Post on 24-Apr-2023
ARTICLE IN PRESS
0142-9612/$ - se
doi:10.1016/j.bi
�CorrespondEngineering, N
Singapore. Tel.
E-mail addr
(F. Chen), g010
(M.A. Woodru
(D.W. Hutmac
Biomaterials 28 (2007) 814–824
www.elsevier.com/locate/biomaterials
Combined marrow stromal cell-sheet techniques and high-strengthbiodegradable composite scaffolds for engineered functional bone grafts
Yefang Zhoua, Fulin Chenb, Saey Tuan Hob, Maria Ann Woodruffb,Tit Meng Lima, Dietmar W. Hutmacherb,c,�
aDepartment of Biological Sciences, Faculty of Science, National University of Singapore, Singapore 119260, SingaporebDivision of Bioengineering, Faculty of Engineering, National University of Singapore, Singapore 119260, Singapore
cDepartment of Orthopaedic Surgery, Yong Loo Lin School of Medicine, National University of Singapore, Singapore 119260, Singapore
Received 21 June 2006; accepted 20 September 2006
Available online 11 October 2006
Abstract
In this study, cell sheets comprising multilayered porcine bone marrow stromal cells (BMSC) were assembled with fully interconnected
scaffolds made from medical-grade polycaprolactone–calcium phosphate (mPCL–CaP), for the engineering of structural and functional
bone grafts. The BMSC sheets were harvested from culture flasks and wrapped around pre-seeded composite scaffolds. The layered cell
sheets integrated well with the scaffold/cell construct and remained viable, with mineralized nodules visible both inside and outside the
scaffold for up to 8 weeks culture. Cells within the constructs underwent classical in vitro osteogenic differentiation with the associated
elevation of alkaline phosphatase activity and bone-related protein expression. In vivo, two sets of cell-sheet-scaffold/cell constructs were
transplanted under the skin of nude rats. The first set of constructs (5� 5� 4mm3) were assembled with BMSC sheets and cultured for 8
weeks before implantation. The second set of constructs (10� 10� 4mm3) was implanted immediately after assembly with BMSC sheets,
with no further in vitro culture. For both groups, neo cortical and well-vascularised cancellous bone were formed within the constructs
with up to 40% bone volume. Histological and immunohistochemical examination revealed that neo bone tissue formed from the pool of
seeded BMSC and the bone formation followed predominantly an endochondral pathway, with woven bone matrix subsequently
maturing into fully mineralized compact bone; exhibiting the histological markers of native bone. These findings demonstrate that large
bone tissues similar to native bone can be regenerated utilizing BMSC sheet techniques in conjunction with composite scaffolds whose
structures are optimized from a mechanical, nutrient transport and vascularization perspective.
r 2006 Elsevier Ltd. All rights reserved.
Keywords: Bone engineering; Bone marrow stromal cells; Mesenchymal stem cell; Composites; Biodegradable polymers; Functional tissue engineering
1. Introduction
Tissue formation within the body, as part of a develop-ment, healing and/or repair process, is a complex event inwhich cell populations in combination with extra cellularmatrix, self-assemble into functional units and ultimatelyinto tissues and organs. There is intense academic and
e front matter r 2006 Elsevier Ltd. All rights reserved.
omaterials.2006.09.032
ing author. Division of Bioengineering, Faculty of
ational University of Singapore, Singapore 119260,
: +65 68741036; fax: +65 67773537.
esses: scip1129@nus.edu.sg (Y. Zhou), cflfx@tom.com
29965@nus.edu.sg (S.T. Ho), biewma@nus.edu.sg
ff), dbsltm@nus.edu.sg (T.M. Lim), biedwh@nus.edu.sg
her).
commercial interest in finding methods to stimulate andcontrol these events, and eventually to replicate these eventsoutside the body as closely as possible [1]. This interest hasresulted in tissue engineering (TE) emerging as a well-recognized research area, in the arena of regenerativemedicine. The most common concept underlying TE is tocombine living cells, biologically active molecules and astructural scaffold to form a ‘‘tissue-engineering construct’’(TEC) to promote the repair and regeneration of tissues [2].In the bone TE literature [3–5], an increasing trend is
seen towards the fabrication of polymer-based compositematerials to obtain matrices with better osteoconductiveproperties. The main approach is to develop composite andbiomimetic scaffolds with nanocrystallites of inorganic
ARTICLE IN PRESSY. Zhou et al. / Biomaterials 28 (2007) 814–824 815
biological compounds such as calcium phosphate (CaP)salts dispersed within polymer matrices. Composites ofpolymer and ceramics are favored by bone engineers due totheir favorable mechanical properties, including strengthvia the ceramic phase, toughness and plasticity via thepolymer phase, and greatly improved mechanical stiffnessof the matrix [5]. Biological advantages of these compo-sites, such as improved biocompatibility arise from anincreased initial flash spread of serum proteins in the invitro phase compared to the more hydrophobic polymersurface. When used in vivo, ceramic particles bothembedded and on the surface of the composites allow forimproved tissue integration by buffering pH change andproviding a suitable microenvironment that mimics thehost tissue’s inorganic phase [6].
In the past, we have successfully reported strategies forbone engineering that are based on immobilizing bonemarrow stromal cells (BMSC) within fibrin glue that is theninjected into medical-grade polycaprolactone (mPCL)-based matrices [7]. However, despite fibrin glue achievinghigher seeding efficiencies, its cell-specific degradationkinetics and cost might mitigate its widespread use. Hencewe have developed our second-generation scaffolds frommPCL–CaP composites whose mechanical properties are inthe range of cancellous bone and that have improved cellattachment and subsequent cell proliferation, in theabsence of fibrin glue [8]. Despite improved attachmentto PCL–CaP composites in the absence of fibrin, approxi-mately 30–40% of the cells fail to attach due to the lowsurface to volume ratio of these scaffolds.
To address this issue, this study investigated the use ofcell-sheet technology [9] in conjunction with pre-seeding asan approach for optimizing the aforementioned techniques.Whilst confluent cultured cells are usually harvested byenzymatic digestion before seeding on a scaffold, thistechnology allowed us to harvest an intact cell sheet.Okano’s group developed the use of single sheets of culturedcorneal epithelial cells and multilayered cardiomyocytessheets for engineering transplantable cornea and myocardialtissues using a smart cell culture plate technology [10,11].Other groups [12] utilized dermal fibroblasts sheets for skinand L’heureux and colleagues reported the engineering ofhuman blood vessels by endothelial and smooth muscle cellsheets [13] using standard tissue culture plates.
