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Reinforcement of hydrogels using three-dimensionally printed microfibres

Reinforcement of hydrogels using three-dimensionally printed microfibres ARTICLE Received 21 Aug 2014 | Accepted 16 Mar 2015 | Published 28 Apr 2015 DOI: 10.1038/ncomms7933 Reinforcement of hydrogels using three-dimensionally printed microfibres 1,2 1,2 2 1,2 3 3 Jetze Visser , Ferry P.W. Melchels , June E. Jeon , Erik M. van Bussel , Laura S. Kimpton , Helen M. Byrne , 1,4 2,5 2,6,7 1,2,4 Wouter J.A. Dhert , Paul D. Dalton , Dietmar W. Hutmacher & Jos Malda Despite intensive research, hydrogels currently available for tissue repair in the musculoskeletal system are unable to meet the mechanical, as well as the biological, requirements for successful outcomes. Here we reinforce soft hydrogels with highly organized, high-porosity microfibre networks that are 3D-printed with a technique termed as melt electrospinning writing. We show that the stiffness of the gel/scaffold composites increases synergistically (up to 54-fold), compared with hydrogels or microfibre scaffolds alone. Modelling affirms that reinforcement with defined microscale structures is applicable to numerous hydrogels. The stiffness and elasticity of the composites approach that of articular cartilage tissue. Human chondrocytes embedded in the composites are viable, retain their round morphology and are responsive to an in vitro physiological loading regime in terms of gene expression and matrix production. The current approach of reinforcing hydrogels with 3D-printed microfibres offers a fundament for producing tissue constructs with biological and mechanical compatibility. 1 2 Department of Orthopaeics, University Medical Center Utrecht, Heidelberglaan 100, 3508 GA Utrecht, The Netherlands. Institute of Health and Biomedical Innovation, Queensland University of Technology, 60 Musk Avenue, Kelvin Grove QLD 4059, Queensland, Australia. Mathematical Institute, University of Oxford, Andrew Wiles Building, Radcliffe Observatory Quarter, Woodstock Road, OX2 6GG Oxford, UK. Department of Equine Sciences, Faculty of Veterinary Medicine, Utrecht University, Yalelaan 112, 3584 CM, Utrecht, The Netherlands. Department of Functional Materials in Medicine and Dentistry, 6 7 University of Wurzburg, Pleicherwall 2, 97070 Wurzburg, Germany. Georgia Institute of Technology, North Avenue, Atlanta, GA, 30332, USA. Institute for ¨ ¨ Advanced Study, Technical University Munich, Lichtenbergstrasse 2a, 85748 Garching, Munich, Germany. Correspondence and requests for materials should be addressed to D.W.H. (email: dietmar.hutmacher@qut.edu.au) or to J.M. (email: j.malda@umcutrecht.nl). NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 1 & 2015 Macmillan Publishers Limited. All rights reserved. ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ydrogels are an important class of biomaterials with many mechanical loading regime in a bioreactor culture, in terms applications, including as cell carriers for the engineering of cartilage gene expression and matrix formation. These 1,2 Hof tissues . In this respect, hydrogels are designed to hydrogel/microfibre composites thus offer a mechanically and provide the cells with a fully hydrated three-dimensional (3D) biologically favourable environment for the engineering of tissues. environment comparable to the extracellular matrix of native 3,4 tissues . When embedded within hydrogels, cells need to form a Results tissue-specific matrix that is capable of repairing or regenerating Fabrication of PCL microfibre scaffolds. PCL scaffolds of 1 mm the damaged tissue. However, hydrogels often have inadequate height were 3D-printed with a reproducible quality using melt- mechanical properties, which are either unfavourable to electrospinning writing (Fig. 1a and Supplementary Movie 1). embedded cells or make them too weak for application in the Scaffolds were fabricated with fibre spacings between 0.2 mm 5–7 musculoskeletal system . (Fig. 1b) and 1.0 mm (Fig. 1c). The fibres were well aligned and The performance of hydrogel implants is expected to be fused at the 0-to90-oriented junctions (Fig. 1d). The fidelity of superior when they closely match the mechanical properties of fibre deposition decreased for constructs higher than 1 mm, 8,9 their target tissue . This is based on the principle that cells in probably because of the electric isolating effect of the scaffold that native tissues are responsive to different types of mechanical 10–12 stresses, such as compression, tension and shear . These properties, for example stiffness, can be improved by increasing PCL 2,3,13 the hydrogel polymer concentration or crosslink density or by the formation of tissue-specific extracellular matrix Heated jacket 2,14,15 before implantation of the graft . However, this generally 2,3 compromises the biological performance of the hydrogel or requires a long and costly period of pre-culture. Boxed structure Tissue engineers have, therefore, been inspired by the High voltage architecture of native tissues, which derive their unique mechan- ical properties from a fibrous protein framework that supports Taylor cone the aqueous component, forming a complex composite . For 17–19 example, hydrogels have been reinforced with nanofibres , 20 8,21,22 23 microfibres and woven or non-woven scaffolds. The mechanical properties of hydrogels in such composites are less demanding; therefore, they can be processed with a low crosslink density, which is beneficial for cell migration and the formation 2,3 Translating collector of neo-tissue . Recent composite systems include non- woven scaffolds manufactured via solution-electrospinning 9,24–27 28–33 techniques and scaffolds fabricated via 3D-printing . Owing to the small fibre diameters that can be obtained, electrospun meshes have the potential to mimic native tissue extracellular matrix structures, including its mechanical properties. However, a disadvantage of traditional solution electrospinning techniques is the limited control over network architecture. Recently, melt electrospinning has been used in a direct writing mode, and may overcome this limitation, with the layer-by-layer assembly of the low-diameter fibres, permitting 34,35 tight control over the network architecture . This study aims to mechanically reinforce gelatin metha- crylamide (GelMA) hydrogels with organized, high-porosity poly(e-caprolactone) (PCL) fibre scaffolds. GelMA is a hydrogel 98% platform with recent use that allows for matrix deposition by 97% embedded cells, for example, vascular networks or cartilage 96% 14,36,37 matrix . Reinforcing scaffolds with different porosities are 95% fabricated from medical-grade PCL by melt-electrospinning in a direct writing mode . Fibrous hydrogel composites are then 94% fabricated by infusing and crosslinking GelMA within the PCL 93% scaffolds. The stiffness and recovery under cyclic compression of 0% these composites are analysed depending on PCL scaffold 0 0.2 0.4 0.6 0.8 1.0 porosity and the degree of crosslinking of the GelMA. A Fibre spacing (mm) mathematical model is developed using the scaffold parameters, and subsequently used to simulate and predict the degree of Figure 1 | Fabrication of microfibre scaffolds. 3D scaffolds were fabricated hydrogel reinforcement. Finally, human primary chondrocytes from PCL by 3D printing, that is, melt-electrospinning in a direct writing are encapsulated in the composite constructs and their viability, mode. (a) Thin PCL fibres were stacked in a 0–90 orientation through combined extrusion and an electrostatic field between the dispensing morphology and chondrogenic differentiation were investigated in an in vitro culture assay under physiological compressive needle and the translating collector plate. Several fibre spacings were applied ranging from (b) 0.2 and (c) 1.0 mm as visualized with loading. We demonstrate that hydrogels are reinforced in a synergistic manner with high-porosity microfibre scaffolds. The stereomicroscopy (scale bar, 1 mm). (d) Detailed image of the fibres that stiffness of the composites approaches that of articular cartilage, fused at the cross-sections (fibre spacing 1.0 mm, scale bar, 200 mm). while maintaining a relevant elasticity. Chondrocytes embedded (e) The porosity of the scaffolds varied between 93 and 98% depending on in the composite constructs respond positively to a repetitive the set fibre spacing. 2 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. Porosity NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE had already been deposited on the collector plate. Therefore, two (APS/TEMED) and 15.8 2.0 kPa for alginate (Fig. 2b). When scaffolds were stacked in the current study to obtain 2-mm-high the compressive strength of the 93% porous PCL/hydrogel constructs. The scaffold porosity could be controlled within the composite was investigated, the construct stiffness was increased range of 93–98% by varying the set fibre spacing between 0.2 30-fold, up to 214 24 kPa for GelMA with 12.5 mM APS/ and 1.0 mm (Fig. 1e). Flow rates of 18, 72 and 180 mlh TEMED, and 54-fold, up to 405 68 kPa, when the GelMA was ± ± corresponded to fibre diameters of 19.4 1.7, 48.5 8.2 and crosslinked with 25 mM APS/TEMED (Fig. 2c,d). Alginate 88.5 7.6 mm, respectively. Larger diameter scaffolds of the same hydrogels also showed synergistic reinforcement (15-fold up to polymer were 3D-printed with traditional melt extrusion-based 240.7 37.8 kPa) that was comparable to GelMA gels (Fig. 2e). methods as a control for the mechanical analyses. These scaffolds The reinforcing effect in the composites was dependent on the had a fibre diameter of 219.7 14.2 and a porosity ranging from porosity of the PCL scaffold, which ranged from 93 to 98%. 