Orthopedic problems such as fractures, osteoporosis, and bone tumors make bone transplantation the second most transplanted tissue in the world. Because of organ donor shortage and the problems associated with allograft transplantation, tissue engineering has emerged as a promising alternative for organ implantation to repair or replace the diseased organs.1 This strategy supports neotissue formation, both in vitro and in vivo, and is based on using cells, growth factors, and biodegradable or nondegradable scaffolds.2 To increase the biological effectiveness of the scaffolds, electrospinning technique is introduced for constructing nanofibers. These nanosized microenvironments mimic the structure of extracellular matrix (ECM) and possess unique features like high porosity with large surface area, proper interconnections, and enhanced cell recruitment, which make them more suitable for tissue engineering.3
Bone regeneration therapy mainly relays on application of adult stem cells, such as mesenchymal stem cells (MSCs), because of their potential to differentiate into different cell lineages.4 Cells in their native microenvironments receive essential information through cell–cell cross talk and ECM—which contains tissue-specific as well as general growth factors—to remain functional. ECM exerts its role in cell attachment and regulating cells functions.5
Biochemical characteristics of the scaffolds exert profound effects on cell fate.6 But it should be noted that not all materials represent suitable properties, for instance polycaprolactone (PCL) and poly-l-lactic acid (PLLA), which have proper mechanical strengths for bone regeneration, do not support proper cell attachment nor induce direct differentiation.7 In this case, surface modification plays a critical role in making this category of synthetic materials suitable for tissue engineering. Particularly, in bone tissue engineering, modification of biochemical as well as mechanical properties of the scaffolds has been under the vast investigation.8 Hydroxyapatite (HA) and b-tricalcium phosphate (b-TCP) are the most widely used bioceramics in bone regeneration studies and clinical applications. These materials enhance the osteogenesis process and bind tightly to the host bone as well.9 Although high porosity with interconnected pores of these bioceramics scaffolds favors the tissue regeneration process, but on the other hand, weakens the mechanical characteristics of the scaffolds; therefore, this issue is a major challenge in bone-tissue engineering.10 Therefore, combination of material such as PCL—which guarantees the required mechanical properties—and HA nanoparticles (nanohydroxyapatite [nHA]) with suitable biological properties makes it possible to exploit these characteristics simultaneously.10
Fibronectin (Fn), an adhesive glycoprotein, exerts its critical role by induction of specific cell morphology11,12 and activation of intracellular signaling pathways.11 It has been revealed by molecular studies that Fn has profound effect on osteogenesis and serves as regulatory checkpoint in mechanotransduction, which is required for normal osteogenic differentiation.13 In addition, it might nucleates mineralization14 and also induces HA crystal formation.15
The objective of the current study was to determine the synergistic effect of Fn and nHA in PCL scaffold on the bone regeneration. It was also investigated whether the simultaneous incorporation of nHA in nanofibrous PCL scaffolds and Fn coating can improve the biocompatibility and osteoconductivity of the scaffolds.
Materials and Methods
Polycaprolactone and PCL/nHA nanofibrous webs were constructed using two-nozzle electrospinning technique. Nanofibrous scaffolds were prepared by dissolving PCL (8% w/v) with 8 × 104 Mn mixed in a solvent of chloroform and dimethylformamide (DMF) with a volumetric ratio of 8:2. For fabrication of PCL/nHA scaffolds, nHA nanoparticles of ~200 nm size (10% w/w of polymer mass) were dispersed in chloroform/DMF and homogenized by ultrasonic homogenizer (Bandelin, Berlin, Germany) followed by addition of PCL granules to this suspension. The electrospinning condition was 2300 rpm rotation, 23 cm distance from the needle, and applied voltage of 22 kV.
