Materials and Design 179 (2019) 107886

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Materials and Design

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Three-dimensional printed electrospun fiber-based scaffold forcartilage regeneration

Weiming Chen a,b,1, Yong Xu c,1, Yanqun Liu d,1, Zongxin Wang d, Yaqiang Li a, Gening Jiang c,Xiumei Mo e,⁎, Guangdong Zhou a,b,d,⁎⁎a Department of Plastic and Reconstructive Surgery, Shanghai Key Lab of Tissue Engineering, Shanghai 9th People's Hospital, Shanghai Jiao Tong University School of Medicine, Shanghai, Chinab National Tissue Engineering Center of China, Shanghai, Chinac Department of Thoracic Surgery, Shanghai Pulmonary Hospital, Tongji University School of Medicine, Shanghai, Chinad Research Institute of Plastic Surgery, Wei Fang Medical College, Shandong, Chinae College of Chemistry, Chemical Engineering and Biotechnology, Donghua University, Shanghai, China

H I G H L I G H T S G R A P H I C A L A B S T R A C T

• Electrospun fibers were processed intothree-dimensional inks for the firsttime.

• Fiber scaffolds with controlled shapesand large pores were successfully fabri-cated.

• Three-dimensional printed fiber scaf-folds exhibited elastic property in thewet condition.

• Three-dimensional fiber scaffolds com-binedwith chondrocytes achieved satis-factory cartilage regeneration in vivo.

⁎ Corresponding author.⁎⁎ Correspondence to: G. Zhou, Department of Plastic andSchool of Medicine, Shanghai, China.

E-mail addresses: xmm@dhu.edu.cn (X. Mo), guangdo1 These authors contributed equally to this work.

https://doi.org/10.1016/j.matdes.2019.1078860264-1275/© 2019 The Authors. Published by Elsevier Ltd

a b s t r a c t

a r t i c l e i n f o

Article history:Received 6 March 2019Received in revised form 22 May 2019Accepted 23 May 2019Available online 24 May 2019

The fabrication of three-dimensional (3D) electrospun fibrous scaffoldswith customizable shapes and large pores isa challenging task. In this study, for thefirst time, one-dimensional (1D) gelatin/poly (lactic-co-glycolic acid) (PLGA)electrospun fibers were processed into inks suitable for 3D printing. By combining 3D printing and freeze drying,electrospun fiber-based inks were successfully fabricated into 3D-printed scaffolds (3DP) with precisely controlledshapes and large pores, in addition to fibrous surface morphologies similar to that of a native extracellular matrix(ECM). The 3DP exhibited good elasticity and water-induced shape memory, and was found to be superior to 3D-printed scaffolds fabricated using pure gelatin fibers and freeze-shaped scaffolds fabricated using non-fibrous gela-tin/PLGA powder. Moreover, 3DP combined with chondrocytes achieved satisfactory cartilage regeneration in vivo.The novel strategy of 3D printing electrospun fibers established in this study provides a research model for the de-sign and fabrication of multiple scaffolds for various tissue regeneration applications.© 2019 The Authors. Published by Elsevier Ltd. This is an open access article under the CC BY-NC-ND license (http://

creativecommons.org/licenses/by-nc-nd/4.0/).

Keywords:Electrospinning fiberBiomimetic scaffold3D printingTissue engineeringCartilage regeneration

Reconstructive Surgery, Shanghai Key Lab of Tissue Engineering, Shanghai 9th People's Hospital, Shanghai Jiao Tong University

ngzhou@126.com (G. Zhou).

. This is an open access article under the CC BY-NC-ND license (http://creativecommons.org/licenses/by-nc-nd/4.0/).

2 W. Chen et al. / Materials and Design 179 (2019) 107886

1. Introduction

Scaffolds play a very important role in the process of tissue regener-ation.With a better understanding of the complex interactions betweencells and their micro-environments in tissues, researchers have fabri-cated scaffolds with structures and chemical compositions similar tothose of an extracellular matrix (ECM) [1,2]. Among existing scaffolds,three-dimensional (3D) fibrous scaffolds have attracted significant at-tention, as their fibrous networks can efficiently mimic the structureof an ECM and regulate cellular biological behaviors, which include ad-hesion, differentiation, and matrix deposition [3]. Electrospinning, asthe most versatile fiber fabrication technology, can be employed forthe preparation of controllable nanofibers, which can accurately mim-icking of ECM structures (such as fibrous collagen) [2]. However,electrospun fibers typically form two-dimensional (2D) membraneswith small pore sizes and low thicknesses rather than bulk 3D scaffolds[4]. These features are apparently unfavorable for cell infiltration andcomplex 3D tissue regeneration. Therefore, the preparation ofelectrospun fibers into 3D porous biomimetic scaffolds with accuratelycontrollable shapes and large pores for tissue regeneration is a challeng-ing task, which has attracted research attention.

