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Modulation of Osteogenic Differentiation ofHuman Mesenchymal Stem Cells by Poly[(L-lactide)-co-(e-caprolactone)]/Gelatin Nanofibers

Nae Gyune Rim, Ji Hye Lee, Sung In Jeong, Bu Kyu Lee, Chun Ho Kim,Heungsoo Shin*1

Developing biomaterial scaffolds to elicit specific cell responses is important in many tissueengineering applications. We hypothesized that the chemical composition of the scaffold maybe a key determinant for the effective induction of differentiation in humanmesenchymal stemcells (hMSCs). In this study, electrospun nanofibers with different chemical compositions werefabricated using poly[(L-lactide)-co-(e-caprolactone)] (PLCL) and gelatin. Scanning electronmicro-scopy (SEM) images showed a randomly arranged structure of nanofibers with diametersranging from 400nm to 600nm. The incorporation of gelatin in the nanofibers stimulatedthe adhesion and osteogenic differentiation of hMSCs. For example, the well-stretched andpolygonal morphology of hMSCs was observed on the gelatin-containing nanofibers, while thecells cultured on the PLCL nanofibers were contracted. The DNA content and alkaline phos-phatase activity were significantly increased onthe PLCL/gelatin blended nanofibers. Expression ofosteogenic genes including alkaline phosphatase(ALP), osteocalcin (OCN), and collagen type I-a2 (ColI-a2) were also upregulated in cells cultured onnanofibers with gelatin. Mineralization of hMSCswas analyzed by von Kossa staining and theamount of calcium was significantly enhancedon the gelatin-incorporated nanofibers. Theseresults suggest that the chemical composition ofthe underlying scaffolds play a key role in regulat-ing the osteogenic differentiation of hMSCs.

N. G. Rim, J. H. Lee, H. ShinDepartment of Bioengineering, Hanyang Fusion MaterialsProgram, College of Engineering, Hanyang University, 17Haengdang-dong, Seongdong-gu, Seoul 133-791, Republic of KoreaFax: þ82-2-2298-2346; E-mail: [email protected]. I. JungDepartment of Chemical Engineering, College of Engineering,Hanyang University, 17 Haengdang-dong, Seongdong-gu, Seoul133-791, Republic of Korea

B. K. LeeDepartment of Oral and Maxillofacial Surgery, Asan MedicalCenter, College of Medicine, Ulsan University, Seoul 138-736,Republic of KoreaC. H. KimLaboratory of Tissue Engineering, Korea Institute of Radiologicaland Medical Sciences, Seoul 139-240, Republic of Korea

Macromol. Biosci. 2009, 9, 795–804

� 2009 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim DOI: 10.1002/mabi.200800358 795

N. G. Rim, J. H. Lee, S. I. Jung, B. K. Lee, C. H. Kim, H. Shin

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Introduction

Adult mesenchymal stem cells (MSCs) can be isolated from

bone marrow and expanded in vitro while maintaining

their multi-potent differentiation capability, resulting in

the formation of multiple mesenchymal tissues including

bone, cartilage, and muscle.[1–4] One of the major strategies

used in bone tissue engineering is the use of a scaffold that

may serve either as a delivery vehicle for ex vivo cultured

MSCs or an active stimulant to recruit MSCs into the scaffold

to enhance tissue repair processes.[5] Control of the

induction and maintenance of osteogenic differentiation

of MSCs by well-defined scaffolds is therefore of crucial

importance. Ideally, the scaffold is required to provide a

regenerative or reparative environment and newly formed

bone should be integrated with native tissue while the

scaffold materials are degraded.[6]

Bone tissue engineering scaffolds have been prepared by

numerous methods and different materials including

synthetic and natural polymers.[7–9] Poly(a-hydroxy esters)

are the most popular source for this purpose because of their

good biocompatibility, tunable biodegradability, and

mechanical properties.[10–12] Nevertheless, use of these

materials often results in poor control of cell–biomaterial

interactions due to lack of bioactivity; this represents a

major hurdle in the use of MSCs in bone repair. Therefore,

incorporation of natural biomacromolecules such as

chitosan, collagen, and gelatin into synthetic polymers

has been widely employed to create scaffolds that mimic

the biological and chemical features of the cellular

environment of the native tissue.[13–15] This ‘‘biomimetic’’

approach may improve the bioactivity of synthetic

material-based scaffolds and thereby control the directed

differentiation of MSCs.

