Materials Science and Engineering C 49 (2015) 746–753
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Calcium phosphate coated Keratin–PCL scaffolds for potential bone tissue regeneration Xinxin Zhao a,1, Yuan Siang Lui a,b,1, Caleb Kai Chuen Choo a, Wan Ting Sow a, Charlotte Liwen Huang a, Kee Woei Ng a, Lay Poh Tan a, Joachim Say Chye Loo a,c,⁎ a b c
Nanyang Technological University, School of Materials Science & Engineering, Division of Materials Technology, N4.1-01-04a, 50 Nanyang Avenue, Singapore 639798, Singapore Institute for Sports Research, Nanyang Technological University, 50 Nanyang Avenue, Singapore 639798, Singapore Singapore Centre on Environmental Life Sciences Engineering (SCELSE), Nanyang Technological University, Singapore 637551, Singapore
a r t i c l e
i n f o
Article history: Received 10 May 2014 Received in revised form 6 January 2015 Accepted 25 January 2015 Available online 27 January 2015 Keywords: Electrospinning Keratin Crosslinking Hydroxyapatite Surface coating
a b s t r a c t The incorporation of hydroxyapatite (HA) nanoparticles within or on the surface of electrospun polymeric scaffolds is a popular approach for bone tissue engineering. However, the fabrication of osteoconductive composite scaffolds via benign processing conditions still remains a major challenge to date. In this work, a new method was developed to achieve a uniform coating of calcium phosphate (CaP) onto electrospun keratin– polycaprolactone composites (Keratin–PCL). Keratin within PCL was crosslinked to decrease its solubility, before coating of CaP. A homogeneous coating was achieved within a short time frame (~10 min) by immersing the scaffolds into Ca2+ and (PO4)3− solutions separately. Results showed that the incorporation of keratin into PCL scaffolds not only provided nucleation sites for Ca2+ adsorption and subsequent homogeneous CaP surface deposition, but also facilitated cell–matrix interactions. An improvement in the mechanical strength of the resultant composite scaffold, as compared to other conventional coating methods, was also observed. This approach of developing a biocompatible bone tissue engineering scaffold would be adopted for further in vitro osteogenic differentiation studies in the future. © 2015 Elsevier B.V. All rights reserved.
1. Introduction The basic building block of natural bone is made up of 69–80 wt.% plate-like hydroxyapatite (HA) Ca10 (PO4)6(OH)2 nanocrystals and 17–20 wt.% collagen nanofibers [1]. In order to mimic the bone microenvironment, HA has been extensively employed as a synthetic alternative to natural bone in applications such as dental [2] and bone tissue engineering (TE) [3]. However, the drawbacks of HA, such as its brittleness and low strength, make it unsuitable for long-term load-bearing applications. As such, extensive research has been conducted to address this issue by incorporating HA into synthetic polymers that possess tougher properties. Among the various scaffold fabricating techniques, electrospinning has emerged as a simple and robust method, capable of generating large quantities of nanofibers from a wide variety of polymers. Its versatility also allows for the fabrication of polymeric-based composites that ⁎ Corresponding author at: Nanyang Technological University, School of Materials Science & Engineering, Division of Materials Technology, N4.1-01-04a, 50 Nanyang Avenue, Singapore 639798, Singapore. E-mail addresses: [email protected] (X. Zhao), [email protected] (Y.S. Lui), [email protected] (C.K.C. Choo), [email protected] (W.T. Sow), [email protected] (C.L. Huang), [email protected] (K.W. Ng), [email protected] (L.P. Tan), [email protected] (J.S.C. Loo). 1 Authors who contributed equally to this work.
http://dx.doi.org/10.1016/j.msec.2015.01.084 0928-4931/© 2015 Elsevier B.V. All rights reserved.
contain inorganic materials, such as HA [4]. However, the direct incorporation of HA into synthetic polymers presents two key problems: (1) a decrease in mechanical properties arising from the inhomogeneous dispersion of HA within the polymer matrix, and (2) a lack of osteoconductivity of the scaffold due to the absence of bone-like HA on the scaffold surface. With these, studies have moved towards achieving uniform HA or calcium phosphate (CaP) coating onto polymeric scaffolds to improve on their osteoconductivity, while minimizing reduction in mechanical strength. Liu et al. reported that the mechanical properties of nanofiber scaffold could be affected by both the thickness and the porosity of the bio-mineral coating [18]. However, establishing a highly scalable, yet benign processing technique still remains to be a major challenge. For instance, conventional CaP coating techniques tend to be harsh and require lengthy processing conditions. To induce Ca2+ chelating sites for subsequent CaP formation, techniques such as O2 plasma treatment [5], laser [6], strong alkali (NaOH or KOH) [7] and acid (HNO3) [8] treatments have been used. These harsh processing techniques can cause defects, such as polymer chain scission or changes to polymer properties that may compromise the mechanical strength of the subsequent scaffolds. Moreover, a long processing time required in some of these techniques may also result in the loss of drugs or biomolecules from scaffolds that are loaded with these agents. For instance, Liu et al. achieved a homogeneous coating of CaP coating onto polymer scaffolds,
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but not after a long period of immersion in simulated body fluid (SBF) [5]. To address the problems above, our approach was therefore to improvise a quick, highly scalable, simple and benign technique that would allow for the uniform coating of CaP onto poly-caprolactone (PCL) scaffolds; without compromising on mechanical properties. PCL is a good candidate for bone tissue engineering because of its biocompatibility, bioresorbability and slow degradation rate [9]. Here, keratin was chosen to be incorporated into the PCL polymer during electrospinning because keratin is autologous and abundant and it could potentially improve mechanical properties due to the presence of free thiol groups in cysteine [10]. Moreover, crosslinked keratin is less water soluble, and has enhanced mechanical strength. Besides, keratin also possesses cell-binding motifs such as Leucine–Aspartic acid– Valine (LDV), which can potentially facilitate cell–matrix interactions [11]. Keratin was also reported recently to be a suitable vehicle for delivering BMP-2 for bone regeneration [12]. It was therefore hypothesized that the presence of crosslinked keratin could aid in improving the mechanical strength of CaP-coated PCL scaffolds, which would otherwise be compromised after coating with CaP. The presence of crosslinked keratin would also introduce strong Ca2+ chelating Schiff base groups onto the fibers and aid in the formation of CaP. These scaffolds were subsequently evaluated for their fiber morphology, mechanical properties, in-vitro apatite forming ability and cell proliferation potential. 2. Materials and methods 2.1. Materials Poly-caprolactone (PCL, Mw = 70,000–90,000 Da), 1,1,1,3,3,3hexafluoro-2-propanol (HFIP), hydroxyapatite nanoparticles (HA: 20–200 nm), calcium nitrite (Ca(NO3)2·4H2O), ammonium phosphate
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(NH4H2PO4) and glutaraldehyde were used as received from SigmaAldrich (St. Louis, MO). Keratin was extracted from the human hair by mixing delipidized hair with 0.125 M of Na2S (Sigma-Aldrich, MO, USA) in Milli-Q deionized H2O (Biocel Ltd., Cork, Ireland). This mixture was incubated for 4 h at 40 °C and subsequently filtered and dialyzed against 5 L of deionized H2O in cellulose tubing (Sigma-Aldrich; molecular weight cutoff = 12,400 Da) for 3 days with 6 changes of H2O. The dialyzed solution was then lyophilized to obtain keratin powder. [13]. 2.2. Scaffold preparation 2.2.1. Electrospinning of PCL, Keratin–PCL and CaP–PCL scaffolds PCL and extracted human hair keratin were blended to a final concentration of 10 wt.%, at a weight ratio of 7:3, in HFIP solvent. Similarly, HA was added into 10 wt.% of PCL in HFIP solution at a ratio of 1:10. Solutions were stirred overnight to ensure complete homogeneous dissolution and were subsequently electrospun. Electrospinning was performed with the Nanon-01A (MECC, Fukuoka, Japan) Electrospinning Setup. The solutions were dispensed from a single nozzle spinneret (22 gauges) at a constant feed rate of 1 mL/h and electrospun at an accelerating voltage of 20 kV. PCL, Keratin–PCL, and CaP–PCL fiber scaffolds were collected on a grounded, flat metallic platform fixed at 14 cm below the tip of the spinneret. 2.2.2. Pre-treatments of PCL and Keratin–PCL scaffolds Oxygen plasma treatment was performed on PCL surface to improve the efficiency and homogeneity of CaP surface coating with a plasma generator consisting of PDC-32G and PDC-FMG Plasmaflo (Harrick Plasma, Ithaca, New York, USA). PCL scaffold was placed in the plasma reaction chamber and exposed to pure oxygen gas under low plasma input power of 10 W for 30 s. For comparison, Keratin–PCL scaffold was
Scheme 1. Experimental design: (A) PCL scaffold was plasma treated followed by a two-step CaP surface coating procedure; (B) Keratin–PCL scaffold was crosslinked by glutaraldehyde and CaP surface coated; (C) CaP aggregated on the surface of scaffolds induced apatite forming.
