Porous scaffolds of poly(lactic-co-glycolic acid) and mesoporous

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Tissue Engineering and Regenerative Medicine

Porous scaffolds of poly(lactic-co-glycolic acid) and mesoporous hydroxyapatite surface modified by poly(#benzyl-L-glutamate) (PBLG) for in vivo bone repair Linlong Li, Xin Cui Shi, Zongliang Wang, Min Guo, Yu Wang, Zixue Jiao, and Peibiao Zhang ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/ acsbiomaterials.8b01614 • Publication Date (Web): 10 Apr 2019 Downloaded from http://pubs.acs.org on April 12, 2019

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ACS Biomaterials Science & Engineering

Porous

scaffolds

of

poly(lactic-co-glycolic

acid)

and

mesoporous hydroxyapatite surface modified by poly(γbenzyl-L-glutamate) (PBLG) for in vivo bone repair Linlong Li a, b, Xincui Shi a, Zongliang Wang a, Min Guo a, Yu Wang a, Zixue Jiao a and Peibiao Zhang a, b, * a

Key Laboratory of Polymer Ecomaterials, Changchun Institute of Applied Chemistry, Chinese

Academy of Sciences, 5625 Renmin Street, Changchun 130022, P. R. China. b

University of Chinese Academy of Sciences, 19 Yuquan Road, Beijing 100049, P. R. China.

* Co-corresponding author: Peibiao Zhang E-mail address: [email protected] (P. Zhang) Tel.: +86-431-85262058; Fax: +86-431-85262058.

ACS Paragon Plus Environment

ACS Biomaterials Science & Engineering 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

Abstract Three-dimensional (3D) composite porous scaffolds containing hydroxyapatite (HA) and polymer matrix showed wide applications in bone repair due to its improved biocompatibility, bioactivity and mechanical property in our previous studies. In this work, mesoporous hydroxyapatite (MHA) surface-modified by poly(γ-benzyl-Lglutamate) (PBLG) with different amounts (from 11 wt% to 50 wt%) was synthesized by the in-situ ring opening polymerization of γ-benzyl-L-glutamate N-carboxy anhydride (BLG-NCA) and then PBLG-g-MHA/PLGA composite films were prepared to illustrate the biological performance of the composites. Furthermore, porous scaffolds of PBLG-g-MHA/PLGA were fabricated through modified solvent casting/particulate leaching (SC/PL) method to demonstrate the ability of in vivo bone defect repair. In vitro cytological assay indicated that enhanced cell expansion on PBLG-g-MHA/PLGA with 11 wt% PBLG amounts and improved osteogenic differentiation on the composites with 33 wt% and 50 wt% PBLG amounts were achieved. And the porous scaffolds exhibited high porosity and interconnected pores. Results of the in vivo rabbit radius defect repair indicated that rapid mineralization and new bone formation could be observed on the composites with 22 wt% and 33 wt% amounts of PBLG. This study revealed that PBLG-g-MHA/PLGA composites might have potential applications in clinical bone repair.

Keywords: hydroxyapatite; poly(γ-benzyl-L-glutamate); porous scaffolds; osteogenic differentiation; bone repair


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1. Introduction Hydroxyapatite (HA) has been widely used as an ideal material in bone repair and replacement because of its similarity to the inorganic parts of natural bone in chemical composition and structure.1-3 Recently, mesoporous hydroxyapatite (MHA) nanoparticles have been synthesized and regarded as excellent candidates for functional protein adsorption due to its special properties, such as high surface area and large pore volume, and it might be useful in bone repair.4-5 But its brittleness and lack of processability limited its use as the implanted materials for the regeneration of bone defects.6 Synthetic polymers, such as poly(lactic acid) (PLA) and poly(lactic-coglycolic acid) (PLGA) have already been used in some biomedical applications such as bone implants, sutures and scaffolds because of their biocompatibility, biodegradability and good processability. While some practical problems still exist, such as the inflammatory responses mainly caused by in vivo degradation products of them, low efficiency of cell seeding and adhesion, and poor mechanical properties.7-9 To overcome the above shortages, inorganic/organic composite materials have been utilized to fabricate bone implanted scaffolds10-12 and they could closely mimic the natural bone and have excellent biocompatibility and bone integration ability.13-15 However, HA nanoparticles were hydrophilic. Thus, they could easily aggregate in hydrophobic polymer matrix and phase separation could easily occur due to the weak interfacial adhesion between the two components. These would induce the bad mechanical performance and early failure of the implanted composites.16 Therefore, it is of vital importance to modify the surface properties of HA nanoparticles to enhance the interfacial adhesion between the two components. Many methods have been utilized to modify the surface properties of HA, including the use of silane coupling agents,17 isocyanates,18 poly(lactide) (PLA)19 and polycaprolactone (PCL)20. However, traditional methods mainly focused on the improvement of the surface properties of HA, and the modification substances did not have other special functions, such as bioactivity and osteoinductivity. Poly(α-amino acids) and their derivatives are a series of natural or synthetic biodegradable and



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biocompatible polymers. Their degradation products are short peptides or α-amino acid monomers which possessed better biocompatibility compared to that of PLGA. Among them, glutamate seems to be associated with bone cell signaling and maturation.21 Poly(γ-benzyl-L-glutamate) (PBLG) is one of the most widely studied poly(α-amino acid) in the field of controlled drug release22 and gene delivery23-24. Besides, Ravichandran

et

al.25

fabricated

PBLG/PLA/collagen

scaffolds

using

the

electrospinning method and then HA was deposited onto the scaffolds. Results demonstrated that adipose-derived stem cells (ADSCs) could be induced osteogenic differentiation on the scaffolds owing to the bioactivities of PBLG and HA. One of the main objectives in bone tissue engineering is the fabrication of biocompatible and biodegradable scaffolds for the treatment of bone defects.26-27 New bone formation comprised a complex series of courses which included cell adhesion, proliferation, differentiation, ECM secretion and mineralization.28-29 Scaffolds served as the matrices for cell attachment and tissue formation, and they should possess some basic requirements, such as high porosity, proper pore size, appropriate surface bioactivities and desirable mechanical support to maintain the original designed structure.30-31 The ideal scaffolds should be easy to fabricate into a pre-designed shape, and a controlled porous structure was also needed to allow sufficient cell ingrowth, vascularization, the transport of oxygen and nutrition and tissue regeneration.32 While, traditional fabrication methods, such as gas foaming, salt leaching and emulsification, possessed one common drawback, which was that the porous structure could not be produced with reliable control of pore size and porosity.33 In recent researches, the frequently used methods for fabricating porous bioceramic/polymer composite scaffolds were the phase separation method and solvent casting/particulate leaching (SC/PL) method.34-35 For the SC/PL technique, the organic solvents were utilized to dissolve the polymer into a homogeneous solution which could be molded into pre-designed shape and the porogen was mixed throughout the polymer solution. Then the polymer matrix was fused together to form a continuous phase and the porogen in the matrix was subsequently leached out to produce interconnected porous scaffolds.36 Porous scaffolds prepared by porogen leaching method possessed the easy


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regulation of the porosity and pore size of scaffolds through controlling the content and size of the porogen.37 Salt and sucrose particles were the most commonly used ones to generate porous structure of scaffolds.38-39 However, residual solvents in the scaffolds during the preparation process may be harmful to the host cells or tissues.40-41 Ellis et al. fabricated PLGA hollow fibre membranes with asymmetric porous structure using 1-methyl-2-pyrrolidinone (NMP) as the organic solvent to dissolve PLGA and water as the nonsolvent to remove NMP.42 NMP is a kind of FDA approved organic solvent which could be used in medical field. Thus it could be considered as a safe chemical solvent.43 In this work, mesoporous HA (MHA) was synthesized through CTAB cationic surfactant template strategy, and different amounts of PBLG were grafted onto the surface of MHA by the in-situ ring opening polymerization (ROP) of γ-benzyl-Lglutamate N-carboxyanhydride (BLG-NCA). Graft amounts were controlled from 11 wt% to 50 wt% by varying the feeding ratio. The result materials possessed enhanced interfacial adhesion with PLGA matrix. In vitro cytological assessments were performed to evaluate the biological properties of PBLG-g-MHA/PLGA composites, and osteogenic activities of the composites would be improved compared to traditional bone implants. The obtained materials were employed to fabricate biodegradable and biocompatible porous scaffolds through the solvent casting/particulate leaching (SC/PL) technique for bone repair. In vivo evaluation of rabbit radius defect repair was conducted to illustrate the properties of osteoinduction and bone formation of PBLGg-MHA/PLGA composites.