89 to 72%, depending on the set fibre spacing. Interestingly, increasing the fibre roughness and surface area by etching of the PCL scaffold did not increase the degree of rein- Stiffness of reinforced GelMA and alginate hydrogels. forcement (Supplementary Fig. 1). In fact, prolonged etching A synergistic effect was observed on construct stiffness (strain times resulted in mass loss of PCL, and hence decreased the rate 25% min ) when GelMA and alginate hydrogels were construct stiffness. Therefore, we concluded that the strength reinforced with the microfibre PCL scaffolds (Fig. 2). The stiffness increase seen in the composites was not due to hydrogel/fibre of PCL scaffolds alone increased from 1.1 0.3 kPa for 98% bonding. porosity to 15.2 2.2 kPa for the 93% porosity scaffold (Fig. 2a). The stiffness of the scaffolds was comparable to that of hydrogels: Comparison of reinforced gels with articular cartilage. The ± ± 7.1 0.5 and 7.5 1.0 kPa for GelMA crosslinked with either average stress–strain curve of native cartilage samples (strain rate 12.5 or 25 mM ammonium persulfate/tetramethylethylenediamide 25% min ; n ¼ 8) that were harvested from the patellofemoral * * 10 10 93% 94% 96% 97% 98% Porosity of PCL scaffold * 500 * 500 * 500 * * * * * 400 400 300 300 300 200 200 200 100 100 100 0 0 0 Porosity of reinforcing scaffold Porosity of reinforcing scaffold Porosity of reinforcing scaffold GelMA Native cartilage reinforced 84% GelMA reinforced 93% GelMA reinforced 98% GelMA PCL scaffold 0% 10% 20% 30% 40% 50% Strain Figure 2 | Reinforcing effect of high-porosity 3D-printed PCL scaffolds incorporated into hydrogels. (a) Compressive moduli of PCL scaffolds were in the same range as those of (b) hydrogels alone. GelMA reinforced with PCL scaffolds and crosslinked with either (c)25mM or (d) 12.5 mM APS/TEMED were one order of magnitude stiffer than the scaffolds or gel alone; (e) a comparable degree of reinforcement for reinforced alginate gels (*Po0.05, one-way ANOVA with Bonferroni correction). (f) Stress–strain curves of GelMA, the PCL scaffold and reinforced GelMA, approaching the curve of native cartilage (yellow). Scaffolds fabricated from thick fibres (that is, 84%) were stiffer than native cartilage, but often disintegrated at a strain of B10% (all groups n ¼ 5, mean s.d.; cartilage n ¼ 8). Here GelMA was crosslinked with 12.5 mM APS/TEMED and the strain rate applied was 25% min . NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 3 & 2015 Macmillan Publishers Limited. All rights reserved. GelMA (25 mM) GelMA (12.5 mM) Alginate Stiffness (kPa) Stress (kPa) Stiffness (kPa) 93% 94% 96% 97% 98% gel Stiffness (kPa) 93% 94% 96% 97% 98% gel Stiffness (kPa) Stiffness (kPa) 93% 94% 96% 97% 98% gel ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 groove of an equine knee joint was plotted in Fig. 2f together with horizontally and places the surrounding fibres under tension. the average (n ¼ 5) stress–strain curves for GelMA, high-porosity We hypothesize that the fibres’ resistance to elongation under PCL scaffolds and GelMA reinforced with 98%, 93% and 84% compression loading results in a pre-stress of the scaffold fibres. porosity PCL scaffolds, respectively. Clearly, GelMA gels were This resistance leads to the observed increase in stiffness, as the much softer than the cartilage. However, the stiffness and stiffness of bulk PCL is approximately four orders of magnitude deformation profile (shape of the stress–strain curve) of GelMA higher than that of the gels. In order to confirm our hypothesis, could be tailored to that of native cartilage by reinforcement with we developed a mathematical model that enables us to make the high-porosity PCL scaffolds. While thick-fibre scaffolds (porosity following prediction for the construct stiffness, C: 72–88%), as fabricated with traditional 3D-printing methods, r EN were stiffer than native cartilage, they broke atB10% strain when C ¼ ð2Þ ðÞ 3=2 large fibre spacings were applied (41.25 mm). This sudden loss 2RðÞ 1  l of integrity was not observed for native cartilage or gels reinforced with high-porosity scaffolds (Fig. 2f). where N denotes the number of fibres in the construct, E the Young’s modulus of the reinforcing polymer, r the fibre radius, R the construct radius and l the axial strain (expressed as a Time-dependent stress response of reinforced gels and cartilage. fraction of the initial height). A derivation and more detailed The stress response to isostrain showed that GelMA reinforced explanation of the model are provided in the Supplementary with a 93% porosity scaffold has a higher modulus on initial Table 1 and Supplementary Note 2. loading than cartilage (Fig. 3). However, relaxation is faster, and to a lower equilibrium modulus, than cartilage. The modulus of GelMA was much lower during the isostrain at any time. Analysis Video analysis of the compression cycle of reinforced gels. The of the stress curves at isostrain indicated that the reinforced GelMA model was further validated by lateral video imaging of the constructs could not be adequately described within an isotropic compression cycle of the GelMA/PCL composites, which revealed poroelastic framework, but that a bi-exponential function: elongation (stretching) of PCL microfibres and a concave profile at the sides of the construct because of lateral displacement K t  K t 1 2 S ¼ S þ S e S e ð1Þ 0 1 2 of water from GelMA (Supplementary Movie 2, 97% porosity provided an accurate, phenomenological characterization of the scaffold). The microfibres stretched 9 1% on 30% axial strain. In time-dependent response of the construct, where S is the stress, t is addition, from a top view it was observed that the PCL scaffold time and the constants S and K have units of stress and time, area expanded 17% and the aqueous component expanded 23% i i respectively. A nonlinear least squares fitting algorithm was used to (including exudation of water; Fig. 4a). obtain the best fitting parameters for S and K from the 9% iso- i i strain curves as detailed in Supplementary Fig. 2 and Microfibre compared with thick-fibre composites. The Supplementary Note 1. We report E ¼ S /strain as elastic moduli, i i reinforcing effect was only observed in high-porosity scaffolds so that E is the equilibrium modulus and E and E are ‘transient’ 0 1 2 fabricated from thin fibres (diameter o48.2 mm), as indicated in moduli that characterize the time-dependent response of the Fig. 4b. The 3D-printing of thicker fibres (diameter 488.5 mm) material so that the ‘peak’ equilibrium modulus is obtained by resulted in a significantly lower scaffold porosity, ranging from 88 summing the E (Table 1). to 72% (PCL fraction of 12% to 28%, respectively), depending on the fibre thickness and spacing between the fibres. The stiffness of Modelling the fibre reinforcement of hydrogels. The mechan- the thick-fibre scaffolds ranged from 1.8 0.2 MPa for 88% ism we propose to explain the synergistic reinforcement is porosity to 16.1 1.7 MPa for 72% porosity. These scaffolds had illustrated in Fig. 4a. Hydrogels can be reasonably described as stiffnesses similar to those of their composite counterparts with incompressible ; therefore, the volume of the hydrogel is crosslinked GelMA. The compression cycle of GelMA reinforced conserved and vertical compression must be accompanied by with an 82% porosity scaffold is shown in Supplementary Movie horizontal expansion. As a composite construct is loaded, each 3. Exudation of fluid was observed; while the relatively thick ‘semi-confined cell’ of hydrogel (right column in Fig. 4a) expands scaffold fibres were compressed, they did not elongate. 3,000 Recovery of GelMA constructs after compression. The recovery GeIMA Fast loading phase of (reinforced) GelMA constructs after repetitive axial compres- Reinforced GeIMA sion was measured. Figure 5a–c shows that GelMA, as well as Cartilage reinforced GelMA constructs, are fully elastic after 20 cycles of 2,000 20% strain. The stress–strain curves of the initial and final loading cycle are shown. GelMA also fully recovered after repetitive 50% axial strain (Fig. 5d). However, GelMA reinforced with 93% Equilibrium phase porosity PCL microfibre scaffolds showed significantly decreased 1,000 resistance when axially deformed by over 35% (Fig. 5e). Chondrogenic differentiation of embedded human chondrocytes. Haematoxylin and eosin-stained sections revealed that chon- 01 2 34 5 drocytes retained their spherical morphology within the fibre- Time (min) reinforced hydrogels after 7 days (Fig. 6c). Chondrocytes were Figure 3 | The time-dependent stress response of (reinforced) GelMA also homogenously distributed throughout the construct as gel and articular cartilage at 9% isostrain. Reinforced GelMA displays a shown by the 4 ,6-diamidino-2-phenylindole (DAPI)-stained cell high modulus compared with cartilage on direct loading. The modulus of nuclei (Fig. 6d). Chondrocytes maintained high cell viability on reinforced gels in the equilibrium phase, however, is closer to gel than to days 1 and 7 when crosslinked with 12.5 mM APS/TEMED cartilage. These plots are representative for the series of isostrains that (Fig. 6e). Following 14 days of culture, hydrogels (GelMA with were consecutively performed for 15 min. 0.5% hyaluronic acid (Lifecore, USA; GelMA-HA) for improved 4 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. Modulus (kPa) NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE Table 1 | Fit of a bi-exponential function to the time-dependent stress response. Equilibrium