing the

biochemical profile of adipocytes, chondrocytes, hema-topoietic supporting cells, myocytes, neurons, and

osteoblasts under appropriate culture conditions

in vitro [9].

Other challenges are associated with the biomaterial

scaffolds designed to deliver the cells and guide tissue

growth and differentiation. These biomaterials must

meet several criteria to maximize the chances of a

successful repair, including biodegradability and/or

biocompatibility, facilitating functional tissue growth,

and appropriate biomechanical properties [10–13].

Biomaterials used for cartilage tissue engineering can

have the form of either hydrogels or porous scaffolds.

Among the hydrogel biomaterials, the seaweed-derived

alginate and agarose are typically thought to be inert

because they lack native ligands that could allow

interaction with mammalian cells [14]. However, these

hydrogels provide a number of advantages for tissue

engineering including the possibility of minimally

invasive injection of hydrogel/cell constructs [15,16].

Various researchers have investigated the ability of

alginate and agarose hydrogels to act as a scaffold

material for chondrocytes to regenerate cartilage tissue

[14,17–25]. While many of these studies demonstrated

that alginate and agarose promote maintenance of

the chondrocytic phenotype in vitro, the successof tissue engineering applications using these hydrogels

may be hindered by their poor biomechanical properties

and handling characteristics. Furthermore, the ability of

these hydrogels to support chondrogenic differentiation

of adult stem cells has been less studied [26,27].

Gelatin, a porous denatured collagen scaffold, has

been recently used as a scaffolding structure for cartilage

tissue engineering [28,29]. The biological origin of

collagen-derived gelatin makes this material an attrac-

tive choice for tissue engineering [10]. However, there is

some concern that type I collagen scaffolds may not

preserve the chondrocyte phenotype of cells as well as

type II collagen scaffolds [30]. Furthermore, very little is

known about the functional (mechanical) properties of

this biomaterial although its ability to support chon-

drogenic differentiation of adult stem cells has been

demonstrated [28].

The goal of this study was to assess the functional

(biologic, biochemical, and biomechanical) properties of

alginate hydrogel, agarose hydrogel, and gelatin (Surgi-

foams, denatured porcine collagen type I) porous

sponge as scaffolding biomaterials for cartilage tissue

engineering using hADAS cells that have been shown to

possess a chondrogenic potential under defined culture

conditions [26,31]. We hypothesized that the functional

properties of tissue engineered cartilage will depend on

the choice of biomaterial scaffold, and that the construct

mechanical properties would be directly correlated to

the synthesis and accumulation of extracellular matrix

components.

2. Materials and methods

2.1. Isolation of hADAS cells

Human adipose derived adult stem (hADAS) cells

were isolated from subcutaneous adipose tissue (n ¼ 3

female donors, 29.777.2 years old (mean7standard

deviation) with a body mass index of 28.672.6 kg/m2)

as previously described [26,32,33]. Briefly, liposuction

waste tissue was digested with 0.25 mg collagenase type I

(200 units/mg) per ml of Krebs-Ringer-Bicarbonate

solution (Sigma, St Louis, MO) for 40 min at 37C with

intermittent shaking. The floating adipocytes were

separated from the precipitating stromal fraction by

centrifugation. The stromal cells were then plated in

tissue culture flasks at approximately 3500 cell/cm2 in

stromal media (DMEM/F-12 with 10% fetal bovine

serum (FBS), 100 units/ml penicillin and 100 mg/ml

streptomycin). The primary cells (P0) were cultured for

4 to 5 days, after which they were harvested by

treatment with trypsin (0.05%)/EDTA, counted, and

then frozen in liquid nitrogen in cryopreservation

medium (80% FBS, 10% dimethylsulfoxide, 10%

Dulbecco’s Modified Eagle Medium (DMEM)) untilthey were used in the following experiments.