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Table 1 Descriptions of various scaffolds. Samples
Descriptions
Bonding type between HA and nano-fibers
PCL Keratin–PCL
PCL electrospun scaffolds Keratin electrospun together with PCL Crosslinked PCL–Keratin electrospun scaffold CaP surface coated PCL–cKeratin scaffold CaP surface coated PCL scaffolds HA electrospun together with PCL
N.A. N.A.
cKeratin–PCL sCaP–cKeratin–PCL sCaP–PCL HA–PCL
and cKeratin–PCL were immersed into Ca(NO3)2·4H2O solution for 10 min and subsequently transferred into NH4H2PO4 solution for another 3 min. Coated samples (sCaP–PCL and sCaP–cKeratin–PCL respectively) were gently washed with deionized water (DI water) to remove non-chelated CaP. Scaffolds were then dried in a vacuum oven maintained at room temperature.
N.A.
2.3. Scaffold characterization Strong chemical bonds Strong chemical bonds Weak physical interaction
crosslinked to produce cKeratin–PCL through overnight exposure to glutaraldehyde vapor. 2.2.3. Surface coating with CaP CaP surface coating was achieved by a two-step approach; induction of Ca2 + chelating sites on the electrospun scaffolds followed by a coprecipitation reaction. Aqueous solutions of 0.4 M Ca(NO3)2·4H2O and 0.24 M NH4H2PO4 were prepared and pH of the solutions was adjusted to pH 8 and 11 respectively. Immediately after surface treatment, PCL
2.3.1. Morphology assessment The morphologies of prepared scaffolds were analyzed with Field Emission Scanning Electron Microscopy (FESEM) (JSM-6340F, JEOL Co., Tokyo, Japan). The samples were coated with platinum at 20 mA for 60 s and imaged at an accelerating voltage of 5 kV. Mean fiber diameters and average pore diameter were determined by measuring independent fibers using Image J software (n = 50). 2.3.2. Chemical composition The CaP surface coatings were further confirmed by Fourier Transform Infrared Spectroscopy (FT-IR), using the reflection technique (ATR-FTIR, Perkin Elmer 2000, Swabian, Burladingen, German). All spectra were measured at a resolution of 16 cm−1.
Fig. 1. SEM micrographs of various electrospun scaffolds. PCL scaffolds (A) have larger fiber diameter than PCL–Keratin scaffolds (D). After crosslinking with glutaraldehyde, PCL–cKeratin matrices (E) exhibited significantly a larger fiber diameter. Both CaP surface coated scaffolds, sCaP–PCL (C) and sCaP–PCL–cKeratin (F) have no significant difference compared to respective scaffolds without CaP. Agglomeration of HA can be observed on HA loaded PCL scaffold, HA–PCL (B, inset), which exhibited a similar fiber diameter with PCL scaffolds.
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2.3.3. Mechanical properties The mechanical properties of the scaffolds were evaluated at room temperature using the Instron Tensile Machine 5567 (Instron, Canton, Massachusetts, USA). ASTM dog bone shaped specimens of dimension 9.53 mm × 3.18 mm were prepared and tested using a load cell of 500 N at a tensile loading rate of 10 mm·min−1. The mechanical properties of the scaffolds were calculated (BlueHill version 2.21) for their respective tensile strength (TS), strain at break (ε) and Young's modulus (E) in triplicates (n = 3). 2.3.4. In-vitro bioactivity of electrospun scaffolds In-vitro bioactivity was examined by immersing the scaffolds in 5 mL of stimulated body fluid (SBF) [14] for 7 days. After which, samples were washed thrice with DI water and vacuum dried for 24 h before taking gravimetric measurements. The initial and final weights of the scaffolds were compared to reflect the formation of apatite. The morphology of these scaffolds was characterized with FESEM (JSM6340F, JEOL Co., Tokyo, Japan). 2.4. Cell culture Human mesenchymal stem cells (hMSCs, Lonza) were used to assess the biocompatibility of the prepared scaffolds. The hMSCs were maintained in 75 cm2 cell culture flasks and passaged whenever confluency
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of 70–80% was attained. hMSCs from passage 5 were used in our study. PCL, sHA–PCL, cKeratin–PCL, and sHA–cKeratin–PCL scaffolds were cut into circular discs of 15.6 cm diameter and fitted into 24 well cell culture plates. The scaffolds were UV-sterilized for 30 min and rinsed three times with phosphate buffer saline (PBS, OHME Scientific) followed by 3 h incubation in complete culture medium consisting of Dulbecco's Modified Eagle's Media (DMEM), high glucose medium (PAA Laboratory), 10% Fetal Bovine Serum (PAA Laboratory), 1% L-Glutamine (200 mM, PAA Laboratory) and 1% Penicillin–Streptomycin (PAA) prior to cell seeding. Cells of density 5000 cells per cm2 were seeded on the scaffolds and cultured in a humidified atmosphere at 37 °C with 5% CO2. Cell culture medium was refreshed every 3 days.
2.5. Cell proliferation assay The proliferation of hMSCs that have attached and grown on the scaffolds was determined by measuring the amount of double stranded DNA (dsDNA) with the PicoGreen assay. Dead cells were removed by complete removal of the cell culture medium and thrice washing with PBS. Remaining adherent cells were lysed through gentle agitation in Triton X-100 solution (0.1% v/v). Subsequently, 100 μL of PicoGreen working solution was added and mixed with 100 μL aliquots of the cell lysate. Fluorescent intensity was measured from black 96-well flat plate using an Infinite200 microplate reader (Tecan Inc., Männedorf,
Fig. 2. ATR-FTIR analyses of various scaffolds: (A) PCL and SCaP–PCL and (B) PCL, Keratin, cKeratin–PCL and sCaP–cKeratin–PCL.
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Fig. 3. SEM micrographs of (A) HA–PCL; (B) sCaP–PCL; (C) sCaP–cKeratin–PCL scaffolds after 7 days in vitro apatite forming study. No significant apatite formation was observed on PCL (A) scaffolds. However, CaP-surface coated scaffolds (B: sCaP–PCL; C: sCaP–cKeratin–PCL) showed superior apatite forming ability, by forming a homogeneous layer of apatite after 7 days of SBF immersion.
Zurich, Switzerland) with an excitation/emission wavelength of 480/ 520 nm. 2.6. Statistical analysis All data are expressed as means ± standard deviation. Statistical differences are determined with one way ANOVA followed by Tukey's HSD post hoc test on SPSS version 11.5 and differences are considered statistically significant at p ≤ 0.05. 3. Results and discussion Bone tissue repair is an area of extensive research, whereby the scaffold material and a combination of other properties, such as adequate structural support and a bio-mimicking cell micro-environment, play
pivotal roles in creating a bio-functional bone-regenerating implant. To achieve such a scaffold, a new method was devised to produce CaPcoated, keratin-incorporated, electrospun PCL scaffolds (Scheme 1). In this paper, the cell differentiating properties of this scaffold would not be investigated. To evaluate and make comparisons with this scaffold, all other formulations and their respective controls (see Table 1) were electrospun under similar conditions. All fibers were found to possess uniform and “beadless” morphologies, as shown in Fig. 1. Here, a “beadless” morphology was achieved through the use of HFIP as solvent. Upon the incorporation of keratin into PCL (Keratin–PCL scaffold), the diameter of the fibers decreased from 552 ± 66 nm (Fig. 1A) to 196 ± 51 nm (Fig. 1D). The reduction in fiber diameter could be attributed to the increased electroconductivity of the electrospinning solution with the addition of keratin. The amino acid groups of keratin conferred negative
Fig. 4. Weight differences of various scaffolds after 7 days of SBF immersion. CaP surface-coated scaffolds, sCaP–PCL and sCaP–cKeratin–PCL have shown 17.36% and 16.02% of weight increment respectively after 7 days of immersion. No significant weight incensement was observed for HA imbedded HA–PCL scaffold.