Several approaches (such as multilayering electrospinning [5],liquid-assisted collection [6], and template-assisted collection [7])have been proposed for the fabrication of electrospinning fibers into3D scaffolds. However, the fabrication of 3D scaffolds with complexshapes and controllable pore structures (large pores) using theabovementioned methods was unsuccessful [8]. By combiningelectrospinning and freeze-shaping techniques, several research groupsrecently fabricated electrospun fibers into 3D porous nanofibrousaerogels or scaffolds [4,9–15]. However, scaffolds with accurately indi-vidual shapes and controllable large pore structures were not obtained.As is common knowledge, the control of individual 3D shapes is neces-sary for the reconstruction of tissue with specific shapes (i.e., ear andnose); whereas proper pore structures with respect to pore sizes havea significant influence on cell distribution and nutrient infiltration, andtherefore influence the cell growth, ECM production, and tissue regen-eration. Hence, the preparation of electrospun fibers into 3D scaffoldswith fibrous biomimetic surfaces and accurately controllable individualshapes and pore structures is required for 3D tissue regeneration.

3D printing is a promising technique for the precise control of the in-dividual 3D shapes and large pores (between strands) of scaffolds. How-ever, most current scaffolds based on 3D printing lack fibrous surfacestructures, and therefore cannot efficiently mimic natural ECM struc-tures. Lahann et al. [16] developed a novel electrospinning process re-ferred to as 3D jet writing, which could be used to precisely controlthe pore geometries of 3D scaffold. However, this method required amodified 3D printing machine, and the complex 3D shapes of the scaf-folds are difficult to control. Theoretically, by combining electrospinningwith 3D printing techniques, an ideal 3D scaffold with a customizableindividual shape, controlled large pores, and fibrous surface structurescan be prepared. However, to date, no studies have been reported inwhich electrospun fibers were fabricated into 3D-printed scaffoldswith fibrous surface structures. To validate this concept, a series of tech-nical problems should be addressed: 1) the transformation ofelectrospun fiber membranes into inks suitable for 3D printing withshort single fiber structures; 2) the realization of the efficient bondingof short single fibers and uniform extrusion of fiber-based inks during3D printing; and 3) the enhancement the mechanical properties of thescaffold for the maintenance of the original structure (i.e., the 3Dshape, pore structure, and fibrous surface topology).

To address the above problems, a series of techniques were inte-grated in this study. Firstly, electrospinning, dehydrating, homogeniz-ing, and evaporation drying were carried out in sequence to transformthe electrospun fiber membranes into a short single fiber powder. Sec-ondly, the fiber powder, hyaluronic acid (HA) solution, and polyethyl-ene oxide (PEO) solution were mixed and stirred to form inks for 3D

printing; and then freeze drying and crosslinking were applied to en-hance the mechanical properties of the scaffold and maintain the origi-nal structure. Thereafter, the mechanical properties, fibrousmorphology, and pore structurewere examined for the characterizationof the scaffold. The feasibility of tissue regeneration was further evalu-ated using a cartilage regeneration model in vitro and in vivo.

2. Materials and method

2.1. Materials

Gelatin was purchased from Aladdin reagent (Shanghai) Co., Ltd.Poly lactic-co-glycolic acid (PLGA) was purchased from Jinan DaigangCo., Ltd. Hyaluronic acid (sodium salt) was purchased from BloomageFreda Biopharm Co., Ltd. Polyethylene oxide (PEO) was purchasedfrom Sigma-Aldrich Co., Ltd. All the other chemicals were of analyticalgrades and were used without further purification.

2.2. Preparation of dispersed fibers

For the preparation of the dispersed fibers, 8 g gelatin and 4 g PLGAwere separately dissolved into 50 mL of 1,1,1,3,3,3-hexafluoro-2-propanol. A precursor solutionwas obtained bymixing 50mL of gelatinsolution with 15 mL of PLGA solution. Gelatin/PLGA fiber membraneswere prepared by electrospinning, which involved the application of ahigh voltage of 16 KV, with a flow rate of 3 mL/h and spinneret-collector distance of 15 cm. The obtained gelatin/PLGA fibermembraneswere dehydrated at a high temperature (180 °C) for 60 min in a dryingoven (Bluepard, China). To explore the influences of temperatures treat-ment on fiber dehydration, fiber membranes were also dehydrated at25 °C or 160 °C for 60 min. The dehydrated fiber membranes werethen cut into small pieces (0.5 cm × 0.5 cm) and placed in tert-butanol.With a homogenizer (IKA T18, Germany), uniform fiber disper-sions were obtained by homogenizing the fiber pieces for 10 min at6000 rpm. Finally, the dispersed fibers in tert-butanol were dried byevaporation for 6 h in a water bath (60 °C) until the tert-butanol wascompletely evaporated. The dry dispersed fibers were set in a vacuumfor storage and used for subsequent printing.

In addition, pure dispersed gelatin fibers were also prepared usingthe electrospun gelatin fibers by the same method. For preparation ofthe non-fibrous gelatin/PLGA powder, the gelatin/PLGA solution wasdried by evaporation to form membranes, which were heated at 180°C for 60min for dehydration. Finally, non-fibrous gelatin/PLGA powderwas obtained by grinding the membranes.

