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JOURNAL OF TISSUE ENGINEERING AND REGENERATIVE MEDICINE
RESEARCH ARTICLE
J Tissue Eng Regen Med 2015; 9: 1286–1297.
Published online 24 January 2013 in Wiley Online Library (wileyonlinelibrary.com) DOI: 10.1002/term.1682
An additive manufacturing-based PCL–alginate–
chondrocyte bioprinted scaffold for cartilage
tissue engineering
Joydip Kundu1†, Jin-Hyung Shim1†, Jinah Jang2, Sung-Won Kim3 and Dong-Woo Cho1,2*
1
Department of Mechanical Engineering, Pohang University of Science and Technology (POSTECH), Kyungbuk, South Korea
Division of Integrative Biosciences and Biotechnology, Pohang University of Science and Technology (POSTECH), Kyungbuk, South Korea
3
Department of Otolaryngology-Head and Neck Surgery, The Catholic University of Korea, College of Medicine, Seoul, Korea
2
Abstract
Regenerative medicine is targeted to improve, restore or replace damaged tissues or organs using a
combination of cells, materials and growth factors. Both tissue engineering and developmental biology
currently deal with the process of tissue self-assembly and extracellular matrix (ECM) deposition. In this
investigation, additive manufacturing (AM) with a multihead deposition system (MHDS) was
used to fabricate three-dimensional (3D) cell-printed scaffolds using layer-by-layer (LBL) deposition
of polycaprolactone (PCL) and chondrocyte cell-encapsulated alginate hydrogel. Appropriate cell
dispensing conditions and optimum alginate concentrations for maintaining cell viability were determined.
In vitro cell-based biochemical assays were performed to determine glycosaminoglycans (GAGs), DNA
and total collagen contents from different PCL–alginate gel constructs. PCL–alginate gels containing
transforming growth factor-b (TGFb) showed higher ECM formation. The 3D cell-printed scaffolds of
PCL–alginate gel were implanted in the dorsal subcutaneous spaces of female nude mice. Histochemical
[Alcian blue and haematoxylin and eosin (H&E) staining] and immunohistochemical (type II collagen)
analyses of the retrieved implants after 4 weeks revealed enhanced cartilage tissue and type II collagen
fibril formation in the PCL–alginate gel (+TGFb) hybrid scaffold. In conclusion, we present an innovative
cell-printed scaffold for cartilage regeneration fabricated by an advanced bioprinting technology.
Copyright © 2013 John Wiley & Sons, Ltd.
Received 27 February 2012; Revised 25 October 2012; Accepted 14 November 2012
Keywords additive manufacturing; cell printing; cell-printed scaffold; cartilage regeneration; tissue
engineering
1. Introduction
Tissue-engineering strategies attempt to improve, restore or
replace damaged tissues or organs using a combination of
specific cell types, scaffold materials and growth factors.
Natural tissues generally comprise an acellular structural
portion which is filled by a much softer matrix containing
the desired amount of cells (Langer and Vacanti, 1993).
To simulate and mimic the native tissue architecture,
*Correspondence to: Dong-Woo Cho, Division of Biosciences
and Biotechnology and Department of Mechanical Engineering, Pohang University of Science and Technology (POSTECH),
San 31, Hyoja-dong, Nam-gu, Pohang, Kyungbuk, 790-784,
South Korea. E-mail: dwcho@postech.ac.kr
†
These authors contributed equally to this study.