2. Experimental 2.1 Materials Calcium nitrate tetrahydrate (Ca(NO3)2·4H2O), diammonium hydrogen phosphate [(NH4)2HPO4], sucrose and ammonium hydroxide (NH3·H2O) were purchased from Beijing Chemical Works (China). Cetyltrimethyl-ammonium bromide (CTAB) was obtained from Yili Fine Chemical Co., Ltd. (China). γ-Aminopropyl triethoxysilane (APS) was purchased from Tokyo Chemical Industry (Japan). γ-Benzyl-L-glutamate acid was purchased from Sinopharm Chemical Reagent Co., Ltd. (China). Ethanol, tetrahydrofuran (THF), dioxane, and acetone were purchased from Beijing Chemical Works. N-Methyl-pyrrolidone (NMP) was purchased from Aladdin Biochemical Technology Co., Ltd. (China). 2.2 Synthesis of MHA and APS-modified MHA MHA was synthesized based on our previous study 44 and APS-modified MHA (MHAAPS) was prepared according to published methods.16 Briefly, 6.34 g of (NH4)2HPO4 and 17.49 g of CTAB were dissolved in 200 mL of deionized water at 40℃ for 1 h and the solution was adjusted to pH=10 using NH3·H2O. Subsequently, 18.89 g of Ca(NO3)2·4H2O was dissolved in 120 mL of deionized water. The Ca(NO3)2·4H2O solution was then added dropwise to the (NH4)2HPO4 and CTAB mixed solution. After the addition of Ca(NO3)2·4H2O solution, the reaction was continued for 6 h at 70 ℃. The precipitate was obtained through sedimentation and then washed several times with ethanol and deionized water, respectively. The product was dried at 70°C for 24 h and then calcined in a box furnace (Thermo Fisher, USA) at 550°C for 6 h to totally remove the CTAB template. Subsequently, 0.44 g of APS was added to the alcohol/water solution (90/10, V/V). The solution was stirred for 0.5 h at room temperature and 2 g of MHA was then added into the above solution. The mixture was subjected to ultrasonic treatment for 15 min and then stirred for 6 h at 25℃. The product was collected by centrifugation and washed with ethanol for three times. After that, the powder was dried at 25°C and then cured at 130°C for 2 h in a box furnace. 2.3 Synthesis of BLG-NCA


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γ-Benzyl-L-glutamate N-carboxyanhydride (BLG-NCA) was synthesized according to the previous work.45 Briefly, 20 g of BLG (0.084 mol) and 14 g of triphosgene (0.047 mol) (recrystallized from CHCl3) were added to a flame-dried 250 mL round-bottom flask, and then 180 mL of dried THF was injected into the flask. The solution was heated to 55°C under a nitrogen atmosphere and stirred for approximately 40 min until the solution became transparent. The solution was then poured slowly into 1500 mL of petroleum ether and white precipitate appeared immediately. The white product was filtered and re-dissolved in ethyl acetate. The product solution was gently washed 3 times using cooled distilled water and then dried using anhydrous magnesium sulfate (MgSO4). After that, the solution was concentrated using rotary evaporator and then ethyl acetate was removed. The obtained white solid was recrystallized 3 times from anhydrous ethyl acetate and petroleum ether (3:1, volume ratio), successively. And then white precipitate was filtered, vacuum dried to obtain the product BLG-NCA (Yield: 69.6%). 2.4 Synthesis of PBLG-g-MHA MHA-APS (denoted as M-A) (0.5 g) and various amounts of BLG-NCA (see Table 2) were added to a flame-dried 100-mL round-bottom flask, and the air in the flask was extracted and replaced with water-free high-purity nitrogen. Subsequently, 60 mL of dried dioxane was injected into the flask. The mixture was dispersed by ultrasonic treatment for 15 min at room temperature and then stirred for 32 h at 25°C. The resulting materials were collected and washed through ultrasonic treatment with dioxane and acetone, respectively. The products were dried at room temperature. 2.5 Preparation of PBLG-g-MHA/PLGA thin films and porous scaffolds To investigate the effects of PBLG-g-MHA with different PBLG graft amounts on cell behaviors, thin materials films on cover glasses were prepared according to the previous method.46 Briefly, round siliconized coverslips (Φ14 mm) were prepared through treating with 2% trimethylchlorosilane/CHCl3 solution (V/V), then baking at 180°C for 4 h. After that, composites of the synthesized materials and PLGA (weight ratio, 1 : 9) were dissolved into CHCl3 with a concentration of 3% (w/V). The system was treated with ultrasonic until the solution became homogeneous. Subsequently, 20 μL of the



solution was added to a coverslip and another one was put onto the previous one, then the two slides were separated parallelly to form thin film on each of the coverslips by solvent evaporation in the air. The coverslips were dried under vacuum for 48 h at 25°C. Then the obtained cover glasses were sterilized with 75% (V/V) alcohol and ultraviolet ray for 60 min and washed with sterile PBS (pH = 7.4) for three times. PBLG-g-MHA/PLGA porous scaffolds were fabricated through the modified solvent casting/particulate leaching method (SC/PL method) based on our previous work.15 The sucrose particles were used as porogen with diameter of 200-450 μm obtained by sieving. The mould utilized for the fabrication of scaffolds were self-made with the dimensions of Φ9 x 55 mm (used for mechanical properties tests) and Φ4 x 45 mm (used for in vivo bone repair). The composites of the synthesized materials and PLGA (weight ratio, 1 : 9) were dissolved into NMP with a concentration of 20% (w/V) to form a homogeneous solution. Then the sieved sucrose particles were mixed into the composite solution (2.2 g/mL). The mixture was cast into the pre-designed moulds and they were frozen in the ultra-low temperature refrigerator (-80℃) for 1 h. After that, the moulds were immersed into the deionized water for 72 h and the water was replaced twice a day. During the immersion period, the sucrose particles and organic solvent NMP were gradually removed from the scaffolds by leaching in distilled water. After the thorough exchange of the organic solvent and the solvation of porogen, the obtained porous scaffolds were air-dried under room temperature. 2.6 Characterization 2.6.1 Materials characterization The X-ray diffraction (XRD) tests for the synthesized materials from 20° to 70° were collected through X-Ray Diffractometer (Bruker D8 Advance, Germany) with a Cu tube anode. Fourier-transform infrared spectroscopy (FT-IR, Bio-Rad Win-IR Spectrometer, UK) spectra were recorded using the traditional KBr slice method. The porous structure of hydroxyapatite samples was observed using FEI Tecnai G2 S-Twin transmission electron microscope (TEM). The TEM samples were prepared by depositing a drop of hydroxyapatite/ethanol suspension on a carbon-coated copper grid, drying at room temperature. N2 adsorption/desorption isotherms were measured with a


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Quantachrome Autosorb-1 gas adsorption analyzer at 78 K after degassing the mesoporous HA and nano-HA samples at 453 K for 24 h. The relative pressure P/P0 of the isotherm was collected from 0.01 to 1.0. The surface areas of samples were calculated through the Barrett–Emmett–Teller (BET) equation. The pore parameters were calculated from the desorption branches of the isotherm according to the Barrett– Joyner–Halanda (BJH) model. The amounts of surface-grafted PBLG on MHA were characterized using a TGA500 thermal gravimetric analysis instrument (TA Instruments, USA) by heating the samples from 25°C to 800°C at a rate of 10°C/min under air condition. The graft amount of PBLG (GA%) and the graft efficiency (GE%) were calculated according to the following equations (1) and (2): GA (%) = W(PBLG-g-MHA) – W(MHA-APS)

