Bioinspired Design of Polycaprolactone Composite Nanofibers as

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Bio-inspired design of polycaprolactone composite nanofibers as artificial bone extracellular matrix for bone regeneration application Xiang Gao, Jinlin Song, Yancong Zhang, Xiao Xu, Siqi Zhang, Ping Ji, and Shicheng Wei ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b10417 • Publication Date (Web): 30 Sep 2016 Downloaded from http://pubs.acs.org on October 1, 2016

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ACS Applied Materials & Interfaces

Bio-inspired design of polycaprolactone composite nanofibers as artificial bone extracellular matrix for bone regeneration application Xiang Gaoa, b, c, Jinlin Songa, b, c, Yancong Zhangd, Xiao Xud, Siqi Zhange, Ping Ji*a, b, c

,and Shicheng Wei*b, c, d, e

a

College of Stomatology, Chongqing Medical University, Chongqing 401147, China

b

Chongqing Key Laboratory of Oral Diseases and Biomedical Sciences, Chongqing

Medical University, Chongqing 401147, China c

Chongqing Municipal Key Laboratory of Oral Biomedical Engineering of Higher

Education, Chongqing 401147, China d

Department of Oral and Maxillofacial Surgery, Central Laboratory, Peking University

School and Hospital of Stomatology, Beijing, 100081, China e

Center for Biomedical Materials and Tissue Engineering, Academy for Advanced

Interdisciplinary Studies, Peking University, Beijing, 100871, China Corresponding Author *Prof. Shicheng Wei, Chongqing Key Laboratory for Oral Diseases and Biomedical Sciences, Chongqing Medical University, Chongqing 401147, China Tel & Fax: +86 23 88860222. E-mail: [email protected] *Prof. Ping Ji, Chongqing Key Laboratory for Oral Diseases and Biomedical Sciences, Chongqing Medical University, Chongqing 401147, China Tel & Fax: +86 23 88860222. E-mail: [email protected]

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Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Notes The authors declare no competing financial interest.

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ABSTRACT: The design and development of functional biomimetic systems for programmed stem cell response is a field of topical interest. To mimic bone extracellular matrix, we present an innovative strategy to construct drug-loaded composite nanofibrous scaffolds in this study, which could integrate multiple cues from calcium phosphate mineral, bioactive molecule and highly ordered fiber topography for control of stem cell fate. Briefly, inspired by mussel adhesion mechanism, a polydopamine (pDA)-templated nano-hydroxyapatite (tHA) was synthesized and then surface functionalized with bone morphogenetic protein-7-derived peptides via catechol chemistry. Afterwards, the resulting peptide-loaded tHA (tHA/pep) particles were blended with polycaprolactone (PCL) solution to fabricate electrospun hybrid nanofibers with random and aligned orientation. Our research demonstrated that the bioactivity of grafted peptides was retained in composite nanofibers. Compared to controls, PCL-tHA/pep composite nanofibers showed improved cytocompatibility. Moreover, the incorporated tHA/pep particles in nanofibers could further facilitate osteogenic differentiation potential of human mesenchymal stem cells (hMSCs). More importantly, the aligned PCL-tHA/pep composite nanofibers showed more osteogenic activity than randomly-oriented counterparts even under non-osteoinductive conditions, indicating excellent performance of biomimetic design in cell fate decision. After in vivo implantation, the PCL-tHA/pep composite nanofibers with highly ordered structure could significantly promote the regeneration of lamellar-like bones in a rat calvarial

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critical-sized defect. Accordingly, the presented strategy in our work could be applied for a wide range of potential applications not only in bone regeneration application, but also in pharmaceutical science. KEYWORDS: biomimetic, apatite, peptide, nanofiber, bone extracellular matrix

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1.

Introduction Bone loss caused by trauma, disease and congenital anomalies remain a serious

fast-growing challenge in the medical area, and the associated annual healthcare expenditures are estimated to exceed tens of billions of dollars during the next decades.1-2 Usually, the application of autologous graft and transplantation of allograft are regarded as the standard procedures for management of large bone defects. However, some problems, such as limitation of donor source, immunologic rejection and risk of cross infection, limit their clinical potential.3 Recently, with the development of stem cell-based tissue engineering (STE), the combination of stem cells with a bioactive scaffold provides a promising therapeutic alternative for reconstruction and repair of serious bone defect.4 Nevertheless, the clinic applications of STE therapy still faces many obstacles. One of the glaring challenges for scientists and engineers is that the conventional scaffold design is short of biomimicry and mismatches in the physical and chemistry properties of natural microenvironment where stem cells reside in, ultimately leading to the failure in the precise regulation of programmed stem cell response for regenerative medicine.5 To develop biomimetic scaffolds, many processing approaches, such as phase separation, extrusion, porogen leaching, and electrospinning, have been employed to fabricate scaffolds with different constituents and three dimensional structures.6-7 Among these scaffolds, electrospun nanofibers have attracted tremendous interest in

