Constructing an Anisotropic Triple-Pass Tubular Framework within a

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Constructing Anisotropic Triple-Pass Tubular Framework within Lyophilized Porous Gelatin (GEL) Scaffold Using Dexamethasone (DEX)-Loaded Functionalized Whatman Paper (FP) to Reinforce Its Mechanical Strength and Promote Osteogenisis Jiabing Ran, Hao Zeng, Janak Lal Pathak, Pei Jiang, Yi Bai, Pan Yan, Guanglin Sun, Xinyu Shen, Hua Tong, and Bin Shi Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.7b00673 • Publication Date (Web): 09 Oct 2017 Downloaded from http://pubs.acs.org on October 10, 2017

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Constructing Anisotropic Triple-Pass Tubular Framework within Lyophilized Porous Gelatin (GEL) Scaffold Using Dexamethasone (DEX)-Loaded Functionalized Whatman Paper (FP) to Reinforce Its Mechanical Strength and Promote Osteogenisis Jiabing Rana, 1, Hao Zengb, 1, Janak Lal Pathakc, Pei Jianga, Yi Bai,b , Pan Yana, Guanglin Suna, Xinyu Shena, Hua Tonga, *, and Bin Shib, * a

Key Laboratory of Analytical Chemistry for Biology and Medicine, Ministry of Education, College of

Chemistry and Molecular Sciences, Wuhan University, Wuhan, 430072, China. b

The State Key Laboratory Breeding Base of Basic Science of Stomatology (Hubei-MOST) & Key

Laboratory of Oral Biomedicine Ministry of Education, School & Hospital of Stomatology, Wuhan University, 237 Luoyu Road, Wuhan 430079, PR China. c

School of Pharmaceutical Science and Technology, Health Sciences Platform, Tianjin University,

A-304/Building 24, 92 Weijin Road, Nankai District, 300072, Tianjin, China. *

Corresponding authors:

Hua Tong (E-mail address: [email protected]; Tel: +86 027 68752136; Fax: +86 027 68752136) Bin Shi (E-mail address: [email protected]; Tel: +86 027 87873260; Fax: +86 027 87873260) 1

contributed equally to this paper

ACS Paragon Plus Environment

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Abstract In bone tissue engineering (BTE), most of the currently developed scaffolds still lack the ability to demonstrate high porosity and high mechanical strength simultaneously or the ability to maintain bioactivity and sustained release of loaded biofactors. In this work, we constructed an anisotropic triple-pass tubular framework within a lyophilized porous GEL scaffold using FP, which was prepared by coating DEX-covered Whatman paper (WP) using silk fibroin (SF) membrane with β-sheet conformation. This novel structural design endowed the functionalized paper frame (FPF)/scaffold implant high porosity, high mechanical strength, and sustained DEX delivery capability. Specifically, its porosity was as high as 88.2%, approximating to that of human cancellous bone. The pore diameters of the implant ranged from 50 to 350 µm with an average pore diameter of 127.7 µm, indicating proper pore sizes for successful diffusion of essential nutrients/oxygen and bone tissue-ingrowth. Owing to the construction of double-network-like structure, the FPF/scaffold implant demonstrated excellent mechanical properties both in dry (174.7 MPa in elastic modulus and 14.9 MPa in compressive modulus) and wet states (59.0 MPa in elastic modulus and 3.3 MPa in compressive modulus), indicating its feasibility for in vivo implantation. Besides, the FPF/scaffold implant exhibited long-term DEX releasing behavior (over 50 days) with constant release rate in phosphate buffered saline (PBS). Murine osteoblasts MC3T3-E1 cultured in the porous FPF/scaffold implant had excellent viability. Furthermore, the cells co-cultured with the FPF/scaffold implant showed positive proliferation, osteogenic differentiation, and calcium deposition.


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Twenty-eight days after implantation, extensive osteogenesis was observed in the rats treated with the FPF/scaffold implants. The Anisotropic triple-pass tubular framework of the FPF/scaffold implant demonstrates structural similarities to the long bone. Therefore, this novel FPF/scaffold implant could be a better alternative for long bone defect repair. Key words: Bone tissue engineering; scaffold; porosity; mechanical strength; silk fibroin, dexamethasone


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1. Introduction In BTE, scaffold serves as an adhesion substrate for the localization and delivery of implanted cells as well as a carrier of drugs or growth factors.1, 2 An ideal scaffold should firstly have appropriate porosity, biocompatibility, and mass-transport system of nutrients, oxygen, and metabolic wastes. Secondly, it should provide provisional mechanical function until the regenerated tissue can support mechanical loads. Thirdly, it should be able to interact with cells biomolecularly and locally deliver biofactors to guide cell differentiation and tissue regeneration.3-5 However, most of the currently developed scaffolds are still far from satisfactory, which lack the ability to demonstrate high porosity and high mechanical strength simultaneously or the ability to maintain bioactivity and sustained release of loaded biofactors.4,

6, 7

Directly incorporating biofactors or biofactor-loaded carriers into

porous scaffolds can endow scaffolds high porosity and controlled biofactor delivery capability, but such scaffolds are incapable of supporting heavy external loads because higher porosity and bigger pore size result in diminished mechanical strength, which has long been regarded as an inherent contradiction for a porous scaffold.8-11 3D printing (3DP) and solid free-form fabrication (SFF) based scaffolds have high porosity and high mechanical strength, but they lack the ability to load biofactors as well as to release them in a controlled manner.12-18 Biofactors coated on the outer surface of SFF scaffold could be exhausted within few days.19 WP, a cellulose fiber, has been extensively used in tissue engineering, since it is biocompatible, foldable, creaseable, and shapeable to form desired 3D structures at