(1)

(W(PBLG-g-MHA) represented the weight loss of the PBLG-g-MHA sample and W(MHA-APS) was the weight loss of MHA-APS), GE (%) = M(PBLG) / M(NCA)

(2)

(M(PBLG) was the mass of surface-grafted PBLG calculated by the graft ratio and M(NCA) was the mass of BLG-NCA used in the correlative synthesized reactions according to feed ratio (shown in Table 2)). The wetting properties of PBLG-g-MHA/PLGA thin films were determined based on the measurements of the static water contact angle analysis. The thin films were firstly dried under vacuum for 12 h at room temperature. The static contact angle measurements were collected using DSA 10 instrument (Krüss, Germany) following the manufacturer’s standard sessile drop method with ultrapure water. At least five droplets were utilized on the same membrane and their contact angles were calculated using the relevant software provided by the manufacturer. 2.6.2 In vitro cell culture, viability and morphology assay All the cell experiments were conducted using mouse pre-osteoblast cells (MC3T3-E1) obtained from the Institute of Biochemistry and Cell Biology, Shanghai Institutes for Biological Sciences, Chinese Academy of Sciences. 2 × 104 cells were seeded onto the films of different materials (on coverslips of Φ14 mm) in 24-well culture plates (Costar,



Corning, USA) and the cells were cultured with Dulbecco’s modified Eagle’s medium (high-glucose) (H-DMEM, Gibco, USA) supplemented with 10% (V/V) fetal bovine serum (Zhejiang Tianhang Biotechnology Co., Ltd., China), 63 mg/L penicillin (Solarbio, China), and 100 mg/L streptomycin (Solarbio, China) under humidified atmosphere of 95% air and 5% CO2 at 37°C. The culture medium was changed every other day. To investigate the effects of the introduction of PBLG and the grafting ratios on cell proliferation, cell viability was measured using 3-(4, 5-dimethyl) thiazol-2-yl-2, 5dimethyl tetrazolium bromide (MTT, Biosharp, China) assays. At culture period of 1, 3 and 7 days, 60 μL of MTT solution (5 mg/mL in distilled water) was added to each well. The cells were then cultured for another 4 h. During this period, MTT could be reduced to formazan by viable cells, and then 600 μL of dimethyl sulphoxide (DMSO, Aladdin, China) was added into each culture well after the removal of the original culture medium. Subsequently, 150 μL of the above mentioned DMSO solution was transferred to a 96-well culture plate (Costar, Corning, USA). The experimental data was recorded using a microplate reader (Infinite M200, Tecan, Switzerland) at absorbance of 492 nm. The results were obtained from four parallel samples. The obtained optical density could reflect the total cell amounts cultured on different substrates. Besides, the cell proliferation rate was calculated by the normalization of the data at 3d and 7d of MTT assay to those at 1 d. To illustrate the effects of PBLG and the grafting ratios on MC3T3-E1 cell attachment and cell morphology, after cultured for 1 day, the cells were investigated by fluorescence staining using fluorescein isothiocyanate (FITC, Aladdin, China). The original culture medium was removed and the culture wells were washed three times with PBS. Then the cells cultured on different materials were fixed with 4% paraformaldehyde (PFA, Aladdin, China) (m/V, PBS solution) at room temperature for 15 min, then the PFA solution was removed and the cells were rinsed three times with PBS. After that, the staining was conducted using FITC staining solution (0.1mg/mL). A total of 500μL solution was added into each well and stained for 10 min at 37°C. After the solution was discarded the cells were washed with PBS for three times at room


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temperature. Subsequently, nuclear staining assay was performed using 4’, 6diamidino-2-phenylindole (DAPI, Sigma, USA). The cells were incubated in DAPI staining solution (2 μL/mL in PBS) for 1 min at room temperature, and then washed three times with PBS. The stained cells were observed using the TE2000-U (Nikon, Japan) inverted fluorescence microscope. The fluorescence images of cells grown on different materials were further quantitively analyzed using color-threshold analyzer of NIH Image J software. And the cellular area fraction and cell numbers in a field of view could be obtained. Briefly, for the FITC stained images, the color channels were firstly split and the green channel of the image was reserved. Then the color-threshold was adjusted to a suitable value. After that, the area fraction could be obtained automatically using “Analyze” function in the software. The cell number could also be obtained. 2.6.3 Alkaline phosphatase activity assay (ALP assay) ALP could be seen as an early marker of osteogenic differentiation of osteoblasts and its activity was closely related to the biomineralization abilities of the cells.47 After cultured for 7 and 14 days, the ALP activities of osteoblast cells grown on different substrates were investigated by ALP staining and quantitative assay. For the ALP staining, after the cells were fixed with 4% paraformaldehyde for 30 min at room temperature, a BCIP/NBT alkaline phosphatase staining kit (Beyotime, China) was utilized according to the manufacturer’s operational protocol. The staining was continued for 24 h at room temperature and then observed using inverted optical microscope (TE2000-U, Nikon, Japan). With regard to quantitative assay, at the designated culture time, the cells were lysed using 200 μL/well RIPA Lysis buffer through repeating freezing at -80℃ and thawing at 37℃ for 3 times. Afterwards, the original cell lysates were transferred to eppendorf tubes and were centrifugated at 1000 g for 10 minutes, then the supernatant solutions were collected. 150 μL of the supernatant was utilized for ALP assay and 30 μL was used for bicinchoninic acid (BCA) assay. For the ALP assay, the absorbance at 405 nm was recorded after adding 150 μL of lysis supernatant into 150 μL of pNPP substrate solution (Sigma, USA), following by continuous incubation in the dark at 37℃ for 30 min. BCA kit (Thermo Fisher, USA) was used to measure the total protein amount in the lysis supernatant by



adding 25 μL of the supernatant into 200 μL of the BCA working solution, following by continuous reaction for 30 min at 37 °C, and the absorbance of reaction solution was read using a microplate reader at 562 nm. The relative ALP activities values were calculated from the two above mentioned absorbance values to reflect the relative variation of ALP activities of MC3T3-E1 cells cultured on different materials. 2.6.4 Alizarin Red S staining assay (ARS staining assay) Calcium deposits in the extracellular matrix of cells could be considered as a phenotypic marker of the late stages of osteoblast differentiation. After 14 d and 21 d of culture, calcium deposition of MC3T3-E1 cells was analyzed by Alizarin Red S staining. Briefly, the culture medium was removed and washed for 3 times with PBS. Then the cells were fixed with 4% PFA solution for 15 min at room temperature and washed with acidic PBS (pH 4.2) for three times. Afterwards, the cells were incubated in 1% Alizarin Red S (Solarbio, China) solution (w/V, pH 4.2, in PBS) at 37℃ for 30 min. After the removal of the staining solution, the stained cells were rinsed with PBS (pH 4.2) 3 times and then observed under inverted optical microscope (TE2000-U, Nikon, Japan). Calcium deposits quantification was conducted via previous reported methods.29 Briefly, ARS stained cells after optical microscope observation were washed with distilled water for 3 times and then treated with 1 mL/well of 10% CPC solution (w/V, in distilled water) for 1 h at 37℃. The absorbance of the solution was recorded at 540 nm using microplate reader. 2.6.5 Quantitative Real-Time PCR The expression of osteogenesis-related genes was quantitatively assessed by real-time PCR technique. Briefly, MC3T3-E1 cells cultured on various materials were cultured for 7 and 14 days. Total RNA of the cells grown on different groups was extracted using TRIzol Reagent (Thermo Fisher, USA) according to the product instruction. The purity and concentration of extracted RNA were measured through Nanodrop Plates Reader (Infinite M200, Tecan, Switzerland). The mRNAs of all the samples were obtained by reverse transcription using the PrimeScript RT Reagent Kit (RR047A, TaKaRa, Japan). The expression of osteogenesis-related genes was quantified using SYBR Premix Ex Taq Kit (RR420A, TaKaRa, Japan). Gene-specific primers of glyceraldehyde-3-