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recent years, because its structure is close to that of bone tissue extracellular matrix (ECM).8 Furthermore, the porous electrospun nanofibers have a large surface area, and thus could also provide a desirable microenvironment for early cell adhesion and proliferation. Recently, several studies have reported that the surface topography of scaffold can contribute to the decision of stem cell fate.9-10 Many synthetic or natural degradable polymers have been explored for the fabrication of nanofibers with various fiber diameter or nanopattern,11 of which uniaxially aligned nanofibers have been widely investigated as a scaffold for bone regeneration application,12-13 due to its surface topographic similarity to the well-aligned collagen fiber bundles present in native bone tissue.14 Compared to those on randomly oriented nanofibers, stem cells cultured on aligned nanofibers can be arranged along the direction of nanofibers, and present enhanced osteogenic potential, because some osteo-related gene expression may be triggered via mechanotransduction after the cytoskeleton of cells was elongated.15 Therefore, uniaxially aligned nanofibers as a scaffold material hold a great potential in the field of bone regenerative medicine. It is widely accepted that in native microenvironment, stem cells are under the regulation of a variety of physical and chemical cues emanated from surrounding extracellular matrix.16 As far as bone ECM is concerned, it is a typical organic– inorganic nanocomposite material.17 Apart from collagenous fibers with specific surface topography, inorganic mineral nanocomponents and various bioactive factors like BMP

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proteins are also present in the bone matrix, which play a critical role in bone formation.17 Thus, these two factors should also be taken into account when designing ECM-mimicking materials. To simulate the mineral constituent of bone ECM, a traditional strategy is the incorporation of bioceramics, such as nano-hydroxyapatite,18 β-tricalcium phosphate19 and silica,20-21 into nanofibers due to the versatility of the electrospinning technique. Given the intrinsic osteoconduction and osseointegration potential, those bioinorganic materials endow the polymeric nanofibers with enhancd osteogenic bioactivity. However, because of the absence of growth factors and cytokines, the important regulators of stem cell functions, in polymer matrix, the composite nanofibers produced by purely incorporating bioceramics into nanofibers are still far from the ideal artificial osteogenic niche, resulting in limited success. Therefore, recent research efforts have been focused on developing an approach to recapitulate the complexities of native bone ECM in the nanofibrous materials.8 To this end, one reasonable strategy is to load the target biomolecules surface into the inorganic nanoparticles before co-electrospinning with polymers. For instance, previous studies have shown that nano-hydroxyapatite (nano-HA; Ca10(PO4)6(OH)2), the main mineral component of bone matrix,22 can not only be used to improve physicochemical and biological properties of polymer nanofibers via blending approach due to its high surface activity and good biocompatibility,17 but also be applied as a drug carrier via direct physical adsorption to allow the controlled drug release in degradable

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fiber matrix, because the high surface area endows the nanoparticles with strong adsorbability.23-24 However, the interaction between physically adsorbed drugs and nano-HA particles is weak, the problem of burst release is likely to arise and thereby the long term effectiveness of adsorbed biomolecules can not be guaranteed. Moreover, the physical adsorbed bioactive factors may also denature in organic solvents and lose their bioactivity during blend electrospinning.8 To address these issues, covalent immobilization of bioactive factors associated with osteogenesis onto the surface of nano-apatite particles is a promising strategy for further development of advanced biomimetic materials targeted at bone regeneration application. Bone morphogenetic protein-7 (BMP-7), which belongs to the transforming growth factor-β superfamily, has been proved to be involved in the regulation of the growth and osteogenic differentiation of hMSCs at the biomaterial-tissue interface.25-26 Nevertheless on account of the complex secondary structure, BMPs is highly vulnerable to degradation in aqueous physiological environment, ultimately leading to the loss of bioactivity.27 Therefore, short functional peptides derived from BMPs have attracted widespread interest because of their better structural stability and less immunogenicity in comparison with their natural counterparts.28 Recently, kim et al. reported a new peptide sequence (GQGFSYPYKAVFSTQ), named bone forming peptide-1 (BFP-1), from the immature region of BMP-7, which showed more osteogenic activity than its derived protein in vitro and in vivo.29 Therefore, the combination of BFP-1 loaded