2.2. Preparation of the biomaterial scaffolds

Cryopreserved cells were thawed and plated in

stromal media for 5 to 7 days until the cultures became

confluent. Cells were harvested using trypsin/EDTA,

counted, and then loaded onto alginate, agarose, and

gelatin scaffolds as described. For the alginate scaffolds,

cells were suspended in 2% (w/v) low viscosity alginate

(Sigma) in 0.9% NaCl at a concentration of 107 cells/ml.

The cell suspensions were cast in custom molds (25 mm

diameter and 2 mm thickness). The alginate molds were

placed into a bath of 102 mm CaCl2 and allowed to gel

for 10 min. The CaCl2 was removed and the molds

were washed three times in PBS. Similarly, cells were

suspended in 2% (w/v) low-melting point agarose

(Type VII, Sigma) at a concentration of 107 cells/ml.

The agarose molds were allowed to gel at 4C for

10 min. Smaller alginate and agarose disks were then cut

out using a 6 mm diameter biopsy punch and placed in

the appropriate culture conditions.

Porous, absorbable gelatin (Surgifoams, Ethicon,

Inc., Somerville, NJ) disks (8mm diameter) were

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pre-wetted in culture medium in flat bottom tubes.

Aliquots (200ml) of the cell suspension (107 cells/ml)

were pipetted on top of each scaffold. The disk and cell

suspension were then centrifuged at 50 g for 30s to

seed the cells into the scaffold. The tubes containing the

seeded scaffolds were then incubated on an orbital

shaker at 5% CO2 and 37

C for 2h to enhance cellinfiltration into the scaffold. The disks were then

incubated undisturbed overnight to allow for cell

attachment. The following day, all disks were removed

from the tubes and grown in the appropriate culture

media. Acellular blank scaffolds were also prepared and

incubated in identical conditions.

2.3. Culture conditions

The scaffolds were grown in control or chondrogenic

culture media in a humidified environment at 5% CO2and 37C for up to 28 days, with culture media

replenished every 3 days. Control culture media

comprised Dulbecco’s Modified Eagle Media–high

glucose (DMEM-hg), 10% fetal bovine serum,

100 units/ml penicillin, and 100 mg/ml streptomycin.

The chondrogenic culture media comprised the control

media, 1 insulin-transferrin-selenium supplement

(ITS+, Collaborative Biomedical, Becton Dickinson,

Bedford, MA), 0.15 mm ascorbate 2-phosphate (Sigma),

100nm dexamethasone (Sigma), and 10 ng/ml rh-TGF-

b1 (R & D Systems, Minneapolis, MN) [26,34,35].

A broad array of biological, biochemical, and

mechanical analyses were used to assess the functional

properties of the scaffolds (Table 1).

2.4. Biological properties

Cell viability and cellular morphology were examined

in situ on days 1, 7, 14, and 28 using a confocal laser

scanning microscope (LSM 510, Carl Zeiss Microima-

ging, Inc., Thornwood, NY ) and the fluorescent Live-

Dead probes (Calcein AM and Ethidium homodimer,

Molecular Probes, Eugene, OR). To quantify the

protein and proteoglycan biosynthetic activity, con-

structs were dual-labeled with 10mCi/ml [3H]-proline

and 5mCi/ml [35S]-sulfate for 24 h on days 1, 7, 14 and

28. Afterwards, the scaffolds were washed 4 times to

remove unincorporated free label and then digested in1ml of a 50 mg/ml papain solution in glass scintillation

counting tubes at 65C overnight. Aliquots (100ml) of

the scaffold digests were sampled from each vial, diluted

to 1 ml, and stored at –80C for later DNA quantifica-

tion as described below. To the remaining 900 ml of the

construct digests, 4.5 ml of Hionic-Fluor Scintillation

Fluid was added (Packard Instrument Company,

Meriden, CT) and the [3H]-proline and [35S]-sulfate

disintegrations per minute were measured on a Model

1900TR Liquid Scintillation Analyzer (Packard).

2.5. Biochemical composition

The DNA content in the scaffold digests was

determined using the fluorescent picoGreen dsDNA

quantification assay (Molecular Probes, Eugene, OR).