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phosphate dehydrogenase (GAPDH), osteopontin (OPN), osteocalcin (OCN), and collagen-I (Col-I) were designed by Comate Bioscience Co., Ltd (China) and shown in Table 1. Real-time PCR was performed using Stratagene Mx3005P Real-time quantitative PCR System (Agilent Technologies Inc., USA). The replication cycles n was set to 40 and the gene expression levels were recorded by the operational software as the threshold cycles (Ct). Relative expression quantities were calculated using the ΔΔCt method as previously reported.48 2.6.6 Scaffolds characterization The morphology and surface elements analysis of the porous scaffolds were characterized by scanning electron microscopy (XL-30 ESEM FEG, FEI Company, USA) with Energy Dispersive X-ray Detector (X-MAX, Oxford Instruments, UK). The cross sections of the scaffolds were obtained as brittle fractures after frozen in liquid nitrogen and were observed after gold coating using a sputter-coater. The porosity of the prepared scaffolds was measured by modified liquid displacement method according to a previous work.49 Briefly, a porous scaffold sample with the weight of w1 was immersed in a measuring cylinder containing a certain volume of distilled water, and evacuation process was conducted to ensure that the pores of the scaffold were thoroughly filled with water. The water-impregnated scaffold was taken out from the cylinder and the water absorbed on the outer surface of the scaffold was removed carefully using filter paper. The weight of the water-impregnated scaffold was recorded as w2. The density of distilled water was marked as ρ. The volume of the scaffold could be calculated easily and was marked as V. Thus, the porosity of the sample could be calculated according to the following equations (3): porosity = (w2-w1) / ρV

(3)

Dried porous scaffolds with the diameter of 9 mm and height of 20 mm were utilized for compressive strength tests measured by the universal mechanical testing machine (Instron 1121, USA) at room temperature. The compressive strength was measured at a compression speed of 2 mm/min and the tests were conducted until 50% of the total deformation relative to original dimensions was achieved. Five samples were measured for each kind of material.



2.6.7 In vivo repair of the rabbit radius defect, X-ray and micro-CT Examination Bilateral critically sized radius defects of rabbits were created using surgical saws on each rabbit forelimb by removing 20 mm of midshaft diaphyseal bone. The porous scaffold (Φ4 x 21 mm, similar in size of the defect) of different materials were implanted into the defect areas without any fixation. The blank radius defects were used as control group. The wounds were then closed with surgical sutures. After surgery, the rabbits were placed back into cages and were fed in the Institute of Experimental Animal of Jilin University, in accordance with the institutional guidelines for care and use of laboratory animals. All rabbits were injected with Benzylpenicillin Sodium for Injection in a dose of 400, 000 units for each rabbit for 7 days. All the wounds were healed gradually and the rabbits had no postsurgical complications. X-Ray Digital radiographs of each forelimb were taken using CR-400 plus Filmless Radiology System (Kodak, USA) at 0, 2, 4, 8 and 12 weeks after surgery to evaluate the bone healing process. Animals were sacrificed at 12 weeks after surgery, and the forelimbs were collected and then fixed with 4% paraformaldehyde (n = 4 forelimbs in each experimental group). For the observation of the regeneration of radius defects, micro-computed tomography (micro-CT) analysis was obtained using a micro-CT scanner (SkyScan 1172, Bruker) with a 0.5 mm aluminum filter at a voltage of 80 kV, current of 100 μA, exposure time of 900 ms, rotation step of 0.6 deg and 180° rotation. After scanning, the raw images were optimized using NRecon software (SkyScan, Bruker) and 3D reconstruction of samples were performed using CTvox software (SkyScan, Bruker). After reconstruction, the quantitative analysis was performed using CTAn software. The bone volume faction (bone volume/tissue volume, BV/TV) was automatically determined to evaluate new bone formation,50 using the manufacturer’s operational protocol. 2.7 Statistical Analysis All the data presented in this study were expressed as the mean ± standard deviation. Statistical comparisons were carried out using one-way analysis of variance (one-way ANOVA). The value p < 0.05 was considered to be a statistically significant difference.


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3. Results and discussion The primary target of this study was to fabricate biodegradable and biocompatible scaffolds consisting of PLGA and PBLG-g-MHA synthesized through in-situ ring opening polymerization of BLG-NCA on the surface of MHA nanoparticles. The preparation route for PBLG-g-MHA/PLGA porous scaffolds were shown in Scheme 1. 3.1 Characterization of PBLG-g-MHA 3.1.1 Characterization of BLG-NCA and MHA As our previous work reported,51 the successful synthesis of BLG-NCA could be confirmed through 1H-NMR analysis (dissolved in CDCl3), which was shown in Figure 1 (a). Mesoporous hydroxyapatite (MHA) was successfully synthesized according to the results of XRD, TEM and N2 adsorption/desorption isotherms analysis, which was shown in Figure 1 (b), Figure S1 and Table S1. It could be seen from Figure 1 (b) that MHA exhibited characteristic diffraction peaks of hydroxyapatite crystal according to the standard diffraction peaks of hydroxyapatite (JCPDS #09-0432) and no peaks of other kinds of apatite phases were detected. The TEM images of nano-hydroxyapatite and mesoporous hydroxyapatite were shown in Figure S1. Compared to nano-HA, numerous irregular mesopores were spread over the nanorods, indicating the mesoporous structure of the synthesized sample. For the mesoporous properties, as shown in Table S1, the BET surface area (SBET) of MHA was 59.099 m2/g, average pore size (dP) and pore volume (VP) were 27.282 nm and 0.437 cm3/g, respectively, which were much higher than that of nano-HA. The above results indicated that mesoporous hydroxyapatite was successfully synthesized. 3.1.2 FT-IR analysis The MHA nanoparticles was surface-modified by APS to obtain MHA-APS, which has amino groups on the surface. These groups could be used to initiate the ring opening polymerization of BLG-NCA. The modification reaction was conducted in dried dioxane at 25°C. The products were thoroughly washed by repeated dispersion in organic solvent, ultrasonic treatment and centrifugation to remove any unreacted matters. Successful preparation of PBLG-g-MHA could be confirmed by FTIR, which



was shown in Figure 1 (c). In comparison with pure MHA, all the different PBLG-gMHA materials showed characteristic peak at 3290 cm-1, which represented the N-H stretching vibration, and peaks at 1652, 1547 cm-1 indicated the existence of amide groups, a characteristic functional group of polypeptides. The peak at 1735 cm-1 was attributed to the carbonyl bond on the side chain of PBLG, and peaks at 697 and 749 cm-1 both confirmed the existence of benzene rings on the side chain of PBLG, which could further indicate that PBLG was successfully grafted onto MHA. 3.1.3 XRD analysis X-Ray diffraction data could reflect the crystal structure of a crystalline material. As shown in Figure 1 (b), the characteristic diffraction peaks of PBLG-g-MHA products with different PBLG graft amounts were almost the same as the XRD patterns of original MHA, like the diffraction peaks at (002), (211), (310), (222), (213) and (004). Meanwhile, the intensity of the above diffraction peaks declined, which was mainly due to the increasement of polymer content. In another word, the content of apatite crystal decreased. The result demonstrated that the surface modification did not affect the crystallographic properties of MHA nanoparticles. 3.1.4 TGA analysis The graft amount of PBLG was measured through TGA analysis. The calculated results were shown in Table 2. As results shown, by adjusting feed ratio, the surface-modified MHA with different amounts of grafted PBLG could be obtained. The graft amount of 1.80 wt% could be attributed to the APS on the surface of MHA. Different weight loss was ascribed to different amounts of PBLG grafted onto the surface of MHA. By varying the feed ratio, PBLG-g-MHA products with different PBLG amounts, ranging from 11.82 wt% to 50.30 wt%, were successfully prepared. 3.1.5 Water contact angle With the increasement of the hydrophilicity of materials, the attachment and proliferation of cells could be effectively facilitated. Namely, hydrophilic materials were more suitable for successful cell adhesion and growth. Thus, static water contact angle tests were introduced to investigate the hydrophilicity of the PBLG-gMHA/PLGA composite materials films. As the results shown in Figure 2, PLGA was