Our first hypothesis was that partially mineralized andstrong cell sheets could be grown on normal tissue cultureplates under appropriate cell culture conditions. We furtherhypothesized that not only structural, but also functionalbone could be engineered in vivo by using mineralized cellsheets in combination with novel composite scaffolds.
2. Methods
2.1. Scaffold fabrication and characterization
Flakes of mPCL (Birmingham Polymers, Alabama, USA) and tri-
calcium phosphate (TCP) (Biobone Inc, China) (mPCL–TCP) (80:20%)
were prepared into +1.7070.10mm monofilaments via a filament
extrusion process using an extruder built in-house. A fused deposition
modeling (FDM) 3000 rapid prototyping system (Stratasys Inc, USA) was
used to fabricate composite scaffolds with a bulk dimension of
50� 50� 4mm3 (Length, breadth and height, respectively), a lay down
pattern of 01/601/1201 and a porosity of 50%. Details of scaffold
fabrication and characterization can be found elsewhere [14].
mPCL–TCP scaffolds were cut into 5� 5� 4mm3 (length, breadth and
height) or 10� 10� 4mm3 blocks. Five molar NaOH treatments were
used to increase surface roughness and hydrophilicity. Scaffolds were
washed and sterilized using 70% ethanol prior to use. The scaffold
porosity, interconnectivity, distribution of CaP within the mPCL matrix,
compressive modulus and surface/volume ratio were measured as
previously reported [15]. Mechanical compression tests were conducted
using an Instron 4302 Material Testing Sytsem in accordance with the
ASTM D695-96 guidelines. The specimens were compressed at a rate of
1mm/min up to a strain level of approximately 60% at room temperature
under 75% humidity. The stress–strain curves were obtained to evaluate
the compressive stiffness (Young’s modulus) and compressive strength of
scaffolds (n ¼ 4). Phase contrast imaging and tomography (PCIT) were
used to provide a 3D reconstruction of the matrix of a mPCL–CaP
scaffold at the Singapore Synchrotron Light Source. One thousand
projections were collected during the experiment with 2 s per projection.
Each projection was a 2D image with 1392� 1040 pixels and 12-bit
accuracy and the image comprised 976� 976� 728mm3 and a spatial
resolution of �1mm was achieved.
2.2. Cell isolation and culture
Porcine BMSC were isolated and cultured as reported previously [16].
Pigs were obtained from the animal holding unit of the National
University of Singapore (NUS) and samples of bone marrow were
harvested in accordance with the IACUC approval of NUS. Briefly,
BMSC were aspirated from the bone marrow, gradient centrifuged and
plated into flasks containing Dulbecco’s Modified Eagle’s medium
(DMEM) low glucose (GIBCO, CA, USA) containing 10% fetal bovine
serum (GIBCO, USA) and 2% antibiotics (200mg/ml penicillium and
200mg/ml streptomycin). Only passage two to four cultures were used for
all the experiments. At confluence, culture media was changed to
osteogenic media consisting of standard media plus L-ascorbic acid-
2-phosphate (50mg/ml), b-glycerophosphate (10mM) and dexamethasone
(100 nM) (Sigma, USA). Control cultures (uninduction group) were
maintained in standard media. Media was changed every 3 days.
2.3. Fabrication of sheet–scaffold constructs
Confluent cells in flasks with density of 50,000 cells/cm2 were cultured
in osteogenic medium for 1–2 weeks, until BMSC sheets were formed and
could be detached intact from the substratum using a cell scraper (Fig.
1A). The sheets were then wrapped around the pre-seeded scaffolds (1
million cells per scaffold and 25 cm2 sheet for 5� 5� 4mm3 scaffolds, or 4
million cells per scaffolds and 150 cm2 sheet for 10� 10� 4mm3
scaffolds), (Fig. 1B–C). The pre-seeded scaffolds were prepared by
delivering cells in 20–100ml volume and the scaffolds containing the cell
suspensions were kept in an incubator for 2 h before the media was topped
up. The pre-seeded constructs were cultured in DMEM media with 10%
FBS. Cell sheets were wrapped on 5� 5� 4mm3 and 10� 10� 4mm3
scaffolds, respectively. The assembled sheet-scaffold construct was held in
culture medium until implantation (Fig. 1D).
2.4. Construct characterization in vitro
Cell attachment on scaffolds was performed by phalloidin–propidium
iodide (PI) staining according to the manufacturers’ protocol (Molecular
Probes Inc., Oregon, USA). Qualitative cell viability was assessed using
fluorescein diacetate–propidium iodide (FDA–PI) staining. To observe cell
ARTICLE IN PRESS
Fig. 1. Experimental procedures for the assembly of scaffold–cell constructs with cell-sheet construction and implantation. (A) Porcine bone marrow
stromal cell (BMSC) sheets were harvested using cell scraper. (B) The BMSC sheet was wrapped around the pre-seeded scaffold. (C) The constructs after
assembly demonstrated dense sheet coverage over the pre-seeded scaffolds (Right) compared with the unassembled one (Left). (D) Constructs were
implanted into subcutaneous pockets of 6-month-old nude rats and the implants were accepted by the host tissue with no detectable foreign body reaction
over the entire study period.
Y. Zhou et al. / Biomaterials 28 (2007) 814–824816
growth inside the scaffolds, specimens were cut at the middle and the depth
of visualized specimen was 200–300mm. The specimens were viewed using
an Olympus IX70-HLSH100 confocal laser-scanning microscope [17].
For SEM, samples were fixed in paraformaldehyde, dehydrated in a
gradient series of ethanol and sputter-coated with gold. Images were taken
by XL30SEM (FEI Inc, OR, USA) at 15Kv. Von Kossa staining was
utilized to assess the degree of mineralization throughout the scaffold–cell
construct, as referred to in previous reports [18].