2.3. Preparation of electrospun fiber-based inks

For the preparation of the electrospun fiber-based inks, 0.7 g of HAand 0.7 g of PEO were completely dissolved in 10 mL of deionizedwater. Thereafter, 5 g of dry dispersed fibers were stirred into 2 mL of7% HA solution and 3 mL of 7% PEO solution, and then kneaded likedough to form a stable semifluid.

2.4. Fabrication of 3D-printed scaffold (3DP) and freeze-shaping scaffold(3DF)

The3DPwere fabricated by extrudingmixtures through nozzles (theinner diameter was 500 μm or 610 μm) using a 3D plotting system(MAM-II FreeForm Fabrication System, Fochif Mechatronics TechnologyCo., Ltd., China). Themoving speed of the plotting headwas 0.5 mm s−1,and the dosing speed was 0.0018 mm s−1. The strand spacing in soft-ware (the distance between the middle lines of the adjacent strands)was set as 900 μm, 1000 μm, or 1100 μm to obtain different pore sizes.Layer slicing wasmaintained at 350 μm. The strand angle of orientationbetween subsequent layers was set at 60° or 90°to obtain different poreshapes. After the scaffoldswere plotted in air, theywere freeze-dried. To

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improve the stability of the scaffolds, they were crosslinked using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC)/N-hydroxysuccinimide (NHS) (EDC and NHS were dissolved in 95% etha-nol, and the concentrations of the EDC andNHSwere 5% and 3%, respec-tively) for 12 h. Thereafter, the scaffolds were washed using distilledwater and immersed in water at 37 °C for 24 h to remove free salts,then freeze-dried. In addition, Fiber Gel and Non-fiber Gel/PLGA scaf-folds were also printed using dispersed gelatin fiber powder and non-fibrous gelatin/PLGA powder, respectively; and the contents of thesetwo scaffolds were the same as that of the 3DP.

In this study, a freeze-shaped 3D scaffold (3DF) was prepared as fol-lows. The samemixtures composed of dispersed gelatin/PLGAfibers, HAsolution, and PEO solution were poured into a mold and formed into arectangular-shaped scaffold. The scaffold was freeze-dried andcrosslinked using EDC/NHS, washed using water, immersed in waterat 37 °C for 24 h, and then freeze-dried.

2.5. Characterization

2.5.1. Morphology of inks and 3D scaffoldThe Morphologies of the electrospun gelatin/PLGA fibers, dispersed

gelatin/PLGA fibers, dispersed gelatin fibers, non-fibrous gelatin/PLGApowder, 3DP, 3DF and decellularized cartilage matrix (DCM)were all an-alyzed by scanning electronmicroscopy (SEM,Hitachi TM-1000, Japan) atan accelerating voltage of 15 kV. The Homogenized fibers were dispersedin ethanol and analyzed using optical microscopy. The average fiberlength of the dispersed fibers was determined by measuring a minimum150 random fibers from optical micrographs using ImageJ software (Na-tional Institutes of Health, Bethesda, MD). The strand diameter, poresize, and strand spacing of the 3DP were determined by measurementscarried out on the SEM images using Image J software.

2.5.2. Mechanical properties of the 3D scaffoldRectangular scaffolds (10 mm × 10 mm × 2.5 mm) were prepared

for a compressive strength analysis using a mechanical test machine(HY-940FS, China). The compression strain-stress curves of the FiberGel scaffold, Non-fiber gel/PLGA scaffold, and 3DP in the dry statewere measured at a constant displacement rate of 5 mm/min. In thewet state, 100 cycles loading-unloading tests were also carried outwith 50% (3DP) and 60% (Fiber Gel scaffold, Non-fiber Gel/PLGA scaf-fold, and 3DP) compression strains at a strain rate of 10 mm/min.

2.5.3. Water absorption capacity of 3DPThe water absorption capacity (W) of the 3DP was measured as de-

tailed in a previous study [14], and calculated using the followingequation:

W = (Ww − Wd)/Wd × 100%.where Wd and Ww represent the weights of the dry and wet 3DP,

respectively.

2.5.4. Preparation of cell-scaffold constructsScaffolds for cell-scaffold constructs were printed according to the

following parameters: the inner diameter was 600 μm, the movingspeed of the plotting head was 0.5 mm s−1, and the dosing speed was0.0018 mm s−1. The strand spacing in software (the distance betweenthemiddle lines of the adjacent strands)was set as 900 μm. Layer slicingwas maintained at 350 μm. The strand angle of orientation betweensubsequent layers was set at 90°.

All the experimental animals, which included five rabbits and tennude mice, were treated in accordance with the standard guidelines ap-proved by the Shanghai Jiao Tong University Ethics Committee. The isola-tion and culture of chondrocytes were carried out as detailed in apreviously reported study [17]. Auricular cartilage samples harvestedfrom New Zealand white rabbits were digested using 0.15% type II colla-genase (Gibco), and the isolated chondrocytes were cultured in DMEM,which contained 10% FBS, 100 μg/mL of streptomycin, and100 U/mL of

penicillin. The chondrocytes at Passage 2 were harvested for further ex-periments. Prior to cell seeding, the scaffolds were disinfected using eth-ylene oxide. The harvested chondrocytes were resuspended in DMEM,which contained 10% FBS, and a final concentration of 1.0 × 109 cells/mlwas obtained. Thereafter, 400 μL of chondrocyte suspension was seededevenly onto the 3DF and 3DP. These cell-scaffold constructs were incu-bated for 4 h at 37 °C in a humidified atmosphere; and then DMEM thatcontained 10% FBS was added for further culturing at 37 °C in 5% CO2.