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a typical hydrophobic polymer matrix with the water contact angle of 104.07±1.95°. With the addition of hydroxyapatite, the contact angle was sharply decreased. This was mainly because that hydroxyapatite was a kind of hydrophilic biomineral. While, PBLG is a kind of hydrophobic polypeptides. Therefore, with the increasing graft amount of PBLG, the hydrophobicity of the PBLG-g-MHA/PLGA composite films was tended to increase gradually. Traditionally, if the water drop spread on the surface of a kind of material with contact angle of no more than 90 degrees, the material would be regarded as a hydrophilic matter. The results indicated that most of the obtained samples were hydrophilic, except the composites with PBLG graft amount of 50.30 wt% (P-M4/PLGA group). But even for that group, the contact angle was only 90.64±1.46°. From this point of view, the PBLG-g-MHA/PLGA composite materials would act as an effective matrix for cell attachment. 3.2 Biological assessment 3.2.1 Cell proliferation The efficient proliferation of osteoblast cells played a key role in new bone formation. The biological properties of the matrix materials have an influence on the promotion of cell adhesion and growth.52 To illustrate the biocompatibility of PBLG-g-MHA/PLGA composites, cell viability assessment was performed via the MTT method and the results were shown in Figure 3. Generally, from Figure 3 (a), the total cell amounts for all the groups increased from 1 d to 7 d, which indicated that all the samples were biocompatible and suitable for cell growth. At 1 day, cell viability was slightly higher for MHA/PLGA and P-M-1/PLGA (PBLG graft amount of 11 wt%) groups and lower for PLGA and P-M-4/PLGA (graft amount of 50 wt%) groups, but there were no significant differences among those groups. On the third day, the P-M-1/PLGA group exhibited higher cell viability compared to M/PLGA and M-A/PLGA groups (p < 0.05), which demonstrated that the MHA surface grafted with 11 wt% amount of PBLG exhibited improved cell attachment and proliferation abilities. However, cell amounts decreased at a certain extent with increasing graft amount of PBLG. After 7 days, similar results were obtained. The results from Figure 3 (a) demonstrated that P-M1/PLGA group possessed a better cell expansion ability compared to M/PLGA and M-



A/PLGA groups, indicating that PBLG possessed better biological properties. While, from Figure 3 (b), after cultured for 7 days, the cell proliferation rate increased with the increasing graft amount of PBLG. We deduced that hydroxyapatite and PBLG both had influence on cell behaviors. With the existence of glutamate, bone cell signaling and intercellular communication would be enhanced,53 and higher graft amounts of PBLG showed enhanced cell proliferation ability. While higher amounts of PBLG might also hinder the exposure of HA in the composites, which might influence the bioactivity of HA. We also speculated that the increasement of hydrophobicity of the composites might also influence the adhesion of the cells. Thus, the early cell attachment on substrates with higher PBLG amount was influenced. And with culture time increased, the bioactivity of PBLG come into play, exhibiting higher promotion on cell proliferation. The above results could indicate that PBLG could effectively promote cell proliferation ability, while the early cell attachment on composite substrates with higher PBLG amount was impeded. 3.2.2 Cell attachment and morphology To visually investigate the effect of PBLG on cell adhesion and proliferation, the cellular morphology of MC3T3-E1 cells grown on different materials after cultured for 1 day was observed through fluorescence staining using fluorescein isothiocyanate (FITC) and 4', 6-diamidino-2-phenylindole (DAPI). The FITC staining could effectively show cellular morphology and the DAPI staining could intuitively indicate the number of cells grown on the surfaces of materials (images not shown). As shown in Figure 4 (a), overall, osteoblast cells exhibited polygonal outline which indicated that most of the samples were suitable for bone cells attachment and adhesion. And the osteoblast cells distributed more densely and evenly on the control (cell culture coverglass), M/PLGA and P-M-1/PLGA groups (Figure 4 (a) (1), (3) and (5), respectively), which was nearly consistent with the results of Figure 3 (a) at 1 d. For PLGA and M-A/PLGA groups (Figure 4 (a) (2) and (4)), fewer cells were observed and the cellular morphology was not as well as the above groups. Besides, with the PBLG graft amounts increased, the biocompatibility of the composites was decreased. Especially for P-M-4/PLGA, the cells expressed less cell quantities and worse cell


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morphology (Figure 4 (a) (8)). The results further demonstrated that PBLG-gMHA/PLGA composite materials could effectively promote the early attachment and proliferation of osteoblast cells. Subsequently, the area fraction and cell numbers of MC3T3-E1 cells cultured on different materials were obtained using NIH Image J software. As shown in Figure 4 (b) and 4 (c), the cellular area fraction represented the cell growth situation in the manner of the cell spreading and attaching area on the materials, and cell numbers could reflect the quantities and distribution of cells grown on the materials. Results were in accordance with the fluorescence staining images, which also supported that with the addition of PBLG, promoted attachment and proliferation abilities were achieved. 3.2.3 Alkaline phosphatase activity and mineralization Alkaline phosphatase (ALP) is a kind of efficient enzyme expressed in live cells that could transform organic phosphate p-nitrophenyl phosphate into p-nitrophenol and phosphate anions. The ALP activity could be regarded as an early-stage osteogenic differentiation marker. Figure 5 and 6 showed the ALP staining images and quantitative assay results of MC3T3-E1 cells cultured on different materials at 7 and 14 d. Our previous work reported that HA could induce the osteogenic differentiation.54 As shown in Figure 5 (a), with the addition of hydroxyapatite, the ALP activity of M/PLGA group (Figure 5 (a) (3)) was higher than PLGA group (Figure 5 (a) (2)). Besides, an increase in ALP coloration could be found on PBLG-g-MHA/PLGA groups (Figure 5 (a) (5-8)) compared to M/PLGA and M-A/PLGA groups (Figure 5 (a) (3, 4)), indicating that the cell differentiation toward osteogenesis was improved with the addition of PBLG. While there were no obvious variation tendencies for PBLG-g-MHA/PLGA groups with different PBLG graft amounts. It could be observed that with the increasing graft amount, the ALP activities rose at first, and then decreased. And from quantitive data in Figure 6 (a), similar results could be obtained. After cultured for 14 days, the ALP activities of all the groups decreased compared to the data at 7 d. But the variation tendency among different groups didn’t changed to some extent. Overall, the increasing ALP activity of cells cultured on PBLG-g-MHA/PLGA composites could be achieved