Sulfated glycosaminoglycan (S-GAG) content of scaf-

fold digests was measured using a modified dimethyl-

methylene blue (DMB) assay in 96 well plates. Alginate

disk digests were analyzed using a pH 1.5 dye, as

previously described [36], while agarose and gelatin

digests were analyzed using a pH 3.0 dye.

Hydroxyproline (OHP) content was measured using

the Ehrlich’s reaction assay previously described [37,38].

Aliquots (50ml) of scaffold digests, after proper dilu-tions, were hydrolyzed in 6n HCl (Pierce) at 110C for

18h and then lyophilized. The samples were then

reconstituted in 200 ml of the assay buffer (5 g/l citric

acid (monohydrate), 12 g/l sodium acetate (trihydrate),

3.4 g/l sodium hydroxide, and 1.2 ml/l glacial acetic acid

in distilled water, pH 6.0). The reconstituted sample

solutions were subsequently filtered through activated

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Table 1

Summary of the techniques used to assess the functional properties of the scaffold materials

Functional assay Days in culture Sample sizea

Cellular viability andmorphology

Fluorescent calcein-AM and ethidium labeling, andImaging using confocal laser scanning microscopy (CLSM)

1, 7, 14, 28 days n ¼ 3

Protein and proteoglycan

biosynthesis rates

Radioactive [3H]-proline incorporation

Radioactive [35S]-sulfate incorporation

1, 7, 14, 28 days n ¼ 9

Biochemical composition DNA content (PicoGreen assay) 1, 7, 14, 28 days n ¼ 9

Sulfated glycosaminoglycan content (DMB assay)

Collagen content (hydroxyproline assay)

Immunohistochemistry Type II collagen

Chondroitin sulfate

28 days n ¼ 9

Biomechanical properties Equilibrium compressive modulus (step-wise stress-relaxation in

unconfined compression)

0, 14, 28 days n ¼ 9

Equilibrium shear modulus (step-wise shear stress relaxation)

Rheological properties (frequency sweep in dynamic shear)

aSample size per culture condition, per time point. Scaffolds were prepared from cells of three different donors.

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charcoal. Added to 50ml of each filtered sample in a 96

well plate was 50ml of 62mm chloramine-T (Sigma). The

mixture was incubated at room temperature for 15 min

to allow oxidation reaction. Oxidized samples were then

mixed with 50 ml of 0.94m dimethylaminobenzaldehyde

(p-DMBA) colorimetric solution and incubated at 37C

for 30 min. The optical densities of the assayed sampleswere measured using a plate reader at 540 nm, and

the OHP content of the samples was computed relative

to a standard curve of trans-4-hydroxy-L-proline

(0–300mg/ml).

For immunohistochemical analysis, scaffolds were

fixed in 10% buffered formalin for 24h at room

temperature. The fixed scaffolds were dehydrated in a

gradient of alcohols and then embedded in paraffin

blocks. Sections of 7 mm thickness were obtained from

each scaffold and mounted on microscope slides.

Following deparaffinization, rehydration, and endogen-

ous peroxidase activity quenching, the sections were pre-

digested for 60 min at room temperature with 0.25 units/

ml chondroitinase ABC, or for 10 min at 37C with

pepsin to allow for the retrieval of the chondroitin

sulfate (CS) or collagen type II (Coll-II) antigens,

respectively. Sections were then incubated with the

2B6 monoclonal antibody specific to CS (kind gift from

Dr. Virginia Kraus, Duke University Medical Center)

overnight at 4C or with the II-II6B3 antibody specific

to Col II (Developmental Studies Hybridoma Bank,

Iowa City, IA) for 1 h at 37C, respectively. Immunos-

taining was detected using Histostain-Plus Kit for AEC

(Zymed Laboratories Inc., San Francisco, CA).