Plasma Treatment and Fibronectin Coating
Plasma treatment was carried out to generate hydrophilic surface on the scaffolds, using a plasma generator (Diener Electronics, Ebhausen, Germany) with a cylindrical quartz reactor set on 44 kHz and by O2 introduction into the reaction chamber (0.4 mbar) for 5 minutes. For Fn coating on the scaffolds, plasma-treated samples were immersed in 500 µl of a mixture of 5 g/l N-hydroxysuccinimide (NHS) and 5 g/l N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC), for 24 h at room temperature (RT) to generate functional groups on the scaffold. Finally, samples were placed in 1 mg/ml Fn (1% v/v) for 24 h and dried at RT.
To determine the tensile properties of the scaffolds, section of 10 mm × 60 mm × 0.12 mm was used. Tensile test was performed in parallel direction at 20 mm/min using SANTAM testing equipment (Iran).
Contact angles of water droplets were measured at RT and ambient pressure using G10 contact angle goniometer (Kruss) before and after plasma treatment.
Fourier transform infrared spectroscopy (FTIR) was used to confirm the incorporation and grafting of HA and Fn, respectively. FTIR spectra were obtained by an Equinox 55 spectrometer (Bruker Optics, Germany) with 2 cm−1 resolution.
Scanning electron microscopy (SEM, LEO 1455VP, Cambridge) analysis was done for evaluation of morphology and cells biocompatibility. Before these modifications, the cell-loaded mats were fixed in glutaraldehyde 2.5% (v/v) for 45 min. Dehydration was carried out by immersing the scaffolds in increasing gradients of alcohol concentration.
In Vitro Cell Culture
Stem cell expansion.
Mouse MSCs (mMSCs) were obtained from Stem Cells Technology Research Centre cell bank (Tehran, Iran) and then were expanded in the high-glucose dulbecco’s modified eagle’s medium (DMEM) supplemented with 10% (v/v) fetal bovine serum (FBS) and 1% (v/v) penicillin/streptomycin (all from Gibco, Waltham, MA) as basal medium.
Cell seeding and differentiation.
Before cell seeding, the scaffolds were sterilized using 70% ethanol and after that kept in DMEM for sterility check and enhancing the cell attachment.16 Cell seeding was carried out at 105/cm2 and cultured under basal medium incubated with 95% air and 5% CO2 at 37°C. For osteogenic induction, the differentiation medium (250 nM dexamethasone, 10 mM beta-glycerol phosphate and 50 μg/ml ascorbic acid-2-phosphate, 10% FBS, all from Gibco) was used.
Evaluation of the Osteogenic Differentiation of Mouse Mesenchymal Stem Cells
Real-time reverse transcription polymerase chain reaction.
For monitoring the possible changes in gene expression, mMSCs mRNA extraction was carried out using RNA extraction kit (CinnaGen, Tehran, Iran), and cDNA was synthesized using cDNA kit (Takara, Japan). Changes in gene expression were quantified by an RT PCR analyzer (Corbett, Sydney, Australia). The analysis of results was done by Rotor Gene software. Relative gene expression was calculated by ΔΔCt against hypoxanthine phosphoribosyltransferase 1 (HPRT1). Specificity of the signals was confirmed by considering the melting curve for each gene. The primers are presented in Table 1.
Alkaline Phosphatase Activity and Calcium Content.
Cell lysate was used to measure the total (BCA kit Thermoscientific) and ALP activity (Pars Azmun, Tehran, Iran). Optical density (OD) was measured at 405 nm, and then ALP activity was normalized against total protein.
For measuring the calcium content, cell were lysed by 0.6 N HCl, and based on the applied kit (Pars Azmun, Iran), the OD was measured at 570 nm. The amount of calcium content of unseeded PCL/nHA scaffold was considered as baseline and subtracted from all corresponding groups.
In Vivo Subcutaneous Transplantation and Tissue Staining.