and that might induce the high osteogenic differentiation ability and high bone formation ability in vitro and in vivo. Calcium deposition in the extracellular matrix of bone cells could be regarded as a phenotypic marker of the late stages of osteoblast differentiation, which could be used to test the bone induction and formation abilities of the matrix materials. Alizarin Red S staining and calcium quantification were conducted to evaluate the mineralized nodule formation of osteoblast cells cultured on different materials at 14 and 21 days. As shown in Figure 7 (a), after 14 days of culture, nearly no staining nodules could be observed in PLGA group, which indicated that PLGA did not possessed favorable osteoinduction ability. For MHA/PLGA group, enhanced red nodule staining could be obtained (Figure 7 (a) (3)), which demonstrated that bone induction could be effectively promoted with the addition of HA. Meanwhile, PBLG-g-MHA/PLGA groups exhibited stronger capacity of calcium deposition, especially for the composites with higher graft amount of PBLG (P-M-3/PLGA and P-M-4/PLGA, Figure 7 (a) (7) and (8)). At 21 d, the variation tendency was similar to the results at 14 d, and much more stain intensity of calcium deposition was observed in the P-M-3/PLGA and P-M-4/PLGA groups (Figure 7 (b) (7) and (8)). The calcium quantification of mineral deposits was conducted through extracting Alizarin Red S staining nodules with 10% hexadecylpyridinium chloride monohydrate (w/V, in distilled water) to quantitatively evaluate the bone mineralization ability of different materials. As shown in Figure 7 (c), the total calcium content was higher for M/PLGA and all the PBLG-g-MHA/PLGA groups at 14 days. And at 21 days, the calcium contents of P-M-3/PLGA and P-M-4/PLGA groups were significantly higher than M/PLGA and M-A/PLGA groups (p < 0.05). The above results suggested that with a relative higher graft amounts of PBLG, the composites of PBLGg-MHA/PLGA exhibited enhanced bone induction and calcium deposition ability. 3.2.4. Quantitative Real-Time PCR analysis The main growth stages of osteoblast cells after cultured on a material matrix were cell attachment and adhesion, proliferation, ECM secretion and mineralization.28-29 These processes were regulated and controlled by various genes in each stage. The results of the above biological assays revealed that the PBLG-g-MHA/PLGA composites were


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suitable for enhancing the proliferation and osteoindution of osteoblast cells. However, the influence mechanism of the enhanced biological behaviors was unclear. Therefore, the expression levels of osteogenic genes (OPN, OCN and Col-I) of MC3T3-E1 cells cultured on different materials for 7 and 14 days were quantitatively analyzed by realtime PCR and the results were shown in Figure 8. OPN expression could be detected at the middle stage and OCN was highly expressed during late differentiation. Col-I was firstly expressed during the initial ECM secretion and enhanced expression could be observed along with the expansion and maturation of osteoblast cells.29,

55-56

The

expression levels of OPN were higher on all the PBLG-g-MHA/PLGA groups than others at 7 d, and statistically significant differences were observed between P-M3/PLGA group and control, PLGA groups (p < 0.05). After cultured for 14 days, the expressions of OPN for all the materials were slightly increased, and for PBLG-gMHA/PLGA composites with higher PBLG contents (P-M-3/PLGA and P-M-4/PLGA), significant differences were observed compared to M/PLGA and M-A/PLGA groups (p < 0.05). At 7 d of culture, the expression levels of OCN on all the materials were similar and lower than control group to a certain extent (p > 0.05). At 14 d, compared to M/PLGA and M-A/PLGA, all the PBLG-g-MHA/PLGA composites presented higher OCN levels, and there were significant differences for P-M-3/PLGA and P-M4/PLGA groups (p < 0.05), which reflected that higher amount of PBLG might enhance the expression of OCN. For the expression of Col-I, P-M-1/PLGA group showed higher level than other groups at 7 d, but there were no significant differences (p > 0.05). While after cultured for 14 days, except the P-M-1/PLGA, enhanced expression was found on all the other PBLG-g-MHA/PLGA groups compared to M/PLGA and M-A/PLGA groups (p < 0.05). The natural bone matrix is a kind of biological composites, which mainly consisted of collagens (mainly Col-I) and hydroxyapatite nanocrystals (nHA). Osteocalcin (OCN) is a kind of functional protein expressed during bone matrix maturation and is usually viewed as the specific marker for bone formation.57 The results showed that the expression levels of OCN and Col-I for the P-M-3/PLGA and P-M-4/PLGA composites groups were increased compared to M/PLGA and M-



A/PLGA groups, indicating the enhanced osteoinduction and bone formation abilities of PBLG. 3.3 Scaffold characterization and evaluation of in vivo repair of the rabbit radius defect 3.3.1 Scaffold characterization Figure 9 (8) showed the macroscopic image of the porous scaffolds fabricated through the modified solvent casting/particulate leaching (SC/PL) method. Porous surface and regular shape could be achieved through our fabrication technique. The cross-section microstructures of the porous scaffolds of different materials were observed using SEM. As shown in Figure 9 (1-7), all the porous scaffolds presented relatively homogeneous pore structure and distribution, and the pores of all the scaffolds were well interconnected.

Table 3 showed the results of the porosity of porous scaffolds

measured by liquid displacement method and average pore size analyzed through NIH Image J software. All the porous scaffolds possessed high porosity and similar average pore size, and there were no significant differences in average pore size and porosity among all the scaffolds (p > 0.05). The results indicated that the porous structure could be effectively achieved and there was no obvious influence of different materials on the porous structure of the scaffolds. EDX analysis was utilized to assess the exposure amounts of calcium (Ca) and phosphor (P) elements on the surface of porous scaffolds. As shown in Table 4, elements of Ca and P could be effectively detected, which indicated that HA exposed on the surface of the scaffolds. After surface treatment by APS, exposure of Ca and P elements enhanced and the standard deviations decreased at the same time. And for PM-1/PLGA group, the amounts of the two elements was elevated. While, with the graft amounts of PBLG increased, the levels of Ca and P exposure decreased. Results indicated that surface modification by APS and PBLG could effectively improve the dispersion of HA in the PLGA matrix. While with the amounts of PBLG increased, the exposure of HA might be hindered by the polymer chains of PBLG. The compressive strength of the porous scaffolds was summarized in Figure 10. After the addition of hydroxyapatite, the mechanical property of M/PLGA scaffolds was not changed obviously. But after surface modified by APS and PBLG, the compressive


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strength of composite scaffolds was gradually increased. Especially for P-M-3/PLGA composites, the compressive strength was up to 2.93 MPa. But for P-M-4/PLGA scaffolds, the strength was sharply decreased to 1.68 MPa. As previously mentioned, HA nanoparticles were hydrophilic and could easily aggregate in hydrophobic PLGA matrix. Thus, the compressive strength of M/PLGA was slightly decreased compared to pure PLGA scaffolds. After the surface modification by APS and appropriate amounts of PBLG, the dispersion of HA in PLGA matrix could be effectively improved. So the compressive strength was increased gradually. But for P-M-4/PLGA group, the overmuch amount of PBLG exhibited negative effects on the improvement of relevant properties. 3.3.2 X-ray and micro-CT examination of in vivo repair of the rabbit radius defect The bone healing was observed through macroscopic observation of X-Ray Digital radiographs and 3D reconstruction images of micro-CT. Generally, the critical size of bone defect for rabbits is 15 mm in length, which could not be healed spontaneously without bone implantation or scaffolds replacement.58 As shown in Figure 11 (a-1 to h1), after surgery, bone defect areas could be observed distinctly. At 2 weeks postsurgery, there was slight bone callus formed at the end of the bone defect in the P-M1/PLGA, P-M-2/PLGA and P-M-3/PLGA groups (Figure 11 (e-2), (f-2) and (g-2)). After 4 weeks, the bone defects in M/PLGA and M-A/PLGA groups were started to heal (Figure 11 (c-3) and (d-3)). At the same time, more quantities of new formation bone were observed in different PBLG-g-MHA/PLGA groups except for P-M-4/PLGA composites. Specially, the defects implanted with the P-M-2/PLGA and P-M-3/PLGA scaffolds were bridged to some extent (Figure 11 (f-3) and (g-3)). At 8- and 12-week post-surgery, the densities of the repaired areas in P-M-2/PLGA and P-M-3/PLGA groups increased gradually (Figure 11 (f-4, 5) and (g-4, 5)). While uncompleted new bone formation was observed in the groups of M/PLGA, M-A/PLGA, P-M-1/PLGA and P-M-4/PLGA even after 12 weeks. As shown in Figure 12, 3D reconstruction images of radius defect area at 12 weeks post-operation could clearly indicate that new formation bone were observed for P-M-2/PLGA and P-M-3/PLGA groups. From Figure 13, significant increments in the bone volume fraction (BV/TV) were found in



