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Journal of Biomaterials Science, Polymer Edition

ISSN: 0920-5063 (Print) 1568-5624 (Online) Journal homepage: http://www.tandfonline.com/loi/tbsp20

Development and evaluation of cross-linkedcollagen-hydroxyapatite scaffolds for tissueengineering

Niladri Nath Panda, Sriramakamal Jonnalagadda & Krishna Pramanik

To cite this article:Niladri Nath Panda, Sriramakamal Jonnalagadda & Krishna Pramanik

(2013) Development and evaluation of cross-linked collagen-hydroxyapatite scaffolds fortissue engineering, Journal of Biomaterials Science, Polymer Edition, 24:18, 2031-2044, DOI:

10.1080/09205063.2013.822247

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Published online: 01 Aug 2013.

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Development and evaluation of cross-linked collagen-hydroxyapatitescaffolds for tissue engineering

Niladri Nath Pandaa, Sriramakamal Jonnalagaddab and Krishna Pramanika*

aDepartment of Biotechnology and Medical Engineering, National Institute of Technology,Rourkela 769008, Odisha, India; bPhiladelphia College of Pharmacy, University of the Sciences,

600 S 43rd Street, Philadelphia PA 19104, USA

(Received 27 April 2013; accepted 2 July 2013)

This study examines the tissue engineering potential of type I collagen cross-linkedin the presence of hydroxyapatite (HAp). Scaffolds were prepared by controlledfreezing followed by lyophilization of composite mixtures of collagen and HAp inacetic acid, followed by cross-linking with 0.3% glutaraldehyde. Scaffolds of threeratios were prepared, corresponding to collagen/HAp ratios of 1:2, 1:4, and 1:6. Thescaffolds were evaluated for their microstructure, chemical and physical properties,swelling behavior, mechanical strength, biodegradability hemocompatability, cyto-compatibility, and histopathology following subcutaneous implantation in SpragueDawley rats. The collagen/HAp matrices showed a smaller pore size of 1040 mcompared to 50100 m for pure collagen scaffolds. Pure collagen showed amechanical strength of 0.25 MPa, and the value almost doubled for cross-linkedcomposites with collagen/HAp ratio 1:6. The improvement in mechanical strengthcorresponded to a decrease in swelling and enzymatic degradation (measured by

resistance to collagenases). FTIR spectra results in conjunction with scanning elec-tron micrographs showed that cross-linking in the presence of HAp did not signi-cantly alter the structure of collagen. MTT assay and calcein AM staining revealed

prominent and healthy growth of mesenchymal stem cells in both the pure collagenas well as collagen:HAp composites of ratio 1:2. In vivo implantation in SpragueDawley rats showed an initial acute inammatory response during days 3 and 7, fol-lowed by a chronic, macrophage-mediated inammatory response on days 14 and28. Overall, a cross-linked collagen/HAp composite scaffold of ratio 1:2 was identi-ed as having potential for further development in tissue engineering.

Keywords: collagen; hydroxyapatite; scaffolds; tissue engineering

1. Introduction

Tissue engineering presents an alternative approach to repair or replace lost or damaged

human tissue resulting from injury or disease.[1] The discipline encompasses an under-

standing of the attachment and proliferation of progenitor cells [2,3] or growth factors

[4] onto scaffolds. Skeletal tissue is a highly organized structure comprising of mainly

type I collagen, nanometre-sized carbonate-substituted hydroxyapatite (HAp) crystals,

water, and various other noncollagenous proteins. Collagen type I presents natural bind-

ing sites such as Arg-Gly-Asp (RGD) and Asp-Gly-Glu-Ala (DGEA) peptide sequences

*Corresponding author. Email: [email protected]

Journal of Biomaterials Science, Polymer Edition, 2013

Vol. 24, No. 18, 20312044, http://dx.doi.org/10.1080/09205063.2013.822247

2013 Taylor & Francis


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that modulate the adhesion of osteoblasts and broblasts.[5,6] A major limitation for

the use of collagen alone is its low mechanical strength, high swelling capacity, and

rapid degradation. HAp is a common excipient used to reinforce collagen in matrices

intended for tissue engineering.[710] HAp is a mineral (a form of calcium apatite),

and, therefore, lacks the natural binding sites presented in collagen. It also has certain

drawbacks such as limited biodegradability and osteointegration properties.[11]