All in vivo experiments were carried out in accordance with the Stem Cell Technology Research Center (Tehran, Iran) ethics. Cell seeded and unseeded scaffolds were grafted subcutaneously in bulb C mouse (Stem Cell Research Center, Tehran, Iran) with 25 to 30 g weight. Three mice were used in the study for each group. Animals were fasted a minimum of 12 hours before being anesthetized. Decisions concerning the choice of surgical instruments, suture material, and intravenous fluids for surgery as well as the need for antibiotics and analgesics administration postoperatively were made upon the institutional animal care and use committee (IACUC) protocol. Mice were fainted using ketamine (20 mg/kg weight of mouse) and xylazine (2 mg/kg weight of mouse) intraperitoneal injection. After hair cutting and skin disinfection, skin of mouse was slivered, and a proper pocket was created for scaffold transplantation. Eight weeks after transplantation, scaffolds were explanted and then fixed in 10% formalin and placed in paraffin. Von Kossa and Masson trichrome staining was performed on 2 µm-sectioned scaffolds.
Statistical analyses were carried out by one-way analysis of variance using Prism software, and the differences were regarded as significant if p value was <0.05. All samples were run in triplicate otherwise is stated. Data are shown as mean ± standard deviation (±SD).
Scaffold Fabrication, Cell Seeding, and Characterization
A schematic diagram of Fn coating on plasma-treated PCL scaffold is shown in Figure 1. Scanning electron microscopy micrographs in Figure 2A show the porosity and pore interconnectivity of the aligned nanofibrous PCL with an average nanofiber diameter of 470 nm. As shown in Figure 2B, nHA were uniformly distributed without aggregation in the matrix. PCL- and nHA-loaded PCL scaffolds showed tensile modulus of 47.40 ± 0.04 and 52.37 ± 0.71 MPa and break strain of 72.95% and 56.92%, respectively. The mechanical properties of the plasma-treated scaffolds did not change significantly. Contact angle of the plasma-treated scaffold (0°) confirmed the hydrophilicity of the surfaces. Based on FTIR results shown in Figure 2C, strong characteristic peaks of PCL were detected at 1723.43 cm−1 for C=O group and at 1046 cm−1 for CO stretching. In addition, the presence of nHA (vibrations in) and Fn (amide I and II bands) was confirmed by specific FTIR peaks at 630, 1661, and 1550 cm−1, respectively.
Scanning electron microscopy micrographs in Figure 3A–D show cells attachment and spreading, which confirm the biocompatibility of the studied scaffolds.
In Vitro Cell Differentiation Analysis
Possible effects of scaffolds modification on osteogenesis were evaluated by monitoring the osteogenic markers. Microscopic analysis of seeded cells after 7 and 21 days of differentiation in Figure 4A–H shows that organic aggregates were deposited upon differentiation, and longer duration of differentiation led to significantly higher calcium deposition. The size and amount of deposited aggregates were the highest for PCL/nHA/Fn scaffolds. Scanning electron micrograph in Figure 5 indicates the porous structure of the scaffold with deposited organic aggregates.
The amount of deposited calcium was measured quantitatively by calcium content method. As shown in Figure 6A, the calcium content of PCL/nHA and PCL/nHA/Fn scaffolds was significantly higher than that of PCL and PCL/Fn after 7 days of cell seeding. After day 21, all stem cells cultured on the scaffolds increased calcium deposition compared with tissue culture polystyrene plates (TCPS), and among scaffolds, PCL/nHA and PCL/nHA/Fn demonstrated significantly highest calcium content. The monitored alkaline phosphatase activity, as shown in Figure 6B, indicates that the difference between ALP activities of PCL/nHA and PCL/nHA/Fn scaffolds at days 7 and 21 compared with the other groups was significant. As ALP activity is an early osteogenic marker, at day 21, ALP activity of all groups decreased in comparison to the day 7.