2.5.5. In vitro cartilage regeneration and in vivo implantationAfter 24 h of culturing, the viabilities of seeded cells were detected

using a Live and Dead Cell Viability Assay (Invitrogen, USA), and thenexamined using a confocal microscope (Nikon, Japan). After 1 weeksof cultivation, the constructswere eithermaintained in vitro for an addi-tion five weeks or subcutaneously implanted into nude mice for eightweeks (n = 5 per group). For in vivo cartilage regeneration, Balb/cnude mice were anesthetized with 0.3 mL of 1% pentobarbital sodium,and then a pocket was formed by separating the subcutaneous tissue,which was implanted with the cell-scaffold construct. After culturedin vitro culturing for six weeks, the morphologies of the neocartilagewere observed using SEM.

2.5.6. Histological evaluation and quantitative analysis of regeneratedcartilage

Cell-scaffold constructs were fixed in 4% paraformaldehyde, embed-ded, and cut into tissues. To evaluate the structure and ECM depositionof the constructs, frozen sections were stained with hematoxylin andeosin (HE) and Safranin-O. Type collagen II was also detected by immu-nohistochemical staining of tissue sections, for the evaluation ofcartilage-specific phenotypes. All the stained tissue sections were ob-served using an optical microscope.

A quantitative analysis of the neocartilages constructed by 3DF and3DP in vivo, in addition to the native auricular cartilage, was carriedout. The thicknesses of the samples were analyzed using a vernier cali-per. The sulfated glycosaminoglycan (GAG) contents of the sampleswere quantified using Alcian blue method [18]. The deoxyribonucleicacid (DNA) contents were measured using a nucleic acid protein quan-titation detector (Nanodrop 2000, Thermo Fisher, Waltham, MA). Thetotal collagen contents were detected using a hydroxyproline assay[17]. The Young's modulus of the cell-scaffold constructs and native tis-sue were determined using a biomechanical analyzer (Instron-5542,Canton, USA), as previously described [17]. All the samples of the vari-ous groups (n=5per group)were subjected to a continuous planar un-confined strain rate of 1 mm/min until 80% of the maximumdeformation was achieved. The Young's moduli of the tested sampleswere calculated based on the slope of the stress-strain curve.

2.5.7. Statistical analysisAll the datawere obtained in triplicate. A statistical analysiswas con-

ducted using one-wayANOVA and significant datawere indicated usingan asterisk (*). Moreover, the criterion for statistical significance was *pb 0.05.

3. Results

3.1. Preparation and morphology of 3DP

The detailed preparation processes of the conventional 3D fibrousscaffolds and 3DP are shown in Fig. 1. Gelatin/PLGA was first fabricatedinto an electrospun fiber membrane (Fig. 1a). After dehydration at ahigh temperature (180 °C), which was necessary for the maintenanceof the fiber structure stability during subsequent processes, the mem-brane still retained its uniform fiber structure (Fig. 2a) and was easilydispersed into short fibers (Fig. 2b) using a high-speed homogenizerin tert-butanol. The thermal dehydration temperature had a significantinfluence on the fiber homogenization, and higher dehydration

Table 1Compositions, processing methods, and surface morphologies of different scaffolds.

Scaffold Fiber Gel 3DP Non-fiber Gel/PLGA 3DF

Main Materials Gel Fibers Gel/PLGA Fibers Gel/PLGA Powder Gel/PLGA Fibers

Additive components HA PEO HA PEO HA PEO HA PEOProcessing Method 3D Printing 3D Printing 3D Printing Freeze-shapingSurface morphology Fiber Fiber Non-Fiber Fiber

The compositions, processing methods, and surface morphologies of the Fiber Gel, 3DP, Non-fiber Gel/PLGA, and 3DF were shown in Table 1.

Fig. 1. Schematic of various electrospun fiber scaffolds. (a) Traditional fiber scaffold-electrospun fiber membrane. (b) 3D fibrous scaffold constructed by dispersed electrospun fibers viafreeze-shaping. (c) The synthetic steps of 3D-printed fiber-based scaffold.

Fig. 2. Electrospun fibers, dispersed fibers, and 3D-printed scaffold. SEM image of (a) electrospun gelatin/PLGA fibers and (b) dispersed fibers. (c) Average lengths of different dispersedfibers obtained from treatment fiber membranes at different temperatures (25 °C, 160 °C, and 180 °C) (*p b 0.05, n = 3). (d) Image of dispersed fiber powders. (e) Image showing inkscomposed of fibers (treated at 180 °C), HA solution, and PEO solution, which could be squeezed out from the needle. (f) 3D-printed cuboid scaffolds.