2.6. Biomechanical properties

The elastic compressive modulus was determined

from equilibrium stress–strain curves generated from

stepwise stress relaxation tests in unconfined compres-

sion at strains of 4%, 8%, 12%, and 16%. Similarly, the

elastic shear modulus was determined from equilibrium

shear stress–strain curves from stepwise shear stress-

relaxation (pure torsion) experiments at shear strains

of (0.03, 0.04, and 0.05 rad). Following the stress-

relaxation tests, the rheological properties of the

constructs were determined by subjecting the samples

to oscillatory shear strain gðtÞ ¼ go:sinðotÞ of afixed amplitude (go ¼ 0:05 rad) and varying frequency(1–100 rad/s). The resultant oscillatory shear stress

sðtÞ ¼ so:sinðot þ dÞ was recorded and the rheologicalproperties such as the complex shear modulus

jG ðoÞj2 ¼ ½G 0ðoÞ2 þ ½G 00ðoÞ2 and the loss angle d were

determined for each of the applied frequencies; where G 0

is the storage modulus [G 0ðoÞ ¼ so:cosðdðoÞÞ=go] and G 00

is the loss modulus [G 00ðoÞ ¼ so:sinðdðoÞÞ=go]. Biome-chanical tests were performed in a bath of DMEM at

room temperature using an ARES Rheometrics System

(Rheometric Scientific, Piscataway, NJ).

2.7. Statistical analysis

Analysis of variance with Student–Newman–Keuls

(SNK) multiple ranges tests were used to compare the

different biomaterials and culture conditions (a ¼ 0:05).Data was further examined in a multiple linear

regression context to determine which of the biochem-ical parameters (DNA, OHP, S-GAG) or combination

of parameters correlated with the biomechanical proper-

ties (E ; G ; d; |G |). Statistical analyses were performedusing Statistical Analysis Software (SASs, Cary, NC)

and S-Pluss (Insightful Corp. Seattle, WA).

3. Results

3.1. Biological properties

Cell viability and morphology in the different scaffold

materials were visualized using confocal laser scanning

microscopy and the Live-Dead fluorescent probes. All

scaffolds showed relatively uniform distributions of cells

with viability greater than 95% at all time points. Cells

in agarose (Fig. 1a) and alginate (Fig. 1b) scaffolds

displayed a spherical morphology that persisted

throughout the culture period, regardless of culture

conditions. In contrast, the cells in the gelatin scaffolds

(Fig. 1c) displayed a distinct ‘‘fibroblastic’’ morphology.

By day 28, the cells in the gelatin scaffolds proliferated

and became confluent with notable cell-to-cell contact

that was associated with the significant cell-mediated

contraction of the gelatin disks, with reduction of up to70% and 87% their initial diameters under chondro-

genic and control culture conditions, respectively

(Fig. 1d). Alginate and agarose disks containing cells

and acellular gelatin disks did not exhibit any contrac-

tion. Furthermore, culture conditions did not appear to

affect cell morphology, but cells were sparse in control

conditions compared to chondrogenic conditions.

Protein and proteoglycan biosynthesis rates were

quantified by [3H]-proline (Fig. 2a) and [35S]-sulfate

(Fig. 2b) incorporation, respectively. Biosynthesis rates

in the hydrogel (agarose and alginate) scaffolds were

significantly greater in chondrogenic conditions com-

pared to control conditions (1.2–20 fold greater; data

not shown). However, for scaffolds grown in chondro-

genic conditions, protein and proteoglycan biosynthesis

rates were significantly lower for the agarose scaffolds

than those of the alginate and gelatin scaffolds

throughout most of the culture ( po0:05). Whennormalized by the DNA content (Figs. 2c and d),

protein and proteoglycan biosynthesis rates in the

gelatin scaffolds were significantly greater than agarose

(31%) and alginate (68%), respectively, on day 1

( po0:05). However, the differences between biosynth-esis rates in the different scaffolds diminished by days 14

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and 28. Regardless of culture conditions or scaffold

biomaterial, incorporation rates decreased significantly

with time ( po0:05).