Copyright © 2013 John Wiley & Sons, Ltd.
composites are engineered using organized scaffolds with
the desired mechanical strength and an appropriate
amount of cells (Hutmacher, 2001). Several uses for
this type of hybrid scaffold are documented, varying from
musculoskeletal tissues, such as cartilage, bone and
ligaments, to blood vessels (Ma and Elisseeff, 2006;
Hollister, 2005). A conventional scaffold-based tissueengineering approach is based on random cell-seeding
and growth factor administration (Langer and Vacanti,
1993). In this case, cells can only attach to the scaffold
surface; their distribution inside and the inner composition
of the product cannot be controlled (Silva et al., 2006). In
addition, the concentration of local growth factors inside
is uncontrollable. As a result, it is very difficult to control
and facilitate the invasion of native vasculature from the
host tissue (Rouwkema et al., 2008). For these reasons,
PCL–alginate–chondrocyte bioprinted scaffold for cartilage tissue engineering
conventional scaffolds have considerable disadvantages
that affect their clinical applicability, including limited
cell-seeding efficiency and control over spatial distribution
and localization (Mironov et al., 2003).
Considering the inherent shortcomings of conventional
scaffold-based tissue repair, a new biofabrication approach,
termed ‘three-dimensional (3D) bioprinting’, has been introduced in regenerative medicine (Mironov et al., 2009).
Several additive manufacturing (AM) technologies enabling
the fabrication of customized parts and devices that are
geometrically complex, using a layer-by-layer process, have
been applied to direct biofabrication approaches (Rosen,
2007). These include fused deposition modelling (FDM)
(Hutmacher et al., 2004; Jung et al., 2012), stereolithography (Lee et al., 2010a, 2010b, 2011), cell sheet
lamination (Xu et al., 2010; Haraguchi et al., 2011), inkjet
printing (Boland et al., 2006; Nakamura et al., 2006),
bioplotter technique (Pfister et al., 2004; Peltola et al.,
2008), biolaser printing (Guillermot et al., 2010; Barron
et al., 2004) and selective laser sintering (Wiria et al., 2010;
Simpson et al., 2008). Each of these methods has advantages
and disadvantages. Scaffolds made of synthetic biomaterial
only are unable to provide a suitable environment for cell
attachment, due to the high order of hydrophobicity of the
synthetic biomaterials used in these scaffolds. However, direct biofabrication approaches using living cells and biological materials are reasonable means for developing organs,
because cells and extracellular matrix (ECM) proteins are
the major components of biological tissues, and tissues are
directly fabricated with those histological components by
controlling their spatial positions (Mironov et al., 2009). In
addition, for synthetic biomaterial scaffolds with the cellseeding method, specific spatial arrangements of specific cell
types cannot be achieved, despite the fact that the distribution of the cells may be influenced by variations in the local
scaffold porosity. To overcome this limitation, multihead deposition system (MHDS) and three-dimensional (3D) printing, which are able to print various biological components
including synthetic and natural biopolymers, proteins and
cells, are beneficial (Mironov et al., 2009; Shim et al., 2011).
Bioprinting strategies using the principles of AM to
deposit the cells, matrix or both by a layer-by-layer (LBL)
process allow for the development of porous 3D cell-printed
scaffolds with a detailed specific cellular arrangement
(Shim et al., 2011; Mironov et al., 2007). Bioprinting
technology facilitates the fabrication and realization of a
pretissue scaffold, based on the zonal design of native
articular cartilage. Development of the biohybrid technology
based on bioprinting concepts combines the dispensing of
specific cell types with the simultaneous deposition of
biomaterials and will aid the further development of zonal
cartilaginous tissue (Shim et al., 2011). With the advancement and applications of AM-based printing technologies,
a certain degree of organization can be engineered with a
highly ordered interconnected porous polymer network
structure (Fedorovich et al., 2011). The order of alignment
and organization developed via both scaffold design and
controlled deposition of cells at predefined locations can
potentially accelerate the organization of cells into a
Copyright © 2013 John Wiley & Sons, Ltd.
1287
functional tissue (Kim et al., 2003). Moreover, controlled
dispensing layers of isolated zone-specific chondrocytes
from the deep, middle and superficial regions of cartilage
encapsulated within natural polymer hydrogels allows the
production of cartilage pretissue constructs with controlled
architecture (Klein et al., 2009). Thus, the innovative cellprinting technology facilitates mimicking the topographical
and spatiotemporal organization of cells within the
native tissue.
