Preparation and characterization of foamy poly(γ-benzyl-L-glutamate-co-L-phenylalanine)/bioglass composite scaffolds for bone tissue engineering

Abstract

In this study, novel foamy scaffolds of poly(γ-benzyl-L-glutamate) (PBLG) and poly(γ-benzyl-L-glutamate-co-L-phenylalanine) (P(BLG-PA)) were fabricated via a combination of a sintered NaCl templating method and ring-opening polymerization of α-amino acid N-carboxyanhydrides. The obtained polypeptide-based scaffolds displayed a foamy structure and had a desirable glass transmission temperature (23.2–44.6 °C), contact angle (78.5–88.5°), compression modulus (56.3–110.2 kPa) and degradation time (>8 weeks). MC3T3-E1 cells were used to test the in vitro biocompatibility, and we found that PBLG scaffolds could significantly promote cell viability and proliferation compared to P(BLG-PA) scaffolds. In addition, micro-scale bioglass particles (BG) were incorporated into PBLG to form a porous scaffold, namely PBLG/BG composite scaffolds, which further enhanced the cell viability and adhesion, as confirmed by live-dead staining and SEM observation. The results of the ALP activity assay suggested that PBLG/BG composite scaffolds could promote the osteogenic differentiation of MC3T3-E1 cells. The results of histology analysis showed that the PBLG/BG composite scaffolds rather than PBLG scaffolds significantly promoted connective tissue and vascular ingrowth. The results obtained using digital radiography technology showed that PBLG/BG composite scaffolds significantly improved osteogenesis in vivo. In summary, our results indicated that PBLG/BG composite scaffolds could be a promising bioactive substance for bone regeneration.

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1. Introduction

The gold standards of scaffold materials for tissue engineering include excellent biocompatibility, biodegradability, appropriate physicochemical properties and nontoxic degradation products. Currently, scaffold materials for tissue engineering mainly originate from natural polymers (e.g. hyaluronic acid,¹ chitosan,² collagen,³ alginate⁴ and silk⁵) and synthetic polymers (e.g. polylactic acid,⁶ poly(lactic-co-glycolic acid)⁷ and polycaprolactone⁸). The scaffolds developed from natural polymers may have potential unpredictable and changeable enzymatic degradation, weak mechanical properties and risk of immunological rejection.⁹ On the other hand, most scaffolds derived from synthetic polymers exhibit controllable physicochemical properties and predictable biodegradability,¹⁰ but their acidic breakdown products show cytotoxicity in vivo.¹¹ Hence, developing an idea scaffold material that can better mimic the structure, morphology, chemical and biological properties of natural bone extracellular matrix is still a big challenge.

As an alternative of classical synthetic biopolyesters, polypeptides and their derivatives have been developed to mimic natural proteins because they possess excellent biocompatibility and biodegradability. In recent years, synthetic polypeptides, which are composed of α-amino acids that are linked together through peptide bonds,¹³ have emerged as an inspiring platform for biomedical applications.¹² The synthetic polypeptides possess many merits, including inherent stability of the polyamide skeleton against hydrolysis, specific enzyme degradation into natural α-amino acids, non-toxic products with a near neutral pH, facile synthetic routes, flexible functionality, excellent biocompatibility, non-immunogenicity, and unique properties different from conventional synthetic polymers.¹⁴ Synthetic polypeptides have attracted increasing attention in the field of drug and gene delivery,^15–19 such as poly(γ-glutamic acid),²⁰ poly(L-lysine),²¹ poly(L-phenylalanine)²² and polyalanine.²³ Till now, the ring-opening polymerization (ROP) of α-amino acid N-carboxyanhydrides (NCAs) has become the most widely used synthetic method of polypeptides with high molecular weights and well-fined structures.²⁴ The ROP of α-amino acid-NCAs can be efficiently initiated by amino and hydroxyl groups as well as water molecules. However, few reports on preparing polypeptide derivative scaffolds for bone tissue engineering by the ROP method have been published.

Commonly serving as a temporary support, polymeric scaffolds cannot promote bone regeneration. Therefore, bioactive inorganic materials, particularly bioactive glasses/bioglasses (BG), need often be incorporated into the scaffolds because of their abilities to convert to hydroxyapatite (HA, the mineral constituent of bone) and to release ions that stimulate osseous healing and vascularization.²⁵ This kind of material has emerged as a promising and versatile candidate for biomedical applications and models that require the infiltration and colonization of a blood vessel network.²⁶

Porous scaffolds can be fabricated by many methods, such as salt-leaching, gas foaming, electrospinning and rapid prototyping.²⁷ However, most of them are unable to produce scaffolds with ideal high inter-pore connectivity or adjustable pore size. Some researchers found that the porous scaffolds prepared from a sintered negative NaCl template display some advantages, such as easy-preparation, no residual salt, controllable pore size and high pore connectivity.²⁸ To the best of our knowledge, polypeptide derivative/BG composite scaffolds with such pore characteristics have never been reported.