Several additional excipients have, therefore, been explored to limit the use of high

ratios of HAp in collagen/HAp matrices. Examples include gelatin,[12] poly(-lactide-

co-verepsilon-caprolactone),[3] Poly L-lactic acid,[13] and Poly-benzyl-L-glutamate.[14]

This research proposes to strengthen a collagen/HAp matrix that is reinforced via cross-

linking with glutaraldehyde as the cross-linking agent. Strengthening via cross-linking is

expected to enhance mechanical strength, decrease swelling and degradation rate at rela-

tively low concentrations of HAp. Glutaraldehyde has previously been used to strengthen

collagen scaffolds.[15] The microstructural mechanism for reinforced strength has been

reported to be due to increased resistance to bril rotation.[16]

2. Materials and methods

2.1. Collagen isolation

The use of rat tail tendon for tissue engineering applications is well established in litera-

ture.[17] In this study, type I collagen was extracted from Sprague Dawley rat tail ten-

dons in a manner similar to that described by Liu et al. [18] with modications.

Sprague Dawley rat tails were obtained from the animal house of Indian Institute of

Chemical Biology, India. HAp was prepared by microwave synthesis using the method

used by Liu et al. [19]

2.2. Determination of collagen purity

The purity of isolated collagen was analyzed using sodium dodecyl sulfate poly acrylam-

ide gel electrophoresis (SDS-PAGE) according to the method described by lammelli.[20]

Briey, a 3% stacking gel was prepared from a stock solution of 30% w/w acrylamide and

0.8% w/w of N,N bis methylene acrylamide. The nal concentration of gel prepared for

separation was 0.375 M TrisHCl (pH 8.8) and 0.1% SDS. Tetramethylenediamine and

ammonium persulfate were used for gel polymerization, and the casted gel was cut to a

dimension of 15 cm 6 cm. The sample was immersed in boiling water for 1.5 min toenable protein dissociation. The electrophoresis was performed with a current of 3 mA per

gel until the bromophenol blue marker reached the bottom of the gel (about 78 h). Theproteins were xed in the gel with 50% TCA overnight, stained for 1 h at 37 C with a

0.1% coomassie brilliant blue solution made freshly in 50% TCA. The gels were then

washed repeatedly in 7% acetic acid prior to taking the autoradiograph.

2.3. Fabrication of scaffolds

A 0.5% w/v solution of collagen in acetic acid was used as the stock solution for scaffold

preparation. HAp was added to correspond to a collagen:HAp ratio of 1:2, 1:4, or 1:6.

The solution mixtures were gradually frozen and subsequently lyophilized to obtain

scaffolds. The HAp-containing scaffolds were cross-linked by immersion in a

glutaraldehyde bath of concentration 0.3% and pH 8.0 for 2 h at room temperature. Thescaffolds were then washed repeatedly with distilled water, and then lyophilized for an

additional 14 h at40 C to eliminate water without loss of structural integrity (Table 1).

2032 N.N. Pandaet al.


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2.4. Characterization of scaffolds

(1) Scanning electron microscopy: The surface microstructures of the scaffolds

before and after cross-linking with glutaraldehyde were observed under scanning

electron microscope (SEM, JeolJsm 5200) operated at an accelerating voltage of

2

15 kV. The mean pore size of the scaffolds was determined from SEM imagesby importing data to image-J software. Ten pores were measured for each mea-

surement, in a perpendicular direction through the long axis and the minor axis.