2.5. Alkaline phosphatase (ALP) activity
Cellular ALP activity was determined using a kinetic assay, based on
measuring the rate of p-nitrophenol formation from p-nitrophenyl
phosphate (kit 104-LL, Sigma). Briefly, cell lysates and culture super-
natant were combined with 50ml of ALP reagent and the activity measured
in a 96-well plate following 30min incubation at 37 1C. ALP activity was
read at 405 nm (Bio-Rad, USA) as per the manufactures instructions. ALP
activity in the lysates was expressed as nanomoles of p-nitrophenol
produced per minute per microgram of protein [17]. The ALP activity in
the culture media was expressed as unit/L in accordance with manufac-
turer’s protocol since total proteins in the media were mainly from serum
in media and could not be used to normalize all samples.
2.6. RNA isolation and RT-PCR
Total cellular RNA was extracted weekly using Trizol reagent
(Invitrogen Corp., Carlsbad, California, USA) and the cDNA synthesis
was performed from 2mg total RNA using Superscript II and Oligo dT
(Invitrogen) according to the manufacturer’s instructions. The expression
of bone matrix proteins was semi-quantitated by PCR using specific
primers as follows: (1) GAPDH, glyceraldehyde-3-phosphate dehydro-
genase, NM_002046, 192 bp sense 50-GTTCCAATATGATTCCACCC-30,
and anti-sense 50-TGAGTCCTTCCACGATACC-30; (2) Collagen type I
(Col I), NM-000088.2, 501 bp, sense 50-CTGGCAAAGAAGGCGGC-
AAA-30 and anti-sense 50-CTCACCACGATCACCACTCT-30; (3) Runx2,
AY406695, 381 bp, sense 50-CCAGATGGGACTGTGGTTAC-30 and anti-
sense 50-ACTTGGTGCAGAGTTCAGGG-30; (4) Osteopontin (OPN),
AF052124, 347 bp, 50-CCAAGTAAGTCCAACGAAAG-30 and anti-sense
50-GGTGATGTCCTCGTCTGTA-30; (5) Osteocalcin (OCN), X53698.1,
315 bp, sense 50-CATGAGAGCCCTCACACTC-30 and anti-sense 50-
AGAGCGACACCCTAGACCG-30; (6) Osterix (Osx), AF477981, 566 bp,
sense 50-TGGGGTATCTCTTGATTAGG-30 and anti-sense 50-TTGGGG-
CCTTGAGACAGCA-30. Target gene expression values were semi-
quantified by densitometry and standardized against GAPDH. The
reaction products were also cloned into pGEM-TEasy vector (Promega)
and sequenced for confirmation.
2.7. Western blot
Cell lysates were prepared using ice-cold lysis buffer (5mM MgCl2,
150mM NaCl, 1% triton-100, pH 7.5) containing protease inhibitors
(Calbiochem, UK). The protein concentrations were determined using a
Protein Assay Kit (Bio-Rad). Cell lysates (40mg) were resolved by 6–12%
SDS-PAGE (polyacrylamide gel electrophoresis) gels and the proteins
were transferred to nitrocellulose membranes (Amersham, UK). After
blocking with 5% low fat milk in tris-buffered saline (TBS), membranes
were then incubated with either mouse anti-OCN (Biodesign, ME, USA),
OPN (DSHB*, IA, USA) or –actin (Santa Cruz, CA, USA). The primary
antibody was diluted 1:1000 in TBS with 0.1% Tween (TBST) overnight at
4 1C, then incubated for 1 h with secondary antibodies diluted 1:1000 in
TBST. The specific signals were visualized by chemiluminesence (Super-
signal west pico kit, Pierce, USA) and the intensity of bands was assessed
using densitometry.
2.8. Implantation
The animal research protocol was reviewed and approved by the
IACUC of National University of Singapore (NUS) (NIDCR 00–113).
Nude rats originally obtained from Harlan Sprague Dawley (Indianapolis,
USA) were bred and maintained in pathogen-free conditions. Twelve
constructs consisting of cell-sheets assembled with scaffolds measuring
5� 5� 4mm3, were cultured in osteogenic medium for 8 weeks and then
transplanted subcutaneously into the dorsal pockets of 4-month-old nude
rats weighing between 110 and 130 g. Four constructs with cell-sheets
assembled with scaffolds measuring 10� 10� 4mm3 were also implanted
immediately without further culture. For each group, two mPCL–CaP
scaffolds without cells and cell sheets were implanted as a control. Prior to
implantation, cells in two constructs per group were labeled with 2 mM of
the cFDA-SE (carboxyfluorescein diacetate succinimidyl ester) (Molecular
Probes). Transplants were recovered 4, 8 and 12-weeks post-transplanta-
tion, fixed in 4% formalin and decalcified in 10% formic acid for paraffin
embedding. Paraffin sections (10mm) were deparaffinized, hydrated, and
stained with hematoxylin and eosin (H&E), safranin-O/fast green (S/F).
Immunohistochemical analysis was performed by using monoclonal
antibodies against OCN (Sigma) and Col I (Sigma), Col II (Chondrex),
which had been biotinylated through Animal research kits (DAKO, USA).
The signals were visualized using Envision System Horseradish Peroxidase
kits (DAKO) and samples were counterstained with hematoxylin.
2.9. Micro CT (mCT) scan and X-ray analysis
A mCT scanner (Skyscan, Belgium) was used to determine bone growth
occurring in the engineered constructs. Specimens were scanned with a
resolution of 35 mm and approximately 500 scan slices were taken and
reconstructed according to the manufactures recommendations (Skyscan,
Belgium). The output was reconstructed into 3D stacks using Mimics 7.3.
The thresholds used in this study were 68–1732 Housefield units (HU) for
cortical bone and �70–67 HU for cancellous bone based on the threshold
calculations for samples of porcine femur bone. With conventional X-ray
analysis, samples were also analyzed using a Mammomat 3000 (Siemens)
X-ray machine at the condition of 25KV and 2.8Mas. After mCT
ARTICLE IN PRESSY. Zhou et al. / Biomaterials 28 (2007) 814–824 817
scanning and X-ray analysis, the specimens underwent compression
testing as described in Section 2.1 (n ¼ 4).
2.10. Statistics analysis
All values were presented as means7standard deviations. All data was
subjected to one-way analysis of variance (ANOVA) of post-hoc testing
(SPSS Version 11.02). Significance levels were set at po0:05. Data were
the average of three replicates performed under identical conditions.