Herein, we synthesized poly(γ-benzyl-L-glutamate) (PBLG) and poly(γ-benzyl-L-glutamate-co-L-phenylalanine) (P(BLG-PA)) with different molar ratios via ROP of γ-benzyl-L-glutamate-N-carboxyanhydride (BLG-NCA) and L-phenylalanine N-carboxyanhydride (PA-NCA) in anhydrous chloroform, using triethylamine (TEA) as an initiator. These polypeptides were fabricated into foamy scaffolds by the sintered negative NaCl templating method. The in vitro response of MC3T3-E1 cells to the PBLG and P(BLG-PA) scaffolds was evaluated, and the effects of glass transition temperature (T_g), contact angle and compression modulus on the adhesion and proliferation activity of MC3T3-E1 cells were discussed. Subsequently, both PBLG and poly(γ-benzyl-L-glutamate)/bioglass (PBLG/BG) composite scaffolds were implanted subcutaneously into SD rats, and PBLG-based scaffolds were implanted into rabbit tibia defects. The results from histology analysis confirmed that PBLG/BG composite scaffolds displayed better tissue and vessel ingrowth as well as good in vivo osteogenesis. In summary, PBLG/BG composite scaffolds were a favorable substrate for MC3T3-E1 cells and showed good in vitro and in vivo biocompatibility, rendering great potential for bone tissue engineering application.

2. Materials and methods

2.1 Materials

L-Glutamic acid hydrochloride, L-phenylalanine, bis(trichloromethyl) carbonate (BTC) and TEA were purchased from Aladdin Reagent Inc. (Shanghai, China). 3-(4,5-Dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) was purchased from Sigma-Aldrich. The LIVE/DEAD viability/cytotoxicity kit for mammalian cells was purchased from Invitrogen (Carlsbad, CA, USA). Minimum essential medium α (MEMα) and fetal bovine serum (FBS) were purchased from Procell (Wuhan, China). The alkaline phosphatase (ALP) kit, hematoxylin and eosin (H&E) staining kit and Masson's trichrome staining kit were purchased from BOSTER (Wuhan, China). CD34 monoclonal anti-rat antibody was purchased from Abcam (Cambridge, MA, USA). Both MC3T3-E1 cells and Sprague Dawley (SD) rats were supplied by Medical Center of Xi'an Jiaotong University (Xi'an, China). Tetrahydrofuran (THF) and n-hexane were dried by refluxing over sodium wire and distilled prior to use. Chloroform, diethyl ether and TEA were dried by refluxing over calcium hydride and distilled before use. All other chemicals were of analytical grade, purchased from Sinopharm Chemical Reagent (Xi'an, China) and used without further purification. All aqueous solutions were prepared using ultrapure water with a resistance of 18.25 MΩ.

2.2 Synthesis of PBLG, P(BLG-PA) and PBLG/BG composite scaffolds

2.2.1 Synthesis of BLG-NCA, PA-NCA and P(BLG-PA) copolymer. BLG-NCA was synthesized by reacting γ-benzyl-L-glutamate (BLG) with BTC according to our previously reported procedure.²⁹ Briefly, BLG (10 g, 42.2 mmol), which was obtained from L-glutamic acid and benzyl alcohol following a literature method²⁹ (44.6% yield, mp 173–173.5 °C), was suspended in 100 mL of anhydrous THF in a flame-dried flask under nitrogen atmosphere. BTC (5.0 g, 16.9 mmol) dissolved in 20 mL of THF was added in one portion via a syringe. The mixture was stirred at 50 °C until a clear solution was obtained. Residual BTC was removed by bubbling dry nitrogen gas through the solution, and then the solution was concentrated under vacuum. The crude product was precipitated with anhydrous n-hexane and purified by recrystallization twice with n-hexane, resulting in white BLG-NCA crystals (10.2 g, yield 92.2%, mp 91–92 °C). Proton nuclear magnetic resonance (¹H NMR) (CDCl₃, δ ppm): 2.15–2.3 (m, CH₂), 2.63 (t, CH₂), 4.39 (m, CH), 5.16 (s, CH₂), 6.3 (s, NH), 7.38 (m, C₆H₅). PA-NCA was synthesized by the same method. Yield: 95.1%, mp 92–93 °C. ¹H NMR (CDCl₃, δ ppm): 3.0–3.3 (m, CH₂), 4.56 (m, CH), 5.78 (s, NH), 7.38 (m, C₆H₅).

PBLG and P(BLG-PA) were synthesized as follows. BLG-NCA and PA-NCA at different mole ratios (1 : 0, 7 : 3 and 5 : 5) were dissolved in anhydrous chloroform under nitrogen atmosphere. Initiator TEA, 0.1% of the number of moles of monomers, was added to the solution. After 72 h of reaction, most of the solvent was removed under reduced pressure and the solution was poured into ice-cold diethyl ether. A faint yellow solid was collected and dried under vacuum at 30 °C for 48 h. The polypeptide copolymers synthesized from BLG-NCA and PA-NCA at mole ratios of 7 : 3 and 5 : 5 were designated as P(BLG7-PA3) and P(BLG5-PA5), respectively.

2.2.2 Synthesis of micro-scale BG particles. BG was prepared according to a modified literature process.³⁰ Briefly, tetraethyl orthosilicate ((C₂H₅O)₄Si) and triethyl phosphate (O=P(OC₂H₅)₃) were mixed with a 1 M HNO₃ solution and hydrolyzed for 60 min under vigorous stirring. Then, calcium nitrate tetrahydrate (Ca(NO₃)₂·4H₂O) and sodium nitrate (NaNO₃) were gradually added into the mixture. After stirring for 6 h, BG sol was allowed to age in a sealed container and then dried for 12 h, resulting in formation of the gel. Next, the gel was heat-treated at 550 °C for 2 h in order to remove the organic substances. The obtained BG was further processed by a ball-milling method to obtain micro-scale particles. Since the content of SiO₂ in BG was about 45 wt%, the BG was called 45S5 BG.