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Fig. 5. Stress–strain curves of various scaffolds. The Young's modulus of sCaP–cKeratin–PCL (25.92 MPa) N cKeratin–PCL (15.38 MPa) N PCL (9.61 MPa) N sCaP–PCL (7.55 MPa) N HA–PCL (6.86 MPa).
charges to the protein and hence increased the solution's electroconductivity [15]. The formation of such nano-sized fibers was nonetheless beneficial at such size range, as these nano-fibers were reported to improve cell adhesion and growth kinetics [16]. In place of keratin, HA was also electrospun together with PCL polymer, as another control (HA–PCL). Not surprisingly, the HA–PCL scaffold (Fig. 1B) only showed minute amounts of HA on the surface of the fibers, even at high HA concentrations (10 wt.%), implying that most of this bone-like mineral was embedded within the fibers. PCL scaffolds were subsequently coated with CaP; i.e. pre-treated PCL and Keratin–PCL scaffolds (sCaP–PCL & sCaP–cKeratin–PCL — ‘s’ denotes ‘surface’), as shown in Scheme 1. Pre-treatment was first conducted for both sample scaffolds, and also for the non-keratin PCL scaffolds for comparison (Scheme 1A). Pre-treatment was essential to enhance the coating efficiency of CaP through the introduction of electrondonating groups for the adsorption of Ca2+ to initiate the nucleation of apatite. The most common pre-treatment methods used for biodegradable polyesters were either oxygen plasma or alkaline and acid treatment for polyester scaffolds. [17] Oxygen plasma treatment introduces oxygen free radicals, while alkaline and acid treatments increase the number of carboxyl and hydroxyl end groups. These functional groups aid in chelating with Ca2 + ions and initiate the formation of CaP as a coating. However, both techniques can be detrimental to the polymer, and may cause chain scission in PCL. To minimize polymer degradation of PCL, a mild oxygen plasma condition, such as low power (10 W) and short exposure time (30 s), was used. This mild process did not result in any observable morphology changes to the PCL fibers (data not shown). For the keratin-based PCL scaffolds, keratin was first crosslinked to introduce multidentate electron-donating groups on the fiber surface (Scheme 1B). Keratin–PCL electrospun scaffolds was pre-treated via a condensation reaction with glutaraldehyde to create the multidentate Schiff base ligands, which are well-known metal ion chelating ligands [18,19]. The increase in fiber diameter from 196 ± 51 nm (Fig. 1D) to 253 ± 54 nm after crosslinking (Fig. 1E) could be
due to the absorption of water vapor during the crosslinking process [20]. Pre-treated scaffolds were subsequently coated with CaP through sequential immersion of the scaffolds (i.e. plasma-treated PCL scaffold and crosslinked Keratin–PCL scaffold), in Ca(NO3)2 and NH4H2PO4 solutions. A thick and homogeneous layer of CaP (Fig. 1C and F) was observed within 10 min. It is worth mentioning that this is the first report in which CaP surface coating was achieved through the chelation reaction of crosslinked keratin and Ca2+ on the surface of electrospun polymeric fibers, and within such a short period. This process largely improved the homogeneity and efficiency of CaP coating by shortening the coating period from days [5] to just minutes. The coating of CaP on the scaffolds was qualified by ATR-FTIR spectroscopy. From Fig. 2A, comparing against PCL, the spectrum of sCaP– PCL scaffolds displayed the characteristic triplet absorption bands of the (PO4)3− groups at 1000–1100 cm−1. The presence of an absorption band at 1579 cm−1 for the spectrum of sCaP–PCL scaffold is contributed from the red shift of Ca2+ chelated ester groups (–C=O: 1722 cm−1). The strong multi-band absorptions from 1200–1400 cm−1 may be attributed to Ca2+ chelated C–O–C groups that were introduced by oxygen plasma treatment. In Fig. 2B, compared with the spectrum of PCL, the extra absorbances at 1643.9 and 1542.8 cm−1 were typical amide 1 and amide II absorptions, which were attributed by the presence of keratin. The red shift of amide II from 1520.5 cm− 1 of keratin into 1542.8 cm−1 further proved the crosslinking reaction between amide and aldehyde groups. Similar to sCaP–PCL, typical CaP absorption bands were observed at 1000–1100 cm−1 for sCaP–cKeratin–PCL scaffolds, representing the typical (PO4)3− adsorption of surface CaP. Moreover, the amide II absorbance was found to be blue-shifted from 1542.8 cm−1 to 1591.3 cm−1, suggesting the chelation of Ca2 + ions with the amide groups of keratin. The coating of the scaffold surface with CaP is aided through the interaction of the exposed hydroxyl groups from the CaP [Ca10 (PO4)6 (OH)2] with the Ca2 + ions present in bodily fluid [21]. The influence
Table 2 Comparison of mechanical properties of various scaffolds in terms of tensile strength (TS), strain at break (ϵf) and Young's modulus (E). Properties
PCL
HA–PCL
sCaP–PCL
cKeratin–PCL
sCaP–cKeratin–PCL
Tensile strength (MPa) Strain at break (%) Young's modulus (MPa) Pore size (μm)
8.00 ± 0.43 279.06 ± 15.38 9.61 ± 0.56 2.64 ± 0.21
3.33 ± 0.41 164.59 ± 31.58 6.86 ± 0.41 2.35 ± 0.51
5.54 ± 1.17 251.58 ± 69.93 7.55 ± 0.70 3.12 ± 0.21
13.41 ± 0.94 226.44 ± 29.44 15.38 ± 0.57 2.11 ± 0.30
16.53 ± 1.16 152.78 ± 19.86 25.92 ± 0.96 2.66 ± 0.41
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Fig. 6. Cell viability study of various scaffolds was conducted through Picogreen Assays after 4 days of culturing. Keratin containing scaffolds, cKeratin–PCL & sCaP–cKeratin– PCL, showed better cell viability than controls (PCL & sCaP–PCL).
of CaP-containing scaffolds was subsequently evaluated, i.e. electrospun HA–PCL; sCaP–PCL and sCaP–cKeratin–PCL. Scaffolds were determined for their rate of mineralization of the scaffold in SBF medium [22], after 7 days of immersion (Fig. 3). From SEM images, sCaP–PCL (Fig. 1C) and sCaP–cKeratin–PCL (Fig. 1F) scaffolds were covered with a thick layer of calcium phosphate (CaP) after 7 days, as represented by Fig. 3B and C respectively. To quantify the amount of CaP formed, weight measurements were conducted on these scaffolds. The results from the analysis (Fig. 4) indicated ~ 16 wt.% of CaP on both sCaP–PCL and sCaP–cKeratin–PCL scaffolds, with the lowest wt.% of CaP (~2%) observed for HA–PCL scaffold. This implied that apatite formation was not significant in HA–PCL scaffold due to a lack of surface CaP. These findings therefore highlighted the importance of having a significant concentration of surface CaP on the fibers.