3. Results
3.1. In vitro characterization of tissue-engineered constructs
BMSC sheet-scaffold constructs were cultured for up to8 weeks and were evaluated based on morphologicalappearance, cell adhesion and viability (Fig. 2). Thecomposite scaffolds (mPCL–CaP, 80:20) with a lay-downpattern of 01/601/1201 presented a fully interconnectedhoney-comb-like pore morphology with a porosity of 50%(Fig. 2A). The surface/volume ratio was 8.5mm2/mm3 andthe scaffolds had a compression modulus of 124MPa andcompressive strength of 4.2MPa. Phase Contrast Imagingand Tomography (PCIT) demonstrated that the CaPparticles with sizes in the mm3 range were uniformlyembedded within PCL matrix (Fig. 2A inset). During invitro expansion, cells were well attached on the surface ofscaffolds and exhibited mesenchymal cell-like morphologywith actin alignment on cell bodies (Fig. 2B,C). After 1week in culture, viable cells (�90%) formed bridges overthe scaffold and only a relatively small proportion of dead,PI-stained cells (Fig. 2D) were observed. After 3–5 weeks in
Fig. 2. BMSC–scaffold constructs as observed by SEM and confocal microscop
image of the matrix with visible particles of CaP (white) in the mm3 range withi
were aligned with the cell body after 3 days. (C) Strong actin fiber staining was
center at day 21 (Phalloidin-PI). (D) BMSCs formed bridges in scaffolds after
scaffold after 35 days, with limited cells stained red. (F) BMSCs in cell sheets w
PI staining). (G) The top view of the constructs shows BMSCs sheets covering
the pores with an even distribution. (I) Distinct mineral-like nodules formed
showing nodules formed at 35 days under induction. Scale bar: (A) 500mm, (
1000mm, (I) 50mm and (J) 100mm.
culture, bars of the interior constructs where covered bylive cells (�70%) that had spread evenly on the surface ofscaffolds and throughout its pores (Fig. 2E). When cell-sheets were wrapped on the pre-seeded mPCL-TCPscaffolds, a high proportion of cells remained viable forup to 8 weeks (Fig. 2F). Assessment of the cell-sheetconstructs after 3 weeks culture, via SEM, showed auniform covering of neo-tissue inside and around the pre-seeded mPCL-TCP scaffolds (Fig. 2G, H) and the calciumdeposition was confirmed by the presence of mineralizednodules (Fig. 2I, J).
3.2. ALP activity
Measurements of ALP activity from cell lysates andmedium following osteogenic media supplementation wereundertaken to assess the ability of the constructs toaugment in vitro osteogenesis. Following induction,cellular ALP activity significantly increased to day 21(Po0:05), and then maintained a plateau to day 56(Fig. 3A). In comparison, cultures from the uninducedgroup expressed minimal ALP activity and failed to showan increase over time. Similarly, ALP activity in themedium following induction was significantly higher thanthe uninduced group reaching a peak at day 49 (Po0:05)that was 4-fold higher than the uninduced group (Fig. 3B).
3.3. Expression of osteogenic biomarkers
To confirm the constructs were able to support in vitroosteogenic differentiation, RNA from cells cultured on the
y. (A) mPCL–CaP scaffold architecture. Inset shows the reconstructed 3D
n the PCL (yellow). (B) Actin fibers visualized using Phalloidin-PI staining
observed at the BMSC–scaffold interface while comparatively weak in the
7 days (FDA-PI staining). (E) The BMSCs maintained viability inside the
rapped onto scaffolds remained viable after 56 days in vitro culture (FDA-
the scaffolds. (H) Dense BMSCs grew inside the scaffolds and nearly filled
on the BMSCs sheet after osteogenic induction. (J) Von kossa staining
B) 50 mm, (C) 80 mm, (D) 200mm, (E) 80 mm, (F) 40 mm, (G) 500mm, (H)
ARTICLE IN PRESS
0
1
2
3
4
5
56
49
4942
35
3528
21
2114
7
Construct non-induction
Construct under induction
Construct non-induction
Construct under induction
7
1
1
0
5
10
15
ALP
act
ivity
(U
nit/L
)
20
25
30
Days
Days
(nm
ol/u
g/m
in)
p-ni
trio
phen
yl p
hosp
hate
6
7
8
9
(B)
(A) 10
Fig. 3. ALP activity of BMSC sheet–scaffold constructs both in cell
lysates (A) and the medium (B). (A) Lysates of constructs demonstrated
significantly elevated ALP activity after osteogenic induction. ALP levels
remained high over the whole culture period. (B) ALP activity in the
medium. Constructs under induction demonstrated higher extracellular
ALP activity than constructs without induction over the whole culture
period (po0:05).
Y. Zhou et al. / Biomaterials 28 (2007) 814–824818
constructs was extracted and RT-PCR was performed tomonitor the temporal expression of the osteo-relatedmRNA transcripts, Runx2, Osx, OPN, Col I and OCN.Fig. 4 shows that Osx and OCN expression levels weresignificantly up-regulated at least 10 and 5 times, respec-tively after induction and maintained this elevated levelthroughout the culture period. OPN expression levels werealso up regulated, and the levels of Runx2 and Col I wereslightly increased in induced constructs. To further confirmosteogenesis on a protein level, OCN and OPN synthesiswere also measured using western blots. As shown inFig. 5, OCN was specifically expressed with inducedconstructs and its expression remained stable over 7 weeksculture. OPN expression increased around 3–4 times atweek 3 and then decreased slightly.
3.4. In vivo bone formation
To assess the in vivo bone-forming properties of theengineered constructs, samples were implanted under theskin flank of nude rats.