3.2. Biochemical composition

Biochemical analysis of the scaffolds was performed

to quantify the DNA, S-GAG, and OHP content of

the scaffolds at different times in culture. The DNA,

S-GAG, and OHP content of scaffolds grown in

chondrogenic conditions were significantly higher than

their control conditions counterparts on days 7, 14,

and 28. For scaffolds grown in chondrogenic conditions,

the DNA content of gelatin scaffolds was 37–51%

greater than agarose and alginate scaffolds on days 14

and 28 (Fig. 3a, po0:05). DNA content increased,reaching peak values of nearly 4.3 mg/scaffold on day 7

for the agarose and alginate and 6.5mg/scaffold on day

14 for the gelatin with insignificant declines afterwards.

For scaffolds grown in chondrogenic conditions, there

were no significant differences in S-GAG content among

the scaffold materials (Fig. 3b), whereas the OHP

content in gelatin was 28–47% greater than agarose

and alginate on days 14 and 28 (Fig. 3c, po0:05).

Furthermore, the S-GAG and OHP content increased

significantly by 2.5–9 fold between days 1 and 28 for all

scaffold materials grown in chondrogenic conditions.

When normalized by DNA content, S-GAG (Fig. 4a)

and OHP (Fig. 4b) content increased significantly

between days 1 and 28 for all scaffold materials grown

in chondrogenic conditions ( po0:05). However, ingeneral, there were no significant differences between

the different scaffold materials.

The accumulation of cartilage matrix macromolecules

was evident in the agarose and alginate hydrogel in disks

cultured under chondrogenic conditions, as demon-

strated by the positive immunohistochemical staining

against the 2B6 epitope of chondroitin sulfate and type

II collagen (Fig. 5). The staining was most intense in the

pericellular matrix, characteristics associated with cells

found in native cartilage. Positive staining against the

same antigens was also observed in the gelatin scaffolds.

In areas of sparse cell density where little contraction

had occurred, the staining was confined to regions of

neomatrix within the folds of the scaffold (Fig. 5c). By

contrast, in regions where significant contraction has

occurred, there was intense staining of both antigens in a

hyper cellular matrix (Fig. 5f ).

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Fig. 1. Cell viability and morphology in the different scaffold materials visualized using confocal laser scanning microscopy and the Live–Dead

fluorescent probes. Cells in agarose (a) and alginate (b) scaffolds displayed a spherical morphology that persisted throughout the culture period,

regardless of culture conditions. By contrast, the cells in the gelatin scaffolds (c) displayed a distinct ‘‘fibroblastic’’ morphology at day 7. By day 28,

the cells in the gelatin scaffolds proliferated and became confluent with significant cell–cell contact as they exerted considerable contraction of the

scaffolds (d).



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3.3. Biomechanical properties

The stress–strain behavior of all scaffold materials

was linear in both compression and shear stress-

relaxation experiments over the range of strains used.

There were no significant effects of culture conditions on

the mechanical properties of the hydrogel materials

(agarose and alginate), however, after 28 days gelatin

scaffolds grown in chondrogenic conditions had equili-

brium compressive and shear moduli that were 86%,

and 160%, respectively, greater than gelatin scaffolds

grown in control conditions ( po0:05; data not shown).The equilibrium compressive modulus (Fig. 6a) showed

a 48% and 53% softening between days 0 and 14 for the

agarose and alginate disks, respectively ( po0:05). Bycontrast, the equilibrium compressive modulus of

gelatin scaffolds increased by 3.9 and 4.6 fold of day

zero values by days 14 and 28, respectively ( po0:05).The equilibrium shear modulus (Fig. 6b) increased

significantly over time in chondrogenic culture by 2.6,

1.8 and 6 folds for the agarose, alginate, and gelatin

scaffolds, respectively. Likewise, the complex shear

modulus at o ¼ 10rad/s and go ¼ 0:05 (Fig. 6c)increased over time by 2.5, 1.8 and 8.3 folds for the

agarose, alginate, and gelatin scaffolds, respectively. By

the end of 28 days in chondrogenic culture, the

equilibrium shear modulus of alginate and gelatin was

22% and 67% of agarose ( po0:05). Similarly, thedynamic shear modulus at o ¼ 10 rad/s (|G (10 rad/s)|)

of alginate and gelatin was 22% and 79% of agarose,

respectively ( po0:05) (Fig. 7a). Values for the complexshear modulus jG ðoÞj showed linear trends with the

logarithm of frequency. The loss angle (d) showed no

specific trends with frequency (Fig. 7b). The loss angle

for all scaffold materials was less than 15, indicating

that all scaffolds tested behave like viscoelastic solids.