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temperatures yielded shorter average lengths of the dispersed fibers(Fig. 2c and Supplementary Fig. S1). After the removal of tert-butanolby a rapid dry evaporation process, dry dispersed short fiber powderwas obtained (Fig. 2d). The fiber powder was then mixed with smallamounts of HA and PEO solutions to form inks for 3D printing. Asshown in Fig. 2e, the inks could be smoothly squeezed out from the nee-dle, and they formed continuous strands. Finally, through subsequent3Dprinting, freeze drying, and crosslinking processes,mechanically sta-ble 3D scaffolds (3DP) were obtained (Fig. 1c and Fig. 2f). Moreover, in-dividual shapes and patterns such as cuboid, ear, letters (A and B),honeycomb, and spider web also could be easily controlled by the accu-rate design of a 3D digital model (Fig. 2f and Supplementary Fig. S2).

By adjusting the appropriate printing parameters, the structuresproperties of the 3D scaffolds such as the strand diameters, pore sizesand pore shapes between strands could be well controlled (Fig. 3a–c),which is very important for cell ingrowth and tissue regeneration. Sim-ilar to the conventional preparation method of the 3D fibrous scaffold,the inks containing the same components were directly freeze driedand crosslinked to form a 3DF as a control (Fig. 1b). 3DF exhibitedsmall and uncontrolled pores (Fig. 3d) and fiber structure (Fig. 3g),which were apparently unsuitable for cell ingrowth and tissue regener-ation. Furthermore, the fiber structure of 3DP (Fig. 3e, f, and h) could bestably maintained after all the above treatments; thus, the surfaces ofthe strands in the 3DP (Fig. 3h) exhibited typical fibrous biomimeticstructures, which were similar to the surfaces of the decellularized car-tilagematrix (DCM, Fig. 3i). Besides, the strands in 3DP possessedmanysmall interconnected pores (Fig. 3e and f), which were different fromthe strands of conventional 3D-printed scaffolds [19]; and it is probablethat they enhanced of the nutrient transport during tissue regeneration.

Fig. 3.Microstructures of 3DP, 3DF andDCM. (a–c) SEM images of 3D-printed scaffoldswith diffimages of 3DP at various magnifications corresponding to the image in (a). (i) SEM image of D

3.2. Mechanical properties and water absorption

Themechanical strength of scaffolds play an important role inmain-taining structural stability during tissue regeneration [20]. According tothe current results, the printed scaffolds printed based on gelatin fibers(Fiber Gel) exhibited poor elasticity in thewet state andwere easily bro-ken after deforming several deformations using tweezers (Supplemen-tary Fig. S3(a–e)). However, the scaffolds printed based on gelatin/PLGAfibers (3DP) exhibited high elastic in the wet state, which could befolded repeatedly and recover its original shape after the externalforce was released (Supplementary Fig. S3(f–j)). In addition, the scaf-folds printed based on the non-fibrous gelatin/PLGA powder (Non-fiber Gel/PLGA) were brittle property in the wet state (SupplementaryFig. S3(k–o)).

The results of the quantitative analysis of the mechanical propertiessupported the above results presented above. In the dry state, the 3DPexhibited the highest compressive strength among the three groups,as indicated by the compressive stress-strain curves (Fig. 4a) andYoung's modulus analysis (Fig. 4b). In the wet state, 3DP exhibitedhighly nonlinear and closed curves during cycle compressions, whichindicates a good compressive elasticity and the ability to recover itsoriginal shape (Fig. 4c and Supplementary Fig. S3q). Furthermore, the3DP maintained preserved the maximum stress value (Fig. 4d) withtheminimum stress loss (20.0%, Fig. 4e) after 100 cycles of compressionat 60% strain, which is similar to the plastic deformation (typically 20%–30% loss at 60% strain) of polymeric foams [21]. However, the Fiber Geland Non-fiber Gel/PLGA scaffolds exhibited significantly decreasedmaximum stresses (Fig. 4d) with significant stress losses (Fig. 4e, Sup-plementary Fig. S3p, and Supplementary Fig. S3r). These results

erent pore structures. (d, g) SEM images of 3DF at differentmagnifications. (e, f, and h) SEMCM.

Fig. 4.Mechanical properties of different scaffolds andwater absorption of 3DP. (a) Compressive stress-strain curves and compressive (b) Young'smodulus of Fiber Gel scaffold, Non-fiberGel/PLGA scaffold, and 3DP in dry condition. The compressive Young's modulus was determined from the slope of the stress-strain curve at a strain range of 0%–10% (*p b 0.05, n = 6).(c) Compressive stress-strain curves of wet 3DP under 100 cycles in the compressive test at 60% strain. The history of (d) maximum stress and (e) stress loss with respect to numberof compressive test cycles. (f) Reversibility of water absorption for 3DP.

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indicated that the compositing synthetic material (PLGA) and fiberstructure had a significant influence on the mechanical properties of3D printed scaffold: 1) the composite scaffold exhibited a higher me-chanical strength than thepure gelatin scaffold; and2) the Fiber scaffoldexhibited a better elasticity than non-fiber scaffold.