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ease.20-24 More importantly, cellulose microwave paper has biodegradability, which is of great significance for an ideal BTE scaffold.25 For instance, in the work of Whitesides et al., paper was shaped into 3D scaffold by folding/rolling methods and the as-prepared scaffold induced template-guided mineralization by osteoblasts.23 Park et al., prepared a paper-based bone scaffold platform by depositing repellent layer onto paper and then grafting cell-adhesive polymer onto the repellent layer. The resultant paper scaffold significantly enhanced in vivo bone regeneration of human adipose-derived stem cells (hADSCs) in a critical-sized calvarial bone defect. In addition, stacking the paper scaffolds with osteogenically differentiated hADSCs and human endothelial cells resulted in vascularized bone formation in vivo.24 Kim et al., fabricated a cylindrically constructed scaffold system by using hydrogel-laden paper. When the scaffold system with chondrocytes was applied into a three-ring defect trachea in rabbits, it replaced the native trachea without stenosis after 4 weeks. 26 However, WP alone is incapable of supporting heavy load and wet WP demonstrates much lower mechanical strength than the dry one. Dehydration followed by water-insoluble silk fibroin (SF) membrane coating onto the WP, reinforces its mechanical strength and improves its waterproof capability.27 In addition, many studies have reported that SF membrane can provide good preservation and regional delivery of drugs or growth factors.28-30 Osteogenic factors, such as BMP-2 and DEX have been frequently loaded in BTE scaffolds to accelerate bone defect healing.19, 31 Therefore, FP prepared by coating biofactor-covered WP using SF membrane with β-sheet conformation, which has high mechanical strength and can deliver drugs and


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growth factors efficiently, could be an alternative biomaterial for novel 3D scaffold design. However, FP cannot be shaped into highly porous architecture while ideal 3D scaffold for BTE should be stable, porous, and cell adhesive. GEL, a derivative of collagen, forms hydrogels that are naturally non-toxic and amenable to adhesion of endothelial cells and chondrocytes.32 In addition, lyophilized GEL hydrogel has high porosity with big pore size and can serve as adhesion substrate for localization of seeded cells. Genipin (GNP), which has high crosslinking efficiency and low toxicity, cross-links GEL to improve the robustness and stability of the GEL hydrogel.33 Therefore, lyophilized, GNP cross-linked GEL hydrogel could be a potential filling material for a FP based 3D scaffold. In this regard, we hypothesized that artificial construction of a well-designed FPF, which could support heavy external load as well as locally deliver biofactors, within a lyophilized porous GEL scaffold might be an alternative strategy for ideal bone implant design. In this work, the FPF was shaped into anisotropic triple-pass tubular architecture to create high stability and structural similarity to long bone (Figure.1). Herein, axially aligned channels can mimic the anisotropic structure of long bone.34 In addition, scaffolds with axially aligned channels have better mechanical properties than isotropic scaffolds and support homogeneous cell seeding, sufficient nutrient supply, and neo-vascularization.35 Herein, DEX was used as a model of osteogenic drug to support osteogenic differentiation of human bone mesenchymal stem cells. The as-prepared material was termed the FPF/scaffold implant. Figure.1 exhibits the


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schematic diagrams for the preparation processes of the FP, FPF, and FPF/scaffold implant. The Whatman paper framework (WPF)/scaffold implant was used as a control in compression tests while the SF coated paper framework (SPF)/scaffold implant without DEX was used as a control in cell culture assay. The morphology, porosity, pore size, mechanical strength, and DEX delivery capability of the FPF/scaffold implant were investigated. The cytocompatibility and osteogenic ability of the implant were analyzed by a series of in vitro cell culture assay. Besides, a preliminary in vivo animal experiment was carried out, as well. 2. Experimental section 2.1 Materials WP (Hangzhou Wohua Filter Paper Co, Ltd., Hangzhou, China), DEX (Wuhan Keri Biological Technology Co., Ltd., Wuhan, China), Methanol and sodium carbonate (Na2CO3) (Sinopharm Chemical Reagent Co, Ltd., Shanghai, China), bombyx mori silkworm cocoons (Nanyang, Henan, China), radiopaque reinforced glass ionomer luting cement (GC Fuji I) (GC dental (Suzhou) Co, Ltd., Suzhou, China), medical-grade high-purity GEL from bovine skin type B (CAS 9000-70-8) (Aladdin Industrial Co., Ltd., Shanghai, China) and lithium bromide (LiBr) (Aladdin Industrial Co., Ltd., Shanghai, China), and GNP (Chengdu ConBon Biotech Co., Ltd., Chengdu, China) were used in this study. All chemicals and reagents were used without any purification. Deionized ultrapure water was used throughout the experiment. 2.2. Preparation of regenerated SF solution