2.2.3 Preparation of polypeptide-based scaffolds. Foamy polypeptide-based scaffolds were prepared via a sintered NaCl templating method. In a typical process, the porous NaCl templates were first fabricated by sintering NaCl particles in cylindrical alumina moulds at 550 °C for 2 h. The NaCl particles were 100–400 μm in diameter. A 10% wt/v solution of polypeptide in CHCl₃ was obtained by dissolving PBLG, P(BLG7-PA3) or P(BLG5-PA5) in anhydrous CHCl₃. To the solution the sintered NaCl templates were added, and the system was kept under −0.08 MPa for 10 min to allow the complete infiltration of the solution into the templates. Afterwards, the NaCl/polypeptide composites were taken out and air-dried at 30 °C for 24 h. Polypeptide scaffolds were obtained after the NaCl templates were removed by leaching with distilled water and dried under vacuum at 25 °C for 72 h. Foamy PBLG/BG composite scaffolds were prepared via the same process. Briefly, micro-scale BG particles were dispersed in the solution of PBLG in CHCl₃, and the mass ratio of BG particles and PBLG was fixed at 1 : 10. The following procedure was the same as described above except that the polypeptide solution was replaced by a PBLG/BG particle suspension.

2.3 Measurements

¹H NMR spectra of polypeptide derivatives were recorded on a Bruker 400 MHz spectrometer (Bruker, Germany) using CDCl₃ as the solvent and tetramethylsilane as the internal reference at 25 °C. Chemical shifts were recorded in ppm. Differential scanning calorimetry (DSC) measurement was performed on a Netzsch DSC200 F3 differential scanning calorimeter (Germany) in nitrogen atmosphere with a gas flow of 40 mL min⁻¹. The DSC sample was firstly cooled from room temperature to −30 °C at a rate of 10 °C min⁻¹ and kept for 10 min. The sample was subsequently heated to 120 °C at the same rate and held isothermally for 5 min. The compressive behaviour was measured using an Instron 5943 universal testing instrument at a constant rate of 1 mm min⁻¹. The slopes of the stress–strain curves from 10% to 20% deformation were used to calculate the compressive modulus values. The morphology of scaffolds was observed by scanning electron microscopy (SEM, S-3400 N, Hitachi, Japan). The wettability of polypeptide films was evaluated by measuring the contact angle of a 5 μL drop of deionized water using a contact angle goniometer (JC2000D2, Shanghai Zhongchen Digital Technology Apparatus Co., Ltd, China). Three samples were used for each test to obtain an average value. Degradation tests were carried out in PBS (pH 7.4) at 37 °C, and the PBS was replaced every day.

2.4 In vitro cell culture

MC3T3-E1 cells were used to assess the responses of osteoblastic cells to the scaffolds. MC3T3-E1 cells were cultured in MEMα supplemented with 10% FBS and 1% penicillin/streptomycin in a humidified atmosphere with 5% of CO₂ at 37 °C for 1, 3 and 5 days. Prior to cell seeding, all the scaffolds were sterilized with ethylene oxide gas and placed in a 24-well culture plate. MC3T3-E1 cells suspended in 0.2 mL of media were seeded onto the surface of scaffold samples at a concentration of 3 × 10⁴ cells per well. After 4 h of culture, 1 mL of culture medium was added. The culture medium was replaced everyday.

2.4.1 MTT assay. The MTT assay was used to evaluate the metabolic activity of cells cultured on the scaffolds. After 1, 3 and 5 days of incubation, the culture medium was removed and 100 μL of MTT solution was added to each well of the 24-well plate. The scaffold/cell constructs were incubated for 4 h at 37 °C in a humidified atmosphere with 5% of CO₂. After the culture medium had been removed, 200 μL of DMSO was added to each well to solubilize the blue formazan crystals. The solutions were then transferred to the wells of a 96-well plate and the optical absorbency of the solutions was measured at 490 nm using a microplate reader (PerkinElmer, EnSpire, America). Non-seeded scaffolds were used as a negative control. Each experiment was done in triplicate.

2.4.2 Live/dead cell viability and proliferation assays. The viability of MC3T3-E1 cells cultivated on PBLG scaffolds and PBLG/BG composite scaffolds was evaluated using the live/dead viability/cytotoxicity assay. On days 1, 3 and 5, the scaffold/cell constructs were clearly rinsed with PBS and stained with calcein AM and ethidium homodimer (ethD-1). The living cells were stained green and the dead cells were stained red. The concentration of both calcein AM and ethD-1 was fixed at 10 μM. The scaffold/cell constructs were visualized using a confocal laser scanning microscope (Leica, TCS SP5 II, Germany).