Cells in native tissue reside within an extracellular matrix

(ECM) environment where many biomacromolecules are

aligned on a nanometer scale with complex structure.[16] In

addition to chemical signals in the ECM milieu, the physical

features of the ECM may also affect cell adhesion,

differentiation, and programmed death. Therefore, several

techniques to fabricate scaffolds with a controlled archi-

tecture on a nanometer scale have been employed to control

MSC function. In particular, the electrospinning process has

recently been used in tissue engineering to produce a

nanofibrous scaffold comprising a nanoscaled, non-woven,

and porous structure with a high surface area to volume

ratio.[17–19] Studies using nanofibers from poly[(D,L-lactide)-

co-glycolide] (PLGA) and poly(e-caprolactone) showed osteo-

genic differentiation of human MSCs.[10,15,20] Despite several

promising results, the ideal scaffold with defined chemical

and structural features to actively control directed differ-

entiation of MSCs has not yet been identified.

The overall objective of this study was to prepare

biomimetic nanofibrous scaffolds that can modulate the


� 2009 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

osteogenic differentiation of human MSCs (hMSCs). First,

the scaffolds were fabricated by electrospinning the

polymer blend of gelatin and poly[(L-lactide)-co-(e-capro-

lactone)] (PLCL) in various blending ratios. The resulting

scaffolds were used to test the response of hMSCs derived

from human mandible. Specifically, the effect of the

chemical composition of the nanofibrous scaffolds on the

maintenance of osteogenic differentiation of hMSCs was

investigated, as measured by gene expression, enzymatic

activity, and mineral deposition.

Experimental Part

Materials

Gelatin and 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide

hydrochloride (EDC) were purchased from Sigma Chemical (St.

Louis, MO, USA). PLCL was synthesized as described previously.[21]

Dulbeccos’s modified eagle medium (DMEM), phosphate-buffered

saline (PBS), trypsin-ethylenediaminetetraacetic acid (EDTA), and

penicillin–streptomycin were purchased from Gibco BRL (Rockville,

MD, USA). Fetal bovine serum (FBS) was purchased from Hyclone

(Logan, UT, USA). Ascorbic acid and sodium dodecyl sulfate (SDS)

were purchased from Amresco (Solon, OH, USA). Glutaradehyde,

NaCl, NaOH, and formaldehyde were purchased from Junsei

Chemical (Japan). Tri-reagent, rhodamine–phalloidin, and Hoechst

22358 were purchased from Invitrogen (Carlsbad, CA, USA). Water

(DW) was distilled and deionized using a Milli-Q System (Waters,

Millipore, MA, USA). Unless otherwise specified, all other chemicals

and solvents were obtained from Sigma (St. Louis, MO, USA).

Preparation of Composite Nanofibrous Scaffolds

PLCL/gelatin nanofibers were prepared using an electrospinning

technique as described previously with minor modification.[13]

Briefly, PLCL and gelatin were separately dissolved in 2,2,2-

trifluoroethantol (TFE) at a concentration of 8% (w/v) and then

mixed in four different volume ratios (10:0, 7:3, 3:7, and 0:10 of

PLCL:gelatin). The PLCL/gelatin blends were loaded in a syringe

fitted with a 23-gauge needle and directly electrospun to the

collector that was covered with aluminium foil. The diameter of the

round collector was 9 cm and the spinning rate of the collector was

200 rpm. The distance between the needle and the collector was

18 cm. The electrospinning conditions including the flow rate,

applied voltage, and designation of corresponding nanofibers are

shown in Table 1. Following the electrospinning process, the

nanofibers were crosslinked with 0.35% EDC for 10 min in an

acetone/water mixture (9:1 by volume). The crosslinked nanofibers

were rinsed with DW to remove any residual chemicals, and freeze-

dried for 2 d. A scanning electron microscope (SEM, JEOL JSM-6300,

Japan) was used to examine the morphology of the nanofibers. We

took pictures of 3 independent nanofibers for each group using SEM

and selected 15 fields from each image for measurement of the fiber

diameter using NIS-elements AR 3.0 software (Nikon, Tokyo,

Japan).

DOI: 10.1002/mabi.200800358

Modulation of Osteogenic Differentiation of Human Mesenchymal Stem Cells . . .

Table 1. Sample conditions for PLCL/gelatin nanofibers.

Sample code Gelatin in PLCL/gelatin solutiona) Voltage Flow rate

% kV mL �h�1

PG10 0 14 2.0

PG73 30 17 2.0

PG37 70 15 2.0

PG01 100 18 1.5

a)PLCL and gelatin (8%) were separately dissolved in 2,2,2-trifluoroethanol, and a 23-gauge needle was used for electrospinning.

In vitro Culture of hMSCs

hMSCs extracted from the mandibular bone marrow were kindly

donated from Asan Medical Center (Seoul, Korea). For examination

of the in vitro cellular responses of the hMSCs, the nanofibers were

cut into a round shape (3 cm in diameter) and sterilized by

immersion in 70 vol.-% ethanol under ultraviolet (UV) light for

30 min. After sterilization, the nanofibers were rinsed with PBS

three times and pre-wetted in serum-free media overnight at 37 8C.