Average values were reported as the mean pore size. The tensile strength of the

scaffolds was evaluated on an Instron Material tensile testing machine 4204

(UK) using pneumatic grips with load cell 500 N at room temperature. The

cross-head speed was set at1 mm min1. Four samples were tested for each com-

position. The specimens were of rectangular shape and the dimensions were

measured with dial calipers. Three readings were taken at different positions,

and mean values were used to calculate the cross-section. Infrared spectra were

obtained in absorption mode using a FTIR spectrometer (NicoletTM Thermo

380) from collagen-HAp scaffolds.

(2) Swelling behavior: The water absorption capacity of collagen-HAp scaffolds was

determined by equilibrium swelling studies.[21] Four scaffolds were taken from

each specimen, and their dry weights recorded. The scaffolds were then dipped

in 30 mL PBS solution in Petri dishes, and weights recorded at 10 min intervals

for up to 2 h. At each immersion interval, the samples were removed, dabbed

with a lter paper to remove excess water, and weighed on a microbalance with

a sensitivity of 0.01 mg. The sample weight before and after immersion was

denoted as W0 and We, respectively. The percent swelling at equilibrium, Esw,

was calculated from the Flory-Huggins swelling equation shown below:

Esw We W0

W0

100 1

(3) Water permeation: Water permeation studies were performed following protocol

described by Sato et al. (2010).[22] Briey, a water reservoir was attached with

a polyethylene tube. The collagen and its composite scaffold of about 4 cm in

length were connected to the other end of polyethylene tube. The water reservoir

was set to a height of about 165 cm. A pressure transducer was set to verify a120 mm Hg pressure constantly at its distal end. The permeated water through

Table 1. Collagen-HAp composites compositions.

SpecimenCollagen:

hydroxyapatite (w/w)Concentration of collagen

solution (w/v) in acetic acid, %

Collagen 0.5

CH1 1:2 0.5CH2 1:4 0.5CH3 1:6 0.5

Journal of Biomaterials Science, Polymer Edition 2033


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the scaffold structure was collected with respect to time(s) in a graduated cylin-

der in order to calculate in mL/min/cm2.

(4) Biodegradability: In vitro degradation tests of cross-linked collagen-HAp

scaffolds were performed using bacterial collagenase obtained from Clostridium

histolyticum (EC 3.4.24.3, Sigma Chemical Co., St. Louis, MO, USA) with a

collagenase activity of 300 units/mg. The in vitro degradation studies for

collagen and its composite scaffolds were performed at 37 C in 10 mL phos-

phate buffered solution (pH 7.4) containing collagenase in CO2 incubator

(Lishen HF151UV). The dynamic degradation of scaffolds was monitored for

four compositions with known dry weights by incubated at varying collagenase

concentrations (0.1, 0.5, 1, and 2 mg) for 24 h.

2.5. Preliminary biocompatibility testing

(1) Hemo-compatibility testing: Human peripheral blood was collected in a 15 mL

tube containing 3.8% sodium citrate to avoid coagulation.[23] Blood diluted

with PBS (pH 7.4) at a ratio of 1:20 was used as the negative control, while

blood mixed with triton X was the positive control. Punched scaffolds of 5 mm

diameter were sterilized by exposure to UV light in a sterile culture hood, fol-

lowed by rinsing with 70% ethanol. The sterile scaffolds were immersed in

100 l of blood diluted with PBS solution followed by incubation at 37 C for

60 min. The samples were then centrifuged at 3000 rpm for 10 min. The absor-

bance of the supernatant was measured at 545 nm.

(2) Calcein AM staining: Pure and composite scaffolds of 6 mm diameter wereadded individual wells of a 12 well-tissue culture plate. MSCs in DMEM media

were added to these wells at the concentration of 105 cells/mL. The cells were

placed in a 5% CO2 incubator for 1012 h to enable attachment to scaffolds, fol-

lowing which the media was discarded, and calcein green AM added at a con-

centration of 10 M. The stain was prepared by dissolving in DMSO (less than

0.01%), and adding to fresh DMEM media.