For assembled constructs which had undergone 8 weeksin vitro culture, bone formation without fibrous tissuecapsule formation was observed after 4 weeks of implanta-tion with the mineralized tissue observed around theconstruct (Fig. 6A, B). Strong mineralization was verifiedby X-ray radiography and mCT with average bone volume
of 49mm3 (Fig. 6B–C). In contrast, control specimens(mPCL–CaP scaffolds without cells) showed mild fibrousencapsulation with structured and vascularised fibroustissue throughout the scaffold architecture (Fig. 6D). In allcell sheet construct specimens, stages of neo-bone forma-tion appeared to mainly follow a path akin to endochon-dral ossification, as observed by the presence ofhypertrophic chondrocytes within the constructs (Fig. 6E,F). Calcified cartilage tissue was detected at the interface ofneo-bone and neo-cartilage with weak glycosaminoglycan(GAG) staining and eosin staining were also observed(Fig. 6F, G). Intramembranous trabecular bone formationwas also observed with the typical osteocyte embedded indense matrix accompanied by both blood vessel andmarrow cavities next to scaffold bars (Fig. 6H, I). Bone-related extracellular matrix biomarkers were detected.Specifically, the neo-bone stained positive for OCN andCol I and negative for Col II and safranin-O (Fig. 6J–L).The neo cartilage-like tissue, which was detected mainly inthe interior part of construct, stained positive for Col IIand safranin-O but very weakly for Col I and OCN(Fig. 6I, J–L). These implants, with 8 weeks in vitro cultureperiods, demonstrated improved compressive modulusover the control implants but there was no statisticallysignificant difference between them (data not shown).For the constructs implanted immediately after assem-
bly, dense neo tissue was formed around the constructswith a well vascularised and non-fibrotic gross appearance,while the control demonstrated only membrane-like fibroustissue formation (Fig. 7A). Extensive bone formation wasobserved after 4, 8 and 12 weeks in constructs with strongradiograph signal similar to fully mineralized compactbone (Fig. 7B–E). mCT showed that a mineral densitysimilar to cortical and cancellous bone was observed at theperiphery and the interior of the constructs, respectively.(Fig. 7F, G). The volume of neo bone accounted for about40% of the whole construct for all time points with amaximal volume of 403mm3 and the cells within the neo-bone demonstrated positive cFDA staining, previouslylabeled before implantation (Fig. 7H, I). No scaffolddegradation was detected for up to 12 weeks of implanta-tion (data not shown). The bone formation followed bothintramembranous ossification and endochondral ossifica-tion while only limited safranin-O staining and sparselydistributed chondrocytes were observed in constructs(Fig. 7J–M). Most neo-bone tissues demonstrated lamellarbone characteristics, resulting from the rapid manner inwhich bone was laid down (Fig. 7J, K). The alignedosteoblasts were observed on the surface of the mPCL-CaPstruts with osteocytes inside the pore architecture accom-panied by blood vessel and marrow cavities (Fig. 7L).Specifically, the neo-bone stained extensively for OCN andCol I with limited signals detected for Col II in somechondrocyte locations (Fig. 7N–Q). Compression testingrevealed that the constructs had modulus of 184MPa andcompressive strength of 5.2MPa, which was 30% higherthan the control scaffolds (without cells).
ARTICLE IN PRESS
Non-induction(a)
(b)
Days
OCN
OPN
Col I
Runx2
GAPDH
14
12
Rel
ativ
e ex
pres
sion
leve
l
10∗
∗
∗∗
8
6
4
2
Osterix Osteocalcin Osteopontin
Day 1 Day 7 Day 14
Day 28 Day 42 Day 49
0
Osx
Induction
1 7 14 28 42 49 1 7 14 28 42 49
Fig. 4. In vitro RT-PCR profiles to analyze expression of osteogenic marker genes. (A) Representative gel electrophoresis figures showing different gene
expression levels at different time points. OCN, Osx and OPN expression level were up-regulated while Runx2 and Col I levels did not show distinct
variation after osteogenic induction. (B) Intensity levels showing Osx, OCN and OPN mRNA expression levels in sheet-scaffolds constructs significantly
increased after induction (po0:05).
Y. Zhou et al. / Biomaterials 28 (2007) 814–824 819
4. Discussion
In this study we have demonstrated the feasibility ofengineering large (400mm3) structural and functional bonetissue grafts by combining pre-seeded novel mPCL–CaPcomposites and partially mineralized cell-sheets. Theengineered constructs could be candidates for mediumload-bearing applications, or, if used in combination withfixation devices could be utilized in high load-bearingapplications. The scaffold design is based on a fullyinterconnected large pore network which supports goodnutrient support and fast vascularization, slow degradationkinetics and can sustain higher mechanical loadingthan those PCL-based constructs used in low load-bearingareas [19].
The composite scaffolds in this study showed mechanicalproperties in the range of cancellous bone and havedemonstrated good support of BMSC attachment andproliferation as well as differentiation and subsequently invitro and in vivo bone formation. Design and fabricationof composite scaffolds are currently an area of intenseinterest in bone TE, and the main idea is to producescaffolds with inorganic biological compounds such as CaP
salts dispersed within polymer matrices [20,21]. Compositescan confer favorable mechanical properties, includingstrength via the ceramic phase, toughness and plasticityvia the polymer phase, as well as a increased mechanicalstiffness [22]. Most importantly, composite surfaces pro-vide a bioactive environment that leads to improved cellattachment, elevation of certain osteogenic biomarkerexpression levels and better integration with host tissue,when compared to contemporary polymer surfaces. BMSCdemonstrated a classical osteogenic differentiation processon the composites with a clear increase in ALP activity,followed by an increase in OCN expression, which was inline with results from various other studies [23,24].Interestingly, the two key transcriptional factors (Osxand Runx2) revealed different kinetic patterns at thetranscription level. Osx expression was highly upregulatedby induction while Runx2 did not change significantlyduring the entire culture period. The results confirmed thatRunx2 is involved in osteogenesis by modulation of itsactivity not by quantificational change in gene or proteinexpression [25,26].This study demonstrated that cell-sheets with high cell
density could be combined with scaffolds that possess good
ARTICLE IN PRESS
Non-induction Induction(A)
(B)
Days
OCN
OPN
Actin
4.5
4 1 14 28 42
∗
∗
49
3.5
3
2.5
Fol
d ch
ange
2
1.5
1
0.5
0
Osteocalcin Osteopontin
1 14 42 49 14 28 42 49
Fig. 5. Western blot assay for OCN and OPN expression. (A) OCN was specifically detected in induced constructs and OPN expression significantly up-
regulated after induction. (B) Analysis of intensity of bands indicated that OCN expression from constructs under induction peaked at day 28 while OPN
peaked at 14 days. The non-induction samples show no OCN expression and no statistical difference in OPN expression.