We used seventh-passage cells for the study. hMSCs were

enzymatically detached using 0.25% trypsin–EDTA and then

resuspended with the growth media and counted with a

hemocytometer. The prepared cell suspension was added to

scaffolds that had been placed within 6-well tissue culture dishes

at a density of 6.4�103 cells � cm�2. After 1 day, the media was

refreshed with osteogenic differentiation media, which contained

50mg �mL�1 of ascorbic acid, 0.01 M of glycerol-2-phosphate,

and 10�7M of dexamethasone in complete medium. The hMSCs

cultured on the nanofibers were maintained in standard culture

conditions for up to 14 d with media changes every 2 or 3 d.

Analysis of in vitro Differentiation of hMSCs on the

Nanofibers

DNA Contents

The cellularity was measured using a Picogreen ds DNA assay kit

(Molecular Probes, OR, USA). After 7 and 14 d of culture, cell-scaffold

constructs were washed with PBS and placed in a sterile 1.5-mL

microcentrifuge tube. Cells were lysed in RIPA buffer (150�10�3M

NaCl, 1% Triton X-100, 1% sodium deoxycholate, 0.1% SDS,

150�10�3M Tris, pH 7.2, and protease inhibitors) and homo-

genized with a homogenizer (Ultra Turbax T10, IKA Lab, Germany)

on ice. The cell lysate was mixed with working solutions according

to the manufacturer’s instructions. A standard curve was obtained

using a diluted lambda DNA solution with a known concentration

and the fluorescence intensity was measured using a spectro-

fluorometer at an excitation wavelength of 480 nm and an

emission wavelength of 520 nm (SpectraMax M2e, Molecular

Devices, CA, USA).

Alkaline Phosphatase Activity

After 7 and 14 d of culture, alkaline phosphatase (ALP) activity was

analyzed using p-nitrophenyl phosphate (p-NPP, Sigma Chemical,



St. Louis, MO, USA). After centrifugation of the cell–scaffold lysate

at 12 000 g for 10 min, aliquots of the supernatants were incubated

with p-NPP at 37 8C for 30 min. NaOH solution (3 N) was added to

stop the reaction and the absorbance of the samples was measured

at 405 nm. A p-nitrophenol standard solution was diluted and used

to generate a standard curve.

Real Time Reverse Transcription-Polymerase ChainReaction

The mRNA expression of glyceraldehyde 3-phosphate dehydro-

genase (GAPDH), core binding factor alpha 1 (Cbfa 1), and collagen

type I-a2 (Col I-a2) was determined by reverse transcription-

polymerase chain reaction (RT-PCR). After total RNA was obtained

from the cell–scaffold constructs using a tri-reagent, the RNA was

reverse transcribed to cDNA and 150 ng of cDNA was used for each

PCR reaction (in 20mL). PCR reaction conditions were 95 8C for 30 s,

55 8C (Cbfa 1: 57 8C, Col I-a2: 58 8C) 30 s, and 72 8C for 60 s. A final

extension was performed for 5 min at 72 8C. Reaction products

(9mL) were separated by gel electrophoresis using a 1% agarose gel

and stained with ethidium bromide. Bands were captured using a

Gel Logic 100 Imaging System (Kodak, Rochester, NY, USA) under

UV illumination.

The quantitative mRNA expression ofGAPDH,ALP, andOCNwas

characterized by real time RT-PCR using SYBR Green dye (Takara,

Japan) and Thermal Cycler Dice (Model TP800, Takara, Japan). The

reaction was performed with 40 ng of the reverse transcribed RNA

under reaction conditions of 95 8C for 5 s and 60 8C for 30 s. All gene

expression was normalized by GAPDH using a comparative cycle

threshold (ddCt) method, and the primer sequences used in the

experiments are listed in Table 2.

Calcium Assay and von Kossa Stain for Analysis ofMineralization

After 14 d of culture, the cell–scaffold constructs were washed with

PBS, then immersed in 0.6 N HCl, and homogenized. Calcium was

extracted overnight at 4 8C and the amount of calcium in the acidic

supernatant was quantified using a Quantichrom Calcium Assay

Kit (Bioassay systems, CA, USA) according to the manufacturer’s

instruction. The absorbance was measured at 612 nm, and the

concentration of calcium in the samples was calculated using

diluted calcium chloride standard solutions. For von Kossa staining,

the cell–scaffold constructs were fixed with 3.7% formaldehyde,

stained with 2% silver nitrate solution, and exposed to a 60-W lamp

www.mbs-journal.de 797


Table 2. Primer sequences used for RT-PCR and real time RT-PCR.