Multiple linear regression analysis suggested that

increases in G ; and jG j; were significantly associatedwith increases in SGAG content (Table 2, po0:05).Increases in E and d were associated with increases in

OHP content, though not significantly (Table 2,

p ¼ 0:09).

4. Discussion

Cellular based tissue engineering approaches have

increasingly used adult stem cells from different sources

including bone marrow [39–42], trabecular bone [43],

muscle [44], and adipose tissue [26,31–33,45–48]. Despite

the many advantages of these abundant and accessible

cells, progress in their utility in tissue engineering has

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Fig. 2. Protein and proteoglycan synthesis rates in chondrogenic culture conditions, quantified by [3H]-proline (a) and [35S]-sulfate (b) incorporation,

respectively, were significantly higher for the gelatin scaffolds compared to agarose and alginate cultured in chondrogenic conditions early in culture.

When normalized by the DNA content (c and d), the differences between biosynthesis rates in the different scaffolds diminished especially at later

times in culture. Data presented are mean7standard deviation. po0:05; po0:01:



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been limited by our ability to exercise precise control

over the cells’ differentiation potential. While much of

the stem cell based-tissue engineering research has been

focused on controlling differentiation using soluble

chemical factors (such as growth factors) and/or

manipulating the mechanical signals to which the cells

are exposed, less attention has been paid to the

importance of the biomaterial scaffold in regulating

differentiation and tissue growth [10,11]. In this study,

we corroborated our previous findings that hADAS cells

can differentiate into a chondrocytic phenotype under

defined culture conditions [26,49]. We further demon-

strated that the functional properties of tissue engi-

neered-cartilage vary with culture conditions, culture

time, and the choice of the scaffolding biomaterial.

Culturing the hADAS cell-laden constructs in chon-

drogenic media containing TGF-b1 significantly in-

creased DNA, S-GAG, and OHP content over control

conditions by nearly 3 fold in the hydrogel materials and

1.5 fold in the gelatin disks. In general, there were no

significant differences in the S-GAG content of all

scaffolds cultured in chondrogenic conditions, although

the DNA and OHP contents in the gelatin scaffolds was

greater than those in the hydrogels late in culture.

Furthermore, the biosynthesis rates of proteins and

proteoglycans were significantly higher for gelatin disks

compared to the agarose and alginate hydrogels. While

these results are similar to the previous observation that

TGF-b1 stimulates chondrogenic differentiation of

adult stem cells [49], they also imply that the cells’

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Fig. 3. Biochemical analysis of the scaffolds at different times in chondrogenic culture conditions. The DNA content (a) of gelatin scaffolds was

significantly higher than agarose and alginate on days 14 and 28. DNA content in all scaffolds increased initially reaching peak values on day 7 for

the agarose and alginate and on day 14 for the gelatin with insignificant declines afterwards. There were no significant differences in s-GAG content

among the scaffold materials (b). OHP content in gelatin was significantly higher than agarose and alginate on days 14 and 28 (c). Data presented are

mean7standard deviation. po0:05; po0:01:



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response to chondrogenic mediators may depend on the

physical and biological properties of the biomaterial

scaffold. These properties may include the diffusion

rates that regulate nutrient and metabolic waste trans-

port, the ability to regulate the cellular morphology that

is thought to affect differentiation, and the presence of

bioactive ligands that can provide anchorage sites forcell attachment.