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Bombyx mori silkworm cocoons were firstly degummed in 0.5% (w/v) Na2CO3 aqueous solution at 100 ℃ for 30 min and then rinsed with deionized water for 3 times. The as-prepared SF fibers were dissolved in 9.3 M LiBr solution (Liquor ratio = 10/1) at 37 ℃. The resultant solution was dialyzed against deionized ultrapure water for 3 days. The solid content of the regenerated SF solution was 3.0%, which was determined by weighing method as described previously.36 2.3. Preparations of the FP and FPF Firstly, pristine WP was tailored into two types: One has a length of 24.5 mm and a width of 1.5 mm, while the other one has a length of 50.5 mm and a width of 1.5 mm. At both the left and right margins of a customized WP, two identical rectangular areas (1.5 mm × 0.5 mm) were left as "adhesion areas" for subsequent construction of frameworks. Then, methanol solution of DEX (100 µg/mL) was dripped onto a customized WP (100 µL/cm2) drop by drop, which was placed in a constant temperature oven (37 ℃). After the methanol was evaporated, a mixed solution of SF/methanol (volume ratio = 10/1) was homogeneously deposited onto the WP (200 µL/cm2) at 37 ℃. After the moisture was totally evaporated, a water-insoluble SF membrane was coated on the WP. The rationale behind this process was as follow: Transforming SF α-helix and random coil conformation to β-sheet conformation through methanol treatment induced the formation of SF micelles, in which the larger terminal hydrophilic blocks (i.e., random coil structure) defined the outer edges while the smaller hydrophilic blocks (i.e., α-helix structure) and hydrophobic blocks (i.e. β-sheet structure) constituted the inner part.36, 37 SF micelles were accumulated on the


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WP surface, and formed a water-insoluble layer (Figure.S1).38 Then, the operation was repeated to deposit another SF layer on the WP. Ten layers were deposited in total. The same operation was carried out on the other side of the WP. In order to determine the DEX encapsulation efficiency, DEX-covered WP (10 µg/cm2) without SF encapsulation was synthesized. In addition, SF-coated Whatman paper (SP) without DEX was prepared, as well. The resultant FP and SP were then shaped into anisotropic triple-pass tubular frameworks using radiopaque reinforced glass ionomer luting cement, which was prepared by mixing one level tablespoon of powder and two drops of liquid. Briefly, the two identical "adhesion areas" (1.5 mm × 0.5 mm) on the opposite sides of each FP or SP were homogeneously smeared with radiopaque reinforced glass ionomer luting cement. Then the FP or SP was rolled into a tube and fixed by attaching an adhesion area to the other one. The resultant tubes have two types: One has a diameter of ~7.5 mm and a height of 1.5 mm, while the other one has a diameter of ~16.2 mm and a height of 1.5 mm. Three parallel small tubes were placed next to each other, and then embedded in a big tube (Figure.S2). The three small tubes were tightly surrounded by a big tube. Resultantly, the FPF or SPF was obtained. WPF was also fabricated and utilized as control. 2.4. Construction of the FPF/scaffold implant 0.8 g GEL was added into 40 mL aqueous solution and vigorously agitated at 40 ℃ for 30 min. After the GEL was totally dissolved, 0.04 g GNP was added into the GEL solution and the resultant solution was stirred for another 20 min. The FPF, WPF, and SPF were totally immersed into the viscous GEL solutions and kept in a


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refrigerator of 4 ℃ for 12 h. At this time, the GEL solution would transform into solidified GEL hydrogel. Then, the framework/hydrogel constructs were placed in a refrigerator of -20 ℃ for 24 h and then lyophilized at -60 ℃ for 3 days (Free Zone, Labconco, USA). As has been noted, lyophilization has been widely used to prepare porous scaffolds especially hydrogel-derived scaffolds. In this scheme, water present in swollen hydrogel acts as a porogen for shape template and transforms into ice crystals during the pre-freezing process. After lyophilization process, these ice crystals are sublimated, forming a porous scaffold. Because the ice crystals are uniformly dispersed in the hydrogels, the resultant scaffolds are inter-connective and porous. In addition, the porosity and pore size of the lyophilized scaffolds can be controlled by the size of ice crystals, which is modulated by pre-freezing conditions.39 The FPF/scaffold implant designed in this study was shown in Figure.S2. 2.5. Materials characterizations 2.5.1. Attenuated total reflection Fourier transformed infrared spectroscopy (ATR FT-IR) The as-prepared WP, DEX-covered WP, and FP were subjected to ATR FT-IR measurements with a Nicolet5700 (USA) spectrometer, equipped with a horizontal ATR diamond crystal accessory. The sample compartment was purged with dry air. Each IR spectrum recorded was averaged over 64 scans at a spectral resolution of 4 cm-1. 2.5.2. X-ray diffraction (XRD) The crystal phases of the WP, DEX-covered WP, and FP were analyzed by using


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XRD analysis (X'pert PRO, Panalytical, Holland). The working condition of XRD was CuK0 radiation via a rotating anode at 40 kV and 40 mA. The data were collected in step of 0.1° and range of scattering angles (2θ) from 5° to 40 °. 2.5.3. X-ray photoelectron spectroscopy (XPS) The surface atomic compositions of the WP, DEX-covered WP, and FP were investigated by using X-ray photoelectron spectroscopy (XPS; ESCALAB 250 Xi, Thermo Fisher, America). XPS spectra were recorded for C1s, O1s, N1s, and F1s. 2.5.4. Scanning electron microscope (SEM) The surface morphology of the WP, the DEX-covered WP, the FP, and the FPF/scaffold implant were observed by using a field emission scanning electron microscope (FE-SEM) (Sigma, Zeiss, Germany). It was performed in high vacuum and a voltage of 15 kV and a working distance of 5.2 mm were used. Prior to SEM imaging, the specimens were sputtered with gold. 2.5.5. Contact angle assay The surface wettability of the WP, DEX-covered WP, and FP was investigated by using a Model 200 video-based optical system (Future Scientific Ltd. Co., Taiwan, China). Briefly, a specimen was placed on a glass slide and then drops were deposited onto the specimen and photographed by a camera. The contact angle was measured three times for each sample. 2.5.6. Tensile and compression tests The Young's modulus and tensile strength of the FP in both dry state and wet state were measured at room temperature using a universal testing machine