For further assessment, the effect of scaffold composition on transcription level of the osteogenic markers (ALP, osteocalcin, osteopontin, and Runx2) was monitored at days 7 and 21 of differentiation. As depicted in Figure 7, ALP gene transcription in cells seeded on PCL/nHA/Fn showed the highest expression at day 7 compared with the other groups. At day 21, ALP expression in all groups was downregulated as ALP is an early osteogenic marker, and expression of this gene was significantly higher in PCL/nHA and PCL/nHA/Fn scaffolds than that in PCL, PCL/Fn, and control groups. Osteocalcin gene expression was also significantly higher in PCL/nHA/Fn and PCL/nHA scaffolds, with an increasing trend toward day 21. In addition, gene expression of Runx2 at days 7 and 21 in PCL–nHA and PCL/nHA/Fn was significantly upregulated compared with the other groups. Osteopontin transcription level in PCL/nHA/Fn scaffold at both days was significantly higher than that in other groups.
In Vivo Analysis
The results of Masson trichrome staining on the sectioned explants, presented in Figure 8A–D, indicate that collagen production enhanced when nHA was incorporated in the scaffolds. It is also shown that the presence of Fn exerted further improvement on collagen production. Figure 9A–E shows that calcium deposition was observed only for nHA containing scaffolds, which increased by surface coating of Fn on the scaffolds. To exclude the false-positive effects of calcium content of nHA, unseeded nHA containing scaffolds were taken as control groups. As shown in Figure 9A, no calcium deposition was observed for these control groups. In addition, the size of nHA-containing scaffolds after 2 months of implantation was less than that of PCL and PCL/Fn scaffolds because of nHA degradation.
Because of various diseases or injuries, bone tissue regeneration has been in the focus of intense research.17 Different strategies have been presented for tissue reconstruction,18 and various studies confirmed the beneficial effects of mimicking the microenvironment of the tissues by appropriate scaffolds.19 For enhanced regeneration process, proper scaffolds with promising results for bone-tissue engineering have been used in combination with HA.20 As well as organic materials, application of the bone-specific growth factors further promotes the regeneration process.21,22 The interconnected pores and the nanosized microenvironment provide a suitable supports for cell attachment, proliferation, and differentiation.23 Mesenchymal stem cells due to their high potency in differentiating into different lineages are commonly used for bone-tissue engineering.4 Attachment of these cells to the modified scaffolds confirmed the scaffolds’ biocompatibility.22 Calcium deposition is one of the end-stage osteogenic markers, which used to assess the osteogenic differentiation. Among the mineral materials, HA is the widely used calcium phosphate implant substitute for bone repairs. Hydroxyapatite exerts its beneficial effects on bone formation through improving cell attachment/differentiation, lack of toxicity, and the direct binding to the bone tissues.24 Because of HA solubility, effect of HA-released ions on osteoblast is still unclear. Similar to other studies,25 calcium deposition reached its maximum at day 21. Although nHA itself induces the calcium deposition as confirmed by many studies,25,26 it was found that Fn coating on nHA scaffolds further increased calcium deposition. Alkaline phosphatase is another biochemical factor that is widely used for evaluation of osteogenic differentiation.27 This enzyme initiates the mineralization process28 by hydrolyzing phosphate esters with consequent increase of local phosphate concentration and ultimate enhanced mineralization of ECM.29 The activity of the enzyme increased in all nHA-containing scaffolds. Among nHA containing scaffolds, PCL/nHA/Fn showed significantly higher ALP activity at day 7. Accordingly, ALP gene expression analysis revealed that during the first week of osteogenic differentiation, expression of this gene significantly upregulated in PCL/nHA/Fn group. Osteocalcin is considered as osteoblast-specific marker and one of the most abundant proteins in bone tissues as well. This protein plays an important role in osteoblast progenitor differentiation and induces the matrix synthesis and mineralization.30 In this study, expression of this gene was significantly increased in PCL–nHA and PCL/nHA/Fn. Runx2 is considered as the central control gene for inducing osteoblast phenotype. This gene exerts its major role through regulating a wide spectrum of osteoblast-specific genes.31 The presence of Fn resulted slight increase in Runx2 gene expression. Although osteopontin is not considered as bone-specific marker, it exerts important bone-related functions such as adhesion, migration, and survival.32 As demonstrated, expression of this gene was significantly upregulated in PCL-nHA and PCL/nHA/Fn groups. A weak collagen production monitored by Masson trichrome staining of the sectioned explants on PCL and PCL/Fn scaffolds and increased by incorporation of nHA in the scaffolds. Calcium deposition did not occur in PCL and PCL/Fn groups, but significant calcium deposition was observed for nHA containing scaffolds. PCL/nHA/Fn had almost higher calcium deposition than PCL/nHA. Based on in vitro and in vivo data analysis, the present of Fn and nHA in the scaffold had synergistic effect on the increase of osteogenic induction, conduction, and differentiation.