Gene Primer sequences Size Accession number

bp

GAPDH Forward: ACC ACA GTC CAT GCC ATC AC 452 NM_002046

Reverse: TCC ACC ACC CTG TTG CTG TA

CBFA1 Forward: CCG CAC GAC AAC CGC ACC AT 289 NM_004348

Reverse: CGC TCC GGC CCA CAA ATA TC

COL1 Forward: GGA CAC AAT GGA TTG CAA GG 461 NM_000089

Reverse: TAA CCA CTG CTC CAC TCT GG

GAPDHa) Forward: GCA CCG TCA AGG CTG AGA AC 142 NM_002046

Reverse: ATG GTG GTG AAG ACG CCA GT

OCNa) Forward: CCC AGG CGC TAC CTG TAT CAA 112 NM_199173

Reverse: GGT CAG CCA ACT CGT CAC AGT C

ALPa) Forward: GGA CCA TTC CCA CGT CTT CAC 137 NM_000478

Reverse: CCT TGT AGC CAG GCC CAT TG

a)Real time RT-PCR.

798

for 1 h. After the exposure, the unreacted silver was removed from

the samples with 5% sodium thiosulfate for 2 min.

Scanning Electron Microscopy

After 14 d of culture, the morphology of the cell–scaffold constructs

was examined by SEM. The samples were washed twice with PBS,

then fixed with 1% glutaraldehyde for 30 min and 1% formaldehyde

for 1 day. The fixed specimens were dehydrated by serially diluted

ethanol, coated with gold using a sputter coater (Eiko IB3, Japan),

and observed by SEM.

Immunofluorescence Staining

To investigate the morphology of the cells on the nanofibers

scaffolds, the cells were fixed in 3.7% formaldehyde in PBS for

10 min, and permeabilized with cold cytoskeleton (CKS) buffer

(50� 10�3M NaCl, 150�10�3

M sucrose, 3� 10�3M MgCl2,

50� 10�3M tris-Base, 0.5% Triton X-100 20mg �mL�1 aprotinin,

1mg �mL�1 leupeptin) for 5 min at 4 8C. The samples were then

incubated in a blocking buffer (5% FBS, 0.1% Tween-20, 0.02%

sodium azide in PBS) for 60 min at 37 8C. After washing with PBS, the

samples were subsequently incubated for an additional 60 min in

rhodamine–phalloidin (1:200) and Hoechst dye (1:10 000). Samples

were visualized on an upright microscope equipped with the

appropriate fluorescence filters. Digital images were acquired using

fluorescence microcopy (TE 2000E, Nikon, Japan).

Statistical Analysis

Quantitative data were obtained in triplicate (n¼3) and reported

as the mean� standard deviation, where indicated. Statistical

analysis was performed using Student’s t-test and a P value of less

than 0.05 was considered statistically significant.



Results and Discussion

Morphology of Electrospun PLCL/Gelatin Nanofibers

The PLCL/gelatin composite nanofibers were generated as a

blend of the synthetic polymer, PLCL, and the natural

polymer, gelatin. Each polymer was separately dissolved

with TFE at 8 wt.-%, and each solution was then mixed at the

given ratio and electrospun to form nanofibers according to

the conditions shown in Table 1. The continuity and

morphology of nanofibers are affected by several working

parameters including external humidity, temperature, flow

rates, concentration of polymer solutions, and applied

voltage.[17,18] Under the given electrospinning conditions,

we were able to facilitate the formation of continuous fibers

with nanoscale diameters devoid of bead or polymer

aggregates. Because gelatin dissolves easily in water, the

composite nanofibers were subsequently crosslinked before

further use. For the crosslinking, we used EDC to trigger an

intramolecular reaction between activated carboxylic acid

and free amine functional groups within gelatin, thereby

resulting in the formation of stable amide bonds. This

process has been used widely in crosslinking natural

polymers because carbodiimine coupling built up by EDC

is a one-step reaction and is relatively non-toxic.[22–24]