The mechanical properties of these scaffolds were subsequently evaluated through tensile mechanical testing as shown in Fig. 5 and summarized in Table 2. From the stress–strain plots shown in Fig. 5, PCL scaffolds had a tensile strength (TS) of 8.0 ± 0.4 MPa and strain at break (ε) of 279%. The presence of CaP, regardless of surface coating (sCaP–PCL) or incorporated during electrospinning (HA–PCL), was accompanied with a significant decrease in both TS and ε values. Comparatively, despite a slight decrease of ε for cKeratin–PCL scaffold, TS value increased significantly to 13.4 ± 0.9 MPa compared to pristine PCL scaffold. The highest TS value was observed on the sCaP–cKeratin–PCL scaffold (16.5 ± 1.2 MPa). The Young's modulus of the scaffolds as ranked in descending order was as follows: sCaP–cKeratin–PCL N cKeratin– PCL N PCL N sCaP–PCL N HA–PCL. The poor mechanical properties of HA–PCL could be due to the possible agglomeration of HA nanoparticles in the polymer matrix, causing micro-defects to the fiber scaffold. In order to overcome this issue, Prabhakaran et al. reported the incorporation of collagen into PLLA scaffold to improve HA dispersion [23]. Nonetheless, TS was substantially reduced from 4.7 to 2.1 MPa upon the incorporation of both HA and collagen. In our study, cKeratin–PCL scaffold was surface-coated with CaP through a chemical Ca2+ chelation. The increase in TS in cKeratin–PCL scaffold could be due to the presence of strong and rigid crosslinkedkeratin structure [24]. With surface CaP, the strong chemical bonds formed between CaP nanoparticles and the cKeratin–PCL scaffold [25] greatly enhanced the TS of sCaP–cKeratin–PCL scaffold, with tensile strength that meets the mechanical requirements for bone TE applications [26]. The biocompatibility of the scaffolds was finally evaluated by using PicoGreen assay after 4 days of culturing of hMSCs on these scaffolds. From Fig. 6, the cell viability of hMSCs on keratin containing scaffolds: cKeratin–PCL and sCaP–cKeratin–PCL, was significantly higher than those on PCL and sHA–PCL (p ≤ 0.05). This may be attributed to the presence of LDV cell binding motifs on keratin in promoting integrin α4β1 mediated cell adhesion [27,28], and consequently improving cell
Fig. 7. SEM micrographs showing that cell adopted spindle-shaped morphologies on PCL scaffolds after 4 days of culturing (A), while wide-spread morphologies were observed on sCaP– cKeratin–PCL scaffolds (C) after 7 days of culturing. Cell on sCaP–cKeratin–PCL scaffolds was well integrated with Keratin-contained fibers on 4 days (B), 7 days (C) and 14 days (D) of culturing. Mineralization was clearly observed especially on 14 days of culturing.
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viability. The presence of HA has been shown to enhance cell attachment at the early stages of cell culture and improved cell proliferation [29]. However, in our study, cell viability data did not show any significant differences between scaffolds containing HA and their respective controls (without CaP coatings). This is probably because CaP and cell interactions are influenced by surface roughness and chemical composition of scaffold, as well as the crystal and grain size of the CaP nanoparticles. For example, Wang et al. [30] and Ribeiro et al. [31] reported that micro-sized HA could improve the performance of SaOS-2 and MC3T3E1 cells respectively but Smith et al. [32] found that MC3T3-EI cells proliferated faster on nano-sized HA scaffolds. Our findings nevertheless indicated higher cell proliferation on sCaP–cKeratin–PCL scaffold as compared to that on sCaP–PCL scaffold. This could be due to better initial cell attachment on keratin containing scaffolds rather than the effect of HA nanoparticles or the lateral apatite forming rate. The SEM images of cell-scaffold constructs are shown in Fig. 7. hMSCs exhibited an elongated morphology on PCL scaffold after 4 days of culture (Fig. 7A). In contrast, well-spread hMSCs with polygonal morphology that integrated well with the nanofibers of the sCaP– cKeratin–PCL scaffold were observed on day 4 (Fig. 7B), day 7 (Fig. 7C) and day 14 of culture (Fig. 7D). In summary, scaffolds containing crosslinked keratin such as cKeratin–PCL and sCaP–cKeratin–PCL are suitable to support cell attachment and proliferation. The osteogenic differentiation of hMSCs cells on sCaP–cKeratin–PCL would be one aspect of this work that would be studied in the future. 4. Conclusions A two-step approach to coat CaP onto the surface of electrospun PCL scaffolds was proposed. Keratin was incorporated into the electrospun PCL scaffold and used to induce the formation of a HA surface coating after a crosslinking reaction. The crosslinking of keratin was able to promote Ca2 + chelation and a homogeneous CaP coating onto PCL scaffolds. Our study suggested that the mineralization rate was related to the amount of surface available CaP. In addition, the presence of crosslinked keratin and CaP further increased the mechanical strength of the resultant scaffolds. Moreover, higher proliferation rate of hMSCs cells was observed on keratin containing scaffolds. In conclusion, the optimized scaffold, sCaP–cKeratin–PCL had improved mechanical properties and good biocompatibility. Therefore, sCaP–cKeratin–PCL scaffold can serve as a potential biocomposite material for bone tissue engineering. Acknowledgments The authors would also like to acknowledge the financial support from the Agency for Science, Technology and Research (A*STAR) (Project No: 102 129 0098), the National Medical Research Council, Ministry of Health (NMRC/CIRG/1342/2012, MOH), the NTU-National Healthcare Group (NTU-NHG) grant (ARG/14012), the Singapore Centre on Environmental Life Sciences Engineering (SCELSE) (M433C70000.M4220001.C70), and the School of Materials Science and Engineering (M020070110), Nanyang Technological University, Singapore. Appendix A. Supplementary data Supplementary data to this article can be found online at http://dx. doi.org/10.1016/j.msec.2015.01.084. References [1] M.J. Glimcher, Molecular biology of mineralized tissues with particular reference to bone, Rev. Mod. Phys. 31 (1959) 359.