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(SHIMADZU, AGS-J, Japan). The WP was utilized as control and a cross-head speed of 0.5 mm min-1 was applied. The width of the sample was 10 mm, and the length between the jaws was 30 mm (Figure.S3A-3B). The WPF, FPF, WPF/scaffold implant, and FPF/scaffold implant in both dry state and wet state were compressed in a direction perpendicular to the horizontal planes at a cross-head speed of 0.5 mm min-1 (Figure.6a). 2.5.7. Porosity and pore size distribution measurements An AutoPore IV 9500 mercury porosimeter (Micromeritics Instrument, Norcross, GA, USA) was used to determine the porosity and pore size of the FPF/scaffold implant. The average pore size and porosity were directly obtained from the experimental data. The log differential specific intrusion volume ("dV/dlogD") versus pore diameter ('D") was also plotted to demonstrate the pore size distribution. 2.5.8. DEX release study of the FP and the FPF/scaffold implant The FP (1.5 cm × 1.5 cm) was placed in a dialysis tubing (3,500 molecular weight cut-off) containing 1 mL phosphate buffered saline (PBS, pH 7.4). The dialysis tubing was then immersed into 14 mL PBS and incubated at a constant temperature of 37 ℃ with agitation at 15 rpm. The PBS buffer was changed at preset intervals and the removed buffer was analyzed by using UV-vis spectrophotometer (UV 2550, SHIMADZU, Japan). The FPF/scaffold implant was placed in a dialysis tubing containing 3 mL PBS then the dialysis tubing was immersed into 22 mL PBS. All the other operations were identical. 2.6. In vitro cell culture assay


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The MC3T3-E1 cells (China Center for Type Culture Collection, Wuhan, China) were cultured in α-modified Eagle medium (α-MEM, Hyclone, China) containing 10% fetal bovine serum (FBS, PAN-Biotech, Germany) and 1% penicillin/streptomycin (Hyclone, China) at 37 ℃ in a 5% CO2 incubator. The medium was replaced every 3 days during culturing and the cells were harvested by using trypsin solution (Gibco) and resuspended in fresh α-MEM medium. Prior to cell seeding, the SPF/scaffold implant and the FPF/scaffold implant were fumigated in 75% ethanol steam for 3 days and then immersed in PBS for 24 hours at ambient temperature. Next, these samples were sterilized under ultraviolet light for 1 h. 2.6.1. Cell morphology The MC3T3-E1 cells were seeded into the SPF/scaffold implant and the FPF/scaffold implant and then cultured in CM. After 24 hours of culture, the implants with adherent cells were rinsed with PBS for twice and then soaked into 2.5% glutaraldehyde solution at 4 ℃ overnight. Resultant samples were dehydrated by ethanol in an increasing concentration gradient (30%, 50%, 70%, 90%, and 100%) and then lyophilized for SEM observation. Besides, immunostaining technology was also used to observe the cell morphology. If the sample to be observed were too thick, the exciting light wouldn't penetrate the samples because of barrier effect of the porous GEL scaffold and the images wouldn't be clearly captured. In addition, if the sample were too thick, it would be hard to focus. Herein, the FPF/scaffold implant and the SPF/scaffold implant were tailored with a height of ~3 mm to facilitate microscopic observation. Briefly, the cell-implant constructs were fixed with 4.0%


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paraformaldehyde (Wuhan Good Biotechnology Co., Ltd, China) and washed for 3 times in PBS. Next, the resultant constructs were permeabilized using 0.1% (v/v) Triton X-100 (Beyotime, China) and then washed for 3 times using PBS. Afterwards, the rinsed constructs were incubated in FITC-phalloidin solution (Sigma, USA) for 30 min and then rinsed with PBS twice. The cell nuclei were stained with 4', 6-diamidino-2-pheylindole (DAPI, Sigma, USA) for 10 min. Afterwards, the resultant constructs were visualized using a confocal laser scanning microscope (CLSM, Leica, TCS-SP8, Germany). 2.6.2. Cell proliferation The MC3T3-E1 cells were pre-cultured in a 24-well plate (4 × 104 cells/well) for 12 h. After the cells totally attached to the bottom of the plate, the FPF/scaffold implants or the SPF/scaffold implants was placed into the well. Then, CCK-8 assay was performed at day 1, 4, 7, 10, and 14. In order to avoid dye adsorption, the implants were removed and the CCK-8 solution (Beyotime, China) was added to the wells and incubated for 1 h at 37 ℃. The optical density (O.D) was measured at 450 nm with an ELX808 Ultra Micro Plate Reader (Bio-Tek Instrument, Inc., America). Three parallel replicates were read for each sample. The cells cultured in culture plate were used as a control. 2.6.3. ALP activity The MC3T3-E1 were seeded in a 12-well plate (2 × 105 cells/well) and then cultured in an incubator for 12 h. After the cells totally attached to the bottom of the plates, the FPF/scaffold implants or the SPF/scaffold implants were placed into the