2.4.3 ALP activity assay. To assess the ALP activity of MC3T3-E1 cells grown on the scaffolds, 3 × 10⁴ cells were seeded on each scaffold in a 24-well plate and cultured for 1, 3 and 5 days. At predetermined time points, culture medium was decanted, and the cell layer was washed gently three times with PBS and washed once in cold 50 mM Tris buffer. Afterwards, the cells were lysed in 200 μL of 1% Triton X-100 and the lysates were sonicated after being centrifuged at 14 000 rpm for 30 min at 4 °C. 50 μL of supernatant was mixed with 150 μL of a working solution according to the manufacturer's protocol (Beyotime, China). The conversion of p-nitrophenylphosphate into p-nitrophenol in the presence of ALP was determined by measuring the absorbance at 520 nm with a microplate reader (Bio-Rad 680, USA). The ALP activity was calculated from a standard curve after normalizing to the total protein content, which was determined by a MicroBCA protein assay kit (Pierce, USA) at 562 nm with the microplate reader.

2.4.4 The morphology of MC3T3-E1 cells on scaffolds. To observe the attachment and morphology of MC3T3-E1 cells on the PBLG and PBLG/BG composite scaffolds, the scaffold/cell constructs after 1, 3 and 5 days of incubation were washed three times with PBS, fixed in 2.5% glutaraldehyde for 24 h, dehydrated in a graded series of ethanol solutions (50, 60, 70, 80, 90 and 100%), and dried by a critical point dryer (Quorum/Emitech K850, UK). The samples were sputter-coated with gold prior to SEM observed.

2.4.5 Release kinetics of Ca²⁺ and Si⁴⁺ of PBLG/BG composite scaffolds. Release of Ca²⁺ and Si⁴⁺ ions from scaffolds soaked in PBS under quasi-dynamic conditions was measured using the ICP-AES (ICAP700, Thermo, USA). Each scaffold was immersed in 50 mL of PBS and enclosed in a test tube that was incubated and shaken in a 37 °C water bath and they were monitored for 28 days. PBLG/BG composite scaffolds were analyzed at predetermined time points for concentrations of Ca²⁺ and Si⁴⁺. Five point calibrations (25, 10, 5, 1 and 0.1 ppm) were performed by diluting certified standards (Carl Roth, Germany) with PBS. Given errors are estimated by linear regression analysis.

2.5 In vivo scaffold implantation

2.5.1 Scaffold implantation into SD rat model. To assess the connective tissue ingrowth, angiogenesis and immunogenicity of PBLG and PBLG/BG composite scaffolds in vivo, the scaffolds were embedded subcutaneously in SD rats. Surgery was approved by the Medical Center of Xi'an Jiaotong University. Before surgery, SD rats with a body weight of 190–200 g were used. All rats were purchased from the Medical Center of Xi'an Jiaotong University (Xi'an, China). All experiments involving animals and their care were approved by the local governmental animal care committee and were conducted in accordance with the NIH Guide for the Care and Use of Laboratory Animals. Two groups of grafts (PBLG and PBLG/BG composite scaffolds) were sterilized with ethylene oxide and then implanted into the subcutaneous pockets on the backs of SD rats that were anesthetized by chloral hydrate solution (10% wt/wt, 4 μL g⁻¹ body weight). Three rats were used for each experiment. After the surgeries, cefazolin (25 μg g⁻¹ body weight) and buprenorphine (0.1 μg g⁻¹ body weight) were administered and each rat was individually housed. After being carefully nurtured in a sterilized environment for 4 and 8 weeks, all SD rats were sacrificed for histological analysis.

2.5.2 Histological staining. After being explanted from SD rats, the scaffold implants were fixed overnight at room temperature in buffered formalin (37%, pH 7.2–7.4, Sigma-Aldrich) and embedded in paraffin, and then 7 μm sections were prepared. H&E, Masson's trichrome staining and CD34 were selectively performed on 7 μm thick sections to observe the tissue response to scaffold implants and the ingrowth ability of tissue and microvessels. The histological sections were observed using a Leica SCN400 Slide Scanner and the images were captured using the Leica Digital Image Hub (Leica, Germany).

2.5.3 Scaffold implantation into a rabbit tibia defect model. A rabbit model with a tibia defect was used to evaluate the in vivo osteogenesis of PBLG-based scaffolds. Six adult white New Zealand rabbits (male, 2–2.5 kg body weight) were obtained from the Laboratory Animal Center of Xi'an Jiaotong University (Xi'an, China). The rabbits were anesthetised by injecting 3% Nembutal (30 mg kg⁻¹) via the ear vein, and a longitudinal incision was made using a scalpel in the rabbit tibia under rigorous aseptic conditions. Quadrate holes about 3 mm width, 8 mm length and 8 mm deep were drilled using a surgical electronic drill and thoroughly rinsed with physiological saline to remove shards of bone. Implants of PBLG and PBLG/BG composite scaffolds were sterilized with ethylene oxide and then implanted into the defect model. The wound was sutured with nylon thread. Three rabbits were used for each experiment. After the surgeries, cefazolin (25 μg g⁻¹ body weight) and buprenorphine (0.1 μg g⁻¹ body weight) were administered and each rabbit was individually housed. After being carefully nurtured for 1 and 8 weeks, the healing state of all rabbits were observed using digital radiography. Sequential radiographs of the tibia bone were recorded at regular intervals (1 and 8 weeks post-operatively) in a digital X-ray imaging system (5000CPlus, Mednova, China) to study the status of the implant and the degree of host bone–material interaction.