modulus Fast decay Slow decay 3  1 3  1 E (kPa) E (kPa) k (10 s ) E (kPa) k (10 s ) 0 1 1 2 2 ± ± ± ± ± GelMA 12 1.7 0.82 0.11 610 310 7.8 0.43 1.5 0.76 ± ± ± ± ± Reinforced GelMA 130 7.8 2700 160 180 9.8 430 49 28 1.1 ± ± ± ± ± Cartilage 440 140 300 160 110 45 430 200 4.9 2 GelMA, gelatin methacrylamide Estimates of the moduli E and the fast and slow decay rates k obtained by fitting the stress response at 9% isostrain of GelMA, reinforced GelMA and cartilage to a bi-exponential function (n ¼ 3, i i mean s.d). the quantitative reverse transcriptase–PCR (qRT–PCR) data indicated that for chondrocytes within the fibre-reinforced hydrogels, ACAN (gene expression for glycosaminoglycan (GAG) synthesis) and COL1A1 (gene expression for collagen type I) mRNA levels were significantly upregulated in the compressed (‘C’) groups compared with the control, non- compressed groups (‘NC’; Fig. 6f–i; Po0.05). Further, chondrocytes in the reinforced gels were more responsive to loading regime than cells in GelMA alone. Stiffness values were B400 kPa in cell-laden fibre-reinforced hydrogels (Fig. 6j). There was no significant effect of reinforcement or compression on construct stiffness and glycosaminoglycan (GAG) content (Fig. 6k). The chondrocytes within the gels exhibited pericellular Axial, unconfined compression collagen type I and II deposition in all groups, with no discernible differences between the groups (Supplementary Fig. 4). Discussion In this study, high-strength composite constructs were fabricated by combining 3D-printed high-porosity scaffolds with a hydrogel. These composite hydrogels can be customized to yield a wide 100,000 White = scaffold range of mechanical properties and, from a biological point of Grey = scaffold + gel view, have the capacity to support cell proliferation and 10,000 Fibre diameter of extracellular matrix production. The PCL scaffolds were built scaffold (µm) by micrometre-scale fibres that were organized and intercon- 1,000 Native articular cartilage 19.4 ± 1.7 nected through the melt-electrospinning writing process. With 48.2 this unique and new 3D-printing technique, pre-set network 88.5 architectures can be realized in a direct writing mode . This 219.7 ± 14.2 technique allows fibres to be printed well below the limits of classical melt-extrusion-based 3D-printing technologies such as fused deposition modelling , with filament diameters as small as 1 35 5 mm, instead of 100 mm or larger . 0% 10% 20% 30% Hydrogels have previously been reinforced with solution- PCL fraction (1-porosity) 9,24,25,27 electrospun nonwovens . However, most traditional Figure 4 | The mechanism of hydrogel reinforcement with organized solution-electrospun meshes have disadvantages from both a high-porosity scaffolds. (a) PCL microfibre scaffolds (blue in schematic) mechanical and biological point of view, for example, fibres are serve as a reinforcing component to GelMA hydrogel (yellow in schematic). not fused and hence slide under loading meshes are usually not When axial compression is applied to the reinforced hydrogels, the stiff thin thicker than 100 m and have a pore size of less than 5 mm and are, 25,27,40 scaffold fibres stretch on lateral displacement of the hydrogel. This therefore, too dense for cell infiltration . To overcome these mechanism provides the composites with a high stiffness and elasticity issues related to traditional solution electrospinning, fibres have (scale bars, 1 mm). (b) Moduli of scaffolds and scaffold/gel composites as a been collected in an earthed ethanol bath, yielding nonwovens 9,24 function of porosity, showing the synergistic increase in stiffness was only with high porosities . Hydrogel stiffness was increased fourfold observed for thin-fibre scaffolds with a high porosity (polymer fraction when reinforced with these randomly organized nonwovens, 24,25 20 2–7% ¼ porosity 98–93% (highlighted in red)), fabricated with melt- irrespective of the porosity . Silk microfibres or carbon electrospinning writing (MEW). Fused deposition modelling (FDM) nanotubes also formed porous reinforcing structures for scaffolds were fabricated from 10-fold thicker fibres, resulting in a higher hydrogels, permitting cellular infiltration and differentiation. stiffness; however, no synergistic reinforcement was observed (mean of However, the stiffness of the gels increased only one- to threefold, n ¼ 5). as the fibres were not fused. In contrast, in our current composite model consisting of a relatively soft hydrogel (7.1–15.8 kPa) and a highly porous and soft PCL scaffold (1.1–15.2 kPa), the maximum chondrogenic differentiation and GelMA-HA with 93% porous stiffness obtained was a factor of 50 larger (405 kPa) than that of PCL mesh) were subjected to physiological loading cycles for the hydrogels alone, on a strain rate of 25% min . another 14 days. The stress response to the 1 Hz dynamic loading The stiffness of cartilage and (reinforced) hydrogels is strongly 41,42 regime is presented in Supplementary Fig. 3. Analysis of strain-rate-dependent . We found that the reinforced NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 5 & 2015 Macmillan Publishers Limited. All rights reserved. Stiffness (kPa) 30% Strain Uncompressed ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 GelMA GelMA + 97% porosity mesh 20% Strain 15 20% Strain Cycle 1 Cycle 1 Cycle 20 Cycle 20 5% 10% 15% 20% 5% 10% 15% 20% Strain Strain GelMA + 93% porosity mesh GelMA 20% Strain 25 50% Strain Cycle 1 Cycle 1 Cycle 20 Cycle 20 0 0 5% 10% 15% 20% 10% 20% 30% 40% 50% Strain Strain GelMA + 93% porosity mesh 20–40% Strain 20% 25% 30% 20 35% 40% 5% 10% 15% 20% Strain Figure 5 | Cyclic compression testing of GelMA and reinforced GelMA constructs. The stress–strain curves of the first and last of 20 cycles of 20% compression are shown for (a) GelMA alone, (b) GelMA reinforced with a 97% porosity PCL scaffold and (c) GelMA reinforced with a 93% porosity PCL scaffold. For all groups no decay in stress was observed. (d) GelMA constructs also recovered after 20 cycles of 50% axial compression; however, (e) GelMA reinforced with a 93% porous PCL scaffold required substantially less force to be compressed to 20% strain, after being compressed to over 30% strain (arrow; mean of n ¼ 3). hydrogels were one order of magnitude stiffer than cartilage on be a combined effect of the osmotic pressure of fixed charges on initial fast loading at 9% strain. However, the modulus of proteoglycans and an organized extracellular matrix. reinforced hydrogels displayed a steep decrease to an equilibrium Etching the high-porosity scaffolds did not result in a further value roughly one-third the equilibrium modulus of cartilage, yet increase in stiffness of the composites. This approach was on the still one order of magnitude larger than the equilibrium modulus basis of our previous finding that covalent attachment of the of hydrogel alone. hydrogel component to a scaffold that was fabricated from The effect of reinforcement can be explained by the highly modified, methacrylated PCL, resulted in increased construct organized fibre architecture of the scaffolds. The mathematical strength . It should be noted that etching results in an increase model we developed demonstrated that the hydrogel places the in van de Waals forces but does not establish covalent bonds PCL fibres under tension on axial compression, and predicts a between the hydrogel and the reinforcing scaffold. theoretical upper bound on the attainable stiffness. However, the Infiltrating hydrogel into scaffolds that had been fabricated theoretical stiffness of the reinforced hydrogels is one order of with traditional 3D melt-printing techniques with thick fibres magnitude larger than that observed experimentally. This is (Z88 mm) did not show a significant mechanical effect. Axial reasonable since an idealized construct was considered, in which loading of these constructs requires compression of the PCL the fibres are initially taut at zero strain, the hydrogel is purely through a strong vertical column of fibre crossings, which is not elastic and there is no slip between the fibres and the hydrogel. In easily deformed. Video imaging of the compression cycle fact, in our experimental set-up, we showed that very little water, confirmed that thick fibres were not stretched and water flowed or hydrogel, was compressed out of the constructs; therefore, the out of these scaffolds without providing a synergistic effect. In fibres did not completely lock the hydrogel. The stress relaxation contrast, the columns of fine fibre crossings are easily deformed on isostrain (Fig. 3) reflects the exudation of the aqueous under axial loading; however, they are supported by the hydrogel component (water or gel) from the reinforcing scaffold. For component in the composites. Supplementary Movie 2 shows cartilage we observed a higher equilibrium modulus, which may some local distortion of the columns, but no large-scale buckling 6 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. Stress (kPa) Stress (kPa) Stress (kPa) Stress (kPa) Stress (kPa) NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE 0.06 0.0012 100% NC NC NC 0.05 C 0.0010 C 25 C 80% 0.04 0.0008 20 60% 0.03 0.0006 15 40% 0.02 0.0004 10 20% 0.01 0.0002 5 0% 0 0 0 Day1 Day7 GeIMA 93% GeIMA 93% GeIMA 93% 0.030 0.05 1,000 NC NC NC C C 0.025 800 0.04 0.020 600 0.03 0.015 400 0.02 0.010 200 0.01 0.005 0 0 GeIMA 93% GeIMA 93% GeIMA 93% Figure 6 | Differentiation of chondrocytes embedded in reinforced GelMA-HA hydrogels. (a) Stereomicroscopy image of a GelMA-HA gel and (b)ofa GelMA-HA gel reinforced with a PCL scaffold (93% porosity) on day 1 (scale bars, 2 