In conclusion, Fn in combination with nHA exerts its beneficial effects on osteogenesis outcomes in a synergistic manner through further enhancement of calcium deposition, collagen synthesis, early ALP activity, and the up-regulation of the bone-specific genes. This construction is a promising candidate for the development of new bone substitute to be applied as bone-tissue implants.
1. Liu Y, Lim J, Teoh S-H. Review: development of clinically relevant scaffolds
for vascularised bone tissue engineering. Biotechnol Adv 2013.31: 688705.
2. Boyan BD, Hummert TW, Dean DD, Schwartz Z. Role of material surfaces in regulating bone and cartilage cell response. Biomaterials 1996.17: 137146.
3. Li LH, Kommareddy KP, Pilz C, Zhou CR, Fratzl P, Manjubala I. In vitro
bioactivity of bioresorbable porous polymeric scaffolds
microspheres. Acta Biomater 2010.6: 25252531.
4. Sun H, Feng K, Hu J, Soker S, Atala A, Ma PX. Osteogenic
differentiation of human amniotic fluid-derived stem cells induced by bone morphogenetic protein-7 and enhanced by nanofibrous scaffolds
. Biomaterials 2010.31: 11331139.
5. Geiger B, Yamada KM. Molecular architecture and function of matrix adhesions. Cold Spring Harb Perspect Biol 2011.3: a005033.
6. Geiger B, Bershadsky A, Pankov R, Yamada KM. Transmembrane crosstalk between the extracellular matrix–cytoskeleton crosstalk. Nat Rev Mol Cell Biol 2001.2: 793805.
7. Yu H, Wooley PH, Yang SY. Biocompatibility of poly-epsilon-caprolactone-hydroxyapatite
composite on mouse bone marrow-derived osteoblasts and endothelial cells. J Orthop Surg Res 2009.4: 5.
8. Dodel M, Hemmati Nejad N, Bahrami SH, Soleimani M, Hanaee-Ahvaz H. Modifying the mechanical properties of silk nanofiber scaffold by knitted orientation for regenerative medicine applications. Cell Mol Biol (Noisy-le-grand) 2016.62: 1625.
9. Hu J, Zhou Y, Huang L, Liu J, Lu H. Effect of nano-hydroxyapatite
coating on the osteoinductivity of porous biphasic calcium phosphate ceramics. BMC Musculoskelet Disord 2014.15: 114.
10. Ramay HR, Zhang M. Biphasic calcium phosphate nanocomposite porous scaffolds
for load-bearing bone tissue engineering. Biomaterials 2004.25: 51715180.
11. Singh P, Schwarzbauer JE. Fibronectin
and stem cell differentiation - lessons from chondrogenesis. J Cell Sci 2012.125(pt 16): 37033712.
12. Schwartz MA. Integrins and extracellular matrix in mechanotransduction. Cold Spring Harb Perspect Biol 2010.2: a005066.
13. Moursi AM, Damsky CH, Lull J, et al. Fibronectin
regulates calvarial osteoblast differentiation. J Cell Sci 1996.109(pt 6): 13691380.
14. Ribeiro N, Sousa SR, Monteiro FJ. Influence of crystallite size of nanophased hydroxyapatite
and osteonectin adsorption and on MC3T3-E1 osteoblast adhesion and morphology. J Colloid Interface Sci 2010.351: 398406.