Hydrogels are water-swellable, yet water-insoluble,
crosslinked networks that provide many advantages as cell
carriers for the engineering of a variety of tissues. The
highly swollen 3D environment maintains a high water
content resembling biological tissues and facilitates cell
proliferation (Peppas et al., 2006). Several natural polymers,
such as collagen, chitosan, hyaluronic (HA) acid, silk
proteins, gelatin and alginates, are widely used as hydrogel
materials for tissue-engineering applications (Hoffman,
2001). Researchers have observed that alginate hydrogels
have good biological properties, and thus they have been
widely used as therapeutic materials in regenerative medicine
(especially in cartilage tissue engineering) and for the delivery of bioactive growth factors (Jeon et al., 2009). Chondrocytes and endothelial cells encapsulated within alginate
hydrogels remain viable and metabolically active. The cellencapsulation strategy can be applied to cell printing, which
is an emerging concept in tissue engineering (Khalil and
Sun, 2009). The principal limitation of hydrogel for tissue
engineering is its inability to maintain a uniform 3D structure. To overcome this problem, hydrogels must be integrated with synthetic biomaterials. Localization of hydrogel
containing cells and growth factors within a synthetic
biomaterial scaffold has led to the engineering of pretissues,
which are critical for more homogeneous tissue formation.
In this study, we utilized the advantages of cell-printing
technology to create pretissue by LBL deposition of polycaprolactone (PCL) and chondrocytes encapsulated by hydrogels (alginate), with and without transforming growth
factor (TGFb), using MHDS, a useful system in AM technology. We established appropriate cell-printing conditions
and the desired alginate hydrogel concentration for maintaining cell viability within the hydrogel. Constructs were
cultured in vitro for the determination of glycosaminoglycan
(GAG), DNA and total collagen contents. To evaluate the
in vivo cartilage tissue formation, 3D cell-printed scaffolds
of PCL–alginate gel were implanted in the dorsal subcutaneous spaces of female nude mice, aged 7 weeks, and
removed after 4 weeks. Histochemical analysis [Alcian blue
and haematoxylin and eosin (H&E) staining] and immunohistochemistry (type II collagen) of the retrieved implants were
performed to observe the cartilage tissue regeneration in vivo.
2. Experimental
2.1. Materials
The following reagents and materials were purchased:
FDA-approved polycaprolactone (molecular weight = 70
J Tissue Eng Regen Med 2015; 9: 1286–1297.
DOI: 10.1002/term
J. Kundu et al.
1288
000–90 000; Polyscience, USA); sodium salt of alginate acid
(Sigma-Aldrich, USA), phosphate-buffered saline (PBS;
Hyclone, USA); calcium chloride (CaCl2; Sigma-Aldrich);
sodium chloride (NaCl; Sigma-Aldrich); collagenase
(Sigma-Aldrich); Dulbecco’s modified Eagle’s medium
(DMEM; Gibco BRL, USA); fetal bovine serum (FBS; Gibco
BRL); penicillin (Gibco BRL); streptomycin (Gibco BRL);
papain (Sigma-Aldrich); chondroitin sulphate (SigmaAldrich); 1,9-dimethylmethylene blue chloride (SigmaAldrich); Hoechst 33258 (Invitrogen, USA); human
transforming growth factor-b (TGFb; BD Biosciences,
USA); a hydroxyproline assay kit (BioVision Research
Products, USA); and a Live/Dead Viability Kit (Lonza,
USA). All reagents were used without further purification.
250 mm polypropylene mesh filter. The filtrate was centrifuged at 6000 rpm and the resulting cell pellet was
washed twice with copious amounts of PBS without Ca2
+
. Collected cells were suspended in DMEM containing
10% FBS, 100 U/ml penicillin and 100 mg/ml streptomycin in humidified air with 5% v/v carbon dioxide (CO2)
at 37 C. The cells were plated at 1 106 cells/cm2 into
75 mm2 cell-culture flasks. These cells were cultured
using the standard monolayer cell-culture methods. The
medium was changed after 72 h and every 3 days thereafter.