mm). (c) Haematoxylin/eosin staining after 7 days of culture shows that chondrocytes remained within the GelMA component and retained a round morphology (scale bar, 500 mm, inlay, 200 mm). (d) DAPI staining confirms a homogenous distribution of the cells throughout the construct (scale bar, 200 mm). (e) Chondrocytes remained viable over 7 days of culture in the gels. Gene expression analysis for (f) ACAN, (g) COL2A1, (h) COL1A1 and (i) COL10A1 on day 14 for compressed (C) and non-compressed (NC) groups (n ¼ 4, mean s.d.; *Po0.05, two-way ANOVA). All gene expression levels were normalized to the housekeeping gene b2 microglobulin (B2M). (j) Stiffness of the constructs following long-term culture (n ¼ 4). (k) GAG/ww values on day 28 (n ¼ 5). of the columns; therefore, the microfibres are put under tension. to native cartilage. The porosity of these woven scaffolds was We further note that this axial loading reduces the horizontal 70–74%, compared with 93–98% in the present study. The pore size, improving gel confinement. To our knowledge, we are stiffness of the woven composites was reported up to 0.2 Mpa (at the first to show the synergistic effect of a well-designed scaffold/ equilibrium stress), which was twice as stiff as the scaffold hydrogel system; the work shown in this paper could therefore without the gel . Both the woven and the melt-electrospun lead to a paradigm shift in the field. composites showed axial recovery after compression for 10% and Reinforced GelMA hydrogels possess stress–strain curves that 20%, respectively. The water that was compressed out of the closely resemble those of healthy articular cartilage. In addition, scaffolds was likely reabsorbed during the relaxation phase, in absolute terms, the stiffness of the biodegradable composites comparable to fluid dynamics in the articular cartilage . Our was comparable to the stiffness of articular cartilage, which has reinforced hydrogels showed decreased resistance if compressed been reported to range from 400 to 800 kPa (refs 43–45). On the by 35% or more, which may be because of fractures or other hand, the stiffness of scaffolds fabricated with traditional delamination in the PCL scaffold junctions. For translational 3D-printing was one order of magnitude larger than that of the purposes, it is important to realize that native cartilage is 30,33 articular cartilage, which is consistent with previous reports . exposed to 15–45% axial deformation under long-term static The porosity of these scaffolds ranged between 72 and 89%, compression . whereas the porosities reported in the literature range from 28.9 Incorporation of organized fibrous PCL scaffolds within a well- and 91.2% (refs 39,46,47). Nevertheless, when aiming to fabricate characterized hydrogel system makes it possible to culture cells in high-porosity scaffolds (480%) from thick fibres, large fibre a customizable, mechanically diversified environment. The spacing (41.5 mm) is required, which causes scaffold hydrogel component of the composite constructs will degrade 14,29 disintegration under strains of B10%. In addition, the thicker within months , allowing the regeneration of new tissue, while fibres occupy a relatively high volume that is inaccessible for the PCL component degrades within years forming a temporary tissue formation until the scaffold has degraded. reinforcing network to the new tissue. Under physiological Hydrogels reinforced with woven scaffolds, composed of either compressive loading of 20% strain and 1 Hz, our gene expression 8 21,22 polyglycolic acid or PCL , have previously been reported to data suggest that chondrocyte expression of matrix mRNAs is possess tensile, compressive and shear properties comparable significantly upregulated in composite hydrogels compared with NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 7 & 2015 Macmillan Publishers Limited. All rights reserved. Cell viability COL10A1 mRNA (normalized to B2M) Stiffness (kPa) ACAN mRNA (day 28) (normalized to B2M) COL2A1 mRNA –1 GAG/ww (µg µg ) (normalized to B2M) (day 28) COL1A1 mRNA (normalized to B2M) ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 non-reinforced, weak hydrogels. These results highlight the Directly after mixing in the TEMED, the GelMA solution as described above was injected in the mold untill all PCL scaffolds were fully infused. Crosslinking of the importance of developing cell-culture platforms and tissue- GelMA was allowed in this mold at 37 C for 1 h. Cell-free samples were then engineered constructs that better mimic the in vivo stored in PBS at 37 C. Chondrocyte-laden samples were kept in a chondrocyte mechanoenvironment of natural articular cartilage. Still, the expansion medium. In order to evaluate the effect of etching time of the scaffolds exact mechanisms of stress transfer to the cells in fibre-reinforced on construct stiffness, GelMA was reinforced with scaffolds etched for 0–16 h (all n ¼ 5) and crosslinked with 25 mM APS/TEMED. In addition, 93–98% porosity hydrogels remain to be investigated. PCL scaffolds (all n ¼ 5) were infused with alginate in the mold. The alginate In conclusion, the current work represents a significant step composite constructs were then crosslinked by immersion in a 102-mmol towards developing biomechanically functional tissue constructs. calciumchloride solution for 30 min. The stiffness of the constructs was significantly enhanced, achieving values similar to those of native articular cartilage, Harvest of equine cartilage. Full-thickness cartilage was harvested from the knee by combining hydrogels with 3D-printed, high-porosity melt- joint from one equine donor (10 years old) with macroscopically healthy cartilage. electrospun PCL scaffolds. This synergistic effect could be This was performed after consent of the horse owner and according to the ethical guidelines of the Utrecht University. Equine cartilage was used because of the modulated by altering the porosity of the reinforcing scaffolds. availability and its similarities to human cartilage . Cylindrical samples with a The composite constructs have a strong elastic component diameter of 5 mm and a height of 2 mm (n ¼ 8) were taken with a biopsy punch and recover after physiological axial strains. Finally, human from the medial patellofemoral groove and stored in PBS for up to 3 h. chondrocytes encapsulated in the GelMA/PCL composites were found to be more responsive to mechanical loading, which led to Mechanical analyses. The stiffness (or compressive modulus) of GelMA and significant changes in gene expression in vitro. alginate hydrogels, PCL scaffolds, the composite constructs and articular cartilage (all n ¼ 5) was measured by single uniaxial unconfined compression in air at room temperature, after 1–3 days of submersion in PBS from their preparation. Methods We confirmed that for single-cycle compression testing, the absence of a PBS Fabrication of PCL microfibre scaffolds. Several scaffolds were 3D-printed from immersion bath does not influence the stress responses of hydrogel samples. The medical-grade PCL (PURAC, Gorinchem, the Netherlands) with a custom-made effect of crosslinking with 12.5 or 25 mM APS/TEMED was tested. The stress–strain melt-electrospinning device . PCL was heated to 103 C in a 2 ml polypropylene curves of (reinforced) GelMA that was crosslinked with 12.5 mM APS/TEMED syringe and extruded at a rate of 18 mlh using a syringe pump. In combination were compared with cartilage samples. Over a period of 2 min, a force ramp (axial with an electrostatic field of 8–10 kV between the syringe needle tip (23G) and the  1 strain rate ca. 25% min ) was applied to the samples employing a microtester collector plate (3 cm distance), a stable PCL jet was obtained. Defined scaffold (Instron, Melbourne, Australia) or a Dynamic Mechanical Analyser (DMA 2980, architectures, with dimensions up to 120  120  1 mm, were realized through TA Instruments, New Castle, DE, USA). The stiffness was calculated from the computer-aided movement of the aluminium collector plate at a speed of linear derivative of the stress–strain curve at 12–15% strain. In the low-porosity 1,400 mm min , using the Mach3 motion control software (Artsoft, USA). scaffolds that were fabricated with traditional 3D-printing, the 6–9% strain region A predefined 0–90 architecture was imposed, with a fibre spacing of 0.2, 0.4, 0.6, was taken, as these constructs often disintegrated when strained beyond 10%. The 0.8 or 1.0 mm. The melt-electrospinning writing was terminated when scaffolds compression cycle of GelMA reinforced with 97 and 82% porosity scaffolds was reached a height of 1.0 mm, to ensure maximum quality of the architecture. To captured from the side (hand-held digital microscope 1.3 MP, Meade instruments, investigate the effect of the PCL flow rate on fibre diameter and scaffold porosity, Europe GmbH & Co, Rhede, Germany) in order to analyse the lateral expansion of we printed constructs at 4  and 10  flow rate, with a fibre spacing of 0.4 and both components of the composite construct. In addition, images from the top of 1.0 mm, respectively. Cylindrical samples were extracted from the scaffolds with a these constructs were taken with a stereomicroscope, when uncompressed and at 5-mm diameter biopsy punch. Two scaffolds were stacked in order to achieve a 30% strain between two glass slides. height of 2 mm, which is comparable to the thickness of human cartilage in the In order to test the time-dependent stress response, GelMA, GelMA reinforced knee joint . The porosity and mechanical properties of the melt-electrospun with 93% porosity scaffolds and articular cartilage samples were subjected to a scaffolds were compared with scaffolds fabricated with traditional 3D-printing series of isostrain steps . Samples were 2% pre-strained for 5 min, followed by technologies. To this end, scaffolds were 3D-printed from PCL with a BioScaffolder strains of 6, 9 and 13% that were consecutively applied for 15 min each. The system (SYS þ ENG, Salzgitter-Bad, Germany) as described previously . Briefly, experiments were performed in PBS and the stress response of all samples was PCL was heated till 70 C and extruded with an Auger screw through a 27-gauge recorded. The modulus on fast initial loading and the equilibrium modulus were needle. Scaffolds measuring 40  40  2 mm were fabricated with a fibre spacing of extracted, and the stress decay rate was estimated by fitting to a bi-exponential 0.75, 1.0, 1.25, 1.5 and 1.75 mm. The quality of the scaffolds was imaged both with function. stereomicroscopy and Scanning Electron Microscopy (SEM, Hitachi TM3000, The resistance to axial deformation of GelMA and GelMA reinforced with PCL Japan and Quanta 200, FEI, Milton, Australia). The fibre diameter was measured constructs (porosity 93 and 97%) was measured after a cyclic (20  ) axial strain of with the ImageJ software (National Institutes of Health, USA). The porosity of the 20% (Allround-line Z020, Zwick Roell, Germany). In addition, the resistance after PCL scaffolds was determined gravimetrically. 