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wells. Afterwards, normal culture medium (CM) or osteogenic medium (OM) was added into the wells. CM was α-MEM containing 10% FBS and 1% penicillin/streptomycin, while OM was α-MEM containing 10% FBS, 1% penicillin/streptomycin, 50 mg/ML L-ascorbic acid and 10 mM glycerophosphate. The medium was changed every two days. The ALP activity was quantitatively measured by ALP activity detection kit (Beyotime, China) at day 4, 7, and 10. Briefly, the cells were collected and resuspended in lysis buffer with 0.2% NP-40 (Beytotime, China). Then diluted samples reacted with detection solution in an incubator of 37 ℃ for 5 min and then the reaction was terminated by adding stop solution. The optical density was detected by microplate reader at 405 nm. Besides, total cellular protein was determined using bicinchoninic acid (BCA) protein assay kit (Thermo, USA). ALP activity was expressed as p-nitrophenol (pNP) (mM) per milligram of total cellular proteins per minute. All the data were normalized against the value of control group at day 4. 2.6.4. ALP and Alizarin Red staining After being co-cultured with the implants in CM or OM for 10 d, the cells were fixed with 4.0% paraformaldehyde for 15 minutes and rinsed with PBS for twice. Then ALP staining was performed using ALP staining kit (Beyotime, China) following the operation manual. In addition, mineralized nodule formation was assessed by using Alizarin Red staining kit (Beyotime, China) at day 21. 2.6.5. OCN immunofluorescence After being co-cultured with the implants in CM or OM for 10 d, the cells were


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fixed with 4.0% paraformaldehyde for 15 minutes and rinsed with PBS for twice. Then the cells were permeabilized by 0.1% Triton-X100 for 10 minutes and then rinsed for twice. Afterwards, the cells were incubated in blocking serum for 1 hours and sequentially incubated with primary antibodies against OCN (Abcam, USA) and Dylight 594-conjugated goat anti-rabbit IgG antibody (Abbkine, USA). Besides, the cytoskeleton was stained with FITC-phalloidin and cellular nuclei were labeled with DAPI. The concentration of primary antibody against OCN, second antibody, and DAPI was 1:50, 1:150 and 0.1µg/mL, respectively. All images were captured by confocal laser scanning microscope (Leica, TCS-SP8, Germany). 2.6.6. Quantitative real-time polymerase chain reaction (RT-PCR) To analyze the gene expression level, the total RNA was isolated by using HP total RNA kit (OMEGA, USA). RNA integrity was determined by electrophoresis on 1.0% agarose gels. Equivalent amount of RNA samples was reversely transcribed into complementary DNA (cDNA) by PrimeScriptTM RT reagent Kit with gDNA Eraser (Takara, USA). The primer sequences of osteogenic genes were listed in Table.1. Then the real-time quantitative PCR for Runt-related transcription factor 2 (Runx2), alkaline phosphatase (ALP), osteopontin (OPN), and osteocalcin (OCN) was performed on Applied Biosystems QuantStudio™ 6 Flex Real-Time PCR System (Thermo, USA) using the real-time PCR kit (SYBR Premix EXII, Takara). The relative expression values were calculated by delta Ct method and all data were normalized against the value of control group. All the operations were performed in triplicates.


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2.6.7. In vivo osteogenesis The cylindrical-shaped samples (7.5 mm in diameter and 5 mm in height) were longitudinally cut off, immersed in 75% ethanol for 6 h, and then washed with PBS for 3 times. The resultant samples were sterilized by UV irradiation for 6 h before implantation. SD male rats at the age of 5 months were purchased from the Laboratory Animal Center of Zhongnan Hospital of Wuhan University (China). All animal experiments were approved by the Animal Research Committee of Zhongnan Hospital, Wuhan University. Before the surgery, the fasting rats were weighed and anaesthetized with 7% chloral hydrate (0.5 mL/100 g). During the surgery, the animals were immobilized in the supine position and placed on a warm plate. Skin around the inner thigh was shaved and disinfected with 5% povidone-iodine (Changdao, Shanghai, China). A skin incision was made to expose the tibia and a defect of tibia was created by cutting the inferior margin of the tibia. Then the FPF/scaffold implant or the SPF/scaffold implant was immobilized into the defect. The tissues were repositioned and sutured with nylon threads. During the postoperative period, the rats were given an appropriate dose of prophylactic antibiotics and fed with water and lab chow regularly. Three weeks after the implantation, the rats were euthanized, and then the implants were retrieved. The obtained implants were fixed in 10% formalin solution for 10 days, decalcified in 10% formic acid for 10 days, dehydrated through a series of graded ethanol, embedded in paraffin, and cut through a vertical section into 2 µm-thick sections. And then the sections were fixed on poly-L-lysine coated glass slides and stained with hematoxylin


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and eosin (H&E). 2.6.9. Statistical analysis Statistical analysis was carried out by one-way analysis of variance (ANOVA) and Tukey procedure for post hoc comparison using Graphpad Prism 5. p < 0.05 was considered statistically significant. Table.1 Base sequences of osteogenic primers for real time PCR Primer name

Forward primer sequence

Reverse primer sequence

Genebank code

β-actin

5’-TGCTATCCAGAAAACCCCTCAA-3’

5’-GCGGGTGGAACTGTGTTACG-3’

NM_009735.3

Runx2

5’-TGCCCAGGCGTATTTCAG-3’

5’-TGCCTGGCTCTTCTTACTGAG-3’

NM_001146038.2

ALP

5’-GCCCTCCAGATCCTGACCAA-3’