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Wei, Influence of pretreatment on the surface characteristics of PLLA fibers and subsequent hydroxyapatite coating, J. Biomed. Mater. Res. B Appl. Biomater. 88 (2009) 220–229. [8] W. Cui, X. Li, S. Zhou, J. Weng, In situ growth of hydroxyapatite within electrospun poly (DL‐lactide) fibers, J. Biomed. Mater. Res. A 82 (2007) 831–841. [9] L.S. Nair, C.T. Laurencin, Biodegradable polymers as biomaterials, Prog. Polym. Sci. 32 (2007) 762–798. [10] Y. Hirao, K. Ohkawa, H. Yamamoto, T. Fujii, A novel human hair protein fiber prepared by watery hybridization spinning, Macromol. Mater. Eng. 290 (2005) 165–171. [11] V. Verma, P. Verma, P. Ray, A.R. Ray, Preparation of scaffolds from human hair proteins for tissue-engineering applications, Biomed. Mater. 3 (2008) 025007. [12] R.C. de Guzman, J.M. Saul, M.D. Ellenburg, M.R. Merrill, H.B. Coan, T.L. Smith, et al., Bone regeneration with BMP-2 delivered from keratose scaffolds, Biomaterials 34 (2013) 1644–1656. [13] W.T. Sow, Y.S. Lui, K.W. Ng, Electrospun human keratin matrices as templates for tissue regeneration, Nanomedicine 8 (2013) 531–541. [14] C. Hengky, B. Kelsen, P. Cheang, Mechanical and biological characterization of pressureless sintered hydroxapatite–polyetheretherketone biocomposite, 13th International Conference on Biomedical Engineering, Springer, 2009, pp. 261–264. [15] A. Aluigi, A. Varesano, A. Montarsolo, C. Vineis, F. Ferrero, G. Mazzuchetti, et al., Electrospinning of keratin/poly(ethylene oxide) blend nanofibers, J. Appl. Polym. Sci. 104 (2007) 863–870. [16] M. Chen, P.K. Patra, S.B. Warner, S. Bhowmick, Role of fiber diameter in adhesion and proliferation of NIH 3T3 fibroblast on electrospun polycaprolactone scaffolds, Tissue Eng. 13 (2007) 579–587. [17] Y. Yang, K.-H. Kim, J.L. Ong, A review on calcium phosphate coatings produced using a sputtering process—an alternative to plasma spraying, Biomaterials 26 (2005) 327–337. [18] S. Sajadi, Metal ion-binding properties of L-glutamic acid and L-aspartic acid, a comparative investigation, Nat. Sci. 2 (2010) 85–90. [19] I. Migneault, C. Dartiguenave, M.J. Bertrand, K.C. Waldron, Glutaraldehyde: behavior in aqueous solution, reaction with proteins, and application to enzyme crosslinking, Biotechniques 37 (2004) 790–806. [20] J. Yuan, Z.C. Xing, S.W. Park, J. Geng, I.K. Kang, J. Shen, et al., Fabrication of PHBV/keratin composite nanofibrous mats for biomedical applications, Macromol. Res. 17 (2009) 850–855. [21] L.C. Bell, A.M. Posner, J.P. Quirk, Surface charge characteristics of hydroxyapatite and fluorapatite, Nature 239 (1972) 515–517. [22] T. Kokubo, H. Takadama, How useful is SBF in predicting in vivo bone bioactivity? Biomaterials 27 (2006) 2907–2915. [23] M.P. Prabhakaran, J. Venugopal, S. Ramakrishna, Electrospun nanostructured scaffolds for bone tissue engineering, Acta Biomater. 5 (2009) 2884–2893. [24] al. Te. Method of making and cross-linking keratin-based films and sheets. In: Patent US, editor. United States: Keraplast Technologies, Ltd.; 1999. [25] L. Li, G. Li, J. Jiang, X. Liu, L. Luo, K. Nan, Electrospun fibrous scaffold of hydroxyapatite/poly (ε-caprolactone) for bone regeneration, J. Mater. Sci. Mater. Med. 23 (2012) 547–554. [26] C.T. Rubin, Skeletal strain and the functional significance of bone architecture, Calcif. Tissue Int. 36 (1984) S11–S18. [27] J.-L. Guan, R.O. Hynes, Lymphoid cells recognize an alternatively spliced segment of fibronectin via the integrin receptor α4β1, Cell 60 (1990) 53–61. [28] M.S. Goligorsky, Regenerative Nephrology: Access Online via Elsevier, 2010. [29] J. Li, Y. Li, X. Liu, J. Zhang, Y. Zhang, Strategy to introduce nano hydroxyapatite/keratin into fibrous membrane for bone tissue engineering, J. Mater. Chem. B 1 (2013) 432–437. [30] C. Wang, Y. Duan, B. Markovic, J. Barbara, C. Rolfe Howlett, X. 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Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Calcium phosphate coated Keratin–PCL scaffolds for potential bone tissue regeneration Xinxin Zhao a,1, Yuan Siang Lui a,b,1, Caleb Kai Chuen Choo a, Wan Ting Sow a, Charlotte Liwen Huang a, Kee Woei Ng a, Lay Poh Tan a, Joachim Say Chye Loo a,c,⁎ a b c
Nanyang Technological University, School of Materials Science & Engineering, Division of Materials Technology, N4.1-01-04a, 50 Nanyang Avenue, Singapore 639798, Singapore Institute for Sports Research, Nanyang Technological University, 50 Nanyang Avenue, Singapore 639798, Singapore Singapore Centre on Environmental Life Sciences Engineering (SCELSE), Nanyang Technological University, Singapore 637551, Singapore
a r t i c l e
i n f o
Article history: Received 10 May 2014 Received in revised form 6 January 2015 Accepted 25 January 2015 Available online 27 January 2015 Keywords: Electrospinning Keratin Crosslinking Hydroxyapatite Surface coating
a b s t r a c t The incorporation of hydroxyapatite (HA) nanoparticles within or on the surface of electrospun polymeric scaffolds is a popular approach for bone tissue engineering. However, the fabrication of osteoconductive composite scaffolds via benign processing conditions still remains a major challenge to date. In this work, a new method was developed to achieve a uniform coating of calcium phosphate (CaP) onto electrospun keratin– polycaprolactone composites (Keratin–PCL). Keratin within PCL was crosslinked to decrease its solubility, before coating of CaP. A homogeneous coating was achieved within a short time frame (~10 min) by immersing the scaffolds into Ca2+ and (PO4)3− solutions separately. Results showed that the incorporation of keratin into PCL scaffolds not only provided nucleation sites for Ca2+ adsorption and subsequent homogeneous CaP surface deposition, but also facilitated cell–matrix interactions. An improvement in the mechanical strength of the resultant composite scaffold, as compared to other conventional coating methods, was also observed. This approach of developing a biocompatible bone tissue engineering scaffold would be adopted for further in vitro osteogenic differentiation studies in the future. © 2015 Elsevier B.V. All rights reserved.
1. Introduction The basic building block of natural bone is made up of 69–80 wt.% plate-like hydroxyapatite (HA) Ca10 (PO4)6(OH)2 nanocrystals and 17–20 wt.% collagen nanofibers [1]. In order to mimic the bone microenvironment, HA has been extensively employed as a synthetic alternative to natural bone in applications such as dental [2] and bone tissue engineering (TE) [3]. However, the drawbacks of HA, such as its brittleness and low strength, make it unsuitable for long-term load-bearing applications. As such, extensive research has been conducted to address this issue by incorporating HA into synthetic polymers that possess tougher properties. Among the various scaffold fabricating techniques, electrospinning has emerged as a simple and robust method, capable of generating large quantities of nanofibers from a wide variety of polymers. Its versatility also allows for the fabrication of polymeric-based composites that ⁎ Corresponding author at: Nanyang Technological University, School of Materials Science & Engineering, Division of Materials Technology, N4.1-01-04a, 50 Nanyang Avenue, Singapore 639798, Singapore. E-mail addresses: [email protected] (X. Zhao), [email protected] (Y.S. Lui), [email protected] (C.K.C. Choo), [email protected] (W.T. Sow), [email protected] (C.L. Huang), [email protected] (K.W. Ng), [email protected] (L.P. Tan), [email protected] (J.S.C. Loo). 1 Authors who contributed equally to this work.
http://dx.doi.org/10.1016/j.msec.2015.01.084 0928-4931/© 2015 Elsevier B.V. All rights reserved.