20 cycles of axial deformation of 50% was measured for GelMA constructs. The recovery measurements were performed in PBS and constructs were allowed to recover for 1 min after every cycle. GelMA constructs reinforced with a 93% porous Etching of PCL microfibre scaffolds. PCL scaffolds were etched in order to scaffold underwent compression with incremental maximum strains of 20–40% increase its hydrophilicity and surface area, which could potentially contribute to (with 5% increments) in order to analyse the maximal strain that could be exerted the stiffness of reinforced GelMA constructs. Cylindrical PCL scaffolds with a before irreversible damage would occur. porosity of 94% were treated with 70% ethanol and subsequently etched with 5 M sodium hydroxide for 0, 1, 2, 4, 8 and 16 h. After etching, the scaffolds were washed in deionized H O until the pH reached 7.4, and then air-dried. The effect of etching Modelling the fibre reinforcement of hydrogels. A mathematical model was was evaluated by measuring the fibre diameter (ImageJ) from SEM images, and by constructed to investigate further the mechanisms by which the 3D-printed scaf- assessing the weight from 2  5 mm diameter scaffolds (expressed in relative weight folds reinforce the hydrogels. There is an extensive literature on the modelling of 41,42 loss). Since mild etching will increase hydrophilicity and may facilitate perfusion of fibre-reinforced biological materials, with recent cartilage-focussed examples . GelMA through the PCL scaffold, 2 h of etching was performed for all other Fibre-reinforced materials are often modelled by assuming that it is reasonable to experiments presented. define a continuously varying fibre density and orientation and then making the material properties a function of these. In this instance, given that we know the arrangement and approximate number of fibres in each plane of the material, we Preparation of GelMA and alginate gels. GelMA was synthesized by reaction of take the more direct approach of considering how each fibre stretches as the type A gelatin (Sigma-Aldrich, St Louis, MO, USA) with methacrylic anhydride at scaffold deforms. Our model takes into account the number of fibres in the scaffold 50 C for 1 h as previously described . GelMA was dissolved in PBS at 60 Cata and the fibre diameter, the Young’s modulus of PCL and the construct dimensions. concentration of 10% (w/v) containing 12.5 or 25 mM APS. TEMED (12.5 or In the model, the composite construct was viewed as an elastic solid, in which the 25 mM) was added followed by 5 s vortexing in order to initiate crosslinking of the PCL fibres are placed under tension by the hydrogel on axial compression. GelMA. Sodium alginate (IMCD, Amersfoort, the Netherlands) was dissolved in PBS at 3% (w/v), and used as a control gel to GelMA for the mechanical analyses. Harvest of human chondrocytes. Macroscopically healthy cartilage was harvested either from a discarded talus bone that was resected from a 7-year-old patient Preparation of reinforced hydrogel constructs. The scaffolds were placed in an undergoing an orthopaedic intervention, or from the femoral condyles of knee injection mold that was custom-made from polymethylmethacrylate in order to fit replacement surgery patients (age: 71.0 4.1; n ¼ 6) with consent. This was in 10 cylinders with a diameter of 5 mm and a height of 2 mm. All cylindrical voids in concordance with the institutional code of conduct regarding the use of discarded the mold were interconnected so that gel could be serially perfused. The mold was tissues in the University Medical Center Utrecht, and ethics approval was also sealed with a sheet of polyethylene that was fixed between two glass histology slides. obtained from Queensland University of Technology and Prince Charles hospital 8 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE before sample collection. Cartilage was cut into small slices and washed with PBS Immunohistochemistry. After 28 days, the reinforced GelMA-HA gels were supplemented with penicillin and streptomycin. The cartilage was digested over- paraffin-embedded and sectioned at 5 mm. For antigen retrieval, ready-to-use night in 0.15% type II collagenase at 37 C. The resulting cell suspension was proteinase K solution (Dako) was used for 10 min at 37 C. Sections were filtered (100 mm cell strainer) and washed three times with PBS. Then, cells were blocked with 2% FBS solution before exposure to primary antibodies: I-8H5 resuspended in chondrocyte expansion medium (DMEM supplemented with 10% (MP Biomedicals), dilution 1:300 for collagen type I; II-II6B3 (DSHB), dilution 1  1 fetal bovine serum (FBS), 100 units ml penicillin and 100 mgml streptomycin, 1:200 for collagen type II. Following incubation in fluorescence-labelled goat and 10 ng ml FGF-2) and expanded for 10 days in monolayer cultures (seeding anti-mouse secondary antibody (Alexa Fluor 488, Invitrogen), sections were density 5,000 cells cm ). For the short-term study (7 day), chondrocytes from the mounted with Prolong Gold (Invitrogen) and visualized using a confocal talus bone were used at passage 2, and for the long-term study (28 day) chon- fluorescence microscope (A1R Confocal, Nikon). drocytes obtained from the femoral condyles of the different donors were pooled and used at passage 1. Statistics. All data were expressed as mean s.d. of five independent samples, unless stated otherwise. One-way analysis of variance (ANOVA) with Bonferroni Cell viability. Chondrocytes were mixed in the GelMA/APS solution at a correction was performed to compare the stiffness of the reinforced hydrogels; the concentration of 10 million ml . Crosslinking was performed with different effect of scaffold porosity was compared for every hydrogel and the effect of the proportions of APS/TEMED (12.5/12.5 and 25/25 mM) in order to evaluate hydrogel was compared for every scaffold porosity. Two-way ANOVA was used to cytotoxicity of the redox crosslinking process. The viability of the chondrocytes was evaluate the effect of compression and hydrogel types on gene expression, GAG/ measured for 4 h (day 1) and 7 days after embedding and crosslinking of the ww and stiffness. An independent samples t-test, not assuming equal variances, was GelMA constructs. To visualize cell viability, a LIVE/DEAD Viability Assay performed to evaluate cell viability (SPSS, IBM software, USA). Differences were (Molecular Probes MP03224, Eugene, USA) was performed according to the considered significant when Po0.05. manufacturer’s instructions. The samples were examined using a light microscope with the excitation/emission filter set at 488/530 nm to observe living (green) cells and at 530/580 nm to detect dead (red) cells. Photographs were taken with an References Olympus DP70 camera. The ratio dead/alive was measured by assessing multiple 1. Seliktar, D. Designing cell-compatible hydrogels for biomedical applications. parts of the sample. Science 336, 1124–1128 (2012). 2. Malda, J. et al. 25th anniversary article: engineering hydrogels for Cell morphology. The morphology and location of chondrocytes in the reinforced biofabrication. Adv. Mater. 25, 5011–5028 (2013). GelMA constructs was assessed 7 days after encapsulation. Reinforced GelMA 3. DeForest, C. A. & Anseth, K. S. Advances in bioactive hydrogels to probe and constructs with chondrocytes (n ¼ 3, 93% porosity) were dehydrated through a direct cell fate. Annu. Rev. Chem. Biomol. 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University of Twente for facilitating mechanical testing on their Zwick universal 34. Brown, T. D., Dalton, P. D. & Hutmacher, D. W. Direct writing by way of melt mechanical tester. We thank professor Harrie Weinans for fruitful discussions on car- electrospinning. Adv. Mater. 23, 5651–5657 (2011). tilage biomechanics. The II-II6B3 monoclonal antibody developed by T.F. Linsenmayer 35. Farrugia, B. L. et al. Dermal fibroblast infiltration of poly(epsilon-caprolactone) was obtained from the Developmental Studies Hybridoma Bank developed under the scaffolds fabricated by melt electrospinning in a direct writing mode. auspices of the NICHD and maintained by The University of Iowa, Department of Biofabrication 5, 025001 (2013). Biology, Iowa City, IA 52242. J.V. was supported by a grant from the Dutch government 36. Schuurman, W. et al. Gelatin-methacrylamide hydrogels as potential to the Netherlands Institute for Regenerative Medicine (NIRM, grant nFES0908); biomaterials for fabrication of tissue-engineered cartilage constructs. Macromol. F.P.W.M. was supported by a Marie Curie grant from the European Commission (PIOF- Biosci. 