In this study, chondrocytes at passage 3 were used.
2.2. Additive manufacturing of scaffolds
Dispensing of chondrocytes was investigated using the
procedure introduced above via the MHDS. When the cellculture flask became near-confluent at 10 days, the cells
were detached with 0.05% trypsin containing 0.53 mM
ethylenediaminetetra-acetic acid (EDTA) for 5 min at
37 C. The cells were collected by centrifugation at
1000 rpm for 5 min, followed by counting the cell number
using a haemocytometer. The chondrocyte cell suspensions
(cell density 106 cells/ml) in 10% FBS were mixed with
sodium alginate solution (4% and 6%) and 10 (10 times
more concentrated) DMEM. The mixture of cells in sodium
alginate was then deposited on the spaces between the lines
previously fabricated with a PCL framework. The deposited
mixtures of cell suspension within the alginate were crosslinked with a 100 mM CaCl2 and 145 mM NaCl solution for
10 min. The constructs were then washed with sterile PBS
three times for 5 min each. The cell-printed constructs were
then placed in DMEM supplemented with 10% FBS, 100 U/
ml penicillin and 100 mg/ml streptomycin in humidified air
with 5% v/v CO2 at 37 C. Cell-printed PCL–alginate
scaffolds containing TGFb were prepared in a similar
manner. The TGFb concentration was held constant at
10 ng/ml, and the TGFb was mixed gently with the alginate
and cell suspension solution before dispensing.
PCL–alginate gel hybrid constructs were fabricated using
AM technology and a MHDS. The MHDS had four dispensing heads, in which motion, temperature and pneumatic
pressure were controlled separately. Therefore, the MHDS
was able to dispense four different biomaterials into a 3D
scaffold or fabricate four scaffolds at the same time. Manufacture was performed at room temperature and consisted
of alternating steps of synthetic polymer (PCL) and hydrogel
(alginate) printing. The PCL polymer was fed into a 10 cc
syringe, melted by heating to 80 C for 10 min and subsequently dispensed through a metal needle (nozzle diameter
250 mm), using a pressure of 650 kPa and a deposition
speed of 80 mm/min. Sodium alginate (4% and 6%) in
PBS was autoclaved and subsequently dispensed in the
spaces between the lines of the PCL layers at room temperature, using a deposition speed of 400 mm/min and a nozzle
diameter of 250 mm. Subsequent layers were deposited at a
90 angle to the underlying layer. After dispensing the
desired number of layers, the alginate was crosslinked with
100 mM CaCl2 and 145 mM NaCl solution for 10 min.
2.5. Chondrocyte printing via MHDS
2.3. Scaffold morphology observation under
scanning electron microscopy (SEM)
2.6. Cell viability using live/dead assay
The surface morphologies of the freeze-dried scaffolds were
examined using SEM (JSM-5300, JEOL, Tokyo Japan).
Lyophilized samples were mounted on an aluminium stub,
using double-sided carbon tape. The samples were coated
with gold by sputtering for 60 s and examined at an acceleration voltage of 15 kV.
2.4. Primary nasal septal cartilage chondrocyte
cell culture
Fresh human cartilage was obtained from a nasal septum
after operation. The cartilage fragments were dissected
under sterile conditions and then were subjected to collagenase digestion at 37 C for 12–18 h (Lee et al., 2006).
The resulting cell suspension was passed through a sterile
Copyright © 2013 John Wiley & Sons, Ltd.
Cell viability studies were conducted using a live/dead
viability kit. The assay dye was prepared by mixing 10 ml
ethidium homodimer (EthD-1) and 2 ml calcein AM in
10 ml PBS and then adding 2 ml solution to cell-printed
scaffolds. The cell-printed scaffolds were incubated for
30 min in the live/dead dye solution. To visualize the live
and dead encapsulated cells in the alginate gel, fluorescent
images were obtained with a fluorescent inverted microscope (Leica DMIL FL). Calcein AM (live) and EthD-1
(dead) were viewed simultaneously with a conventional
fluorescein long-pass filter. The fluorescences from these
dyes were observed separately; calcein was viewed with a
standard fluorescein band-pass filter (420–620 nm), while
EthD-1 was viewed with filters for propidium iodide or
Texas red dye (500–700 nm).