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apatite nanoparticles with aligned nanofibers is believed to simultaneously provide multiple cues for cell fate decision. Guided by the considerations detailed above, a new type of BFP-1 loaded composite nanofibers with enhanced bioactivity was developed in our work. For surface functionalization of apatite nanoparticles with osteogenic peptides, a novel apatite crystal was firstly synthesized using polydopamine (pDA) as template, and the BFP-1 peptides were grafted onto the pDA functionalized nano-apatite via catechol chemistry. Afterwards, the BFP-1-loaded nano-apatite was mixed with polycaprolactone (PCL, a model synthetic polymer) solution for subsequent fabrication of composite nanofibers with highly ordered structure via electrospinning. To our knowledge, this is the first report on the design of bone ECM mimic materials, which could simultaneously provide signaling from osteogenic peptides, apatites and aligned nanofibers for directing the osteogenesis of hMSCs in a sustainable manner. The loading of BFP-1 onto the pDA functionalized nano-apatite and the composite nanofibers were characterized in this study. Additionally, the cytocompatibility and osteogenic activity of the composite nanofibers were also systematically investigated in vitro and in vivo. We believe the design strategy presented in this work could pave a new way to develop advanced biomimetic scaffolds targeted at bone regeneration application. 2. 2.1

Materials and methods Preparation of pDA-templated nano-HA and peptide decorated nano-HA

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For synthesis of pDA-templated nano-HA (tHA), dopamine (Sigma-Addrich, USA) was added to Ca(NO3)2 (Sinopharm, China) solution at a concentration of 2 mg/ml. According to the Ca/P molar ratio of 1.67, an amount of Na2HPO4 (Sinopharm, China) was slowly dropped into Ca (NO3)2-dopamine mix solution with continuous stirring. The pH of mix solution was maintained to about 8.5 via dripping Tris-HCl (10 mM, Aladdin, China) during the synthesis process. The reaction was proceeded for 12 h at 60 ℃. Then, the slurry was aged for 24 h at 37℃, and the impurities were removed by thorough washing with distilled water (DW) and ethanol, respectively, followed by dialysis in DW for at least 3 days to eliminate unreacted dopamine monomers. The obtained precipitate was air dried at 60°C for future use. For surface functionalization of tHA nanoparticles with BFP-1 peptides (KGGQGFSYPYKAVFSTQ sequence, China peptides Co., Ltd), the prepared peptide solution was gradually dripped into an aqueous suspension of tHA powder. After vigorous stirring for 24 h at 37 ℃, the peptide-decorated tHA nanocomposites (tHA/pep) were formed in solution. Then, the products were separated by centrifugation (10000

g,

5 min) and the supernatants were collected to quantify the

immobilized-peptides on tHA nanoparticles by fluorescamine assay. Briefly, the supernatant (75 µl) was transferred into a 96-well plate after reaction. Then, 25 µl of 100 µg/ml fluorescamine solution (Sigma-Addrich, USA) in acetone was added. After 5 min at RT, the fluorescent intensities at 470 nm (excitation at 365 nm) for each solution

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were recorded using a Multilabel Reader (Perkin Elmer, USA). The peptide concentration in the supernatant was obtained from a standard fluorescent intensity-concentration calibration curve. And the grafted peptides was quantified by subtracting the amount of unreacted BFP-1 in the supernatant from the sum of peptides. The loading percentage of peptides was determined based on the following equation: Loading percentage=M1/( M1 + M0)×100%

(1)

Where M1 and M0 stand for the mass of immobilized BFP-1 peptides and the tHA carrier, respectively. The loading percentage of peptides was optimized by changing the concentrations of BFP-1 and tHA. Finally, the tHA/pep particles were lyophilized and stored at -20 ℃ for future use. 2.2

Preparation of electrospun PCL-tHA/pep composite nanofibers PCL

(MW=80000

g/mol,

Sigma-Addrich,

USA)

was

dissolved

in

Hexafluoro-2-propanol (HFIP, Sigma-Addrich, USA) to a final concentration of 10 %. Then, tHA/pep powder (10.8 w.t %) was suspended in PCL solution by ultrasonic and vigorous stirring, and the mixture solution was loaded into a 10 ml syringe with a 23-gauge metal needle for subsequent electrospinning. To obtain randomly oriented composite nanofibers, a flat collector was placed 10 cm away from the needle tip. The feeding rate and the applied voltage was set at 1 ml/h and 14 kV, respectively. On the other hand, aligned composite nanofibers were collected on a metal roller with a rotation speed of 2800 rpm. The distance between roller and the tip of needle was set at 5 cm.