13, 551–561 (2013). GA-272286) and J.M. was supported by the Dutch Arthritis Foundation. This work was 37. Chen, Y. C. et al. Functional human vascular network generated in also supported by the Australia National Health and Medical Research Council and the photocrosslinkable gelatin methacrylate hydrogels. Adv. Funct. Mater. 22, European Community’s Seventh Framework Programme (FP7/2007-2013) under grant 2027–2039 (2012). agreement n309962 (HydroZONES). 38. Hong, W., Zhao, X. H., Zhou, J. X. & Suo, Z. G. A theory of coupled diffusion and large deformation in polymeric gels. J. Mech. Phys. Solids 56, 1779–1793 (2008). Author contributions 39. Hutmacher, D. W. Scaffolds in tissue engineering bone and cartilage. J.V., F.P.W.M., D.W.H. and J.M. designed the study. F.P.W.M., P.D.D. and D.W.H. Biomaterials 21, 2529–2543 (2000). developed the melt electrospinning set-up. J.V., F.P.W.M. and E.M.v.B. fabricated the 40. Pham, Q. P., Sharma, U. & Mikos, A. G. Electrospun poly(epsilon- scaffolds and performed the mechanical experiments. J.E.J. and E.M.v.B. conducted the caprolactone) microfiber and multilayer nanofiber/microfiber scaffolds: dynamic loading experiment. L.S.K. and H.M.B. did the mathematical modelling. characterization of scaffolds and measurement of cellular infiltration. W.J.A.D., P.D.D., D.W.H. and J.M. directed the study. J.V. prepared the manuscript Biomacromolecules 7, 2796–2805 (2006). together with all co-authors. 41. Pierce, D. M., Ricken, T. & Holzapfel, G. A. A hyperelastic biphasic fibre- reinforced model of articular cartilage considering distributed collagen fibre orientations: continuum basis, computational aspects and applications. Comput. Methods Biomech. Biomed. Eng. 16, 1344–1361 (2013). Additional information Supplementary Information accompanies this paper at http://www.nature.com/ 42. Hosseini, S. M., Wilson, W., Ito, K. & van Donkelaar, C. C. How naturecommunications preconditioning affects the measurement of poro-viscoelastic mechanical properties in biological tissues. Biomech. Model Mechanobiol. 13, 503–513 Competing financial interests: The authors declare no competing financial interests. (2014). 43. Athanasiou, K. A., Agarwal, A. & Dzida, F. J. Comparative study of the intrinsic Reprints and permission information is available online at http://npg.nature.com/ mechanical properties of the human acetabular and femoral head cartilage. reprintsandpermissions/ J. Orthopaed. Res. 12, 340–349 (1994). 44. Jurvelin, J. S., Buschmann, M. D. & Hunziker, E. B. Optical and mechanical How to cite this article: Visser, J. et al. Reinforcement of hydrogels using determination of Poisson’s ratio of adult bovine humeral articular cartilage. three-dimensionally printed microfibres. Nat. Commun. 6:6933 J. Biomech. 30, 235–241 (1997). doi: 10.1038/ncomms7933 (2015). 10 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. http://www.deepdyve.com/assets/images/DeepDyve-Logo-lg.png Nature Communications Springer Journals

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Science, Humanities and Social Sciences, multidisciplinary; Science, Humanities and Social Sciences, multidisciplinary; Science, multidisciplinary
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Abstract

ARTICLE Received 21 Aug 2014 | Accepted 16 Mar 2015 | Published 28 Apr 2015 DOI: 10.1038/ncomms7933 Reinforcement of hydrogels using three-dimensionally printed microfibres 1,2 1,2 2 1,2 3 3 Jetze Visser , Ferry P.W. Melchels , June E. Jeon , Erik M. van Bussel , Laura S. Kimpton , Helen M. Byrne , 1,4 2,5 2,6,7 1,2,4 Wouter J.A. Dhert , Paul D. Dalton , Dietmar W. Hutmacher & Jos Malda Despite intensive research, hydrogels currently available for tissue repair in the musculoskeletal system are unable to meet the mechanical, as well as the biological, requirements for successful outcomes. Here we reinforce soft hydrogels with highly organized, high-porosity microfibre networks that are 3D-printed with a technique termed as melt electrospinning writing. We show that the stiffness of the gel/scaffold composites increases synergistically (up to 54-fold), compared with hydrogels or microfibre scaffolds alone. Modelling affirms that reinforcement with defined microscale structures is applicable to numerous hydrogels. The stiffness and elasticity of the composites approach that of articular cartilage tissue. Human chondrocytes embedded in the composites are viable, retain their round morphology and are responsive to an in vitro physiological loading regime in terms of gene expression and matrix production. The current approach of reinforcing hydrogels with 3D-printed microfibres offers a fundament for producing tissue constructs with biological and mechanical compatibility. 1 2 Department of Orthopaeics, University Medical Center Utrecht, Heidelberglaan 100, 3508 GA Utrecht, The Netherlands. Institute of Health and Biomedical Innovation, Queensland University of Technology, 60 Musk Avenue, Kelvin Grove QLD 4059, Queensland, Australia. Mathematical Institute, University of Oxford, Andrew Wiles Building, Radcliffe Observatory Quarter, Woodstock Road, OX2 6GG Oxford, UK. Department of Equine Sciences, Faculty of Veterinary Medicine, Utrecht University, Yalelaan 112, 3584 CM, Utrecht, The Netherlands. Department of Functional Materials in Medicine and Dentistry, 6 7 University of Wurzburg, Pleicherwall 2, 97070 Wurzburg, Germany. Georgia Institute of Technology, North Avenue, Atlanta, GA, 30332, USA. Institute for ¨ ¨ Advanced Study, Technical University Munich, Lichtenbergstrasse 2a, 85748 Garching, Munich, Germany. Correspondence and requests for materials should be addressed to D.W.H. (email: dietmar.hutmacher@qut.edu.au) or to J.M. (email: j.malda@umcutrecht.nl). NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 1 & 2015 Macmillan Publishers Limited. All rights reserved. ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ydrogels are an important class of biomaterials with many mechanical loading regime in a bioreactor culture, in terms applications, including as cell carriers for the engineering of cartilage gene expression and matrix formation. These 1,2 Hof tissues . In this respect, hydrogels are designed to hydrogel/microfibre composites thus offer a mechanically and provide the cells with a fully hydrated three-dimensional (3D) biologically favourable environment for the engineering of tissues. environment comparable to the extracellular matrix of native 3,4 tissues . When embedded within hydrogels, cells need to form a Results tissue-specific matrix that is capable of repairing or regenerating Fabrication of PCL microfibre scaffolds. PCL scaffolds of 1 mm the damaged tissue. However, hydrogels often have inadequate height were 3D-printed with a reproducible quality using melt- mechanical properties, which are either unfavourable to electrospinning writing (Fig. 1a and Supplementary Movie 1). embedded cells or make them too weak for application in the Scaffolds were fabricated with fibre spacings between 0.2 mm 5–7 musculoskeletal system . (Fig. 1b) and 1.0 mm (Fig. 1c). The fibres were well aligned and The performance of hydrogel implants is expected to be fused at the 0-to90-oriented junctions (Fig. 1d). The fidelity of superior when they closely match the mechanical properties of fibre deposition decreased for constructs higher than 1 mm, 8,9 their target tissue . This is based on the principle that cells in probably because of the electric isolating effect of the scaffold that native tissues are responsive to different types of mechanical 10–12 stresses, such as compression, tension and shear . These properties, for example stiffness, can be improved by increasing PCL 2,3,13 the hydrogel polymer concentration or crosslink density or by the formation of tissue-specific extracellular matrix Heated jacket 2,14,15 before implantation of the graft . However, this generally 2,3 compromises the biological performance of the hydrogel or requires a long and costly period of pre-culture. Boxed structure Tissue engineers have, therefore, been inspired by the High voltage architecture of native tissues, which derive their unique mechan- ical properties from a fibrous protein framework that supports Taylor cone the aqueous component, forming a complex composite . For 17–19 example, hydrogels have been reinforced with nanofibres , 20 8,21,22 23 microfibres and woven or non-woven scaffolds. The mechanical properties of hydrogels in such composites are less demanding; therefore, they can be processed with a low crosslink density, which is beneficial for cell migration and the formation 2,3 Translating collector of neo-tissue . Recent composite systems include non- woven scaffolds manufactured via solution-electrospinning 9,24–27 28–33 techniques and scaffolds fabricated via 3D-printing . Owing to the small fibre diameters that can be obtained, electrospun meshes have the potential to mimic native tissue extracellular matrix structures, including its mechanical properties. However, a disadvantage of traditional solution electrospinning techniques is the limited control over network architecture. Recently, melt electrospinning has been used in a direct writing mode, and may overcome this limitation, with the layer-by-layer assembly of the low-diameter fibres, permitting 34,35 tight control over the network architecture . This study aims to mechanically reinforce gelatin metha- crylamide (GelMA) hydrogels with organized, high-porosity poly(e-caprolactone) (PCL) fibre scaffolds. GelMA is a hydrogel 98% platform with recent use that allows for matrix deposition by 97% embedded cells, for example, vascular networks or cartilage 96% 14,36,37 matrix . Reinforcing scaffolds with different porosities are 95% fabricated from medical-grade PCL by melt-electrospinning in a direct writing mode . Fibrous hydrogel composites are then 94% fabricated by infusing and crosslinking GelMA within the PCL 93% scaffolds. The stiffness and recovery under cyclic compression of 0% these composites are analysed depending on PCL scaffold 0 0.2 0.4 0.6 0.8 1.0 porosity and the degree of crosslinking of the GelMA. A Fibre spacing (mm) mathematical model is developed using the scaffold parameters, and subsequently used to simulate and predict the degree of Figure 1 | Fabrication of microfibre scaffolds. 