5’-GCAGAGCCTGCTGGTCCTTA-3’

NM_007431.3

OPN

5’-TCTCCTTGCGCCACAGAATG-3’

5’-TCGTCCATGTGGTCATGGCT-3’

NM_001204201.1

OCN

5’-AGCAGCTTGGCCCAGACCTA-3’

5’-TAGCGCCGGAGTCTGTTCACTAC-3’

NM_007541.3

3. Results and discussions 3.1. The FP with controlled DEX delivery capability and high mechanical strength Figure.2a demonstrates the ATR FT-IR spectra of the WP, DEX-covered WP, and FP. In Figure.2a-A, the band at 1208 cm-1 was attributed to the in-plane bending of OH while the band at 1647 cm-1 was assigned to the adsorbed water in cellulose fibers. The absorbance peaks at 1375 cm-1, 1335 cm-1, and 1280 cm-1 were due to the deformation, in-plane bending, and bending of CH, respectively. Absorption peak at 1106 cm-1 was attributed to the asymmetric stretching of C-O-C in cellulose and the peak at 1309 cm-1 was because of the CH2 wagging in cellulose.41 In Figure.2a-B, the band at 1268 cm-1 was due to the stretching vibration of C-F bond in DEX. The absorbance peaks at 1706 cm-1, 1662 cm-1, and 1621 cm-1 were attributed to the stretching vibrations of -C=O and double bond framework conjugated to -C=O bonds in DEX.41 When SF membrane was coated onto the DEX-covered WP, the absorbance


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peaks of WP and DEX disappeared while the absorbance peaks of SF appeared (Figure.2a-C), which indicated that the DEX molecules had been encapsulated into the FP. In Figure.2a-C, it could be found that the SF membrane mainly took on β-sheet conformation (1699 cm-1, 1620 cm-1, 1514 cm-1, and 1231 cm-1).36 Figure.2b exhibits the XRD spectra of the WP, DEX-covered WP, and FP. The diffraction peaks at 2θ of 14.9°, 16.5°, and 22.9° (Figure.2b-A) corresponded to the (100), (002), and (101) crystal planes of cellulose fibers, respectively.42 In Figue.2b-B, the newly emerging diffraction peaks at 2θ of 7.77°, 14.47°, 15.39°, 17.13°, and 18.77° were attributed to DEX molecules.43 When SF membrane was coated onto the DEX-covered WP, it could be found that the diffraction peaks of DEX disappeared and the diffraction peaks of WP weakened, which also indicated that DEX had been encapsulated into the FP. In addition, the result of XPS spectra (Figure.2c and Figure.2d) was in accordance with the data of IR and XRD. The XPS spectrum of the WP showed two peaks of C1s (285 eV) and O1s (531 eV) (Figure.2c).23 When DEX was deposited on the WP, the characteristic F1s (686 eV) peak showed up (Figure.2d). Besides, the DEX-covered WP showed higher C:O ratio compared with the WP, which was mainly due to the high C content and low O content of DEX. The XPS spectrum of the FP confirmed the appearance of N1s peak at 401 eV, which was attributed to the SF membrane (Figure.2c). The F1s peak in the spectrum of the DEX-covered WP completely disappeared in the spectrum of the FP, indicating that the SF membrane completely covered the WP and DEX had been totally encapsulated into the FP.


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Figure.3A-3F demonstrate the SEM images of the WP, the DEX-covered WP, and the FP. Pristine WP shows the surface geometry of 3D microfibrous architecture (Figure3A and 3D). When DEX was deposited on the WP, a large number of DEX crystals could be found on the surface of the WP (Figure.3B and 3E). Figure.3C and 3F exhibit the SEM images of the FP, from which we found that DEX crystals were completely covered by a dense SF membrane. From the enlarged image (the insert in Figure.3F), it could be found that a large amount of spherical SF micelles constituted the dense SF membrane. Figure.3G-3I show the water contact angles (WCAs) of the WP, the DEX-covered WP, and the FP. The WCAs of the WP and the DEX-covered WP were 19.6° and 17.0° respectively, indicating that the WP was highly hydrophilic. When SF membrane was coated on the WP, the WCA of the FP was increased to 90°，indicating that the hydrophobicity of the FP was greatly enhanced compared to that of pristine WP. The water-insoluble SF membrane was assumed to be able to reinforce the load-bearing capability of the FP as well as slow down DEX release from the FP. Figure.4A demonstrates the DEX release profile of the FP in PBS. The experiments were carried out in triplicate and the error bars indicated the standard deviations. Owing to the barrier effect of the SF membrane, the FP did not show burst release behavior. Instead, the encapsulated DEX molecules were slowly released into PBS in a controlled manner. In addition, the cumulative release curve could be divided into two intervals. From day 0 to day 21, the DEX molecules went through the hydrophilic gaps between the SF micelles and were slowly released into PBS at a