Y. Zhou et al. / Biomaterials 28 (2007) 814–824820
mechanical properties and could potentially be placed inbone defects that undergo significant load bearing. Anyscaffold used in bone regeneration is generally required tohave biological, biomechanical and biomaterial functions.These include appropriate porous morphology for thediffusion of nutrients and the invasion of vascularity froma surrounding tissue, appropriate material surface chem-istry and biocompatibility to allow cells to adhere andexpress their normal phenotypes, sufficient mechanicalproperties as a load-bearing construct during the regenera-tion process, and adequate biodegradability after boneremodeling in the porous part has occurred. Thus, ascaffold with controlled design of the macro and micro-structure as well as biological and biomaterial considera-tions should balance these complex functions.
From the viewpoint of biomechanical and functional TEfor bone regeneration, the mechanical function representedby stiffness and strength, and capabilities for transporta-tion and exchange of nutrients and wastes could be apossible design target to be optimized [27]. Since thestiffness of scaffolds is proportional to Fp
2, where Fp is thevolume fraction of the polymer, it is the design goal toincorporate high-modulus materials. This generally in-volves a hard reinforcement phase within a ductile matrix.Despite the incorporation of ceramics resulting in a highermodulus the major problem faced is the interfacialadhesion between the reinforcement and the polymermatrix. Unless effective stress transfer takes place acrossinterfaces, the compression and shear strength are not
expected to be significantly improved. Certainly, a certaindegree of compromise may be unavoidable to produce acomposite system with acceptable scaffold performance inhigh load-bearing areas. Hence, in our developed compo-site system it is a balance of properties between strengthand porosity.Reviewing the bone-engineering literature, it can be
concluded that scaffolds should not only have a pore sizebigger than 300 mm but also possess pore interconnectivityof a similar size to facilitate good in vitro nutrient/wastetransport and in vivo vascularization [28]. However, thedrawback for scaffolds optimized using these designparameters, like the ones used in this study, include acompromised surface/volume ratio, which leads to sig-nificant loss of cells during seeding. The application ofhydrogels such as fibrin glue [14,29] allows the achievementof up to 100% seeding efficiency. However, this eleganttechnology has associated drawbacks such as cell- andpatient-dependent degradation kinetics, and high cost.Hence, our results show that using the cell-sheet technol-ogy circumvents these disadvantages associated withoptimizing the scaffold morphology from a nutrient/wasteand vascularization point of view.The cell-sheet technique has proven effective for the
engineering of soft tissues, including skin [30], blood vessels[13], corneal [10], and myochardium [11] owing to thesheets effectively preserving the cell–cell contact and ECM.To our knowledge this is the first report detailing the use ofcell sheets for hard TE. Our results show the layered cell
ARTICLE IN PRESS
Fig. 6. Cell sheet–scaffold constructs after in-vitro culture for 49 days were implanted into nude rat and harvested after 28 days. (A) Gross appearance of
implants showing the implants were well vascularized. Inset shows that the implants were not encapsulated by rat fibrous tissue. (B) X-ray images. Distinct
bone formation was observed in the constructs. The inset shows the top view of X-ray image. (C) Micro CT images show the neo tissue in the constructs
had similar mineral density to native bone. (D) Hematoxylin and eosin (H&E) staining of cell free scaffold group. Fibrous tissue was mainly found in
constructs. (E) Growth plate like tissue formed in constructs as shown by H&E staining. (F) High-magnification image indicates the osteocyte (black
arrow) formation in neo bone while chondrocytes like cells (white arrow) were located in calcified cartilage. (G) Safranin-O/fast green staining indicated
endochondral ossification. Cartilage-like tissue composed of chondrocytes (white arrow) was observed in the interior part of constructs with positive
safranin-O staining, and next to neo bone tissue embedded with osteocytes (black arrow). (H) Distinct bone tissue formed inside the pores and scaffold
surface. (I) H&E staining indicated typical cancellous bone tissue formation. Marrow cavities (Ma) were surrounded by neo bone tissue with distinct
osteocytes embedded. Blood vessels (BV) penetrated into tissue with red blood cells stained, indicating vascularization. (J) Osteocalcin staining. Strong
signals (black arrow head) were mainly detected at the periphery of constructs while weak signals were detected in the interior of the calcified cartilage. (K)
Collagen type-I staining. Extensive staining (black arrow head) was observed with weak and no staining observed on chondrocyte-like cells. (L) Collagen
type-II staining. Only limited signals (black arrow head) were detected on some chondrocyte-like cells. BV: blood vessel+red blood cells; Ma: marrow; Bo:
bone; OC: osteocyte; CY: chondrocyte; CC: calcified cartilage. Scale bar: (D, F, G, I–L) 50mm (E,H) 200mm.
Y. Zhou et al. / Biomaterials 28 (2007) 814–824 821
sheets after wrapping on composites contributed to theextensive in vivo bone formation and this indicated that thecell-sheet techniques could have promising potential inhard TE. In addition to the obvious advantages of usingcell-sheet techniques to achieve a highly efficient celldelivery system over the traditional cell suspension systems[9,31], the layered and condensed cell sheets might mimicthe in vivo deposition of bone matrix formation processwhere osteoblasts are attached on the mineralized sheet likeextracelluar matrix [32]. Furthermore, it is postulated thatthe cell sheets underwent significant contraction after beingwrapped around the composites scaffolds. It is reported
that cell-generated contraction forces stimulate a conden-sation-like process between the sheets and these biomecha-nical-related forces are thought to play a critical role in thespecific type of tissue that developed [33,34].The neo-bone formed in constructs demonstrated
endochondral ossification characteristics with the observa-tion of hypertrophy chondrocytes, which may be related tothe specific architecture of the constructs—the interior partof the constructs experienced hypoxia and had limitedmedium diffusion. The assembled constructs had a size ofover 4mm, which is the limitation of oxygen and nutrientdiffusion in vivo [21]. A dense mineralized BMSC sheet was
ARTICLE IN PRESS
Fig. 7. Constructs with BMSC sheet-scaffolds were implanted into nude rat, harvested after 28, 56 and 84 days. (A) Gross appearance of BMSC sheet-
scaffolds constructs with induction (Right) and non-induction control constructs (Left) after 4 weeks. (B) X-ray detected bone like tissue formation in 28
day (C) 56 day (D) 84 day implants. (E) X-ray image of constructs without cell seeding (Control). (F) Micro CT demonstrated the overall highly
mineralized tissue similar to cortical (golden) and cancellous (red) bone in implanted constructs after 28 day. (G) Micro CT images disclosed the hard
tissue formation within constructs. The red color represented mineralized tissue with similar density with cancellous bone while the golden color
represented cortical bone. (H) Fluorescence was detected on the formed bone tissue after 28 days; the fluorescence came from the cFDA labeled BMSC. (I)
Micro CT quantification of implant tissue compositions depicted substantial bone formation in the induction group, accounting for 40% total volume for
the 28, 56 and 84 day implantation, while the control group formed only connective tissue. (J) H&E staining shows lamellar bone-like tissue formed in
both outer part and interior of constructs after 28 days. (K) High-magnification image shows wel-organized lamellar bone like tissue (Bo) with distinct
osteocytes located within bone tissue. (L) The typical osteoblasts (OB, black arrow) located on the surface of neo-mineralized tissue with marrow cavities
and blood vessels in 56 days implants. (M) Safranin-O staining demonstrated that hypertrophic chondrocytes (white arrow) were observed in 56 day
implants in very low numbers. High magnification shows the chondrocytes (white arrow) with weak safranin-O staining surrounded by the osteocyte
(black arrow). (N) OCN staining. Strong signals (arrow head) were detected on neo mineralized tissue while very weak to no-signals were detected on
chondrocyte-like cells (O). (P) Collagen type-I staining. Extensive staining was detected on the neo-mineralized tissue in constructs. (Q) Collagen type-II
staining. Limited signals (arrow head) were detected on chondrocyte like cells while no staining for the mineralized tissues. BV: blood vessel+red blood
cells; Ma: marrow; OB: osteoblast; Bo: bone; OC: osteocyte; CY: chondrocyte. Scale bar: (J) 200mm; (H–Q) 50mm.