3D scaffolds were fabricated hydrogel reinforcement. Finally, human primary chondrocytes from PCL by 3D printing, that is, melt-electrospinning in a direct writing are encapsulated in the composite constructs and their viability, mode. (a) Thin PCL fibres were stacked in a 0–90 orientation through combined extrusion and an electrostatic field between the dispensing morphology and chondrogenic differentiation were investigated in an in vitro culture assay under physiological compressive needle and the translating collector plate. Several fibre spacings were applied ranging from (b) 0.2 and (c) 1.0 mm as visualized with loading. We demonstrate that hydrogels are reinforced in a synergistic manner with high-porosity microfibre scaffolds. The stereomicroscopy (scale bar, 1 mm). (d) Detailed image of the fibres that stiffness of the composites approaches that of articular cartilage, fused at the cross-sections (fibre spacing 1.0 mm, scale bar, 200 mm). while maintaining a relevant elasticity. Chondrocytes embedded (e) The porosity of the scaffolds varied between 93 and 98% depending on in the composite constructs respond positively to a repetitive the set fibre spacing. 2 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. Porosity NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE had already been deposited on the collector plate. Therefore, two (APS/TEMED) and 15.8 2.0 kPa for alginate (Fig. 2b). When scaffolds were stacked in the current study to obtain 2-mm-high the compressive strength of the 93% porous PCL/hydrogel constructs. The scaffold porosity could be controlled within the composite was investigated, the construct stiffness was increased range of 93–98% by varying the set fibre spacing between 0.2 30-fold, up to 214 24 kPa for GelMA with 12.5 mM APS/ and 1.0 mm (Fig. 1e). Flow rates of 18, 72 and 180 mlh TEMED, and 54-fold, up to 405 68 kPa, when the GelMA was ± ± corresponded to fibre diameters of 19.4 1.7, 48.5 8.2 and crosslinked with 25 mM APS/TEMED (Fig. 2c,d). Alginate 88.5 7.6 mm, respectively. Larger diameter scaffolds of the same hydrogels also showed synergistic reinforcement (15-fold up to polymer were 3D-printed with traditional melt extrusion-based 240.7 37.8 kPa) that was comparable to GelMA gels (Fig. 2e). methods as a control for the mechanical analyses. These scaffolds The reinforcing effect in the composites was dependent on the had a fibre diameter of 219.7 14.2 and a porosity ranging from porosity of the PCL scaffold, which ranged from 93 to 98%. 89 to 72%, depending on the set fibre spacing. Interestingly, increasing the fibre roughness and surface area by etching of the PCL scaffold did not increase the degree of rein- Stiffness of reinforced GelMA and alginate hydrogels. forcement (Supplementary Fig. 1). In fact, prolonged etching A synergistic effect was observed on construct stiffness (strain times resulted in mass loss of PCL, and hence decreased the rate 25% min ) when GelMA and alginate hydrogels were construct stiffness. Therefore, we concluded that the strength reinforced with the microfibre PCL scaffolds (Fig. 2). The stiffness increase seen in the composites was not due to hydrogel/fibre of PCL scaffolds alone increased from 1.1 0.3 kPa for 98% bonding. porosity to 15.2 2.2 kPa for the 93% porosity scaffold (Fig. 2a). The stiffness of the scaffolds was comparable to that of hydrogels: Comparison of reinforced gels with articular cartilage. The ± ± 7.1 0.5 and 7.5 1.0 kPa for GelMA crosslinked with either average stress–strain curve of native cartilage samples (strain rate 12.5 or 25 mM ammonium persulfate/tetramethylethylenediamide 25% min ; n ¼ 8) that were harvested from the patellofemoral * * 10 10 93% 94% 96% 97% 98% Porosity of PCL scaffold * 500 * 500 * 500 * * * * * 400 400 300 300 300 200 200 200 100 100 100 0 0 0 Porosity of reinforcing scaffold Porosity of reinforcing scaffold Porosity of reinforcing scaffold GelMA Native cartilage reinforced 84% GelMA reinforced 93% GelMA reinforced 98% GelMA PCL scaffold 0% 10% 20% 30% 40% 50% Strain Figure 2 | Reinforcing effect of high-porosity 3D-printed PCL scaffolds incorporated into hydrogels. (a) Compressive moduli of PCL scaffolds were in the same range as those of (b) hydrogels alone. GelMA reinforced with PCL scaffolds and crosslinked with either (c)25mM or (d) 12.5 mM APS/TEMED were one order of magnitude stiffer than the scaffolds or gel alone; (e) a comparable degree of reinforcement for reinforced alginate gels (*Po0.05, one-way ANOVA with Bonferroni correction). (f) Stress–strain curves of GelMA, the PCL scaffold and reinforced GelMA, approaching the curve of native cartilage (yellow). Scaffolds fabricated from thick fibres (that is, 84%) were stiffer than native cartilage, but often disintegrated at a strain of B10% (all groups n ¼ 5, mean s.d.; cartilage n ¼ 8). Here GelMA was crosslinked with 12.5 mM APS/TEMED and the strain rate applied was 25% min . NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 3 & 2015 Macmillan Publishers Limited. All rights reserved. GelMA (25 mM) GelMA (12.5 mM) Alginate Stiffness (kPa) Stress (kPa) Stiffness (kPa) 93% 94% 96% 97% 98% gel Stiffness (kPa) 93% 94% 96% 97% 98% gel Stiffness (kPa) Stiffness (kPa) 93% 94% 96% 97% 98% gel ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 groove of an equine knee joint was plotted in Fig. 2f together with horizontally and places the surrounding fibres under tension. the average (n ¼ 5) stress–strain curves for GelMA, high-porosity We hypothesize that the fibres’ resistance to elongation under PCL scaffolds and GelMA reinforced with 98%, 93% and 84% compression loading results in a pre-stress of the scaffold fibres. porosity PCL scaffolds, respectively. Clearly, GelMA gels were This resistance leads to the observed increase in stiffness, as the much softer than the cartilage. However, the stiffness and stiffness of bulk PCL is approximately four orders of magnitude deformation profile (shape of the stress–strain curve) of GelMA higher than that of the gels. In order to confirm our hypothesis, could be tailored to that of native cartilage by reinforcement with we developed a mathematical model that enables us to make the high-porosity PCL scaffolds. While thick-fibre scaffolds (porosity following prediction for the construct stiffness, C: 72–88%), as fabricated with traditional 3D-printing methods, r EN were stiffer than native cartilage, they broke atB10% strain when C ¼ ð2Þ ðÞ 3=2 large fibre spacings were applied (41.25 mm). This sudden loss 2RðÞ 1  l of integrity was not observed for native cartilage or gels reinforced with high-porosity scaffolds (Fig. 2f). where N denotes the number of fibres in the construct, E the Young’s modulus of the reinforcing polymer, r the fibre radius, R the construct radius and l the axial strain (expressed as a Time-dependent stress response of reinforced gels and cartilage. fraction of the initial height). A derivation and more detailed The stress response to isostrain showed that GelMA reinforced explanation of the model are provided in the Supplementary with a 93% porosity scaffold has a higher modulus on initial Table 1 and Supplementary Note 2. loading than cartilage (Fig. 3). However, relaxation is faster, and to a lower equilibrium modulus, than cartilage. The modulus of GelMA was much lower during the isostrain at any time. Analysis Video analysis of the compression cycle of reinforced gels. The of the stress curves at isostrain indicated that the reinforced GelMA model was further validated by lateral video imaging of the constructs could not be adequately described within an isotropic compression cycle of the GelMA/PCL composites, which revealed poroelastic framework, but that a bi-exponential function: elongation (stretching) of PCL microfibres and a concave profile at the sides of the construct because of lateral displacement K t  K t 1 2 S ¼ S þ S e S e ð1Þ 0 1 2 of water from GelMA (Supplementary Movie 2, 97% porosity provided an accurate, phenomenological characterization of the scaffold). The microfibres stretched 9 1% on 30% axial strain. In time-dependent response of the construct, where S is the stress, t is addition, from a top view it was observed that the PCL scaffold time and the constants S and K have units of stress and time, area expanded 17% and the aqueous component expanded 23% i i respectively. A nonlinear least squares fitting algorithm was used to (including exudation of water; Fig. 4a). obtain the best fitting parameters for S and K from the 9% iso- i i strain curves as detailed in Supplementary Fig. 2 and Microfibre compared with thick-fibre composites. The Supplementary Note 1. We report