As shown in Figure 1a–d, the nanofibers had a randomly

arranged structure with interconnected pores and smooth

surface. Most importantly, the incorporation of gelatin had

no effect on the continuity and morphology of the

nanofibers. The diameters of PLCL/gelatin nanofibers were

quantitatively characterized, and were within the range of

DOI: 10.1002/mabi.200800358


400–600 nm (Figure 1e). As compared to PLCL-only or

gelatin-only nanofibers, the diameter of composite nano-

fibers decreased as the gelatin content increased in the

polymer blend. The decrease in the fiber diameter may be

due to the reduced viscosity of the blend, which is the most

critical parameter for controlling the morphology of

nanofibers.[25] The addition of the gelatin solution to PLCL

appeared to decrease the viscosity of the blend, which

resulted in smaller fibers. Our observation is similar to that

of a previous report showing that the incorporation of

collagen into PLCL nanofibers decreased the fiber dia-

meter.[26] In the previous study, in addition to the viscosity

of the blend, the electrical property of collagen was

emphasized as a potential parameter for controlling fiber

diameter. In fact, a polymer solution with high conductivity

and dielectric constant facilitates the production of uniform

fibers. As compared to synthetic polymers, natural poly-

mers such as collagen and gelatin (a hydrolyzed form of

collagen) consist of a large number of electrically charged

amino acids, which can increase the electrically conductive

properties of the solution upon dissolution.[26] Therefore, it

Figure 1. Field emission scanning electron microscopy (FE-SEM)images of electrospun nanofibers: (a) PG10, (b) PG73, (c) PG37, and(d) PG01.Mean diameters of the fibers are also plotted (e).Asteriskindicates statistical significance relative to PG10 (P<0.05).



is possible that these electrical properties and the reduced

viscosity may have contributed synergistically to the

change in the fiber diameter. One advantage of electro-

spinning is that it allows fabrication of nanoscale fibers that

can closely mimic the hierarchical structure of native ECM.

The diameter of PLCL/gelatin nanofibers was in an

appropriate range to reconstitute the structural properties

of native ECM because collagen fibers exhibit a diameter

range of approximately 50–500 nm.[27]

Cellularity on Electrospun PLCL/Gelatin Nanofibers

MSCs have emerged as a promising cell source in

regenerative medicine due to their ability for self-renewal

and their capacity for differentiation into various lineages

such as osteoblastic, chondrogenic, adipogenic, and myo-

genic cells.[1,4] There are a number of studies demonstrating

that hMSCs are able to differentiate into osteocytes,

chondrocytes, and adipocytes under an engineered micro-

environment provided by three-dimensional nanofiber

meshes.[10] To induce in vitro osteogenic differentiation

of hMSCs, treatment with a specific cocktail containing

several chemicals such as b-glycerophosphate, dexametha-

sone, and ascorbic acid is frequently utilized.[28] Controlling

the commitment of hMSCs into the osteogenic lineage is

imperative in both human implantation and in in vitro

culture for successful bone regeneration. Large populations

of hMSCs are known to be present in bone marrow, but a

growing body of evidence shows that there are alternative

sources of hMSCs, such as adipose tissue,[29] umbilical cord

blood,[30] peripheral blood,[31] and dental pulp.[32] In our

previous study, we isolated hMSCs from mandible bone and

reported on the osteogenic differentiation and mineraliza-

tion of these cells when cultured on composite nanofibers

composed of poly(L-lactide) and nanosized bovine bone

powders.[20] The use of these primary MSCs may be

clinically more manageable due to easy access to residual

bone fragments associated with oral or maxillomandibular

surgery. To further characterize hMSCs derived from human

mandibles, we examined the responses of these cells

cultured on PLCL/gelatin nanofibers with a focus on how

the chemical composition of the nanofibers affected

proliferation and osteogenic differentiation.

The crosslinked PLCL/gelatin nanofibers were pre-wetted

with PBS for 1 day prior to cell culture. To determine the

cellularity of hMSCs on PLCL/gelatin nanofibers, we

compared DNA content at 7 and 14 d after cell seeding.

As shown in Figure 2, the DNA content from the nanofibers

containing gelatin was significantly greater than that from

the PLCL-only nanofibers (PG10) at both time points. For

example, the DNA content was 1.424� 0.201mg and

1.269� 0.160mg from PG73 and PG37, respectively, which

are higher than 0.260� 0.229mg from PG10 at 7 d of culture.

At 14 d of culture, the DNA contents of PG73 and PG37 were



Figure 2. DNA content of hMSCs after 7 and 14 d of culture.Asterisk indicates statistical significance relative to PG10 at 7 d;double asterisk indicates statistical significance relative to PG10 at14 d (P<0.05).

Figure 3. Immunofluorescence staining for cytoskeletal F-actin(red) and nuclei (blue) of hMSCs after 1 d (a and b) and 14 d(c) of culture; (a) hMSCs on PG10 nanofibers, (b and c) hMSCson PG37 nanofibers.

Figure 4. Scanning electronmicroscopy images of hMSCs culturedon nanofibers for 7 d: (a) PG10, (b) PG73, (c) PG37, (d) PG01.

800

2.088� 0.210mg and 1.777� 0.281mg, respectively, which

is greater than 0.481� 0.729mg from PG10. The gelatin-only

nanofibers (PG01) showed the greatest cellularity with a

DNA content of 1.689� 0.490mg and 2.457� 0.155mg at 7 d

and 14 d of culture, respectively. These results demonstrate

that incorporation of gelatin in the nanofibers increases the

cellularity of hMSCs.