2.6 Statistical analysis

Experimental data were expressed as mean ± standard deviation. The statistical analysis for the differences between the groups was performed by means of Student's t-test using the SPSS13.0 software. A confidence level of 95% (p < 0.05) was considered to be statistically significant. Each experiment was done in triplicate.

3. Results and discussion

3.1 Preparation of polypeptide-based scaffolds

The synthesis route of polypeptide derivatives is shown in Fig. 1a. A mixture of BLG-NCA and PA-NCA at different mole ratios was dissolved in anhydrous chloroform under nitrogen atmosphere. The ROP was performed at 30 °C for 3 days under magnetic stirring, using TEA as the initiator. The solution was precipitated into excess diethyl ether followed by drying in a vacuum, and the product was obtained as a faint yellow solid. The ¹H NMR spectra of the as-synthesized polypeptide derivatives are shown in Fig. 1b. All signals at 2.24 (b and c), 3.89 (a), 5.1 (d) and 7.2 (f) ppm from the protons of PBLG were well identified.²⁹ The characteristic peak at 3.15 ppm originated from PA moieties.³¹ The integrated area ratio of the peaks at 3.15 (e from PA) and 5.1 ppm (d from BLG) of PBLG, P(BLG7-PA3) and P(BLG5-PA5) were 0, 0.42 and 0.73, respectively, which approximate their theoretical molar ratios of 0 : 1, 3 : 7 and 5 : 5, demonstrating the successful synthesis of PBLG and P(BLG-PA).

Fig. 1 Synthesis route (a) and ¹H NMR spectra (b) of P(BLG-PA) scaffolds.

To obtain foamy polypeptide-based scaffolds, a synthetic polypeptide derivative or a mixture of the polypeptide derivative and BG particles was added into anhydrous chloroform to form a solution or suspension, to which the sintered NaCl templates were added. The polypeptide derivative/NaCl and polypeptide derivative/BG/NaCl composites were prepared by the vacuum-pressure infiltration method followed by air-drying at 35 °C for 24 h. NaCl templates were removed by leaching in deionized water, and the final scaffolds were obtained after drying under vacuum at 25 °C for 48 h. The porous scaffold discs of 15 mm in diameter and 3 mm thickness, were seeded with MC3T3-E1 cells for in vitro culture and implanted in the SD rat subcutaneous embedding model and rabbit tibia defect model for in vivo evaluation.

3.2 Morphology

The morphology and architecture of polypeptide-based scaffolds were studied by SEM, and the representative SEM images are displayed in Fig. 2. As revealed in Fig. 2b–d, all polypeptide-based scaffolds retained the foamy characteristics of the sintered NaCl template well (Fig. 2a). The morphology, channels and micropore sizes were determined by the original NaCl template. As reported in the literature,^32,33 such foam characteristics could promote cell attachment, the transport of nutrients and metabolites, and new tissue remodeling. In contrast to the traditional salt-leaching method,^34,35 the sintered NaCl templating method has the advantages of high pore interconnectivity and no salt residue. Therefore, it can be concluded that the porous structure of NaCl templates is crucial to the interpore openings and pore shape of polypeptide-based scaffolds.

Fig. 2 SEM images of the NaCl template (a), PBLG (b), P(BLG7-PA3) (c) and P(BLG5-PA5) (d) scaffolds.

3.3 DSC measurement, wettability assay, and compressive and degradation behaviours

DSC was use to reveal the T_g of the polypeptide derivatives. As shown in Fig. 3a, the DSC curves of all polypeptide derivatives showed a similar trend. As for PBLG, an obvious endothermic peak at 23.2 ± 0.2 °C, representing T_g, was observed in the heating curve. The T_g of P(BLG7-PA3) and P(BLG5-PA5) was 34.5 ± 0.5 and 44.6 ± 0.5 °C, respectively. It was noted that the T_g of P(BLG-PA) increased with the increasing content of PA.³⁶ The reason for this may be related to the steric hindrance of the PA moiety due to the existence of the benzene ring. Since the T_g of PBLG was much lower than the cell culture temperature, PBLG scaffolds were in the thermoplastic elastomeric state during cell culturing. Meanwhile, the T_g of P(BLG7-PA3) was much higher than that of PBLG and a little lower than 37 °C, indicating that P(BLG7-PA3) was in a glass transition region and the corresponding scaffolds were in the viscoelastic state during cell culturing. P(BLG7-PA3) was in the glassy state during cell culturing since its T_g was higher than 37 °C.

Fig. 3 DSC curves (a), water contact angles (b), compression stress–strain curves (c) and degradation kinetics (d) of PBLG, P(BLG7-PA3) and P(BLG5-PA5) scaffolds.

The hydrophilic–hydrophobic property of polymeric materials has a significant effect on the biological properties of scaffolds, including cell adhesion, migration and spreading.^37,38 In this study, the water contact angle of different polypeptide derivative membranes was measured to compare their hydrophilic–hydrophobic property. As shown in Fig. 3b, the contact angle of PBLG, P(BLG7-PA3) and P(BLG5-PA5) membranes was 78.5 ± 0.2, 83.1 ± 0.3 and 88.5 ± 0.9°, respectively. The results indicated that the contact angle of the polypeptide derivative membrane increased with the increasing PA content. The reason is that the PA moiety displayed stronger hydrophobicity than the BLG moiety.³⁹ The ester groups (–COO–) in the side chain of BLG units can interact with water molecules via weak intermolecular hydrogen bonds, which results in their better hydrophilicity than PA units, though these two kinds of units have benzene ring groups.