contain inorganic materials, such as HA [4]. However, the direct incorporation of HA into synthetic polymers presents two key problems: (1) a decrease in mechanical properties arising from the inhomogeneous dispersion of HA within the polymer matrix, and (2) a lack of osteoconductivity of the scaffold due to the absence of bone-like HA on the scaffold surface. With these, studies have moved towards achieving uniform HA or calcium phosphate (CaP) coating onto polymeric scaffolds to improve on their osteoconductivity, while minimizing reduction in mechanical strength. Liu et al. reported that the mechanical properties of nanofiber scaffold could be affected by both the thickness and the porosity of the bio-mineral coating [18]. However, establishing a highly scalable, yet benign processing technique still remains to be a major challenge. For instance, conventional CaP coating techniques tend to be harsh and require lengthy processing conditions. To induce Ca2+ chelating sites for subsequent CaP formation, techniques such as O2 plasma treatment [5], laser [6], strong alkali (NaOH or KOH) [7] and acid (HNO3) [8] treatments have been used. These harsh processing techniques can cause defects, such as polymer chain scission or changes to polymer properties that may compromise the mechanical strength of the subsequent scaffolds. Moreover, a long processing time required in some of these techniques may also result in the loss of drugs or biomolecules from scaffolds that are loaded with these agents. For instance, Liu et al. achieved a homogeneous coating of CaP coating onto polymer scaffolds,
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but not after a long period of immersion in simulated body fluid (SBF) [5]. To address the problems above, our approach was therefore to improvise a quick, highly scalable, simple and benign technique that would allow for the uniform coating of CaP onto poly-caprolactone (PCL) scaffolds; without compromising on mechanical properties. PCL is a good candidate for bone tissue engineering because of its biocompatibility, bioresorbability and slow degradation rate [9]. Here, keratin was chosen to be incorporated into the PCL polymer during electrospinning because keratin is autologous and abundant and it could potentially improve mechanical properties due to the presence of free thiol groups in cysteine [10]. Moreover, crosslinked keratin is less water soluble, and has enhanced mechanical strength. Besides, keratin also possesses cell-binding motifs such as Leucine–Aspartic acid– Valine (LDV), which can potentially facilitate cell–matrix interactions [11]. Keratin was also reported recently to be a suitable vehicle for delivering BMP-2 for bone regeneration [12]. It was therefore hypothesized that the presence of crosslinked keratin could aid in improving the mechanical strength of CaP-coated PCL scaffolds, which would otherwise be compromised after coating with CaP. The presence of crosslinked keratin would also introduce strong Ca2+ chelating Schiff base groups onto the fibers and aid in the formation of CaP. These scaffolds were subsequently evaluated for their fiber morphology, mechanical properties, in-vitro apatite forming ability and cell proliferation potential. 2. Materials and methods 2.1. Materials Poly-caprolactone (PCL, Mw = 70,000–90,000 Da), 1,1,1,3,3,3hexafluoro-2-propanol (HFIP), hydroxyapatite nanoparticles (HA: 20–200 nm), calcium nitrite (Ca(NO3)2·4H2O), ammonium phosphate
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(NH4H2PO4) and glutaraldehyde were used as received from SigmaAldrich (St. Louis, MO). Keratin was extracted from the human hair by mixing delipidized hair with 0.125 M of Na2S (Sigma-Aldrich, MO, USA) in Milli-Q deionized H2O (Biocel Ltd., Cork, Ireland). This mixture was incubated for 4 h at 40 °C and subsequently filtered and dialyzed against 5 L of deionized H2O in cellulose tubing (Sigma-Aldrich; molecular weight cutoff = 12,400 Da) for 3 days with 6 changes of H2O. The dialyzed solution was then lyophilized to obtain keratin powder. [13]. 2.2. Scaffold preparation 2.2.1. Electrospinning of PCL, Keratin–PCL and CaP–PCL scaffolds PCL and extracted human hair keratin were blended to a final concentration of 10 wt.%, at a weight ratio of 7:3, in HFIP solvent. Similarly, HA was added into 10 wt.% of PCL in HFIP solution at a ratio of 1:10. Solutions were stirred overnight to ensure complete homogeneous dissolution and were subsequently electrospun. Electrospinning was performed with the Nanon-01A (MECC, Fukuoka, Japan) Electrospinning Setup. The solutions were dispensed from a single nozzle spinneret (22 gauges) at a constant feed rate of 1 mL/h and electrospun at an accelerating voltage of 20 kV. PCL, Keratin–PCL, and CaP–PCL fiber scaffolds were collected on a grounded, flat metallic platform fixed at 14 cm below the tip of the spinneret. 2.2.2. Pre-treatments of PCL and Keratin–PCL scaffolds Oxygen plasma treatment was performed on PCL surface to improve the efficiency and homogeneity of CaP surface coating with a plasma generator consisting of PDC-32G and PDC-FMG Plasmaflo (Harrick Plasma, Ithaca, New York, USA). PCL scaffold was placed in the plasma reaction chamber and exposed to pure oxygen gas under low plasma input power of 10 W for 30 s. For comparison, Keratin–PCL scaffold was
Scheme 1. Experimental design: (A) PCL scaffold was plasma treated followed by a two-step CaP surface coating procedure; (B) Keratin–PCL scaffold was crosslinked by glutaraldehyde and CaP surface coated; (C) CaP aggregated on the surface of scaffolds induced apatite forming.
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Table 1 Descriptions of various scaffolds. Samples
Descriptions
Bonding type between HA and nano-fibers
PCL Keratin–PCL
PCL electrospun scaffolds Keratin electrospun together with PCL Crosslinked PCL–Keratin electrospun scaffold CaP surface coated PCL–cKeratin scaffold CaP surface coated PCL scaffolds HA electrospun together with PCL
N.A. N.A.
cKeratin–PCL sCaP–cKeratin–PCL sCaP–PCL HA–PCL
and cKeratin–PCL were immersed into Ca(NO3)2·4H2O solution for 10 min and subsequently transferred into NH4H2PO4 solution for another 3 min. Coated samples (sCaP–PCL and sCaP–cKeratin–PCL respectively) were gently washed with deionized water (DI water) to remove non-chelated CaP. Scaffolds were then dried in a vacuum oven maintained at room temperature.
N.A.
2.3. Scaffold characterization Strong chemical bonds Strong chemical bonds Weak physical interaction
crosslinked to produce cKeratin–PCL through overnight exposure to glutaraldehyde vapor. 2.2.3. Surface coating with CaP CaP surface coating was achieved by a two-step approach; induction of Ca2 + chelating sites on the electrospun scaffolds followed by a coprecipitation reaction. Aqueous solutions of 0.4 M Ca(NO3)2·4H2O and 0.24 M NH4H2PO4 were prepared and pH of the solutions was adjusted to pH 8 and 11 respectively. Immediately after surface treatment, PCL
2.3.1. Morphology assessment The morphologies of prepared scaffolds were analyzed with Field Emission Scanning Electron Microscopy (FESEM) (JSM-6340F, JEOL Co., Tokyo, Japan). The samples were coated with platinum at 20 mA for 60 s and imaged at an accelerating voltage of 5 kV. Mean fiber diameters and average pore diameter were determined by measuring independent fibers using Image J software (n = 50). 2.3.2. Chemical composition The CaP surface coatings were further confirmed by Fourier Transform Infrared Spectroscopy (FT-IR), using the reflection technique (ATR-FTIR, Perkin Elmer 2000, Swabian, Burladingen, German). All spectra were measured at a resolution of 16 cm−1.
Fig. 1. SEM micrographs of various electrospun scaffolds. PCL scaffolds (A) have larger fiber diameter than PCL–Keratin scaffolds (D). After crosslinking with glutaraldehyde, PCL–cKeratin matrices (E) exhibited significantly a larger fiber diameter. Both CaP surface coated scaffolds, sCaP–PCL (C) and sCaP–PCL–cKeratin (F) have no significant difference compared to respective scaffolds without CaP. Agglomeration of HA can be observed on HA loaded PCL scaffold, HA–PCL (B, inset), which exhibited a similar fiber diameter with PCL scaffolds.
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2.3.3. Mechanical properties The mechanical properties of the scaffolds were evaluated at room temperature using the Instron Tensile Machine 5567 (Instron, Canton, Massachusetts, USA). ASTM dog bone shaped specimens of dimension 9.53 mm × 3.18 mm were prepared and tested using a load cell of 500 N at a tensile loading rate of 10 mm·min−1. The mechanical properties of the scaffolds were calculated (BlueHill version 2.21) for their respective tensile strength (TS), strain at break (ε) and Young's modulus (E) in triplicates (n = 3). 2.3.4. In-vitro bioactivity of electrospun scaffolds In-vitro bioactivity was examined by immersing the scaffolds in 5 mL of stimulated body fluid (SBF) [14] for 7 days. After which, samples were washed thrice with DI water and vacuum dried for 24 h before taking gravimetric measurements. The initial and final weights of the scaffolds were compared to reflect the formation of apatite. The morphology of these scaffolds was characterized with FESEM (JSM6340F, JEOL Co., Tokyo, Japan). 2.4. Cell culture Human mesenchymal stem cells (hMSCs, Lonza) were used to assess the biocompatibility of the prepared scaffolds. The hMSCs were maintained in 75 cm2 cell culture flasks and passaged whenever confluency
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of 70–80% was attained. hMSCs from passage 5 were used in our study. PCL, sHA–PCL, cKeratin–PCL, and sHA–cKeratin–PCL scaffolds were cut into circular discs of 15.6 cm diameter and fitted into 24 well cell culture plates. The scaffolds were UV-sterilized for 30 min and rinsed three times with phosphate buffer saline (PBS, OHME Scientific) followed by 3 h incubation in complete culture medium consisting of Dulbecco's Modified Eagle's Media (DMEM), high glucose medium (PAA Laboratory), 10% Fetal Bovine Serum (PAA Laboratory), 1% L-Glutamine (200 mM, PAA Laboratory) and 1% Penicillin–Streptomycin (PAA) prior to cell seeding. Cells of density 5000 cells per cm2 were seeded on the scaffolds and cultured in a humidified atmosphere at 37 °C with 5% CO2. Cell culture medium was refreshed every 3 days.