(3) Inside the cells, calcein-AM is hydrolyzed by endogenous esterases into the

negatively charged green uorescent calcein, which is retained inside the cyto-

plasm. After incubation in the stain for an hour, the media was discarded and

the cells thoroughly washed with PBS (pH = 7.2). Around 810 beads were

then taken on a clean microscopic slide and mounted with glycerol: PBS (1:1

ratio), and observed under an upright uorescence microscope (Carl Zeiss

E600, Wavelength = 450 nm) using a green lter (wavelength = 490 nm) at

50 magnication. The cell-scaffold constructs were sectioned and the celladhesion and viability were also observed in the inner regions of the scaffolds.

The images were captured using a Carl Zeiss Axiocam camera attached to the

microscope.

(4) Cell culture studies: Mesenchymal stem cells (MSCs) were isolated from umbili-

cal cord blood (obtained from Ispat General Hospital, Rourkela, India) using

DMEM F-12 medium (Gibco BRL, USA) supplemented with 10% fetal bovine

serum (Gibco BRL, USA) and 1% U/mL streptomycinpenicillin and incubatedat 37C and 5% CO2. The MSCs were sub-cultured every three days by rst

adding 0.25% trypsin and 0.02% EDTA for detachment.



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The prepared scaffolds were cut into cylindrical disks with a 6 mm dermal biopsy

punch (Miltex) and kept in a 96-well plate. They were sterilized with 70% ethyl alcohol

(v/v) followed by UV treatment, followed by hydration in sterile PBS (pH 7.4). The

scaffolds were then conditioned in DMEM medium for 4 h before cell seeding. Cells

were seeded at densities of 1105 cells/well on these scaffolds. The seeded constructswere incubated for 3 or 5 days at 37 C and 5% CO2 to examine for growth and prolif-

eration. The numbers of living cells inside the scaffold were determined using the MTT

assay. After discarding the culture media, the cells were further washed with PBS to

remove the non-adhered cells. One ml/well of fresh medium and 100 l/well of MTT

was added to the wells, and the cells were cultured for an additional 4 h. The same pro-

cedure was repeated for 3 identical samples, except that these were cultured for an addi-

tional 12 h. The medium was then removed, and 150l of dimethyl sulfoxide (DMSO)

was added to each well. The optical density (OD) was measured at wavelength 595 nm

using a microplate reader (Model PerkinElmer 2030 explorer). The percentages of

attached cells were calculated using following formula as:

The OD595of cells on matrix after washing

The OD595of cells in culture wells 100%

2.6. Animal implantation and histology analysis

Sprague Dawley rats were implanted subcutaneously with 1 cm2 ECM scaffold materials

with 1 mm thickness. The rats were sacriced at days 3, 7, 14, and 28 post-implanta-

tion. The tissues along with the scaffold were collected in a neutral buffered formalin

xative. After

xation, the tissues were processed in an automated tissue processor and

tissue sections of 4 m thickness were cut using Leica microtome (Leica RM2255,

Germany). The tissue sections were then stained with routine hematology and eosin

stain and examined under inverted Olympus microscope.

2.7. Statistical analysis

The data are expressed as the mean SEM for three independent experiments and its

statistical signicance was evaluated by one way ANOVA followed by Bonferronis

post hoc test (GraphPad Prism). p < 0.05 was considered statistically signicant.

3. Results

3.1. Characterization of composite scaffolds

3.1.1. Microstructure evaluation

Figure 1 shows SEM photographs of the surface morphology and pore structure of sin-

gle collagen and its composite, i.e. collagen: HAp scaffolds. All scaffolds showed rela-

tively similar macroporous morphology with a high degree of interconnectivity.

Collagen-HAp scaffolds appeared less regular compared to pure collagen scaffolds.

Scaffolds prepared from pure collagen had a larger pore size of approximately

50100 m compared to the composite scaffolds that were considerably smaller at

1040 m. The 3-dimensional interconnectivity that could be seen in all scaffolds isconsidered necessary to enable host cell inltration to guide the differentiation and pro-

liferation of cells toward the targeted functional tissue or organ.[24,25]



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3.1.2. Mechanical analysis

Pure collage showed a tensile strength of 0.251 0.007 MPa. The tensile strength

increased signicantly (p < 0.005) with the increasing ratio of HAp in the composite

scaffold as shown in Figure 2. The mechanical strength almost doubled for scaffolds

prepared with a collagen:HAp ratio of 1:6.