Moreover, the 3DP exhibited an excellent water absorption capacitywith a maximum water absorption of 740%, and a high restorabilitywith no significant decrease in the maximum water absorption within10 cycles (Fig. 4f), This indicates a high porosity and fine hydrophilicity,which is advantageous for cell adhesion and tissue regeneration.

3.3. Water-induced shape memory

Shape memory is an important function for multiple biomedical ap-plications. The 3DP was found to exhibit water-induced shape memoryproperties. As shown in Fig. 5, the square 3DP (Fig. 5a) scaffold could berolled into a tube inwet state, which could be fixed in position by freezedrying (Fig. 5b). The tubular scaffold easily recovered its initial shapeafter absorbing water and the entire shape recovery process requiredapproximately 20 s (Fig. 5c and Fig. 5d). Similarly, the rectangular scaf-fold (Fig. 5e)was folded into awave shape, and the deformed shapewasstably maintained by freeze drying (Fig. 5f). However, upon immersionin water, the original shape was completely recovered within 30 s(Fig. 5g and h). This “intelligentmaterial” characteristic is advantageousfor multiple biomedical applications, given that all human tissues con-tain body fluids.

3.4. In vitro and in vivo cartilage regeneration

The feasibility of cartilage regeneration in vitrowas evaluated usingthe scaffolds combined with chondrocytes. For a comparison betweenthe cartilage-regenerating abilities of the 3D scaffolds with and withoutcontrolled pore structures, the 3DF was selected as a control. As shownin Figs. 6, 24 h after cell seeding, a large number of living cells (green)

could be observed in 3DF (Fig. 6a) and 3DP (Fig. 6g), which indicatedthat the compositions and fibrous structure of both scaffolds were ad-vantageous to cell viability. After six weeks of in vitro culturing, bothgroups formed cartilage-like tissue (Fig. 6b and Fig. 6h), and SEM resultsindicated that chondrocytes produced abundant ECMs to cover the sur-faces of the 3DF (Fig. 6c) and 3DP (Fig. 6i). However, the specimens inthe 3DF group were irregular in shape and formed very thin tissues(Fig. 6b and d–f). In contrast, the specimens in the 3DP group retainedtheir original square shapes and formed thicker tissues (Fig. 6h andFig. 6j-l). Undegraded strands still were observed in the 3DP groupafter six weeks of in vitro culturing, which were dissolved bydimethylbenzene during the staining process, and several holes wereobserved in regenerated tissue (Fig. 6j–l).

Consistentwith in vitro results, after eightweeks of in vivo implanta-tion, the specimens in 3DF group formed very thin tissueswith an irreg-ular shape (Fig. 7a). A histological examination further revealed that thespecimens in the 3DF group exhibited a heterogeneous structure, andthe cartilage-like tissue was only observed in the surface region of thespecimen (Fig. 7b–d). However, the specimens (Fig. 7e) in the 3DPgroup retained their original square shapes and exhibited a good elastic-ity (SupplementaryMovie 1). The specimens in the3DPgroup exhibitedwhite color, which was similar to the native cartilage (Fig. 7l). The his-tological examination revealed the cartilage-like tissue with a typicalcartilage structure and cartilage-specific matrix deposition (Fig. 7f–k)were formed, which were similar to the native cartilage (Fig. 7m-o).

In addition, the thickness of the neocartilage in the 3DP group wassix times thicker than that of the 3DF group (Fig. 8a). The results ofquantitative analysis indicated that the DNA content (Fig. 8b), glycos-aminoglycan (GAG) content (Fig. 8c), total collagen content (Fig. 8d),and Young's modulus (Fig. 8e) in the 3DP group were lower thanthose in the native cartilage, and significantly higher than those in the3DF group. The Young's modulus of the neocartilage in the 3DP group(Fig. 8e) was significantly higher than that of the pure scaffold (Supple-mentary Fig. S4), which indicated that the regenerated tissue

Fig. 5.Water-induced shapememory of 3DP. (a) Square scaffold. (b) Inwet state, a tubular-shaped scaffoldwas obtained by folding the opposite angle of the square, and the scaffold wasfreeze-dried to maintain its deformed shape. (c) Shape recovery process of square scaffold after absorption of water for 12 s. (d) Square scaffold completely recovered its original shapewithin 20 s of water absorption. (e) Rectangular scaffold. (f) Wet scaffold was folded into a wave shape and freeze-dried to maintain its shape. (g) Shape recovery process of rectangularscaffold after water absorption for 12 s. (h) Rectangular scaffold completely recovered its original shape within 30 s of water absorption.

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significantly improved the mechanical strength of the scaffold. In sum-mary, the results indicated that the 3DP combined with chondrocytescan achieve satisfactory cartilage regeneration in vivo.