2.5. Cell proliferation assay The proliferation of hMSCs that have attached and grown on the scaffolds was determined by measuring the amount of double stranded DNA (dsDNA) with the PicoGreen assay. Dead cells were removed by complete removal of the cell culture medium and thrice washing with PBS. Remaining adherent cells were lysed through gentle agitation in Triton X-100 solution (0.1% v/v). Subsequently, 100 μL of PicoGreen working solution was added and mixed with 100 μL aliquots of the cell lysate. Fluorescent intensity was measured from black 96-well flat plate using an Infinite200 microplate reader (Tecan Inc., Männedorf,
Fig. 2. ATR-FTIR analyses of various scaffolds: (A) PCL and SCaP–PCL and (B) PCL, Keratin, cKeratin–PCL and sCaP–cKeratin–PCL.
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Fig. 3. SEM micrographs of (A) HA–PCL; (B) sCaP–PCL; (C) sCaP–cKeratin–PCL scaffolds after 7 days in vitro apatite forming study. No significant apatite formation was observed on PCL (A) scaffolds. However, CaP-surface coated scaffolds (B: sCaP–PCL; C: sCaP–cKeratin–PCL) showed superior apatite forming ability, by forming a homogeneous layer of apatite after 7 days of SBF immersion.
Zurich, Switzerland) with an excitation/emission wavelength of 480/ 520 nm. 2.6. Statistical analysis All data are expressed as means ± standard deviation. Statistical differences are determined with one way ANOVA followed by Tukey's HSD post hoc test on SPSS version 11.5 and differences are considered statistically significant at p ≤ 0.05. 3. Results and discussion Bone tissue repair is an area of extensive research, whereby the scaffold material and a combination of other properties, such as adequate structural support and a bio-mimicking cell micro-environment, play
pivotal roles in creating a bio-functional bone-regenerating implant. To achieve such a scaffold, a new method was devised to produce CaPcoated, keratin-incorporated, electrospun PCL scaffolds (Scheme 1). In this paper, the cell differentiating properties of this scaffold would not be investigated. To evaluate and make comparisons with this scaffold, all other formulations and their respective controls (see Table 1) were electrospun under similar conditions. All fibers were found to possess uniform and “beadless” morphologies, as shown in Fig. 1. Here, a “beadless” morphology was achieved through the use of HFIP as solvent. Upon the incorporation of keratin into PCL (Keratin–PCL scaffold), the diameter of the fibers decreased from 552 ± 66 nm (Fig. 1A) to 196 ± 51 nm (Fig. 1D). The reduction in fiber diameter could be attributed to the increased electroconductivity of the electrospinning solution with the addition of keratin. The amino acid groups of keratin conferred negative
Fig. 4. Weight differences of various scaffolds after 7 days of SBF immersion. CaP surface-coated scaffolds, sCaP–PCL and sCaP–cKeratin–PCL have shown 17.36% and 16.02% of weight increment respectively after 7 days of immersion. No significant weight incensement was observed for HA imbedded HA–PCL scaffold.
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Fig. 5. Stress–strain curves of various scaffolds. The Young's modulus of sCaP–cKeratin–PCL (25.92 MPa) N cKeratin–PCL (15.38 MPa) N PCL (9.61 MPa) N sCaP–PCL (7.55 MPa) N HA–PCL (6.86 MPa).
charges to the protein and hence increased the solution's electroconductivity [15]. The formation of such nano-sized fibers was nonetheless beneficial at such size range, as these nano-fibers were reported to improve cell adhesion and growth kinetics [16]. In place of keratin, HA was also electrospun together with PCL polymer, as another control (HA–PCL). Not surprisingly, the HA–PCL scaffold (Fig. 1B) only showed minute amounts of HA on the surface of the fibers, even at high HA concentrations (10 wt.%), implying that most of this bone-like mineral was embedded within the fibers. PCL scaffolds were subsequently coated with CaP; i.e. pre-treated PCL and Keratin–PCL scaffolds (sCaP–PCL & sCaP–cKeratin–PCL — ‘s’ denotes ‘surface’), as shown in Scheme 1. Pre-treatment was first conducted for both sample scaffolds, and also for the non-keratin PCL scaffolds for comparison (Scheme 1A). Pre-treatment was essential to enhance the coating efficiency of CaP through the introduction of electrondonating groups for the adsorption of Ca2+ to initiate the nucleation of apatite. The most common pre-treatment methods used for biodegradable polyesters were either oxygen plasma or alkaline and acid treatment for polyester scaffolds. [17] Oxygen plasma treatment introduces oxygen free radicals, while alkaline and acid treatments increase the number of carboxyl and hydroxyl end groups. These functional groups aid in chelating with Ca2 + ions and initiate the formation of CaP as a coating. However, both techniques can be detrimental to the polymer, and may cause chain scission in PCL. To minimize polymer degradation of PCL, a mild oxygen plasma condition, such as low power (10 W) and short exposure time (30 s), was used. This mild process did not result in any observable morphology changes to the PCL fibers (data not shown). For the keratin-based PCL scaffolds, keratin was first crosslinked to introduce multidentate electron-donating groups on the fiber surface (Scheme 1B). Keratin–PCL electrospun scaffolds was pre-treated via a condensation reaction with glutaraldehyde to create the multidentate Schiff base ligands, which are well-known metal ion chelating ligands [18,19]. The increase in fiber diameter from 196 ± 51 nm (Fig. 1D) to 253 ± 54 nm after crosslinking (Fig. 1E) could be
due to the absorption of water vapor during the crosslinking process [20]. Pre-treated scaffolds were subsequently coated with CaP through sequential immersion of the scaffolds (i.e. plasma-treated PCL scaffold and crosslinked Keratin–PCL scaffold), in Ca(NO3)2 and NH4H2PO4 solutions. A thick and homogeneous layer of CaP (Fig. 1C and F) was observed within 10 min. It is worth mentioning that this is the first report in which CaP surface coating was achieved through the chelation reaction of crosslinked keratin and Ca2+ on the surface of electrospun polymeric fibers, and within such a short period. This process largely improved the homogeneity and efficiency of CaP coating by shortening the coating period from days [5] to just minutes. The coating of CaP on the scaffolds was qualified by ATR-FTIR spectroscopy. From Fig. 2A, comparing against PCL, the spectrum of sCaP– PCL scaffolds displayed the characteristic triplet absorption bands of the (PO4)3− groups at 1000–1100 cm−1. The presence of an absorption band at 1579 cm−1 for the spectrum of sCaP–PCL scaffold is contributed from the red shift of Ca2+ chelated ester groups (–C=O: 1722 cm−1). The strong multi-band absorptions from 1200–1400 cm−1 may be attributed to Ca2+ chelated C–O–C groups that were introduced by oxygen plasma treatment. In Fig. 2B, compared with the spectrum of PCL, the extra absorbances at 1643.9 and 1542.8 cm−1 were typical amide 1 and amide II absorptions, which were attributed by the presence of keratin. The red shift of amide II from 1520.5 cm− 1 of keratin into 1542.8 cm−1 further proved the crosslinking reaction between amide and aldehyde groups. Similar to sCaP–PCL, typical CaP absorption bands were observed at 1000–1100 cm−1 for sCaP–cKeratin–PCL scaffolds, representing the typical (PO4)3− adsorption of surface CaP. Moreover, the amide II absorbance was found to be blue-shifted from 1542.8 cm−1 to 1591.3 cm−1, suggesting the chelation of Ca2 + ions with the amide groups of keratin. The coating of the scaffold surface with CaP is aided through the interaction of the exposed hydroxyl groups from the CaP [Ca10 (PO4)6 (OH)2] with the Ca2 + ions present in bodily fluid [21]. The influence
Table 2 Comparison of mechanical properties of various scaffolds in terms of tensile strength (TS), strain at break (ϵf) and Young's modulus (E). Properties
PCL
HA–PCL
sCaP–PCL
cKeratin–PCL
sCaP–cKeratin–PCL
Tensile strength (MPa) Strain at break (%) Young's modulus (MPa) Pore size (μm)
8.00 ± 0.43 279.06 ± 15.38 9.61 ± 0.56 2.64 ± 0.21
3.33 ± 0.41 164.59 ± 31.58 6.86 ± 0.41 2.35 ± 0.51
5.54 ± 1.17 251.58 ± 69.93 7.55 ± 0.70 3.12 ± 0.21
13.41 ± 0.94 226.44 ± 29.44 15.38 ± 0.57 2.11 ± 0.30
16.53 ± 1.16 152.78 ± 19.86 25.92 ± 0.96 2.66 ± 0.41
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Fig. 6. Cell viability study of various scaffolds was conducted through Picogreen Assays after 4 days of culturing. Keratin containing scaffolds, cKeratin–PCL & sCaP–cKeratin– PCL, showed better cell viability than controls (PCL & sCaP–PCL).