Fig. 3c shows the representative nominal compression stress–strain curves of PBLG, P(BLG7-PA3) and P(BLG5-PA5) scaffold samples. As displayed in Fig. 3c, the stress increased with the increasing strain for all three scaffolds, and their compression moduli over the 10–20% strain range were 110.2 ± 6.2, 86.5 ± 1.5 and 56.3 ± 2.2 kPa, respectively. The values are reasonable considering their high porosity up to 90%. The three curves exhibited a similar trend. The most plausible reason for this is that the scaffolds with similar porosity and porous structure exhibit a similar mechanical behaviour.^40,41 Moreover, the stress–strain curves exhibited a typical three-stage behaviour: elastic, plateau and densification and depended on the molar ratio of BLG and PA units. The polypeptide-based scaffolds obtained in this study not only can meet the mechanical requirement for in vitro cell culture experiments but also are promising for tissue engineering in non-load-bearing bones.^42–44 Generally speaking, a polymer with a higher T_g has a higher elasticity modulus. The results from Fig. 5 seem to be contradictory to those in Fig. 3. As we know, T_g is mainly decided by the inherent chemical structure of the polymer, and the mechanical properties of porous scaffolds mainly depend on their microstructure, pore size and porous morphology.^45–47 In this study, the mechanical properties of the scaffolds can mainly be attributed to their special porous structure not to T_g. Hence, the above results are reasonable.

Fig. 4 Proliferation of MC3T3-E1 cells incubated on different scaffolds for 1, 3 and 5 days, as indicated by the MTT assay (n = 3).

Fig. 5 CLSM micrographs of MC3T3-E1 cells cultured on PBLG (a–c) and PBLG/BG (d–f) scaffolds for 1 (a and d), 3 (b and e) and 5 (c and f) days.

The polypeptide derivative scaffolds were expected to be degradable. In order to reveal the degradation behaviour of the polypeptide derivative scaffolds, their degradation rates were measured in PBS at 37 °C. At predetermined time intervals, the polypeptide-based scaffolds were taken out from the degradation media, dried and weighed, and the results are shown in Fig. 3d. Overall, all scaffolds showed slow degradation rates and the largest weight loss approximated 23.7% after soaking in PBS for 8 weeks, indicating that all scaffolds had slowly degraded. Other researchers also found that the polypeptide derivatives prepared from NCAs underwent degradation very slowly via the breakdown of amido bonds due to being quite stable against hydrolysis in PBS.⁴⁸ Moreover, the degradation rate of polypeptide derivative scaffolds gradually decreased with the increasing content of PA units. The reason for this may be closely related to the hydrolysis of benzyl ester groups.⁴⁹ As is well known to us, the ester bonds show poor stability and degrade faster than amido bonds under the same circumstance. The lower the content of BLG units in the polypeptide is, the slower the degradation rate of the scaffolds will be.

3.4 MTT assay

To evaluate the in vitro biocompatibility of the polypeptide-based scaffolds, mitochondrial activity of MC3T3-E1 cells grown on the polypeptide derivative scaffolds was measured using the MTT assay, and the results are displayed in Fig. 4. It was found that, for all scaffolds, the absorbance at 490 nm increased as the culture period was prolonged from 1 day to 5 days. The results suggested a significant increase in the cell population, confirming the great proliferative potential of MC3T3-E1 cells on the polypeptide-based scaffolds. After 3 or 5 days of culture, there was a significant difference in the absorbance values between PBLG and P(BLG-PA) scaffolds. Moreover, a PBLG/BG composite scaffold always displayed a higher absorption value than a PBLG scaffold at the same time.

The metabolic activity of MC3T3-E1 cells is related to the physicochemical properties of scaffolds, such as T_g, contact angle, and compression modulus. As indicated in Fig. 3, by comparing the T_g of the polypeptide derivatives with the optimal temperature for cell culture of 37 °C, PBLG, P(BLG7-PA3) and P(BLG5-PA5) scaffolds during the culture period was in the elastomeric state, glass transition region and glassy state, respectively. Some researchers⁵⁰ asserted that the elasticity of scaffolds could effectively manipulate the behaviour of cells. The contact angle has an obvious effect on the cell proliferation and differentiation within a certain range.⁵¹ Compared to P(BLG7-PA3) and P(BLG5-PA5) scaffolds, the PBLG scaffold displayed a lower water contact angle, which facilitated cell adhesion. In contrast to P(BLG7-PA3) and P(BLG5-PA5) scaffolds, the PBLG scaffold with a higher compression modulus showed higher mitochondrial activity, which is consistent with the reported result.⁵² Furthermore, the PBLG/BG composite scaffold displayed a higher optical absorption than the PBLG scaffold. The reason for this is that the addition of BG particles to PBLG could improve the surface roughness of struts and enhance the adhesion, spreading and proliferation of MC3T3-E1 cells.^53,54 Consequently, we can draw the conclusion that the cell growth behaviour on a given substrate depends on multiple factors, including T_g, wettability, stiffness, roughness, etc.