P-M-2/PLGA and P-M-3/PLGA groups compared with blank, PLGA, M/PLGA and M-A/PLGA groups (p < 0.05). The 3D reconstruction images and quantitive analysis results could further illustrate the better bone formation abilities of P-M-2/PLGA and P-M-3/PLGA scaffolds. The in vitro biological assays revealed that PBLG could promote ALP activity and calcium deposition. While for P-M-4/PLGA group, large area of bone defect still existed (Figure 11 (h-5) and Figure 12 (h)). The reason was deduced that the surface modification of HA with higher graft amount of PBLG might decrease the exposure of HA (summarized in Table 4), thus, the abilities of cell attachment and osteoconductivity of the composite scaffolds were weakened to some extent. So the host bone cells could hardly proliferate into the scaffolds when they were implanted into the defect area. The results indicated that the composites of PBLG-gMHA/PLGA with 22 wt% and 33 wt% amounts of PBLG exhibited more rapid mineralization and osteogenesis processes. In general, bioceramic/polymer composites with favorable properties have been widely studied for bone repair applications.12, 40, 59 PLGA was widely used in orthopedic applications as matrix of bone implants and scaffolds because of their good biocompatibility and processability. While unsatisfactory bioactivities, such as bad cell adhesion ability and poor mechanical properties, limited their applications.7-9 Hydroxyapatite nanoparticles have similarities in chemical composition and structure to the mineral phase of natural bone60 and could effectively promote bone binding and osteoconductive abilities.46 While, HA nanoparticles were hydrophilic and could easily aggregate in hydrophobic PLGA matrix. The mechanical and biological performance of the composites might be influenced because of the weak interfacial adhesion between HA and PLGA. Series of methods were utilized to regulate the surface properties of HA, but traditional modification substances did not possess other special biological functions, such as osteoinductivity. Glutamate exhibited enhanced functions of bone cell signaling and intercellular communication,21 thus it was usually utilized as components of biomedical materials through mixing or chemical reactions with other polymers.25, 61 The structure of peptide bonds on PBLG polymer chains was similar to


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natural ECM, which could provide a suitable environment for cells grown on the materials.62 The cell viability assay (as shown in Figure 3) revealed that the PBLG-gMHA/PLGA composites could promote the cell viability. From Figure 3 (b), after cultured for 7 days, the cell proliferation rate increased gradually. While from Figure 3 (a), the quantity of cells cultured on PBLG-g-MHA/PLGA substrate decreased with the graft amounts of PBLG increased. We deduced that the promotion on cell proliferation could be due to the synergetic bioactivities of HA and PBLG, and the osteoconductivity of HA played the early role. Thus, the higher polymer coating amounts might hinder the exposure of the bioceramic (the elements of Ca and P, as shown in Table 4) on the materials surfaces, which would decrease the effect of hydroxyapatite. Thus, the early cell attachment on PBLG-g-MHA/PLGA composites with higher amounts of PBLG might be affected. Besides, PBLG, a kind of derivatives of poly(glutamic acids), exhibited high calcium binding affinity, which could effectively enhance the osteoblast differentiation and new bone formation.25 The in vitro differentiation tests reflected that ALP activity and caicium deposition ability were significantly increased for the composites containing appropriate amounts of PBLG, as shown in Figure 5, 6 and 7. It should be noted that PBLG-g-MHA with relative lower quantities of PBLG could enhance the viability of osteoblast cells while higher amounts of PBLG possessed better bone induction and mineralization abilities. But results from the in vivo repair of the bone defects revealed that large area of bone defect could still be observed for the group of P-M-4/PLGA, compared to other PBLG-g-MHA/PLGA composites. As shown in Figure 2, P-M-4/PLGA exhibited the highest contact angle compared to other groups, except for pure PLGA. Figure 4 (a) demonstrated that the increasement of PBLG amounts of the composites would influence the initial adhesion of MC3T3-E1 cells. But in traditional plain cell culture, the cells would finally converge together on the materials films, except that the biocompatibility of materials was really poor. Figure 3 (b) could clearly indicate that after 7 days of culture, the cell proliferation rate on P-M4/PLGA was higher than other groups. Thus, PBLG-g-MHA/PLGA composites with highest amounts of PBLG could still exhibit promoted calcium deposition ability in



traditional plain culture condition. While after implanted into radius defect area in vivo, P-M-4/PLGA composite scaffolds showed poor cell infiltration, and the new bone formation was hindered (Figure 11 (h)). Tissue engineering therapies and strategies mostly relied on three-dimensional porous scaffolds which could mimic the natural extracellular matrix, and on the scaffolds cells could fulfill attachment, proliferation, differentiation and maturation.63 The scaffolds fabricated through the solvent casting technique were typically made from polymers that were water-insoluble and toxic organic solvents were required to dissolve them. The solvents should be thoroughly removed from the fabricated scaffolds before implantation. Recently, there were several water-miscible organic solvents that had already been approved by the US FDA for utilization in clinical applications, which could be used to dissolve PLGA, such as tetraglycol, N-methyl-2pyrrolidone (NMP) and dimethyl sulfoxide (DMSO).64 When the scaffold solution of PLGA dissolved in NMP was immersed into distilled water, the solvent NMP would diffuse into the surrounding environment, and PLGA would precipitate and form structured scaffold. One of the important features of tissue engineering scaffolds was the sufficient porosity to allow host cells to infiltrate and proliferate throughout the scaffold. Then the new tissue growth could be conducted. In our study, the pores were created using the sucrose as porogen, which could be leached out in water environment. And the sucrose was insoluble in NMP, so it could keep the sieved size range when it was mixed into the PLGA/NMP solution. Table 3 showed the quantitive analysis of porosity and average pore size of the porous scaffolds. The average pore size varied from 206.43 to 258.82 μm, which was in the porogen diameter range of 200-450 μm. And Figure 9 also exhibited that through the modified SC/PL method, high porosity scaffolds with interconnected pores could be successfully fabricated. Although with the incorporation of PBLG, the bone matrix maturation and new bone formation were effectively enhanced, the area of bone defect still existed in part of the experimental groups in vivo. In other words, the cell attachment and infiltration into the scaffolds were still a major problem for the bone tissue engineering. In our future work, the component which could enhance the cell attachment would be introduced to the


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composites. Besides, a new strategy of tissue engineering scaffold would be utilized to enhance the cell ingrowth in the implants.



4. Conclusions In conclusion, PBLG-g-MHA with different PBLG graft amounts were synthesized by the in-situ polymerization of BLG-NCA on MHA nanoparticles, and PBLG-gMHA/PLGA composites films and porous scaffolds with specific biological functions were successfully prepared using film-coating and modified solvent casting/particulate leaching (SC/PL) methods. In vitro MC3T3-E1 cell culture revealed better cell proliferation on the PBLG-g-MHA/PLGA composites with PBLG amounts of 11 wt% and enhanced osteogenic differentiation on the composites with higher PBLG amounts. The fabricated porous scaffolds possessed high porosity and interconnected pores. Results of the in vivo assay of the rabbit radius defect repair indicated more rapid mineralization and new bone formation process on the composites of PBLG-gMHA/PLGA with PBLG graft amounts of 22 wt% and 33 wt%. This study revealed that although there were still problems about the initial cell attachment, PBLG-gMHA/PLGA composites with 22 wt% and 33 wt% amounts of PBLG could have potential clinical applications.

Associated Content Supporting Information TEM images of nano-hydroxyapatite and mesoporous hydroxyapatite. Results of N2 adsorption/desorption isotherms analysis of nano-HA and MHA samples.

Notes The authors declare no competing ﬁnancial interest.

Acknowledgments This research was financially supported by National Natural Science Foundation of China (Projects. 51273195, 51473164, 51673186 and 51403197), the Program of


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Scientiﬁc Development of Jilin Province (20170520121JH and 20170520141JH), the joint funded program of Chinese Academy of Sciences and Japan Society for the Promotion of Science (GJHZ1519), and the Special Fund for Industrialization of Science and Technology Cooperation between Jilin Province and Chinese Academy of Sciences (2017SYHZ0021).