We carried out fluorescence staining of F-actin micro-

filaments of adherent cells to investigate the effect of

gelatin on the morphology of hMSCs cultured on the

nanofibers. As shown in Figure 3, the majority of MSCs

cultured on the PG10 nanofibers maintained a round

shape with a condensed cytoskeletal structure at 1 day of

culture, while the cells on the PG37 nanofibers were well

spread with an elongated and polygonal morphology (the

cells on the other nanofibers containing gelatin exhibited a

similar morphology; data not shown). After 14 d of culture,

we observed the same trend as that for 7 d of culture; the

hMSCs cultured on PG37 nanofibers had proliferated with

stretched F-actin, while a limited number of cells were

found with a round shape on PG10 nanofibers. Collectively,

these results indicate that the addition of gelatin to PLCL

nanofibers contributes to the formation of robust stress

fibers within the cytoskeleton of the hMSCs at an early

adhesion stage, which may have contributed to the

enhancement of cellularity.

To confirm the effect of gelatin content on the prolifera-

tion of hMSCs, we examined the extent of coverage by cells

and their secreted ECM on the nanofiber surface using SEM

(Figure 4). For gelatin-containing nanofibers, the cells

covered the surface of the nanofibers, while the surface

of PG10 nanofibers was devoid of cells, which was

consistent with our results for DNA content and fluores-

cence. These results demonstrate that the addition of

gelatin into the PLCL nanofibers increased cell attachment,

spreading, and growth. Gelatin is a widely used natural


� 2009 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim DOI: 10.1002/mabi.200800358


polymer that is extracted from collagen by hydrolysis, and

natural polymers are usually more biocompatible than

synthetic polymers.[33] The electrospinning process

appeared to have a minimal effect on gelatin bioactivity

in our study and others, resulting in enhanced cell

attachment and spreading. Another explanation for the

positive effect of gelatin on the proliferation of hMSCs is

that the change in polymer surface hydrophilicity may be

improved by the incorporation of the hydrophilic gelatin.

For anchorage-dependent cells such as hMSCs, cell adhesion

is the first event that occurs when the cells contact the

surface of the polymer substrate. Cell adhesion is an

important process because it triggers numerous cellular

responses, including proliferation and differentiation. Cell

adhesion is mediated through the binding of various cell

membrane receptors with ECM components. Most syn-

thetic biodegradable polymers do not have cell-adhesive

motives that can be recognized by cell receptors, and

protein adsorption of proteins such as fibronectin on the

polymer surface is required for cell attachment.[34,35]

Generally, the hydrophilicity of the surface affects protein

adsorption; proteins can be adsorbed more easily on a

hydrophilic surface and previous work has demonstrated

accelerated cell attachment with increased hydrophilicity

of the surface.[36,37] Another study has shown that the

addition of gelatin into a chitosan scaffold increased the

adsorption of fibronectin on the surface of the scaffold,

resulting in concurrent increases in cell attachment and

differentiation.[38] These results indicate that gelatin,

similar to collagen, may be helpful for facilitating cell

adhesion through intermediation of proteins.[39] Our study

demonstrated that the morphology of hMSCs cultured on

PLCL-only nanofibers was round with a condensed actin

structure, which is similar to that of cells cultured on other

hydrophobic surfaces. In contrast, hMSCs were stretched

with defined actin fibrils on the gelatin-containing

nanofibers. Although the concentration of adsorbed pro-

teins on the nanofibers was not measured in our study, our

results indicate that the incorporation of gelatin increased

the hydrophilicity of the nanofibers, thereby improving the

adhesion and proliferation of MSCs.[13] Various methods

can be used to improve cell adhesion on nanofiber

substrates. Immobilization of certain functional groups,

proteins, or peptides can control the attachment of

cells.[40,41] However, these methods usually require com-

plicated processes and are expensive. In this study, we

showed that cell adhesion can be modulated by the simple

process of incorporating gelatin during the electrospinning

process.

Osteogenic Differentiation of hMSCs

To evaluate the osteogenic differentiation of hMSCs on

PLCL/gelatin nanofibers, the cells were cultured under



previously described conditions of osteogenic differentia-

tion.[28] In vitro osteogenic differentiation of MSCs is

initiated with the maturation of secreted ECM proteins

such as collagen, which is mediated by bone-specific

genes. During the maturation of ECM proteins, ALP activity

is concurrently increased and active ALP is involved in the

mineralization of the matrix.[42] Therefore, we evaluated

osteogenic differentiation of hMSCs on the PLCL/gelatin

nanofibers by examining the expression of several differ-

entiation-related genes, ALP activity, and extent of calcium

deposition.