3.1.3. Swelling and water permeability studies

The results of both these studies demonstrated that the inclusion of HAp lowers the

hydrophilic characteristics of the scaffold. Figure 3 shows that equilibrium swelling

decreased signicantly upon incorporation of HAp. The water permeability of the scaf-

fold to water was also consistently lowered upon the inclusion of water. Water ux

through the scaffolds was measured at 57 5 mL/min/cm2 for pure collagen scaffolds,

and 39 4.3, 28 6, and 15 3.5 mL/min/cm2 for collagen composite scaffolds of

composition 1:2, 1:4, and 1:6 ratios, respectively.

Figure 1. Scanning electron micrographs of: (a) collagen scaffold; (b) collagen:HAp (1:2), (c)collagen:HAp (1:4); and (d) collagen:HAp (1:6).



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Figure 2. Effect of HAp ratio on the scaffold strength.

Figure 3. The effect of HAp ratio on water uptake measured as present swelling.



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3.1.4. Infrared analysis

These studies were performed to ensure that the cross-linking with HAp does not affect

the chemistry of collagen. This is important considering that cells are more prone toattach surfaces containing natural binding sites.[26]

Infrared spectrum of a composite scaffold is shown in Figure 4. The FTIR spectra

showed peaks at 3485.6 cm1 and 1529.57 cm1. These were in the range of -NH

stretching and N-H bending of secondary amide on collagen type I, indicating that these

functional groups may still be available to promote cell attachment to collagen. HAp is

characterized by an anti-symmetric phosphate band at around 12001000 cm1.[27] The

peak observed at 1203.5 cm1 in Figure 4 corresponds to phosphate stretching, and

conrms the presence of HAp in the scaffold.

3.1.5. Biodegradability testing

Figure 5 shows that the cross-linked collagen-HAp scaffolds showed greater

resistance to degradation than the pure collagen scaffolds in all collagenase

solutions.

3.1.6. Hemo-compatibility testing

Blood compatibility is a critical requirement for the blood interfacing biomaterials. The

implants should not damage proteins, enzymes, or elements of blood (red blood cells,

white blood cells, and platelets). Figure 6 shows that the % hemolysis increased two to

three times for the cross-linked collagen/HAp scaffolds compared to pure collagenscaffolds.

Figure 4. A representative FTIR spectra of collagen-HAp scaffolds (collagen:HAp ratio 1:2).



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3.1.7. Attachment and proliferation of mesenchymal stem cells onto the scaffolds

The results for MTT and cell adhesion assay are shown in Figure 7. The results wereplotted by measuring the OD and reported as percentage of cells adhered relative to

seeded MSCs. The attachment ratios were found to be more than 50% in all cases when

Figure 6. The effect of HAp ratio on hemo-compatibility, measured as percent% hemolysis inhuman peripheral blood.

Figure 5. The effect of HAp ratio on scaffold degradation.



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MSCs were cultivated for 12 h in the media and seeded over the scaffold. Cell prolifera-tions were performed for 3 and 5 days, and signicant increases were observed.

Figure 7. The effect of HAp ratio on the adhesion of MSCs.

Figure 8. (a) and (b) Two representative uorescent microscopy photographs of calcein AMstaining in collagen/HAp scaffolds of ratio 1:2. (c) The scaffold after 72 h of incubation.



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Morphological analysis of cells was performed through uorescence microscopy

after Calcein AM staining. Figure 8(a) and (b) shows viable green colored cells almost

evenly distributed on surface as well as within the scaffold as shown. Figure 8(c) shows

that after 72 h of incubation, the cells undergo a morphological change from rounded to

spreading type morphological features conrming a healthy microenvironment of the

scaffold for cellular differentiation and proliferation.

3.1.8. Animal implantation and histology analysisThe tissue reaction was assessed by histopathology (Figure 9). Tissue response at days

3 and 7 was that of an acute inammatory reaction mediated by neutrophilic inltration.