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The applied voltage was 10 kV and flow rate was 2 ml/h..The formed electrospun meshes were dried in a vacuum oven at room temperature (RT) for at least two days to remove

residual

organic

solvent,

as

previously

described.30

Moreover,

randomly-oriented PCL (R-PCL) and tHA (10 w.t%)-doped PCL nanofibers (R-PCL-tHA) were fabricated as control. To

visualize

the

tHA/pep

nanoparticles

in

the

PCL nanofibers,

the

rhodamine-labeled BFP-1 peptides were immobilized on the surface of tHA. Then, the rhodamine-labeled tHA/pep nanoparticles were used to prepare the composite nanofibers, and the final samples were observed under laser scanning confocal microscope (LSCM, Carl Zeiss). 2.3

Characterization The morphology of peptide-decorated tHA and the presence of tHA/pep

nanoparticles in PCL nanofibers were characterized by transmission electron microscope (TEM, H-9000, Hitachi) at an accelerating voltage of 100 kV. The surface morphology of nanofibers was also examined by field-emission scanning electron microscopy (FE-SEM, S-4800, Hitachi) after being gold-coated, and 100 fibers in SEM images were randomly selected for the measurement of fiber diameter by ImageJ software. The crystalline phase of tHA nanoparticles was examined by X-ray diffractometer (XRD) equipped with Cu radiation source (λ= 1.54 Å) at 40 kV between 2θ of 10° and 60°. X-ray photoelectron spectroscopy (XPS; AXIS Ultra, Kratos

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Analytical Ltd.) was utilized to analyze the surface chemical composition of samples. A SL200B type contact angle goniometer (Kono, USA) was used to determine the water contact angle (WCA) of nanofibers under ambient temperature and humidity. DW was applied as the media. The elastic modulus of the electrospun meshes was determined by universal testing machine (UTM, Instron 5900) at a crosshead speed of 10 mm/min under a load of 10 N. Protein adsorption assay was performed by incubating FITC-labeled bovine serum albumin (FITC-BSA, 100 µg/ml, Sigma) with samples for 60 min at 37 ℃. To quantify the absorbed proteins on nanofibers, the fluorescent intensity of supernatant for each sample was recorded at 488 nm (excitation at 525 nm) by a Multilabel Reader. The amount of absorbed BSA was calculated from a standard fluorescent intensityconcentration calibration curve.. Afterwards, samples were rinsed with PBS and captured under LSCM to visualize the attached proteins. 2.4

Cell culture For in vitro studies, the prepared nanofibers were punched into round shape.

Afterwards, the circular samples were sterilized with 70% ethanol for 1 hour and then rinsed with disinfected D-Hanks buffer before seeding cells. hMSCs, which was purchased from Sciencell Research Laboratories, were expanded in normal growth media composed of α-MEM media (Gibco), 10 % fetal bovine serum (FBS, Gibco), 1 % penicillin-streptomycin (P-S, Gibco) and 2 mM L-glutamine (Gibco) in a humidified

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atmosphere of 5 % CO2 at 37 °C. After 24 h of cell seeding on nanofibers, the media was replaced with fresh normal growth media or osteoinductive media (α-MEM media, 10 % FBS, 1 % P-S, 2 mM L-glutamine, 10 mM β-glycerophosphate, 50 µg/ml ascorbic acid, 100 nM dexamethasone). The culture media were refreshed every 2-3 days. 2.5