15. Pramatarova L, Pecheva E, Presker R, Pham MT, Maitz MF, Stutzmann M. Hydroxyapatite
growth induced by native extracellular matrix deposition on solid surfaces. Eur Cell Mater 2005.9: 912.
16. Ahvaz HH, Mobasheri H, Bakhshandeh B, et al. Mechanical characteristics of electrospun aligned PCL/PLLA nanofibrous scaffolds
conduct cell differentiation in human bladder tissue engineering. J Nanosci Nanotechnol 2013.13: 47364743.
17. Kneser U, Schaefer DJ, Polykandriotis E, Horch RE. Tissue engineering of bone: The reconstructive surgeon’s point of view. J Cell Mol Med 2006.10: 719.
18. Sloff M, de Vries R, Geutjes P, et al. Tissue engineering in animal models for urinary diversion: A systematic review. PLoS One 2014.9: e98734.
19. Wüst S, Müller R, Hofmann S. Controlled positioning of cells in biomaterials-approaches towards 3D tissue printing. J Funct Biomater 2011.2: 119154.
20. Jeong SI, Ko EK, Yum J, Jung CH, Lee YM, Shin H. Nanofibrous poly(lactic acid)/hydroxyapatite
for guided tissue regeneration. Macromol Biosci 2008.8: 32838.
21. Gdalevitch M, Kasaai B, Alam N, Dohin B, Lauzier D, Hamdy RC. The effect of heparan sulfate application on bone formation during distraction osteogenesis. PLoS One 2013.8: e56790.
22. Mohamadyar-Toupkanlou F, Vasheghani-Farahani E, Bakhshandeh B, Soleimani M, Ardeshirylajimi A. In vitro
and in vivo
investigations on fibronectin
coated and hydroxyapatite
. Cell Mol Biol (Noisy-le-grand) 2015.61: 17.
23. Ferreira L, Karp JM, Nobre L, Langer R. New opportunities: the use of nanotechnologies to manipulate and track stem cells. Cell Stem Cell 2008.3: 136146.
24. Jarcho M. Retrospective analysis of hydroxyapatite
development for oral implant applications. Dent Clin North Am 1992.36: 1926.
25. Bakhshandeh B, Soleimani M, Ghaemi N, Shabani I. Effective combination of aligned nanocomposite nanofibers and human unrestricted somatic stem cells for bone tissue engineering. Acta Pharmacol Sin 2011.32: 626636.
26. Xia Y, Zhou P, Cheng X, et al. Selective laser sintering fabrication of nano-hydroxyapatite
for bone tissue engineering applications. Int J Nanomedicine 2013.8: 41974213.
27. Jaiswal N, Haynesworth SE, Caplan AI, Bruder SP. Osteogenic
differentiation of purified, culture-expanded human mesenchymal stem cells in vitro. J Cell Biochem 1997.64: 295312.
28. Malaval L, Liu F, Roche P, Aubin JE. Kinetics of osteoprogenitor proliferation and osteoblast differentiation in vitro. J Cell Biochem 1999.74: 616627.
29. Prabhakaran MP, Venugopal J, Ramakrishna S. Electrospun nanostructured scaffolds
for bone tissue engineering. Acta Biomater 2009.5: 28842893.
30. Ryoo HM, Hoffmann HM, Beumer T, et al. Stage-specific expression of Dlx-5 during osteoblast differentiation: involvement in regulation of osteocalcin gene expression. Mol Endocrinol 1997.11: 16811694.
31. Viereck V, Siggelkow H, Tauber S, Raddatz D, Schutze N, Hüfner M. Differential regulation of Cbfa1/Runx2 and osteocalcin gene expression by vitamin-D3, dexamethasone, and local growth factors in primary human osteoblasts. J Cell Biochem 2002.86: 348356.
32. Standal T, Borset M, Sundan A. Role of osteopontin in adhesion, migration, cell survival and bone remodeling. Exp Oncol 2004.26: 179184.