At first, expression levels of osteogenic genes were

measured by RT-PCR at 3 and 7 d of culture, using the

primers shown in Table 2. We compared the expression

level of Cbfa 1 at 3 d of culture because Cbfa 1 is a well

known early osteogenic marker and expression of other

osteogenic markers such as Col I-a2, ALP, and OCN

was evaluated at 7 d.[4,43] Gene expression of Cbfa 1 was

significantly higher in the gelatin-containing nanofibers of

PG73, PG37, and PG01 than PG10, while the expression of

collagen type I had no effect, regardless of the type of

nanofiber (Figure 5a). Furthermore, the expression of ALP

and OCN was quantified by real-time RT-PCR (Figure 5b, c).

Expression of ALP increased 1.6, 2.5, and 4.2 times on PG73,

PG37, and PG01 nanofibers, respectively, compared to

expression ofALPon PG10 nanofibers. The expression of one

of the terminal differentiation markers, OCN, was also

increased 1.9 and 7.3 times on PG37 and PG01 nanofibers,

respectively, compared to OCN expression on the PG10

nanofibers.

The ALP activities at 7 and 14 d for all types of nanofibers

are shown in Figure 6. ALP activity on the PG10 nanofibers

increased from 0.036� 0.003 nmol/DNA/30 min at 7 d of

culture to 0.152� 0.053 nmol/DNA/30 min at 14 d of

culture. PG01 showed the highest level of ALP activity as

compared with the other groups: 0.191� 0.022 and

0.296� 0.300 nmol/DNA/30 min at 7 and 14 d of culture,

respectively. ALP activity was 0.108� 0.014 and

0.096� 0.022 nmol/DNA/30 min for PG73 and PG37 nano-

fibers, respectively, which are significantly greater than

that of PG10 at 7 d of culture. After 14 d of culture, the ALP

activity of hMSCs on PG73 and PG37 nanofibers was similar

to that on PG10 nanofibers. These results indicate that

gelatin incorporation stimulates the early initiation of

osteogenic differentiation of hMSCs, but does not enhance

the degree of differentiation.

As a final step toward investigating the effect of gelatin

incorporation on the terminal differentiation of hMSCs, the

amount of calcium deposited by the cell–nanofiber

constructs was measured after 14 d of culture. As shown

in Figure 7, the calcium content of PG10, PG73, PG37,

and PG01 nanofibers was 17� 11.0mg, 246� 76.3mg,

149� 25.3mg, and 239� 37.8mg, respectively. PG10

showed significantly lower levels of calcium than all other



Figure 6. Alkaline phosphatase activity (ALP) normalized by DNAcontent after 7 and 14 d of culture. Asterisk indicates statisticalsignificance relative to PG10 at 7 d; double asterisk indicatesstatistical significance relative to PG10 at 14 d (P<0.05).

Figure 7. Calcium assays after 14 d of culture (a). Images of vonKossa staining after 14 d of hMSCs culture: (b) PG10; (c) PG73; (d)PG37; (e) PG01. Asterisk indicates statistical significance relative toPG10 (P<0.05).

Figure 5. Reverse transcription-polymerase chain reaction (RT-PCR) results after 3 and 7 d of culture for Cbfa 1 and collagentype I-a2 (a) gene expression in hMSCs. Real-time PCR for relativequantitative data: gene expression of (b) ALP and (c)OCN. Asteriskindicates statistical significance relative to PG10 (P<0.05).

802Macromol. Biosci. 2009, 9, 795–804


groups. However, PG73 and PG37 showed the highest

deposition of calcium, similar to that of PG01, suggesting

that the concentration of the incorporated gelatin in the

fibers may not influence the degree of calcium deposition.

To qualitatively confirm the mineral deposition, the

nanofibers cultured with hMSCs for 14 d were analyzed

by von Kossa staining. The PG10 sample was negatively

stained, which is indicative of no mineralization, while

PG37, PG73, and PG01 samples exhibited positively stained

spots of dark brown. To clarify whether the precipitation of

calcium was facilitated by the incorporation of gelatin in

the nanofibers, we used two groups including PG10 and

PG01, as acellular controls that were incubated in the hMSC

osteogenic media under the same conditions for culture of

the hMSCs. We found that both of the nanofibers stained

negatively, indicating that the presence of gelatin in the

nanofibers did not mediate the precipitation of non-

physiologic mineral (data not shown). Collectively, these

results demonstrate that the incorporation of gelatin into

PLCL nanofibers facilitates the osteogenic differentiation of

DOI: 10.1002/mabi.200800358


hMSCs derived from human mandible. As previously

discussed in the results section for cell adhesion and

proliferation, the exposed gelatin on the nanofiber surface

may help with the absorption of ECM components, thereby

positively affecting differentiation, which is downstream

of cell adhesion.[34] Therefore, the hydrophilicity and

bioactivity of gelatin may be a key factor in the stimulation

of osteogenic differentiation of hMSCs.