4. Discussion

It is common knowledge that electrospun fibers are favorable ECMbiomimetic scaffolds for various tissue regeneration applications. How-ever, the fabrication of electrospunfibers into 3D scaffoldswith accurate3D shapes, controllable large pore structures, ECM biomimetic surfaces,and proper mechanical properties is a challenging task. In this study, bycombining electrospinning, 3D printing, freeze drying, and crosslinking,electrospun fibers were successfully fabricated into 3D fibrous scaffoldswith accurate outer shapes, controllable large pore structures, and satis-factory mechanical properties. Moreover, the novel scaffolds exhibitedgood ECM biomimetic surface, elasticity, and water-induced shapememory. Moreover, the scaffold combined with chondrocytes achievedsatisfactory cartilage regeneration in vivo, which verifies its applicabilityto promising application in tissue regeneration.

Fig. 6. In vitro cartilage regeneration. Fluorescencemicrographs of chondrocytes seeded on (a) 3(a) and (g) present images of 3DF and 3DP, respectively. Gross appearance of cell-scaffold conspresent the ECM deposition in the (c) 3DF and (i) 3DP groups after six weeks. Histological analyof in vitro culturing. Sections are stained with (d and i) HE and (e and j) Safranin-O, the cartilacollagen. (For interpretation of the references to color in this figure legend, the reader is referr

The development of methods to transform electrospun fiber mem-branes into inks is a primary problem. Three major limitations shouldbe overcome to obtain an ideal 3D-printed scaffold based on fiberinks: 1) the transformation of electrospun fiber membranes into inksfor 3D printing; 2) the realization of efficient bonding of short single fi-bers and uniform extrusion of fiber-based inks during 3D printing; and3) the enhancement of the mechanical properties of the scaffold for themaintenance of the original structure.

First, to meet the above requirements, firstly, fiber membranesshould be processed into single short fibers to prevent nozzle blocking.In a previous study [14], gelatin-basedfiberswere easily dispersed usinga high-speed homogenizer in tert-butanol. However, tert-butanol wasfound to limit restrained the extrudability of dispersed fibers from thenozzle, which suggests that tert-butanol should be removed from thefiber dispersant. In this study, dispersed fiber powder was successfullyobtained by dry evaporation. Amajor concernwas that themorphologyof gelatin-based fiber would be damaged after immersion in a water-based solution (such as HA solution). Given that physical treatmentssuch as dehydration could crosslink gelatin and render it insoluble

DF and (g) 3DP for 24 h (live and dead cells are dyed green and red, respectively). Insets intructs in (b) 3DF and (h) 3DP groups after six weeks of in vitro culturing. The SEM imagessis of cell-scaffold constructs in (d, j, and f) 3DF and (e, k, and l) 3DP groups after sixweeksge-specific phenotypes were verified by evaluating the expression of the (f and l) type IIed to the web version of this article.)

Fig. 7. In vivo cartilage regeneration. Gross view and histological analysis of regenerated cartilage in (a-d) 3DF and (e-k) 3DP groups after eight weeks of in vivo implantation, (l-o) nativeauricular cartilage was selected as a control. Sections were stained with (b, f, i, and m) HE and (c, g, j, and n) Safranin-O to evaluate the ECM depositions and structures of regeneratedcartilages, whereas the cartilage-specific phenotypes were verified by evaluating the expression of the (d, h, k, and o) type II collagen.

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[22], the dehydrated fiber membranes were dispersed using a high-speed homogenizer in tert-butanol. Second, to transform dispersed fi-bers into inks suitable for printing, small amounts of HA and PEO aque-ous solutionswere added to the dry crosslinked fibers, respectively. Themain role of these additives is to ensure that fibers are glued togetherand smoothly pass through the deposition nozzle, and that the fibrousstructure of the scaffold is unchanged. Hence, the additive contents(HA and PEO) should be appropriately controlled. According to the ex-perimental results, the mass ratio of HA and PEO to the fibers wasabout 0.07, which was confirmed as an appropriate ratio to meet theabovementioned requirements. Third, to generate amechanically stablescaffold, further crosslinking was required. The carboxyl group on theHAmolecule can form a chemical bondwith the amino group on gelatinupon exposure to EDC/NHS [23]; and this bond ensure the joining of fi-bers stuck together and that of strands.

An appropriate mechanical strength of the scaffold is critical for tis-sue regeneration. The 3DP exhibited a better mechanical strength thanthe Fiber Gel scaffold and Non-fiber Gel/PLGA scaffold. Moreover, itwas further confirmed that the 3D printed scaffold based on non-fibers exhibited a poor mechanical strength, which is not appropriatefor tissue regeneration. However, printed scaffold based on compositefiber inks exhibited good mechanical properties. In general, compositeinks (natural and synthetic polymers) are printed using different princi-ples andmethods, i.e., natural polymers are printedmainly based on gel

or inks, whereas synthetic polymers are printed mainly by fused depo-sitionmodeling. Therefore, the printing of composite scaffolds based onnatural and synthetic polymers typically requires multiple nozzles withthe accurate alignment of each nozzle during processing [24]. To over-come this limitation, gelatin and PLGA were first fabricated intoelectrospun fibers and then prepared into inks for 3D printing, thusallowing for the fabrication of a mechanically homogeneous compositescaffold using one nozzle in one-step process at room temperature. The3DP exhibited good elasticity in the wet state, which may be due to thefollowing reasons. First, the chemical bond formed between the HAmacromolecules and gelatin fibers after crosslinking caused the joiningof the HA and fibers stick together [14,23], which formed a stable net-work that significantly enhanced the mechanical strength. Second, thestrands of the 3DP were composed of many fibers (similar to yarn),which were tightly bound and provided improved mechanical robust-ness. Third, due to the fibrous and porous structures of the strands, inaddition to the hydrophilicity of gelatin, the 3DP exhibited an excellentwater absorption capacity and high restorability, which serves as an im-portant support for the elastic property of wet scaffolds.