3.5 Live/dead cell viability assay

The viability and proliferation of MC3T3-E1 cells cultured on the PBLG and PBLG/BG composite scaffolds for 1, 3 and 5 days were further studied by the live/dead cell viability assay. Live cells were stained green with calcein AM while dead cells were stained red with ethD-1. As shown in Fig. 5, green fluorescent points representing viable cells were uniformly distributed on the scaffolds. After 1 day of incubation, a few green points existed in the PBLG and PBLG/BG scaffolds (Fig. 5a and d), which indicated that some cells had been stuck to the scaffolds. More and more green points were observed on the scaffolds as the culture time was prolonged to 3 and 5 days (Fig. 5b, c, e and f), which implied that the seeded cells had begun to proliferate. A greater cell population and a higher viability ratio were observed on the PBLG/BG composite scaffolds compared to the PBLG scaffolds at the same time (3 or 5 days). Quantitatively, after 5 days of culture, the percent of live cells in PBLG and PBLG/BG scaffolds was up to 94.1% and 93.8%, respectively. This result is in accordance with that of the MTT assay, further demonstrating the positive effect of the BG addition on the proliferation ability of cells on the polypeptide-based scaffolds.⁵³

3.6 ALP activity assay

ALP activity was evaluated, and the levels of ALP activity are shown in Fig. 6. ALP is generally considered as an early biomarker for osteoblastic differentiation of pre-osteoblasts.⁵⁵ It can be seen from Fig. 6 that the ALP activity of MC3T3-E1 cells cultured on the polypeptide-based scaffolds increased with the increasing culture time. However, there were no obvious significant differences in ALP activity between PBLG and P(BLG-PA) scaffolds (p > 0.05) as the culture period was prolonged from 1 day to 5 days. After 3 and 5 days of culture, the ALP activity of MC3T3-E1 cells grown on PBLG/BG composite scaffolds was significantly higher than that on PBLG scaffolds (p < 0.05), which is possibly due to the ability of BG to induce osteoblastic differentiation of MC3T3-E1 cells towards osteoblasts.^56,57

Fig. 6 ALP activity of MC3T3-E1 cells cultured on different scaffolds for 1, 3 and 5 days (n = 3).

3.7 Morphology of MC3T3-E1 cells on scaffolds

Based on the results of physicochemical analysis as well as MTT and ALP activity assays, PBLG scaffolds rather than P(BLG-PA) scaffolds showed better elasticity, a more favorable contact angle and higher compression modulus, and could better support the proliferation and osteoblastic differentiation of MC3T3-E1 cells. Consequently, PBLG scaffold and PBLG/BG composite scaffolds were chosen for further in vitro and in vivo evaluation.

To study the morphology of MC3T3-E1 cells on PBLG and PBLG/BG composite scaffolds, the scaffold/cell constructs were fixed, dried, gold sputtered and observed using SEM. Fig. 7 displays the SEM images of the cells cultured on PBLG and PBLG/BG composite scaffolds for 1, 3 and 5 days. As revealed in Fig. 7a and d, after 1 day of culture, the cells attached on the inner wall of the scaffolds and exhibited a spindle shape. After 3 days of culture (Fig. 7b and e), the cells showed a triangular shape. When the culture time was prolonged to 5 days, cells displayed star-like and polygonal shapes (Fig. 7c and f), and even some bridged over the micropores. These results implied that the two kinds of scaffolds could support cell attachment and growth. It was also noted that the cells on the PBLG/BG composite scaffolds exhibited a better spreading morphology than those on the PBLG scaffolds. This phenomenon suggested that cell morphology was closely related to both the good cytocompatibility of BG⁵³ and the great surface roughness caused by BG particle addition.⁵⁴

Fig. 7 SEM images of MC3T3-E1 cells cultured on PBLG (a–c) and PBLG/BG (d–f) scaffolds for 1 (a and d), 3 (b and e) and 5 (c and f) days. Cumulative release curves of Ca²⁺ and Si⁴⁺ for PBLG/BG composite scaffolds soaked in PBS (g).

The change of concentrations of Ca²⁺ and Si⁴⁺ released from PBLG/BG composite scaffolds in PBS was monitored for 28 days (Fig. 7g). It was found that Ca²⁺ and Si⁴⁺ concentrations showed a moderate rise within the first 5 days and a sharp increase after 7 days. After 14 days, the overall ion concentrations became steady. By comparing with the release rates of Ca²⁺ and Si⁴⁺ reported in the literature,^58,59 PBLG/BG composite scaffolds showed a relatively lower release rate. This is attributed to the slower degradation kinetics of PBLG. Only BG particles in the PBLG exposure to PBS solution began to dissolve and the corresponding ions could be detected.