J Tissue Eng Regen Med 2015; 9: 1286–1297.
DOI: 10.1002/term
1289
PCL–alginate–chondrocyte bioprinted scaffold for cartilage tissue engineering
2.7. In vitro cell culture tests on the fabricated
hybrid constructs
Four different cell-printed hybrid constructs [(a) 4% alginate – TGFb (4%); (b) 4% alginate + TGFb (4% + TGF);
(c) 6% alginate – TGFb (6%); and (d) 6% alginate + TGFb
(6% + TGF); –,without; +, with] were screened for in vitro
estimation of sulphated GAGs, DNA and total collagen
content. Chondrocytes were suspended in culture medium
to obtain a cell suspension containing 106 cells/ml. TGFb
was added to the alginate cell mixture at a concentration
of 10 ng/ml. The constructs were cultured in DMEM and
harvested for analysis every week for a total of 4 weeks.
Samples were digested in papain digestion buffer for 16 h
at 60 C and assayed for total sulphated GAG content via
spectrophotometry with 1,9-dimethylmethylene blue chloride
(Farndale et al., 1986). Chondroitin sulphate (0–60 mg/ml)
was used as a standard. The same papain-digested samples
were used for a fluorometric assay of DNA (ng DNA/mg dry
weight) with Hoechst 33258 (Kim et al., 1988). The total
collagen content was measured, using a hydroxyproline
assay kit, by determining the hydroxyproline contents of
the specimens after acid hydrolysis and reaction with
p-dimethylaminobenzaldehyde and chloramine-T, according
to the manufacturer’s suggested protocol. The hydroxyproline contents in samples were calculated based on a
standard curve generated following the kit protocol.
2.8. Cell-printed scaffold implantation and
in vivo analysis
Three types of cell-printed scaffolds [group I, PCL +
alginate gel (no cells); group II, PCL + alginate gel
(chondrocytes); and group III, PCL + alginate gel (chondrocytes + TGFb], with a 24 h residential time in DMEM,
were implanted into the dorsal subcutaneous spaces of
7 week-old female nude mice (n = 5 implants/group) for
4 weeks (Kang et al., 2011). TGFb was added to the
alginate–cell mixture at a concentration of 10 ng/ml. Each
cell-printed scaffold of groups II and III contained 105 cells
encapsulated with the alginate gel. The animal experiments
were performed according to the protocol approved by the
Animal Care and Use Committee of Pohang University of
Science and Technology, South Korea (2011-01-0013). Four
weeks after implantation, the mice were sacrificed and all
implants were retrieved. For histological analysis, implants
retrieved from the mice after 4 weeks were fixed in formalin
for 2 h at 4 C. Prior to sectioning, the implants wereembedded in optimal cutting temperature (OCT) compound
(Tissue-Tek, USA), allowed to permeate for 30 min, frozen
by storing in a deep freezer (–80 C) for 30 min and then
sectioned transversely at a thickness of 20 mm, using a
cryomicrotome. Tissue sections were stained with Alcian
blue to evaluate sulphated GAG and with H&E to evaluate
cartilaginous tissue formation. Immunostaining with collagen type II antibody was used to evaluate the cartilage
matrix tissue formation. Slides were assembled with
resinous medium, and mounted slides were examined
Copyright © 2013 John Wiley & Sons, Ltd.
under a light microscope (Nikon Eclipse E600)/confocal
microscope (Carl Zeiss).
2.9. Statistical analysis
Data are expressed as mean standard deviation (SD)
from several separate experiments. Statistical analysis was
carried out via one-way anal …
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