Moreover, the 3DP with water-induced shape memory is desirablefor biomedical applications. Shape-memory polymers are type ofstimuli-responsive intelligentmaterial that can be packed into a tempo-rary shape, which is recovered upon exposure to an external stimulus[25]. The most common shape memory materials are activated by

Fig. 8. Quantitative analysis of regenerated cartilage. Quantitative evaluation of (a) cartilage thickness, (b) DNA content, (c) GAG content, (d) total collagen, and (e) Young's modulus ofregenerated cartilages of 3DF and 3DP groups after eight weeks of in vivo implantation. Fresh auricular cartilage (native) was used as a positive control. (*p b 0.05, n = 3).

9W. Chen et al. / Materials and Design 179 (2019) 107886

temperature, which should behigher than the switching transition tem-perature (T trans) of the polymer [26]. However, the Ttrans of most shapememory polymers is higher than the average human body temperature,which significantly limits their applicability to tissue engineering. Giventhat all tissues in the human body contain fluids (such as plasma andjoint fluid), large scaffolds can be deformed into smaller shapes fortransplantation into defect positions via microtrauma incision, and thedeformed scaffold then absorbs body fluids in the defect position andrecovers its initial shape for tissue regeneration.

As mentioned in the introduction, proper pore structures have a sig-nificant influence on cell distribution and tissue regeneration. In thisstudy, the precisely controlled pore sizes facilitated the realization of auniform cell distribution throughout the 3D scaffold, and the porousstructure and the interconnectivity between pores also facilitated thediffusion of nutrients and elimination of metabolic wastes in the innerscaffold [27]. The satisfactory cartilage regeneration of the 3DP groupcan be mainly attributed to the controlled pore (between strands)structure of the scaffold, which ensured a uniform cell distribution andnutrient supply, thus forming a relatively uniform cartilage-like tissue.Homogeneous cartilage regeneration is also a critical factor for theshape maintenance. In the 3DP group, regenerated cartilage (like con-crete) and scaffold strands (like steel bar) were uniformly integratedthroughout the sample; therefore, the regenerated cartilage efficientlyenhanced the mechanical strength of the sample and maintained theoriginal shape. In the 3DF group, due to the improper pore structure,the chondrocytes were mainly distributed in the surface region of thescaffold; thus cartilage was only formed in the surface region. Withthe degradation of the scaffold during the in vivo implantation, the me-chanical strength of the scaffold failed to maintain the original shapedue to the lack of neocartilage support.

5. Conclusions

In summary, this study presents a novel strategy for the fabricationof electrospun fiber-based 3D scaffolds with controlled 3D shapes andlarge pores, in addition to ECM biomimetic surface structures, whichovercome the limitations (uncontrollable 3D shaping and small porestructure) of traditional electrospinning fiber scaffolds. Due to thecomposited fiber-based inks, the 3DP exhibited a good elasticity in

wet state, whichwas significantly superior to that of 3D-printed scaffoldfabricated using non-fiber powders. Most importantly, the novel scaf-fold combined with chondrocytes achieved satisfactory cartilage regen-eration and shape maintenance in vivo. This study provides a researchmodel for the design and fabrication of multiple biomimetic scaffolds.

SEM images and opticalmicroscope images of homogenized short fi-bers (Supplementary Fig. S1). Images of 3D-printed scaffolds with di-verse shapes or design patterns (Supplementary Fig. S2). Materialcomposition and mechanical properties of Fiber Gel scaffold, Fiber Gel/PLGA, and Non-fiber Gel/PLGA scaffold (Supplementary Fig. S3). Me-chanical properties 3DF (Supplementary Fig. S4). Elasticity of tissueengineered cartilage (Supplementary Video 1). Supplementary data tothis article can be found online at https://doi.org/10.1016/j.matdes.2019.107886.

Credit author statement

WM Chen, scaffolds design. Y Xu, culture cells. YQ Liu, animal oper-ation. ZX Wang, SEM test. YQ Li, mechanical test. GN Jiang, Review. XMMo and GD Zhou, funding acquisition and Review & editing.

Acknowledgements

This work was supported by China Postdoctoral Science Foundationfunded project (2017M621495), the National Natural Science Founda-tion of China (81571823, 81570089), The National Key Research andDevelopment Program of China (2017YFC1103900), The Key Researchand Development Program of Shandong Province (2016GGB14002),the Shanghai Committee of Science and Technology (15DZ1941600),and the Program for Shanghai Outstanding Medical Academic Leader.

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