We have recently shown that molecular diffusion

coefficients in agarose, alginate, and gelatin constructs

seeded with hADAS cells, measured by fluorescence

recovery after photobleaching (FRAP) of dextran

molecules of equal or greater size than culture nutrients

and growth factors (70 kDa), are at least twice those in

native cartilage even after 28 days in culture with no

significant differences among the constructs [50]. The

implications of these findings are two fold: (1) the

differences in molecular diffusion kinetics do not fully

explain the differences in cell and tissue growth in these

constructs and (2) the transport of nutrients and

metabolites to cells within the constructs is not hindered

in the early stages of tissue generation.

The ability of the scaffolds to biologically interact

with the cells, however, may explain some of the

differences in tissue growth within the constructs. The

entrapment of the cells in the hydrogels imposed a

spherical cellular morphology, whereas the cells seeded

in gelatin displayed various morphologies but were

predominantly of fibroblastic morphology. These find-

ings, together with the biochemical and biological

properties we measured, suggest that the beneficial

effects of spherical cellular morphology in chondrogen-esis may be hindered in the absence of bioactive cell

attachment ligands within the matrix. In addition, the

importance of providing a natural substrate for cell

attachment is manifested by the cell-mediated contrac-

tion of the gelatin scaffolds and the concomitant

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Fig. 4. S-GAG and hydroxyproline content normalized by DNA

content. The normalized s-GAG (a) and hydroxyproline (b) content

increased significantly between days 1 and 28 for all scaffold materials.

However, in general, there were no significant differences between the

different scaffold materials. Data presented are mean7standard

deviation. po0:05:

Fig. 5. Immunohistochemical sections of the agarose (a,d), alginate (b,e), and (c, f) scaffolds cultured up to 28 days in chondrogenic conditions.

Sections were stained for antibodies against the 2B6 epitope of chondroitin sulfate (CS; 20 , Scale bar 50m) and Collagen type II (Coll-II; 40 ,

Scale bar 100 m).



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increases in cell proliferation and collagen synthesis,

which supports previous findings in the literature

[51,52]. Moreover, recent studies have demonstrated

that overcoming the inert nature of the alginate

hydrogel by inserting RGD-containing peptide se-

quences promotes cell multiplication and cellular and

structural organization [53].The growth of the tissue engineered cartilage con-

structs appears to have progressed in two stages that

depend upon the biomaterial scaffold. The first stage, a

cell growth phase, was characterized by increased cell

proliferation and lasted up to 7 days in the hydrogel

materials and continued up to 14 days in the gelatin

scaffolds. The second stage can be described as a cell

differentiation and tissue growth stage, which was

characterized by decreased proliferation and increased

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Table 2

Multiple linear regression correlations between the biomechanical

properties and the biochemical composition of the tissue-engineered

cartilage constructs

Biomechanical Biochemical Slope Intercept R2 p

E OHP 30.56 8.67 0.28 p ¼ 0:08

SGAG 29.2 7.69 0.22 p ¼ 0:12SGAGOHP 109.11 9.51 0.19 p ¼ 0:16

G OHP 10.83 1.56 0.2 p ¼ 0:15SGAG 14.89 0.87 0.32 po0:05

SGAGOHP 56.55 1.67 0.28 p ¼ 0:07d OHP 21.25 5.44 0.26 p ¼ 0:09

SGAG 13.96 5.42 0.10 p ¼ 0:32SGAGOHP 94.88 5.82 0.28 p ¼ 0:08

|G | OHP 14.98 1.83 0.24 p ¼ 0:10SGAG 19.62 0.81 0.36 po0:05

SGAGOHP 76.72 1.99 0.34 po0:05

Fig. 6. Biomechanical properties of the scaffold materials at different times in chondrogenic culture. The elastic compressive modulus (a) of agarose

and alginate showed no improvements between days 0 and 28 following a significant decrease on day 14. By contrast, the elastic compressive modulus

of gelatin scaffolds increased significantly with time reaching values comparable with agarose on day 28. The elastic shear modulus (b) and the

dynamic shear modulus (c), at o ¼ 10 rad/s and go ¼ 0:05; increased significantly with time for all scaffold materials, and was significantly highest foragarose. Data points represent the mean7SEM. po0:01:



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collagen and proteoglycan deposition. An improved

understanding of the interactions between the cell and

the biomaterial scaffold (cell shape, cell–cell contact,

cytoskeletal changes, and cell–matrix interactions) may

improve the ability to control the switch between

cellular proliferation and differentiation [51].