Cytocompatibility evaluation of PCL-tHA/pep composite nanofibers

2.6.1 Early adhesion assay The early adhesion of hMSCs on sample surfaces were examined by SEM and fluorescein diacetate (FDA, Sigma-Addrich) staining. After 2 and 4 h of incubation, samples were retrieved and fixed in glutaraldehyde (2.5 %) for 2 h. Afterwards, the samples were rinsed with DW and dehydrated with graded concentrations of ethanol (30-100 %). Finally, the samples were thoroughly dried by hexamethyldisilazane and visualized under SEM after being gold-coated. For FDA staining, the rinsed cells were immersed in 5 µg/ml fluorescein diacetate (FDA,) solution. After incubation for 5 min, cells on samples were thoroughly washed and viewed under LSCM. To quantify the early adherent cells on each sample, the trypsinized cells from sample surfaces were counted by haemocytometer at 2 h after seeding. 2.6.2 Cell proliferation assay The cell counting assay kit-8 (CCK-8, Dojindo) was applied to evaluate the cell proliferation on sample surfaces according to the manufacturer’s instruction. At 1, 5, 7 days after seeding, 10 % CCK-8 solution in media was added to each well. After 4 h at

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37 ℃, the OD value for each well was recorded with a plate reader (Model 680, Bio-Rad) at 450 nm. 2.7 Osteogenic bioactivity evaluation of PCL-tHA/pep composite nanofibers 2.7.1 ALP assay Intracellular release of Alkaline phosphatase (ALP) on sample surfaces was assessed using a ALP assay kit (Nanjing Jiancheng, China). After 14 days of incubation, cells cultured on sample surfaces were lysed with Triton X-100. Then, cell lysate was transferred to a 96-well plate (30 µl/well), and mixed with working solution following the manufacturer’s instruction. Absorbance of reaction product was recorded using a plate reader at 405 nm. The results were normalized by the total protein amount which was determined by BCA assay kit. The ALP distribution on sample surfaces was visualized using BCIP/NBT ALP color development kit (Beyotime, China) following the manufacturer’s instruction. 2.7.2 ARS staining assay After hMSCs were exposed to the prepared nanofibers in culture media for 21 days, the samples were fixed in ethanol (95 %) and then stained with Alizarin Red S (ARS) solution (2 %, pH 4.2) for 10 min at RT. After thorough rinse with DW, the images of samples were taken via scanner. For quantification, the stained samples were treated with 10% cetylpyridinium chloride solution. After incubation for 1 h, the OD value for each sample was determined at 550 nm. The meshes (without seeding cells) in each

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group were also stained with ARS solution as control. 2.7.3 Quantitative real-time polymerase chain reaction (qRT-PCR) At 14 days, RNA isolation from samples was conducted by Trizol treatment. Then, the extracted RNA was reverse transcribed with RevertAid™ First Stand cDNA Synthesis Kit (Fermentas, Canada) per manufacturer’s protocol. qRT-PCR was performed by ABI PRISM 7500 sequence detection system (Applied Biosystems, CA, USA) using SYBR Green PCR Master Mix. Each specimen was run in triplicate. The primer used in this study were: runt-related transcription factor 2 (Runx2), (forward) 5´-AGGAATGCGCCCTAAATCACT-3´ and (reverse) 5´-ACCCAGAAGGCACAGA CAGAAG-3´; type I collagen alpha 1 (Col1a1), (forward) 5´-AGACACTGGTGC TAAGGGAGAG-3´ and (reverse) 5´-GACCAGCAACACCATCTGCG-3´; osteocalcin (OCN), (forward) 5´-CCTGAAAGCCGATGTGGT-3´ and (reverse) 5´-AGGGC AGCGAGGTAGTGA-3´; (forward)

glyceraldehyde-3-phosphate

dehydrogenase

5´-CGACAGTCAGCCGCATCTT-3´

and

(GAPDH), (reverse)

5´-CCAATACGACCAAATCCGTTG-3´. mRNA folds were calculated using the standard ∆∆Ct method with GAPDH assigned as a endogenous control. Untreated cells in normal growth media were assigned as control. 2.7.4 Immunofluorescence staining For immunofluorescence staining, hMSCs on each sample were fixed with 4 % paraformaldehyde at preset time point. After incubation for 10 min at RT , cells were