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Table 1. Primer sequences for Real-time Polymerase Chain Reaction (Real-time PCR) of GAPDH, OPN, OCN and Col-I. Gene

Forward primer sequence (5' - 3')

Reverse primer sequence (5' - 3')

GAPDH

TGAACTAACACAGAGGAGGATCAG

GCTTAGGGCATGAGCTTGAC

OPN

TCAGGACAACAACGGAAAGGG

GGAACTTGCTTGACTATCGATCAC

OCN

AAGCAGGAGGGCAATAAGGT

TTTGTAGGCGGTCTTCAAGC

Col-I

CGCTGGCAAGAATGGCGATC

ATGCCTCTGTCACCTTGTTCG


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Table 2. Graft amount and reaction efficiency of MHA-APS and PBLG-g-MHA related to reaction conditions analyzed by TGA. Feed ratio(w/w)

Reaction time

Graft amount a

Graft efficiency b

MHA-APS:NCA

(h)

(%)

(%)

M-A

-

-

1.80

-

P-M-1

1 : 0.25

32

11.82

53.61

P-M-2

1 : 0.5

32

22.38

57.66

P-M-3

1:1

32

33.12

49.52

P-M-4

1:2

32

50.30

50.60

Samples

a Graft amount of APS (GA%): Graft amount of PBLG (GA%):

GA(% )=W(MHA-APS) –W( MHA) (%) GA(% )=W(PBLG-g-MHA) –W(MHA-APS) (%)

b Graft efficiency (GE%): GE(% )=M(PBLG ) /M(NCA) (%)



Table 3. The porosity and average pore size of the porous scaffolds. Average

Scaffolds

Porosity (%)

PLGA

89.4 ± 2.3

221.86 ± 59.41

M/PLGA

87.2 ± 1.8

258.82 ± 76.53

M-A/PLGA

90.8 ± 2.9

241.17 ± 63.59

P-M-1/PLGA

88.6 ± 2.5

206.43 ± 60.90

P-M-2/PLGA

91.1 ± 3.2

254.48 ± 53.12

P-M-3/PLGA

89.5 ± 2.6

256.51 ± 53.68

P-M-4/PLGA

88.2 ± 2.7

256.33 ± 63.61


pore size (μm)

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Table 4. Surface Ca/P contents of different porous composite scaffolds analyzed through EDX. Scaffolds

Ca (%)

P (%)

PLGA

-

-

M/PLGA

3.87 ± 2.95

2.45 ± 1.86

M-A/PLGA

5.92 ± 1.57

3.65 ± 1.13

P-M-1/PLGA

6.33 ± 1.46

3.37 ± 1.27

P-M-2/PLGA

5.91 ± 1.83

3.07 ± 1.45

P-M-3/PLGA

5.16 ± 2.29

2.96 ± 1.72

P-M-4/PLGA

3.34 ± 2.04

2.14 ± 1.58



Page 34 of 53

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For Table of Contents Use Only



Scheme 1. (a) Synthesis route of PBLG-g-MHA and (b) preparation of PBLG-g-MHA/PLGA porous scaffolds with the application in bone defect repair.


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Figure 1. (a) The 1H-NMR spectra of BLG-NCA monomer in CDCl3 (300 MHz), (b) XRD patterns and (c) FTIR spectra of MHA (M), MHA-APS (M-A), and PBLG-g-MHA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (P-M-1, -2, -3 and -4), respectively.



Figure 2. Contact angle measurements of PLGA (PLGA), MHA/PLGA (M/PLGA), MHA-APS/PLGA (M-A/PLGA) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (P-M-1, -2, -3 and -4/PLGA), respectively. 279x215mm (300 x 300 DPI)


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Figure 3. (a) Optical Density and (b) Cell Proliferation Rate of MTT assay of MC3T3-E1 cells grown on coverslips (Control), PLGA(PLGA), MHA/PLGA (M/PLGA), MHA-APS/PLGA (M-A/PLGA) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (P-M-1, -2, -3 and -4/PLGA) at different culture times. (*) indicated statistical significance, p < 0.05, n=4.



Figure 4. (a) Fluorescence micrographs, (b) area fraction analysis and (c) cell numbers analysis of MC3T3E1 cells grown on coverslips (1, Control), PLGA (2, PLGA), MHA/PLGA (3, M/PLGA), MHA-APS/PLGA (4, MA/PLGA) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (5-8, PM-1, -2, -3 and -4/PLGA) at 1 d of culture. Scale bars: 100 μm. (*) indicated statistical significance, p < 0.05, n=4.


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Figure 5. ALP staining images of MC3T3-E1 cells grown on coverslips (1), PLGA (2), MHA/PLGA (3), MHAAPS/PLGA (4) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (5-8) at (a) 7 d and (b) 14 d of culture. Scale bars: 200 μm.



Figure 6. Relative ALP activities of MC3T3-E1 cells grown on coverslips (1), PLGA (2), MHA/PLGA (3), MHAAPS/PLGA (4) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (5-8) at (a) 7 d and (b) 14 d of culture analyzed with pNPP method. (*) indicated statistical significance, p < 0.05, n=4.


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Figure 7. Alizarin Red staining images of MC3T3-E1 cells cultured on coverslips (1, Control), PLGA (2, PLGA), MHA/PLGA (3, M/PLGA), MHA-APS/PLGA (4, M-A/PLGA) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (5-8, P-M-1, -2, -3 and -4/PLGA) at (a) 14 d and (b) 21 d of culture and corresponding calcium quantification (c). Scale bar: 200 μm. (*) indicated statistical significance, p < 0.05, n=4.



Figure 8. Real-time quantitative PCR analysis of the osteogenesis-related genes OPN, OCN andCol-I after the cells were cultured for 7 and 14 days on coverslips (Control), PLGA (PLGA), MHA/PLGA (M/PLGA), MHAAPS/PLGA (M-A/PLGA) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (P-M-1, -2, -3 and -4/PLGA), respectively. (*) indicated statistical significance, p < 0.05, n=4.


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Figure 9. SEM micrographs of the porous scaffolds of (1) PLGA, (2) MHA/PLGA, (3) MHA-APS/PLGA, (4-7) PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% and (8) macroscopic phonograph of the porous scaffold. Scale bar: 500 μm.



Figure 10. The compressive strength of porous scaffolds of PLGA (PLGA), MHA/PLGA (M/PLGA), MHAAPS/PLGA (M-A/PLGA) and PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% (P-M-1, -2, -3 and -4/PLGA), respectively. (*) indicated statistical significance, p < 0.05, n=4. 279x215mm (300 x 300 DPI)


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Figure 11. Representative X-Ray computer radiographs of rabbit radius defects implanted with (a) nothing (blank group) and porous scaffolds of (b) PLGA, (c) MHA/PLGA, (d) MHA-APS/PLGA, (e-h) PBLG-gMHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% at post-surgery of (-1) 0 w, (-2) 2 w, (-3) 4 w, (-4) 8 w and (-5) 12 w, respectively.



Figure 12. Micro-CT 3D reconstruction images of rabbit radius defects implanted with (a) nothing (blank group) and porous scaffolds of (b) PLGA, (c) MHA/PLGA, (d) MHA-APS/PLGA, (e-h) PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% at 12 weeks post-surgery.


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Figure 13. Bone volume fraction of rabbit radius defects implanted with (a) nothing (blank group) and porous scaffolds of (b) PLGA, (c) MHA/PLGA, (d) MHA-APS/PLGA, (e-h) PBLG-g-MHA/PLGA with PBLG graft amounts of 11 wt%, 22 wt%, 33 wt% and 50 wt% at 12 weeks post-surgery, calculated from Micro-CT 3D reconstructions. (*) indicated statistical significance, p < 0.05, n=4. 279x215mm (300 x 300 DPI)


Porous scaffolds of poly(lactic-co-glycolic acid) and mesoporous

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