Y. Zhou et al. / Biomaterials 28 (2007) 814–824822
observed in the periphery of the constructs, which mighthave hindered nutrient diffusion and the early establish-ment of blood vessels into the middle of constructs.Consequently a hypoxia environment, promoting carti-lage-like tissue and chondrocyte development was created[35,36]. Intramembranous bone formation was also ob-served at the outer side of neo bone, demonstrating typicaltrabecular bone features as observed in skeletal formationin a developing embryo and during fracture repair; when
osteogenic cells in the presence of a good blood supply andadjacent to a local deposit of calcium salts will differentiateinto bone cells [37]. This is reasonable, since the cells indense mineralized sheets were well contacted with newblood vessel in vivo.The implanted constructs demonstrated improved me-
chanical properties with compression strength of 184MPa,which is at the lower range of cortical bone. This impliesthat the engineered bone tissue integrated well within the
ARTICLE IN PRESSY. Zhou et al. / Biomaterials 28 (2007) 814–824 823
composites. On the other hand, the mechanical propertiesof constructs were still less than native cortical bone sincethe implants were ectopically implanted, and did notexperience biomechanical stimulation comparable with aorthotopic sites [14]. Given a more dynamic environment,the constructs may exhibit superior mechanical strength.Another observation was that constructs with long invitro culture time (8 weeks) demonstrated lower inferiormechanical properties and less bone volume formationthan the immediate implanted constructs without pre-culture. The results might have arisen owing to: (1) Densersheets in the immediate group (9 layers) than in the in vitrogroup (3 layers) and this meant that the immediate groupexperienced stronger condensation processes due to cell-sheet self-contraction [3,34]; (2) The long in vitro culturetime may also have reduced the ability of cells in constructsto synthesize compact mineralized tissue [38].
5. Conclusion
In conclusion, large bone grafts with a volume of up to400mm3 were engineered through combination of BMSCsheets and composite scaffolds. The results presented inthis study show the potential to use the sheet-scaffoldconstructs in medium and high load-bearing defects.Several studies applying the presented technology in load-bearing orthotopic models in immunocompetent animalshave been planned and are currently being undertaken inour laboratory.
Acknowledgements
We thank Mr. Tan Kim Cheng (Temasek Polytechnic,Singapore) for support in the scaffold fabrication andDr. Marian Cholewa (Singapore Synchrotron LightSource) for support with the Phase Contrast Imaging andTomography (PCIT). This work was supported in part bythe Biomedical Research Council (BMRC), Grant numberR-397-000-005-305.
Reference
[1] Langer R, Vacanti JP. Tissue engineering. Science 1993;260:920–6.
[2] Langer RS, Vacanti JP. Tissue engineering: the challenges ahead. Sci
Am 1999;280:86–9.
[3] Davies JE. Bone engineering. em squared incorporated. Toronto,
Canada, 2000.
[4] Hutmacher DW. Scaffold design and fabrication technologies
for engineering tissues—state of the art and future perspectives.
J Biomater Sci Polym Ed 2001;12:107–24.
[5] Vacanti JP, Langer R. Tissue engineering: the design and fabrication
of living replacement devices for surgical reconstruction and
transplantation. Lancet 1999;354(Suppl. 1):SI32–4.
[6] Rezwan K, Chen QZ, Blaker JJ, Boccaccini AR. Biodegradable and
bioactive porous polymer/inorganic composite scaffolds for bone
tissue engineering. Biomaterials 2006;27:3413–31.
[7] Schantz JT, Teoh SH, Lim TC, Endres M, Lam CX, Hutmacher DW.
Repair of calvarial defects with customized tissue-engineered bone
grafts I. Evaluation of osteogenesis in a three-dimensional culture
system. Tissue Eng 2003;9(Suppl. 1):S113–26.
[8] Endres M, Hutmacher DW, Salgado AJ, Kaps C, Ringe J, Reis RL,
et al. Osteogenic induction of human bone marrow-derived mesench-
ymal progenitor cells in novel synthetic polymer–hydrogel matrices.
Tissue Eng 2003;9:689–702.
[9] Yang J, Yamato M, Kohno C, Nishimoto A, Sekine H, Fukai F,
et al. Cell sheet engineering: recreating tissues without biodegradable
scaffolds. Biomaterials 2005;26:6415–22.
[10] Nishida K, Yamato M, Hayashida Y, Watanabe K, Yamamoto K,
Adachi E, et al. Corneal reconstruction with tissue-engineered cell
sheets composed of autologous oral mucosal epithelium. N Engl J
Med 2004;351:1187–96.
[11] Shimizu T, Sekine H, Yang J, Isoi Y, Yamato M, Kikuchi A, et al.
Polysurgery of cell sheet grafts overcomes diffusion limits to produce
thick, vascularized myocardial tissues. FASEB J 2006.