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thoroughly washed with PBS and permeabilized by 0.1 % Triton X-100 for 5 min. To minimize non-specific binding, cells were treated with 3 % BSA solution for 2 h at 37 ℃. After exposure to the primary antibody against OCN (1:100, Abcam), ALP (1:100, Santa Cruz), paxillin (1:200, BD), osteopontin (OPN, 1:500, Abcam) and Runx2 (1:100, CST) at 4 ℃ overnight, cells were rinsed three times with PBS, and stained with FITC-488 goat anti-rabbit (1:1000, CST) or TRITC-543 goat anti-mouse (1:1000, CST) secondary antibodies for 1 h at RT. Finally, the cells were counterstained with DAPI for staining of cell nuclei and visualized under LSCM. 2.7.5 Western blot At 14 days, cells on samples were treated with RIPA buffer containing a protease inhibitor cocktail, and the total protein concentration of each specimen was quantified using BCA assay kit. After SDS-PAGE, the proteins were separated and transferred to a PVDF membrane (Millipore). Afterwards, the membrane was blocked with 5 % skim milk and incubated with Runx2 (1:3200, CST) and Col1a1 (1:1000, Santa Cruz) antibodies overnight at 4℃. Then, the menbranes were treated with HRP-conjugated secondary antibodies and developed using an ECL detection kit (Amersham). GAPDH (1:1000, CST) antibody was used as an internal control. 2.8 In vivo experiment 2.8.1 Establishment of rat calvarial defect model The animal experiment was approved by the the Animal Ethics Committee of the

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Peking University (Approval no: LA2015149). To evaluate the effect of tHA/pep doped nanofibers on in vivo bone formation, 8-week-old Sprague-Dawley rats were applied for in vivo experiment. Prior to operation, the animals were anesthetized with pentobarbital sodium (1%, 50 mg/kg) by intraperitoneal injection. Two critical-sized bone defects (5 mm diameter) were prepared on both sides of the cranium using a dental trephine, as previous literatures reported.31-32 Afterwards, the right defects (n=5 for each group) were implanted with nanofibers, while the untreated left defects were used as negative control. The whole calvarias were harvested for further evaluation after 8 weeks. To label the mineralization front of the regenerated bone tissues, some animals were injected subcutaneously with calcein (10 mg/kg, Sigma) and ARS (20 mg/kg, Sigma) at 2 weeks and 1 week prior to euthanasia, respectively. 2.8.2 Radiographic analysis Following animal sacrifice, the skull bone of each rat was extracted and fixed in 4% paraformaldehyde for two days. Subsequently, Soft X-ray examination was applied to assess the new bone formation at the defect sites under the fixed conditions (24 kV, 2 mA, 90 s). To observe the three-dimensional morphology of bone growth, the specimens were further scanned using Micro-CT (Inveon MM CT, SIEMENS). And the bone volume fraction (Bone volume/total volume, BV/TV) of the defect region was measured by Inveon Research Workplace software (SIEMENS, USA). 2.8.3 Histological analysis

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For processing decalcified paraffin sections, the fixed samples were treated with 10 % EDTA solution (pH 7.4) for 4 weeks. After dehydration, the decalcified samples were embedded in paraffin and 5-µm transverse sections were used for HE and Sirius Red staining. The images were captured under a polarizing microscope. For processing undecalcified resin sections, the fluorochrome-labeled samples were dehydrated through an ascending ethanol series and subjected to resin processing/embedding in methylmethacrylate. Then, the unstained transverse sections (100 µm) were cut for detection of fluorochrome labels under LSCM. The excitation wavelengths used for observation of chelating fluorochromes were 543 nm and 488 nm for ARS (red) and Calcein (green), respectively. 2.8.4 Immunohistochemical analysis After rehydration of sections, 3 % hydrogen peroxide was used to quench endogenous peroxidase activity in the deparaffinised sections. Then, the sections were thoroughly washed and treated with citrate buffer (0.01 M, pH 6.0) for antigen retrieval. Afterwards, the sections were treated with primary anti-body against OCN (1:100, Abcam). After incubation at 4°C overnight, the sections were rinsed and treated with HRP-conjugated secondary antibody for 60 min at RT, followed by the development of color with a DAB assay kit. Finally, all sections were counterstained with hematoxylin and captured under a light microscope (CX21, Olympus). The quantification of OCN was carried out by a web application software ImunoRatio.33

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2.8.5 Mechanical Characterization A nanomechanical test system (TI-900 Tri-boIndenter) was applied to assess the mechanical features of the newly formed bone in defects through nanoindentation experiments, as previously described.34 Before this evaluation, the fixed samples were embedded with acrylic resin. Then, the polished 2-mm sections were prepared and then subjected to nanoindentation tests by applying the loading rates of 10 nm/s. The loading depth was set at 500 nm. The aligned samples were sectioned vertical and horizontal to the fiber orientation. The contact hardness and the elastic modulus of target tissue were obtained from the indentation force-displacement curves. A total of 9 indentations (6 for newly formed bone, 3 for host bone tissue) were recorded for each sample. 2.9 Statistical analysis With the help of SPSS software, the data were analyzed among groups using. One-way ANOVA and Tukey’s post hoc test for multiple comparisons. A value of P < 0.05 was considered statistically significant. All experiments were conducted in triplicate. The data values were expressed as mean ± standard deviations. 2