of CaP-containing scaffolds was subsequently evaluated, i.e. electrospun HA–PCL; sCaP–PCL and sCaP–cKeratin–PCL. Scaffolds were determined for their rate of mineralization of the scaffold in SBF medium [22], after 7 days of immersion (Fig. 3). From SEM images, sCaP–PCL (Fig. 1C) and sCaP–cKeratin–PCL (Fig. 1F) scaffolds were covered with a thick layer of calcium phosphate (CaP) after 7 days, as represented by Fig. 3B and C respectively. To quantify the amount of CaP formed, weight measurements were conducted on these scaffolds. The results from the analysis (Fig. 4) indicated ~ 16 wt.% of CaP on both sCaP–PCL and sCaP–cKeratin–PCL scaffolds, with the lowest wt.% of CaP (~2%) observed for HA–PCL scaffold. This implied that apatite formation was not significant in HA–PCL scaffold due to a lack of surface CaP. These findings therefore highlighted the importance of having a significant concentration of surface CaP on the fibers.
The mechanical properties of these scaffolds were subsequently evaluated through tensile mechanical testing as shown in Fig. 5 and summarized in Table 2. From the stress–strain plots shown in Fig. 5, PCL scaffolds had a tensile strength (TS) of 8.0 ± 0.4 MPa and strain at break (ε) of 279%. The presence of CaP, regardless of surface coating (sCaP–PCL) or incorporated during electrospinning (HA–PCL), was accompanied with a significant decrease in both TS and ε values. Comparatively, despite a slight decrease of ε for cKeratin–PCL scaffold, TS value increased significantly to 13.4 ± 0.9 MPa compared to pristine PCL scaffold. The highest TS value was observed on the sCaP–cKeratin–PCL scaffold (16.5 ± 1.2 MPa). The Young's modulus of the scaffolds as ranked in descending order was as follows: sCaP–cKeratin–PCL N cKeratin– PCL N PCL N sCaP–PCL N HA–PCL. The poor mechanical properties of HA–PCL could be due to the possible agglomeration of HA nanoparticles in the polymer matrix, causing micro-defects to the fiber scaffold. In order to overcome this issue, Prabhakaran et al. reported the incorporation of collagen into PLLA scaffold to improve HA dispersion [23]. Nonetheless, TS was substantially reduced from 4.7 to 2.1 MPa upon the incorporation of both HA and collagen. In our study, cKeratin–PCL scaffold was surface-coated with CaP through a chemical Ca2+ chelation. The increase in TS in cKeratin–PCL scaffold could be due to the presence of strong and rigid crosslinkedkeratin structure [24]. With surface CaP, the strong chemical bonds formed between CaP nanoparticles and the cKeratin–PCL scaffold [25] greatly enhanced the TS of sCaP–cKeratin–PCL scaffold, with tensile strength that meets the mechanical requirements for bone TE applications [26]. The biocompatibility of the scaffolds was finally evaluated by using PicoGreen assay after 4 days of culturing of hMSCs on these scaffolds. From Fig. 6, the cell viability of hMSCs on keratin containing scaffolds: cKeratin–PCL and sCaP–cKeratin–PCL, was significantly higher than those on PCL and sHA–PCL (p ≤ 0.05). This may be attributed to the presence of LDV cell binding motifs on keratin in promoting integrin α4β1 mediated cell adhesion [27,28], and consequently improving cell
Fig. 7. SEM micrographs showing that cell adopted spindle-shaped morphologies on PCL scaffolds after 4 days of culturing (A), while wide-spread morphologies were observed on sCaP– cKeratin–PCL scaffolds (C) after 7 days of culturing. Cell on sCaP–cKeratin–PCL scaffolds was well integrated with Keratin-contained fibers on 4 days (B), 7 days (C) and 14 days (D) of culturing. Mineralization was clearly observed especially on 14 days of culturing.
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viability. The presence of HA has been shown to enhance cell attachment at the early stages of cell culture and improved cell proliferation [29]. However, in our study, cell viability data did not show any significant differences between scaffolds containing HA and their respective controls (without CaP coatings). This is probably because CaP and cell interactions are influenced by surface roughness and chemical composition of scaffold, as well as the crystal and grain size of the CaP nanoparticles. For example, Wang et al. [30] and Ribeiro et al. [31] reported that micro-sized HA could improve the performance of SaOS-2 and MC3T3E1 cells respectively but Smith et al. [32] found that MC3T3-EI cells proliferated faster on nano-sized HA scaffolds. Our findings nevertheless indicated higher cell proliferation on sCaP–cKeratin–PCL scaffold as compared to that on sCaP–PCL scaffold. This could be due to better initial cell attachment on keratin containing scaffolds rather than the effect of HA nanoparticles or the lateral apatite forming rate. The SEM images of cell-scaffold constructs are shown in Fig. 7. hMSCs exhibited an elongated morphology on PCL scaffold after 4 days of culture (Fig. 7A). In contrast, well-spread hMSCs with polygonal morphology that integrated well with the nanofibers of the sCaP– cKeratin–PCL scaffold were observed on day 4 (Fig. 7B), day 7 (Fig. 7C) and day 14 of culture (Fig. 7D). In summary, scaffolds containing crosslinked keratin such as cKeratin–PCL and sCaP–cKeratin–PCL are suitable to support cell attachment and proliferation. The osteogenic differentiation of hMSCs cells on sCaP–cKeratin–PCL would be one aspect of this work that would be studied in the future. 4. Conclusions A two-step approach to coat CaP onto the surface of electrospun PCL scaffolds was proposed. Keratin was incorporated into the electrospun PCL scaffold and used to induce the formation of a HA surface coating after a crosslinking reaction. The crosslinking of keratin was able to promote Ca2 + chelation and a homogeneous CaP coating onto PCL scaffolds. Our study suggested that the mineralization rate was related to the amount of surface available CaP. In addition, the presence of crosslinked keratin and CaP further increased the mechanical strength of the resultant scaffolds. Moreover, higher proliferation rate of hMSCs cells was observed on keratin containing scaffolds. In conclusion, the optimized scaffold, sCaP–cKeratin–PCL had improved mechanical properties and good biocompatibility. Therefore, sCaP–cKeratin–PCL scaffold can serve as a potential biocomposite material for bone tissue engineering. Acknowledgments The authors would also like to acknowledge the financial support from the Agency for Science, Technology and Research (A*STAR) (Project No: 102 129 0098), the National Medical Research Council, Ministry of Health (NMRC/CIRG/1342/2012, MOH), the NTU-National Healthcare Group (NTU-NHG) grant (ARG/14012), the Singapore Centre on Environmental Life Sciences Engineering (SCELSE) (M433C70000.M4220001.C70), and the School of Materials Science and Engineering (M020070110), Nanyang Technological University, Singapore. Appendix A. Supplementary data Supplementary data to this article can be found online at http://dx. doi.org/10.1016/j.msec.2015.01.084. References [1] M.J. Glimcher, Molecular biology of mineralized tissues with particular reference to bone, Rev. Mod. Phys. 31 (1959) 359.
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