The reaction then shifted to chronic inammation predominated by macrophages by

14 days. The chronic inammation persisted at 28 days, but the reaction was less intense

compared to response at day 14. There was evidence of healing by brosis from day

seven, progressing until the 28th day.

4. Discussion

Pore size is an important characteristic of membrane-based scaffolds as it determinesthe overall surface area available for cell-attachment. The composite scaffolds reported

in this study have pore diameters in the range of about 10 to 40m and, therefore,

Figure 9. Tissue reaction to cross-linked scaffolds of ratio 1:2 at 3 days (a), 7 days (b), 14 days(c), and at 28 days (d) after subcutaneous implantation in Sprague Dawley rats. The representative

scaffold shown here represents a ratio of collagen/HAp of 1:2.



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possess signicantly greater surface area compared to pure collagen scaffolds. The pore

sizes obtained in our study were similar to that of Chen et al. [28] Such dissimilar mor-

phologies consisting of a highly porous structure and high mechanical strength may

have potential for the design of dermal implants. The porous surface can facilitate the

recruitment and proliferation of dermal broblasts to promote wound healing, while

composite structure can contribute to mechanical strength and barrier efcacy to prevent

infection and dehydration.[29] The roughly spherical shape of the pores as well as their

presence on the surface suggests that pore diameters may be inuenced by the size of

the evaporating solvent droplets as well as particle size of HAp.[30] Pores over the sur-

face of the scaffold are most common sites of stress concentrator and crack initiator.

But incorporation of HAp increases the density of the composite material which retards

crack propagation. Moreover, the decrease in pore size for the composite improves the

circularity index of the pores and hence decreases crack propagation.[31] This property

has been veried with measurement of ultimate strength (Figure 2) where increase in

addition of HAp improves compressive strength of the scaffold. These observations are

consistent with previous literature.[32] The possible reason may be explained by eitheran increase in total binding area of collagen with HAp or reduction of defects in case

of composite scaffold, both of which may lead to an increase in mechanical strength.

This conclusion is further supported by our decrease in swelling ratio for composite

scaffold. Okoshi et al. proposed that porous grafts with water permeability of about 10

40 mL/min/cm2 induced endothelization and kept patency.[33] The porosity of both col-

lagen and composite scaffolds was in a comparable range.

A decrease in availability of interactive surface area in the composite scaffold

results lowers swelling, water permeation, and therefore exposure to enzymes responsi-

ble for degradation. The collagen/HAp scaffolds may therefore be expected to have a

lower degradation rate and a greater residence time following implantation.While the mechanical and degradation properties of the collagen/HAp scaffolds

seem favorable, a drawback of the glutaraldehyde cross-linking is the lower hemocom-

patibility compared to pure collagen scaffolds. As both collagen and HAp were consid-

ered biocompatible, the decrease may be a consequence of residual glutaraldehyde

entrapped in the collagen-HAp composites. It is possible that the amount of residual

glutaraldehyde is higher in scaffolds containing greater amount of HAp. This outcome

is further supported by the fact that the collagen:HAp (1:2) composite showed nearly

same proliferation rate as that of pure collagen, while the composite with higher ratio

showed decreased cell proliferation and attachment.

5. Conclusion

This study shows that cross-linked collagen-HAp scaffolds may present a valuable bone

tissue engineering platform. Advantages include a relatively slow degradation rate, high

mechanical strength, and a collagenous macroporous structure with potential to trigger

cell attachment and proliferation. The biocompatibility of composite scaffold though

found to be less than those of pure collagen scaffold, the collagen:HAp with 1:2 com-

posite has comparable degree of cell adhesion and viability with those of pure collagen.

However, there is decrease in haemocompatibility with the increasing concentration of

HAp which may be due to increased amount of glutaraldehyde retention in the

composite.



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Acknowledgment

The Indian Institute of Chemical Biology, Kolkata for providing the facilities for conducting thevarious tests, especially Dr T Chakraborty.

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