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decreasing release rate. Figure.S1 exhibits the schematic diagram. From day 22 on, the release rate increased sharply. This might be because of the degradation of the SF membrane. Figure.S3 exhibits the mechanical behaviors of the WP and the FP. Pristine WP demonstrated Young's modulus of 475.0 MPa and fracture stress of 7.5 MPa. Through depositing SF membrane on the WP, the Young's modulus and fracture stress of the FP were increased to 800.0 MPa and 15.9 MPa, respectively. The fracture strain of the FP slightly decreased compared to that of the WP. From Figure.S4, it could be found that the Young's modulus and fracture stress of the wet WP were decreased by ~92.6% and ~97.3% respectively compared to those of the dry WP. Thus, pristine WP is incapable of supporting heavy load in vivo because of its low mechanical strength in wet state (35 MPa in Young's modulus, 0.010 MPa in fracture stress). From Figure.S5, we found that the Young's modulus and fractures stress of the wet FP decreased by ~75% and ~47% respectively compared to those of the dry FP. Even so, the Young's modulus and fracture stress of the wet FP were still as high as 161.0 MPa and 8.8 MPa respectively, which were much higher than those of the wet WP. 3.2. Construction of the FPF/scaffold implant with high porosity, high mechanical strength, and sustained DEX releasing properties In vivo, higher porosity and bigger pore size of a scaffold render greater bone ingrowth. Too small pores limit cell migration, resulting in the formation of a cellular capsule around the edges of the scaffold, which can limit nutrients diffusion and waste metabolism in necrotic regions within the implant. But excessively large pores reduce


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the surface area of scaffolds, limiting cell adhesion.44 Briefly, the minimum requirement for pore size is ~100 µm due to cell size, migration requirement, and transport. In order to enhance osteogenesis and angiogenesis, the optimum pore size for an ideal scaffold is in range of 200-300 µm while the optimum porosity is ~90%.45, 46

Figure.5 shows the SEM images of the FPF/scaffold implant (A, B) and the inner

porous GEL scaffold (C, D). From Figure.5A-5B, it could be found that the FPF was tightly embedded within the GEL scaffold. The lyophilized GEL scaffold was highly porous and inter-connective and the pore size was in the range between ~50 µm and ~350 µm (Figure.5C and 5D). Figure.5E demonstrates the profile of the log differential specific intrusion volume "dV/dlogD" versus pore diameter "D" for the FPF/scaffold implant, which was obtained by using a mercury intrusion technique. The porosity of the implant was as high as 88.2%, which approximated to that of human cancellous bone.47 The pore diameter of the implant ranged from 50 to 350 µm with an average pore diameter 127.7 µm, from which it could be concluded that these pores were favorable not only for successful diffusion of essential nutrients, oxygen, and metabolic wastes but also for bone tissue-ingrowth.48 Biofactors can facilitate osteogenic differentiation and osteogenesis, but uncontrolled biofactor release can cause undesired adverse side effects such as inflammation, swelling, ectopic bone formation, and even carcinogenicity. Especially, uncontrolled DEX delivery can cause hyperglycemia and weakening of immune system.49 Figure.4B demonstrates the DEX release profile of the FPF/scaffold implant in PBS. From day 0 to day 3, the DEX molecules were released into PBS at a low rate


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because of the barrier effects of the SF membrane and the porous GEL scaffold. From day 4 on, the loaded DEX was released into PBS at a constant rate over the following 50 days. In conclusion, this novel FPF/scaffold implant could provide long-term and controlled DEX delivery and maintain local DEX concentration. In addition, the FPF/scaffold implant could be an ideal vehicle for sustained delivery of dual or even multiple biofactors if needed. Ideally, the mechanical strength of a scaffold should match the local biomechanical environment of a specific bone defect site. Mechanical mismatch between the scaffold and the native bone can cause stress shielding or collapse of the implants.50 Generally, the elastic modulus of cortical bone ranges from 15 to 20 GPa and that of cancellous bone ranges from 0.1 to 2 GPa. The compressive strength of cortical bone varies between 100 and 200 MPa while that of cancellous bone varies between 2 and 20 MPa.48 In Figure.6, the compressive stress-strain properties of the WPF, FPF, WPF/scaffold implant, and FPF/scaffold implant were investigated in detail. As to the FPF/scaffold implant and the SPF/scaffold implant, the mechanical probe didn't touch the inner porous scaffold during the compression process, so the external loads were mainly sustained by the WPF and FPF, respectively (Figure.S6). Figure.6c shows the compressive stress-strain profiles of the four specimens. Figure.6d-6f exhibit the histograms of elastic modulus, compressive strength, and fracture strain, respectively. The WPF showed elastic modulus of 80 MPa and compressive strength of 0.99 MPa. When the SF membrane was coated on the WP, the load capacity of the FPF was greatly enhanced (150 MPa in elastic modulus and 12


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MPa in compressive strength). Compared with the WPF and the FPF, the WPF/scaffold implant and the FPF/scaffold implant showed enhanced load capacity correspondingly. In order to illuminate this phenomenon, we got inspiration from the theory of "double-network", which has been extensively applied to construct high-strength hydrogels.51 A typical double-network hydrogel comprises of two networks: the highly stretched and brittle polyelectrolyte as the first network and the flexible and sparsely cross-linked hydrophilic neutral polymer as the second network. In this scheme, the first network sustains the stress throughout the material, and the second network dissipates energy near the crack tip, preventing the facture of gels.52 In this study, the stiff and brittle FPF acted as the first network to sustain external load while the inner porous scaffold acted as the second network to dissipate energy which accumulated at the structural defect sites of the FPF. Figure.S6B exhibits the process of stress accumulation and stress dissipation. The FPF/scaffold implant demonstrated elastic modulus of 174.7 MPa and compressive modulus of 15.0 MPa, which approximated to those of human cancellous bone.53 In addition, from Figure.S7, we found that the wet FPF/scaffold implant still exhibited high mechanical strength (59.0 MPa in elastic modulus and 3.3 MPa in compressive strength), which met the demands for in vivo implantation. Compared to biofactor carriers loaded porous scaffolds, the FPF/scaffold implant demonstrated similar porosity and biofactor delivery behavior but much higher load-bearing capacity. For instance, lyophilized isoniazid (INH) conjugated star