Results and discussion

2.1 Loading of BFP-1 onto the tHA surfaces Figure 1 illustrates the synthesis process of BFP-decorated tHA nanoparticles. First, a novel pDA-templated nano-HA was developed via catechol chemistry. Under weak alkaline (PH = 8.5), dopamine molecules could experience self-polymerization and

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form pDA structure,35 which offers plentiful catecholamine moieties as Ca2+ ion binder.36 With the formation of Ca-P nucleation centers, crystals began to precipitate and grow along the surface of pDA. The ultimate product is named as tHA. As shown in Figure S1a, these nanoparticles presented a plate-like structure. Its diameter was about 31 ± 12 nm and the length was about 98 ± 22 nm, similar to the size and morphology of natural HA in bone tissue.37-38 However, it is noted that because of potent interfacial adhesion effect of pDA, tHA nanoparticles were prone to aggregate into clusters. Moreover, XRD spectra (Figure S1b) showed that most of the indexed diffraction peaks of the tHA sample matched quite well with those of pure nano-HA (PDF #09-0432), even though three intense peaks corresponding to (211), (112) and (300) became broader in synthesized tHA, suggesting the low crystallinity of tHA nanoparticles as that of minerals present in human bone.39 Due to the high chemical affinity of pDA to biomolecules containing thiols or primary amines moieties, pDA-coated biomaterials have been extensively applied in a variety of drug delivery applications. Different from published works associated with the drug encapsulation in nano-HA via physical adsorption,23-24 in this study, tHA nanoparticles were utilized to load osteoinductive peptides BFP-1 on their surface via chemical conjugation. Thus, to optimize the loading percentage of BFP-1 onto tHA, fluorescamine assay was performed. As shown in Figure 2a, it is clearly seen that at the given tHA concentration, the BFP-1 loading rate is enhanced with the increase of BFP-1

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concentrations. However, when the BFP-1 concentration is above 400 µg/ml, there is no obvious enhancement in peptide loading rate. Moreover, we observed that the BFP-1 loading rate dropped as the concentration of tHA was increased. The possible explanation is that at higher concentration, the tHA nanoparticles are prone to aggregate, resulting in the reduction of binding sites for peptides on the surface of tHA nanoparticles. Therefore, the optimal peptide loading rate (8 %) was finally obtained at the concentration of 400 µg/ml BFP-1 and 1 mg/ml tHA. So in the following preparation of composite nanofibers, the percentages of tHA and tHA/pep relative to PCL were set at 10 w.t % and 10.8 w.t %, respectively, to ensure the similar content of the incorporated tHA in both PCL-tHA and PCL-tHA/pep composite nanofibers. The microstructure of tHA/pep particles was also observed via TEM (Figure 2b). it is clearly observed that after peptide loading, the tHA still keeps its primary plate-like morphology with similar size, indicating that the surface modification of peptides does not obviously alter the tHA’s shape. To further verify the successful decoration of BFP-1 onto the surface of tHA, XPS was performed on tHA before and after BFP-1 loading. As shown in Figure 2c, d, tHA displayed peaks for C, O, Ca and P as main atomic elements. The quantification of the elementary composition (Table S1) showed that tHA nanocrystals had a lower Ca/P ratio (1.37), suggesting the calcium-deficient state on the tHA crystalline surface.40 But this type of apatite may be more instrumental for biological applications than stoichiometric ones, because the Ca/P atomic ratio in

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natural bone tissues is lower than 1.67.41 Calcium-deficient HA is able to induce instant precipitation of biologically equivalent apatite on its surface as implanted in vivo, while precipitation on stoichiometric HA needs an induction time.42 Moreover, due to the presence of pDA, a weak signal for nitrogen (N 1s) was recorded at 399 eV in the wide spectrum of tHA. When the BFP-1 peptides were grafted onto the surface of tHA, the signal for nitrogen in tHA/pep particles was significantly enhanced compared to that in tHA, which was clearly revealed in the high-resolution nitrogen spectra (Figure 2d). XPS high-resolution carbon spectra (C 1s) proved these observations as well (Figure S1c, d). Additionally, it is noteworthy that since peptides were dissolved in PBS buffer during experiment, the sodium peaks also appeared in peptide-decorated tHA, providing an indirect evidence for successful loading of peptides. 2.2 Loading of tHA/pep nanoparticles within PCL nanofibers The incorporation of tHA/pep into the PCL nanofibers was confirmed by TEM and LSCM. As shown in TEM image (Figure 3a), the plate-like tHA/pep complexes aggregate and form the nodule-like structures in the PCL nanofibers. So under LSCM (Figure 3a), many dot-shaped green fluorescence were observed in the composite nanofibers containing rhodamine-labeled tHA/pep, further corroborating the loading of BFP-1 in tHA. Moreover, it is also suggested that the immobilization of BFP-1 on the surface of tHA was not compromised during the electrospinning process, which should be ascribed to the strong chemical bonding between tHA and peptides.