Understanding the role of cell–matrix interactions in the

regulation of lineage-specific differentiation of hMSCs is an

active field of research. One recent study conducted by

McBeath et al.[44] highlighted the relationship between cell

spreading and stem cell differentiation; osteogenic differ-

entiation of hMSCs is promoted by contact and culturing on

a larger adhesion area, while adipogenic differentiation of

hMSCs is optimal on a smaller adhesion area. Although a

direct comparison may not be appropriate, our results

support this observation. The hMSCs cultured on the PLCL-

only nanofibers maintained a rounded shape, forming a

condensed actin structure, which resulted in limited

differentiation of hMSCs. However, hMSCs cultured on

the PLCL/gelatin nanofibers demonstrated unlimited

spreading with a defined cytoskeletal structure, which

promoted osteogenic differentiation. Another possible

explanation of our results is related to the presence of

functional groups on the nanofiber surface that may favor

certain differentiation pathways for hMSCs. Keselowsky

et al.[45] investigated the effect of chemical functional

groups on the differentiation of pre-osteoblasts and

demonstrated that surface �OH and �NH2 groups, but

not �CH3 and �COOH functional groups on the surface

increases the osteogenic differentiation of pre-osteoblasts.

A recent study also reported a significant role for small

functional groups, such as phosphorous groups, in the

control of the differentiation of hMSCs.[46] PLCL does not

possess free chemical functional groups, except for the end

carboxyl groups, while gelatin has an abundance of several

functional groups; the relative presentation of certain

functional groups on the gelatin-containing nanofibers

may enhance the osteogenic differentiation of hMSCs.

Although the delivery of MSCs could significantly

enhance the regeneration of damaged tissue, delicate

control over the multi-lineage differentiation of MSCs into

target tissue presents many technical challenges. A number

of recent studies have reported that MSCs seeded on various

types of nanofibers proliferate and undergo differentiation

into several cell types. For example, Li et al.[10] produced

nanofibers from PCL and reported the osteogenic, adipo-

genic, and chondrogenic differentiation of human MSCs.

Other studies have also reported osteogenic and chondro-

genic differentiation of MSCs on PCL and PLGA nanofibers,

respectively.[47,48] However, these previous studies focused

on the maintenance of the multi-lineage differentiation

capability of MSCs, but overlooked how nanofiber composi-



tion affects cell-mediated processes. Using the osteogenic

differentiation of MSCs as our model target, our results

demonstrate that the addition of gelatin to PLCL improves

the proliferation and mineralization of MSCs, indicating

that nanofiber composition significantly regulates cell

responses. From a tissue engineering point of view, the

ultimate goal is to develop an ideal scaffold for stem cells on

which specific differentiation can be controlled. To achieve

this goal, it is imperative to understand how the chemical

composition and structural cues of the nanofibers used as

an underlying artificial substrate can modulate stem cell

function.

Conclusion

In this study, nanofibrous scaffolds were prepared by

blending PLCL and gelatin using the electrospinning

process, and the resulting scaffolds were used as substrates

for culturing hMSCs. After 14 d of culture, hMSCs cultured

on PLCL/gelatin nanofibers showed better in vitro cell

growth and osteogenic differentiation than hMSCs grown

on PLCL-only nanofibers. In particular, the mineralization of

hMSCs and the expression of the osteogenic marker ALP in

hMSCs were significantly increased by the presence of

gelatin. These results reveal that the PLCL/gelatin compo-

site substrate has the potential to control the activity of

bone-forming cells. Therefore, this scaffold design can be

used for future composites to combine the benefits of both

engineered synthetic materials and natural ECM proteins.

Acknowledgements: This work was partially supported by agrant from the Korea Science and Engineering Foundation (KOSEF)funded by the Korea government (MOST) (No. 2008-01224) (to H.Shin) and by Nuclear Research & Development Program of theKorea Science and Engineering Foundation (KOSEF) grant fundedby the Korean government (MEST) (No. 20090062253) (to C. H.Kim).

Received: December 5, 2008; Revised: March 12, 2009; Accepted:March 18, 2009; DOI: 10.1002/mabi.200800358

Keywords: gelatin; human mesenchymal stem cells (hMSCs);nanofibers; osteogenic differentiation; poly[(L-lactide)-co-(e-capro-lactone)]

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DOI: 10.1002/mabi.200800358