The compressive Young’s modulus of the hydrogel

scaffolds (alginate and agarose) decreased in value

between days 0 and 14, suggesting that the hydrogels

have experienced softening, possibly due to a loss of the

cross-linking Ca2+ cations [54] in alginate and a loss in

gel stability due to thermal factors in agarose. However,

the subsequent increase in the compressive moduli in

agarose and alginate between days 14 and 28 is likely

due to increases in new matrix synthesis. The compres-

sive Young’s modulus of the gelatin scaffolds increasedsignificantly over time, reaching values comparable to

agarose by day 28. The mechanism behind the stiffening

of the gelatin scaffolds likely involves the packing of the

scaffold material as a result of the cell-mediated

contraction and the deposition of matrix macromole-

cules within the scaffold. Similarly, the equilibrium and

dynamic shear moduli increased significantly over time

for all scaffold materials, with agarose and gelatin

having shear moduli almost 3 times greater than

alginate. However, it should be noted that the compres-

sive and shear moduli of these scaffolds are on the order

of 5% or less of those of native cartilage [55,56]. These

results are not different from observations reported by

others [17,54,57], although it was recently suggested that

the mechanical properties of agarose disks can be

improved by increasing cell seeding density and the

application of a compressive loading regimen to

stimulate neotissue synthesis [17,58]. Even with the use

of large numbers of chondrocytes and prolonged culture

periods in these studies, the application of mechanical

loading improved the functional properties of the

chondrocyte-seeded agarose constructs to no more than

20% of native cartilage [58]. These studies, while

demonstrating the importance of in vitro mechanical

conditioning and cell seeding density, indirectly under-

score the importance of the inherent mechanical proper-

ties of the biomaterial scaffold and suggest that even the

use of large numbers of cells and prolonged culture

periods might not be sufficient to overcome the

mechanical deficiencies of the hydrogels.

Our results also indicate that increases in G and jG j

for all scaffolds are significantly correlated with

increases in S-GAG content and with the interaction

between S-GAG and OHP (Table 2), but not with OHP

alone. This intriguing finding is similar to previous

studies using chondrocytes in agarose disks [58] and

gives more credence to the presence of a structure–

function relationship in these tissues. Although our

analysis did not examine the structure of the developing

collagen and proteoglycan networks, our data indicatethat manipulating the composition and structure of

these tissue-engineered constructs may have important

implications on the construct’s ability to assume their

mechanical functions.

In conclusion, it is quite apparent that the biomaterial

scaffold of choice influences the growth and differentia-

tion of adult stem cells. Biologically active biopolymers,

such as gelatin, have distinct advantages stemming from

the fact that they modulate cell functions in manners

that can be exploited to create biologically functional

tissue-engineered grafts. While neither of the biomater-

ials studied approached native cartilage mechanical

properties they demonstrated significant composition-

function relationships that could provide important

clues for engineering functional tissues.

Acknowledgements

This study was supported in part by Artecel Sciences,

Inc., the North Carolina Biotechnology Center, the

Kenan Institute, and NIH grant AR49294. We would

like to thank Dr. Lori Setton and Charlene Flahiff for

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Fig. 7. Typical dynamic data for the frequency sweep (shear) response for the different scaffold materials at go ¼ 0:05 and o ¼ 12100rad/s.Scaffolds were cultured 28 days in chondrogenic conditions. Data points represent the mean7SEM.



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help with the biomechanical testing; and Julie Fuller and

Steve Johnson for help with the histology.

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Documents

1.Hani a. Awad -Chondrogenic Differentiation of Adipose-Derived 2003