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The surface morphology of PCL-tHA/pep composite nanofibers with random and aligned orientation was observed and analyzed by SEM and imageJ software (Figure 3b, c). For comparison, the morphologies of randomly oriented PCL and PCL-tHA nanofibers were also investigated. It is clearly seen that the nanofibers in R-PCL group showed uniform and smooth surface morphology without bead defects (Figure 3d). After the incorporation of tHA or tHA/pep, the fiber surfaces became rough, and plenty of nodule-like structures were detected in the composite fibers, in agreement with the observation of TEM. Moreover, the diameter of the PCL nanofibers was also markedly decreased when the nanoparticles were introduced into PCL matrix (P0.05). Considering the potential effect on cell behaviours like adhesion, proliferation, and differentiation,44-45 surface wettability and elastic modulus are two important parameters that should be taken into consideration, when designing scaffolds target at guided bone regeneration. Table S2 shows the WCA and elastic modulus of PCL-tHA/pep composite nanofibers. For comparison, the corresponding data of PCL and PCL-tHA nanofibers

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are also included. It is documented that the hydrophilicity properties of polymeric biomaterials could be improved after pDA coating.35,

46-47

. Moreover, the enhanced

roughness also contributes to the surface wettability of hydrophobic substrates.48 Therefore, due to the combined effect of pDA layers and the enhanced roughness, the surface wettability of R-PCL-tHA nanofibers was improved in comparison with R-PCL nanofibers, which may be beneficial for cell adhesion and spreading. Moreover, it is noteworthy that the water contact angle was further decreased to a certain extent on the surface of R-PCL-tHA/pep nanofibers when compared to R-PCL-tHA nanofibers, suggesting the successful modification of peptides onto tHA nanoparticles. However, R-PCL-tHA/pep and A-PCL-tHA/pep group showed no significant difference, indicating little effect of fiber orientation on the surface wettability of nanofibers. As far as mechanical property was concerned, the elastic modulus of PCL nanofibers was also improved after the incorporation of tHA, although there was no significant difference between R-PCL-tHA and R-PCL-tHA/pep nanofibers (P>0.05). When tHA/pep particles were introduced into aligned PCL nanofibers, the elastic modulus of composite nanofibers was further enhanced compared to R-PCL-tHA/pep group, which may be attributed to the high degree of orientation, similar to the previous publications.49-50 For developing a scaffold for guided bone regeneration, the materials should withstand dynamic mechanical loading in body.51 In our study, the elastic moduli of R-PCL-tHA/pep and A-PCL-tHA/pep nanofibers was in the range of 11 to 30 MPa,

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which endowed them with proper mechanical strength for bone tissue engineering.51 Additionally, the enhanced elastic modulus in tHA/pep reinforced nanofibers may be also beneficial for the osteogenic differentiation of hMSCs.44 Protein adsorption is the first event occurring at the biomaterial-tissue interface when scaffolds are implanted into body, which could significantly influence the subsequent cell behaviors like adhesion, proliferation and differentiation. Therefore, a nonspecific FITC-labeled BSA was employed to assess the bioaffinity of PCL-tHA/pep composite nanofibers to proteins in vitro. Figure 3d presents the adsorbed albumin on the surface of samples after 1 h of incubation. As a result of high specific surface area, pure PCL nanofibers induced uniform protein adsorption and showed weak green fluorescence on surface. With the introduction of tHA particles, some fluorescence aggregates with higher intensity appeared in polymer matrix. According to the quantitative data (Figure 3e), R-PCL-tHA group showed more protein content on fiber surface when compared to R-PCL group (P0.05), hMSCs grew rapidly on R-PCL-tHA nanofibers, with OD value higher than that of R-PCL group at 4 days (P

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