3.8 H&E and Masson's trichrome staining

To study the tissue response to PBLG and PBLG/BG composite scaffolds in a SD rat model, histological analysis after 4 and 8 weeks of implantation was carried out, and both H&E and Masson's trichrome staining images are presented in Fig. 8. The residual material and ingrown microvessels could be easily recognized at different stages. The microvascular structures are indicated with dark arrows. As shown in Fig. 8a1 and b1, massive fibrillar connective tissue that was stained pink existed, and the scaffold material still filled most of the space, indicating that scaffold materials had barely degraded after 4 weeks of implantation. Microvascular structures appeared in both the PBLG and PBLG/BG composite scaffolds. In addition, the presence of large numbers of inflammatory cells suggested the occurrence of an inflammatory response. After 8 weeks of implantation (Fig. 8a3 and b3), the number of microvascular structures in the scaffolds obviously increased, and scaffold materials started to degrade. At the same time, the number of mononuclear inflammatory cells decreased and erythrocytes emerged, indicating a marked decrease in the inflammatory response. It can be found from the Masson's trichrome staining images after 4 weeks of implantation (Fig. 8a2 and b2) that many microvascular structures appeared around and in the scaffolds and that the number of microvascular structures in the PBLG/BG composite scaffolds was much higher than that in the PBLG scaffolds, suggesting that the BG facilitated vascular ingrowth. After 8 weeks of implantation, a large number of erythrocytes were found in the scaffolds, which was accompanied by a decrease in the number of inflammatory cells (Fig. 8a4 and b4).

Fig. 8 H&E (a1, a3, b1 and b3) and Masson's trichrome staining (a2, a4, b2 and b4) of PBLG (a1 and a2 for 4 weeks; a3 and a4 for 8 weeks) and PBLG/BG composite scaffolds (b1 and b2 for 4 weeks; b3 and b4 for 8 weeks).

In a word, after 4 weeks of implantation, both PBLG and PBLG/BG composite scaffolds barely degraded and an obvious inflammatory response was observed. In contrast, after 8 weeks of implantation, the degradation of both PBLG and PBLG/BG composite scaffolds was observed, meanwhile the inflammatory response significantly decreased, suggesting that the degradation scaffolds attracted more erythrocytes and promoted vascular ingrowth more efficiently than the PBLG scaffolds, indicating that the incorporation of BG into PBLG released in vivo inflammation,^53,60 which are consistent to the results of in vitro evaluation.

3.9 CD34 staining

In order to further evaluate the ability of ingrown microvessels into PBLG and PBLG/BG composite scaffolds in vivo, the expression of CD34 was detected by immunohistochemical staining after subcutaneous implantation for 4 and 8 weeks. CD34 is highly expressed on hematopoietic progenitors and endothelial cells. As can be seen in Fig. 9, microvessels appeared in the scaffold/tissue construct after 4 weeks of implantation. And the number of microvessels significantly increased after 8 weeks of implantation, demonstrating good microvascular ingrowth capacity of PBLG and PBLG/BG composite scaffolds. Significantly, the results confirmed that the PBLG/BG composite scaffolds could better promote vascular ingrowth than the PBLG scaffolds, which coincided with the results of H&E and Masson's trichrome staining.

Fig. 9 CD34 staining of PBLG and PBLG/BG composite scaffolds after 4 and 8 weeks of implantation.

3.10 Digital radiography

Fig. 10 exhibited the radiographs of the bony defects with implants of PBLG and PBLG/BG, respectively taken after 1 week and 8 weeks post-operatively. The images in Fig. 10 show that there were still obvious quadrate holes at the tail end of tibia after 1 week of implantation. After 8 weeks of implantation, the bone defects implanted with PBLG scaffold exhibited a partial yet homogeneous filling of bony structure. Moreover, almost complete filling of the bony defects was observed from the border to the center of the holes after implanted with PBLG/BG composite scaffolds. Quantitatively, the mean percentages of newly formed bone filling the segmental defect region was about 50% and 80% for PBLG and PBLG/BG groups, respectively. Therefore, the PBLG/BG group showed more bone filling in the defects compared with the PBLG group. This is related to the osteoinductivity of BG.⁶¹ Taken together, PBLG/BG composite scaffolds showed enhanced angiogenic activity and influence osteogenesis.

Fig. 10 Representative digital radiographs of tibial defects with PBLG and PBLG/BG composite scaffolds at 1 week and 8 week post-operatively.

4. Conclusions

In summary, foamy PBLG, P(BLG-PA) and PBLG/BG composite scaffolds were successfully fabricated via a combination of the sintered NaCl templating method and ROP of α-amino acid N-carboxyanhydrides. The PBLG scaffolds had a lower glass transition temperature and water contact angle, higher compression modulus, a slightly higher degradation rate and better cell viability than the P(BLG-PA) scaffolds. The addition of BG into PBLG enhanced the adhesion, proliferation and differentiation of preosteoblastic MC3T3-E1 cells. After 4 and 8 weeks of in vivo implantation, the results from H&E, Masson's trichrome and CD34 staining revealed that the PBLG/BG composite scaffolds exhibited greater tissue infiltration and vascular ingrowth than PBLG scaffolds. Based on the results of bone defect repair experiments, PBLG/BG composite scaffolds significantly enhanced early-stage bone formation after 8 weeks of implantation. To sum up, the PBLG/BG composite scaffolds displayed good in vitro biocompatibility, little in vivo inflammation and good osteogenesis, and could be a promising bioactive material for enhancing bone regeneration.

Acknowledgements

This work was supported by the National Natural Science Foundation of China (50773062, 50603020) and the Natural Science Basic Research Plan in Shaanxi Province of China (2013K09-27) and the Fundamental Research Funds for the Central Universities (XJJ2014124, XJJ2012146).

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