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poly(lactide-co-glycolide)/β-tricalcium phosphate (β-TCP) scaffold demonstrated high porosity of 84% and could intactly release INH and maintain effective INH concentration in a controlled manner for more than 100 days. But its compressive strength and compressive modulus were only 0.18 MPa and 1.02 MPa, respectively.8 Bone morphogenetic protein-2 (BMP-2)/alginate particles or BMP-2/PLGA particles loaded porous collagen/hydroxyapatite scaffold also exhibited big pore sizes (75-110 µm) and high porosity (> 90%) and the loaded BMP-2 was delivered from the scaffold in a sustained fashion for up to 28 days, but its compressive modulus was lower than 2.0 MPa.9 3DP and SFF based scaffolds exhibited similar mechanical strength as well as slightly lower porosity compared with the FPF/scaffold implant, but they have great difficulty in loading and controlled release of biofactors. For instance, 3DP tricalcium phosphate sintered scaffold presented a maximum strength of 10.95 ± 1.28 MPa and 42% total open porosity.15 3DP tricalcium silicate/mesoporous bioactive glass implant scaffold exhibited 3D interconnected macropores (~400 µm), high porosity (~70%), and high mechanical strength (> 12 MPa).16 However, except immersion/vacuum impregnation method, no other effective methods have been found which could efficiently incorporate biofactors into the 3DP scaffolds because biofactors will lose their biological activity due to degradation during spraying through the nozzles.17, 18 Coating of BMP-2 on outer surface of the SFF poly(ε-caprolactone) (PCL) scaffold endowed the scaffold BMP-2 delivery capability as well as similar pore size (100-300 µm) and higher mechanical strength (~1200 MPa in compressive strength) compared


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with the FPF/scaffold implant. However, the encapsulated BMP-2 was totally released from the scaffold within 50 h and it was hard to homogeneously coat the inner part of the scaffold.19 In conclusion, we reported a BTE scaffold with high mechanical, high porosity, and long-term biofactor delivery capability for the first time. 3.3. In vitro cell culture The MC3T3-E1 cells were implanted in the SPF/scaffold and FPF/scaffold implants for 24 h and their morphology was revealed by a CLSM. Owing to the autofluorescence of GEL molecules, the morphology of the adherent cells was not sufficiently clear but the cytoskeletal organization and fibrous structure could still be observed (the upper part of Figure.7), which indicated the high cytocompatibility of the two implants.54 Moreover, the DAPI staining of cell nuclei demonstrated homogeneous distribution of adherent cells within the two implants. Besides, from the 3D view (the lower part of Figure.7), it could also be found that the cells adhered to the pore walls and were homogeneously distributed in the two implants. From the SEM images (Figure.S8), we also found that the cells showed fusiform and polygonal morphology, typical characteristic of osteoblasts. In order to evaluate the biocompatibility of the SPF/scaffold and FPF/scaffold implants, the cells were co-cultured with the implants in CM and their proliferation capability was quantified at different time points by using CCK-8 colorimetric assays. The cells cultured in tissue culture plastic were used as control. As shown in Figure.8a, the cells co-cultured with either the SPF/scaffold implant or the FPF/scaffold implant


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showed increasing counts by time, indicating the good cytocompatibility of the two implants. However, it was also found that the cell number was not significantly increased after 7 days' culture. Presumably, after 7 days' culture, the cell number reached a threshold and a large number of cells began to die because the culture medium was not enough to support growth and proliferation of all the cells. In addition, cells in the control group demonstrated higher proliferation rate than the cells in experimental groups (p 0.05 at day 1, day 10, and day 14), which might be due to that some cells adhered to the implants during the culture. Besides, it was also found that the cells co-cultured with the FPF/scaffold implant showed lower proliferation rate than the cells co-cultured with the SPF/scaffold implant (no statistic difference, p>0.05), which might be due to the inhibitory effect of DEX in cell proliferation.55 In order to figure out whether the FPF/scaffold implant could induce osteogenesis in vitro, the cells were co-cultured with the implant in CM for a given time and then subjected to ALP activity measurement, ALP and Alizarin Red staining assay, and OCN immunofluorescencce assay. In addition, we further explored the synergistic effect in osteogenic differentiation by combining the implant with the OM. As shown in Figure.8b, the MC3T3-E1 cells cultured in CM (control group) exhibited increasing ALP activity by time, demonstrating the inherent osteogenic tendency of the MC3T3-E1 cells. At given time points (4 d, 7 d, or 10 d), the cells co-cultured with the FPF/scaffold implant in CM showed higher ALP activity than the cells co-cultured with the SPF/scaffold implant in CM as well as the cells in the control


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group (p < 0.05 at day 7 and day 10, p > 0.05 at day 4), which indicated that the DEX molecules released from the FPF/scaffold implant promoted osteogenic differentiation of cells. The cells cultured in OM showed enhanced ALP activity compared to the cells cultured in CM in each group (p < 0.05). At a given time point, the cells co-cultured with the FPF/scaffold implant in OM demonstrated the highest ALP activity among all the other groups (p 0.05), but the expressions of ALP, OCN, and OPN in the FPF/scaffold implant group were much higher than those in the SPF/scaffold implant group and the control group (p

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