[12] Pouliot R, Larouche D, Auger FA, Juhasz J, Xu W, Li H, et al.
Reconstructed human skin produced in vitro and grafted on athymic
mice. Transplantation 2002;73:1751–7.
[13] L’Heureux N, Paquet S, Labbe R, Germain L, Auger FA. A
completely biological tissue-engineered human blood vessel. FASEB
J 1998;12:47–56.
[14] Shao XX, Hutmacher DW, Ho ST, Goh JC, Lee EH. Evaluation of a
hybrid scaffold/cell construct in repair of high-load-bearing osteo-
chondral defects in rabbits. Biomaterials 2006;27:1071–80.
[15] Hutmacher DW, Schantz T, Zein I, Ng KW, Teoh SH, Tan KC.
Mechanical properties and cell cultural response of polycaprolactone
scaffolds designed and fabricated via fused deposition modeling. J
Biomed Mater Res 2001;55:203–16.
[16] Chen F, Chen S, Tao K, Feng X, Liu Y, Lei D, et al. Marrow-derived
osteoblasts seeded into porous natural coral to prefabricate a
vascularised bone graft in the shape of a human mandibular ramus:
experimental study in rabbits. Br J Oral Maxillofac Surg 2004;
42:532–7.
[17] Zhou Y, Hutmacher DW, Sae-Lim V, Lim TM. Effect of collagen-I
modified composites on proliferation and differentiation of human
alveolar osteoblasts. Aust J Chem 2006;59:571–8.
[18] Zhou YF, Sae-Lim V, Chou AM, Hutmacher DW, Lim TM. Does
seeding density affect in vitro mineral nodules formation in novel
composite scaffolds? J Biomed Mater Res A 2006;78:183–93.
[19] Schantz JT, Lim TC, Ning C, Teoh SH, Tan KC, Wang SC, et al.
Cranioplasty after trephination using a novel biodegradable burr hole
cover: technical case report. Neurosurgery 2006;58 ONS-E176.
[20] Griffith LG, Naughton G. Tissue engineering—current challenges
and expanding opportunities. Science 2002;295:1009–14.
[21] Muschler GF, Nakamoto C, Griffith LG. Engineering principles of
clinical cell-based tissue engineering. J Bone Joint Surg Am 2004;86-
A:1541–58.
[22] El-Ghannam A. Bone reconstruction: from bioceramics to tissue
engineering. Expert Rev Med Devices 2005;2:87–101.
[23] Owen TA, Aronow M, Shalhoub V, Barone LM, Wilming L,
Tassinari MS, et al. Progressive development of the rat osteoblast
phenotype in vitro: reciprocal relationships in expression of genes
associated with osteoblast proliferation and differentiation during
formation of the bone extracellular matrix. J Cell Physiol 1990;
143:420–30.
[24] Stringa E, Filanti C, Giunciuglio D, Albini A, Manduca
P. Osteoblastic cells from rat long bone. I. Characterization of their
differentiation in culture. Bone 1995;16:663–70.
[25] Prince M, Banerjee C, Javed A, Green J, Lian JB, Stein GS, et al.
Expression and regulation of Runx2/Cbfa1 and osteoblast pheno-
typic markers during the growth and differentiation of human
osteoblasts. J Cell Biochem 2001;80:424–40.
[26] Shui C, Spelsberg TC, Riggs BL, Khosla S. Changes in Runx2/Cbfa1
expression and activity during osteoblastic differentiation of human
bone marrow stromal cells. J Bone Miner Res 2003;18:213–21.
[27] Adachi T, Osako Y, Tanaka M, Hojo M, Hollister SJ. Framework
for optimal design of porous scaffold microstructure by computa-
tional simulation of bone regeneration. Biomaterials 2006;27:
3964–72.
ARTICLE IN PRESSY. Zhou et al. / Biomaterials 28 (2007) 814–824824
[28] Karageorgiou V, Kaplan D. Porosity of 3D biomaterial scaffolds and
osteogenesis. Biomaterials 2005;26:5474–91.
[29] Risbud M, Ringe J, Bhonde R, Sittinger M. In vitro expression of
cartilage-specific markers by chondrocytes on a biocompatible
hydrogel: implications for engineering cartilage tissue. Cell Trans-
plant 2001;10:755–63.
[30] Cooper ML, Andree C, Hansbrough JF, Zapata-Sirvent RL,
Spielvogel RL. Direct comparison of a cultured composite skin
substitute containing human keratinocytes and fibroblasts to an
epidermal sheet graft containing human keratinocytes on athymic
mice. J Invest Dermatol 1993;101:811–9.
[31] Ouyang HW, Toh SL, Goh J, Tay TE, Moe K. Assembly of bone
marrow stromal cell sheets with knitted poly (L-lactide) scaffold for
engineering ligament analogs. J Biomed Mater Res B Appl Biomater
2005;75:264–71.
[32] Goldstein AS. Effect of seeding osteoprogenitor cells as dense
clusters on cell growth and differentiation. Tissue Eng 2001;7:
817–27.
[33] Stopak D, Harris AK. Connective tissue morphogenesis by fibroblast
traction. I. Tissue culture observations. Dev Biol 1982;90:383–98.
[34] Tacchetti C, Tavella S, Dozin B, Quarto R, Robino G, Cancedda
R. Cell condensation in chondrogenic differentiation. Exp Cell Res
1992;200:26–33.
[35] Pfander D, Cramer T, Schipani E, Johnson RS. HIF-1 alpha controls
extracellular matrix synthesis by epiphyseal chondrocytes. J Cell Sci
2003;116:1819–26.
[36] Robins JC, Akeno N, Mukherjee A, Dalal RR, Aronow BJ,
Koopman P, et al. Hypoxia induces chondrocyte-specific gene
expression in mesenchymal cells in association with transcriptional
activation of Sox9. Bone 2005;37:313–22.
[37] Stevens MM, Marini RP, Schaefer D, Aronson J, Langer R, Shastri
VP. In vivo engineering of organs: the bone bioreactor. PNAS
2005;102:11450–5.
[38] Van den Dolder J, Vehof JW, Spauwen PH, Jansen JA. Bone
formation by rat bone marrow cells cultured on titanium fiber mesh:
effect of in vitro culture time. J Biomed Mater Res 2002;62:350–8.