E ¼ S /strain as elastic moduli, i i reinforcing effect was only observed in high-porosity scaffolds so that E is the equilibrium modulus and E and E are ‘transient’ 0 1 2 fabricated from thin fibres (diameter o48.2 mm), as indicated in moduli that characterize the time-dependent response of the Fig. 4b. The 3D-printing of thicker fibres (diameter 488.5 mm) material so that the ‘peak’ equilibrium modulus is obtained by resulted in a significantly lower scaffold porosity, ranging from 88 summing the E (Table 1). to 72% (PCL fraction of 12% to 28%, respectively), depending on the fibre thickness and spacing between the fibres. The stiffness of Modelling the fibre reinforcement of hydrogels. The mechan- the thick-fibre scaffolds ranged from 1.8 0.2 MPa for 88% ism we propose to explain the synergistic reinforcement is porosity to 16.1 1.7 MPa for 72% porosity. These scaffolds had illustrated in Fig. 4a. Hydrogels can be reasonably described as stiffnesses similar to those of their composite counterparts with incompressible ; therefore, the volume of the hydrogel is crosslinked GelMA. The compression cycle of GelMA reinforced conserved and vertical compression must be accompanied by with an 82% porosity scaffold is shown in Supplementary Movie horizontal expansion. As a composite construct is loaded, each 3. Exudation of fluid was observed; while the relatively thick ‘semi-confined cell’ of hydrogel (right column in Fig. 4a) expands scaffold fibres were compressed, they did not elongate. 3,000 Recovery of GelMA constructs after compression. The recovery GeIMA Fast loading phase of (reinforced) GelMA constructs after repetitive axial compres- Reinforced GeIMA sion was measured. Figure 5a–c shows that GelMA, as well as Cartilage reinforced GelMA constructs, are fully elastic after 20 cycles of 2,000 20% strain. The stress–strain curves of the initial and final loading cycle are shown. GelMA also fully recovered after repetitive 50% axial strain (Fig. 5d). However, GelMA reinforced with 93% Equilibrium phase porosity PCL microfibre scaffolds showed significantly decreased 1,000 resistance when axially deformed by over 35% (Fig. 5e). Chondrogenic differentiation of embedded human chondrocytes. Haematoxylin and eosin-stained sections revealed that chon- 01 2 34 5 drocytes retained their spherical morphology within the fibre- Time (min) reinforced hydrogels after 7 days (Fig. 6c). Chondrocytes were Figure 3 | The time-dependent stress response of (reinforced) GelMA also homogenously distributed throughout the construct as gel and articular cartilage at 9% isostrain. Reinforced GelMA displays a shown by the 4 ,6-diamidino-2-phenylindole (DAPI)-stained cell high modulus compared with cartilage on direct loading. The modulus of nuclei (Fig. 6d). Chondrocytes maintained high cell viability on reinforced gels in the equilibrium phase, however, is closer to gel than to days 1 and 7 when crosslinked with 12.5 mM APS/TEMED cartilage. These plots are representative for the series of isostrains that (Fig. 6e). Following 14 days of culture, hydrogels (GelMA with were consecutively performed for 15 min. 0.5% hyaluronic acid (Lifecore, USA; GelMA-HA) for improved 4 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. Modulus (kPa) NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE Table 1 | Fit of a bi-exponential function to the time-dependent stress response. Equilibrium modulus Fast decay Slow decay 3  1 3  1 E (kPa) E (kPa) k (10 s ) E (kPa) k (10 s ) 0 1 1 2 2 ± ± ± ± ± GelMA 12 1.7 0.82 0.11 610 310 7.8 0.43 1.5 0.76 ± ± ± ± ± Reinforced GelMA 130 7.8 2700 160 180 9.8 430 49 28 1.1 ± ± ± ± ± Cartilage 440 140 300 160 110 45 430 200 4.9 2 GelMA, gelatin methacrylamide Estimates of the moduli E and the fast and slow decay rates k obtained by fitting the stress response at 9% isostrain of GelMA, reinforced GelMA and cartilage to a bi-exponential function (n ¼ 3, i i mean s.d). the quantitative reverse transcriptase–PCR (qRT–PCR) data indicated that for chondrocytes within the fibre-reinforced hydrogels, ACAN (gene expression for glycosaminoglycan (GAG) synthesis) and COL1A1 (gene expression for collagen type I) mRNA levels were significantly upregulated in the compressed (‘C’) groups compared with the control, non- compressed groups (‘NC’; Fig. 6f–i; Po0.05). Further, chondrocytes in the reinforced gels were more responsive to loading regime than cells in GelMA alone. Stiffness values were B400 kPa in cell-laden fibre-reinforced hydrogels (Fig. 6j). There was no significant effect of reinforcement or compression on construct stiffness and glycosaminoglycan (GAG) content (Fig. 6k). The chondrocytes within the gels exhibited pericellular Axial, unconfined compression collagen type I and II deposition in all groups, with no discernible differences between the groups (Supplementary Fig. 4). Discussion In this study, high-strength composite constructs were fabricated by combining 3D-printed high-porosity scaffolds with a hydrogel. These composite hydrogels can be customized to yield a wide 100,000 White = scaffold range of mechanical properties and, from a biological point of Grey = scaffold + gel view, have the capacity to support cell proliferation and 10,000 Fibre diameter of extracellular matrix production. The PCL scaffolds were built scaffold (µm) by micrometre-scale fibres that were organized and intercon- 1,000 Native articular cartilage 19.4 ± 1.7 nected through the melt-electrospinning writing process. With 48.2 this unique and new 3D-printing technique, pre-set network 88.5 architectures can be realized in a direct writing mode . This 219.7 ± 14.2 technique allows fibres to be printed well below the limits of classical melt-extrusion-based 3D-printing technologies such as fused deposition modelling , with filament diameters as small as 1 35 5 mm, instead of 100 mm or larger . 0% 10% 20% 30% Hydrogels have previously been reinforced with solution- PCL fraction (1-porosity) 9,24,25,27 electrospun nonwovens . However, most traditional Figure 4 | The mechanism of hydrogel reinforcement with organized solution-electrospun meshes have disadvantages from both a high-porosity scaffolds. (a) PCL microfibre scaffolds (blue in schematic) mechanical and biological point of view, for example, fibres are serve as a reinforcing component to GelMA hydrogel (yellow in schematic). not fused and hence slide under loading meshes are usually not When axial compression is applied to the reinforced hydrogels, the stiff thin thicker than 100 m and have a pore size of less than 5 mm and are, 25,27,40 scaffold fibres stretch on lateral displacement of the hydrogel. This therefore, too dense for cell infiltration . To overcome these mechanism provides the composites with a high stiffness and elasticity issues related to traditional solution electrospinning, fibres have (scale bars, 1 mm). (b) Moduli of scaffolds and scaffold/gel composites as a been collected in an earthed ethanol bath, yielding nonwovens 9,24 function of porosity, showing the synergistic increase in stiffness was only with high porosities . Hydrogel stiffness was increased fourfold observed for thin-fibre scaffolds with a high porosity (polymer fraction when reinforced with these randomly organized nonwovens, 24,25 20 2–7% ¼ porosity 98–93% (highlighted in red)), fabricated with melt- irrespective of the porosity . Silk microfibres or carbon electrospinning writing (MEW). Fused deposition modelling (FDM) nanotubes also formed porous reinforcing structures for scaffolds were fabricated from 10-fold thicker fibres, resulting in a higher hydrogels, permitting cellular infiltration and differentiation. stiffness; however, no synergistic reinforcement was observed (mean of However, the stiffness of the gels increased only one- to threefold, n ¼ 5). as the fibres were not fused. In contrast, in our current composite model consisting of a relatively soft hydrogel (7.1–15.8 kPa) and a highly porous and soft PCL scaffold (1.1–15.2 kPa), the maximum chondrogenic differentiation and GelMA-HA with 93% porous stiffness obtained was a factor of 50 larger (405 kPa) than that of PCL mesh) were subjected to physiological loading cycles for the hydrogels alone, on a strain rate of 25% min . another 14 days. The stress response to the 1 Hz dynamic loading The stiffness of cartilage and (reinforced) hydrogels is strongly 41,42 regime is presented in Supplementary Fig. 3. Analysis of strain-rate-dependent . We found that the reinforced NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 5 & 2015 Macmillan Publishers Limited. All rights reserved. Stiffness (kPa) 30% Strain Uncompressed ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 GelMA GelMA + 97% porosity mesh 20% Strain 15 20% Strain Cycle 1 Cycle 1 Cycle 20 Cycle 20 5% 10% 15% 20% 5% 10% 15% 20% Strain Strain GelMA + 93% porosity mesh GelMA 20% Strain 25 50% Strain Cycle 1 Cycle 1 Cycle 20 Cycle 20 0 0 5% 10% 15% 20% 10% 20% 30% 40% 50% Strain Strain GelMA + 93% porosity mesh 20–40% Strain 20% 25% 30% 20 35% 40% 5% 10% 15% 20% Strain Figure 5 | Cyclic compression testing of GelMA and reinforced GelMA constructs. The stress–strain curves of the first and last of 20 cycles of 20% compression are shown for (a) GelMA alone, (b) GelMA reinforced with a 97% porosity PCL scaffold and (c) GelMA reinforced with a 93% porosity PCL scaffold. For all groups no decay in stress was observed. (d) GelMA constructs also recovered after 20 cycles of 50% axial compression; however, (e) GelMA reinforced with a 93% porous PCL scaffold required substantially less force to be compressed to 20% strain, after being compressed to over 30% strain (arrow; mean of n ¼ 3). hydrogels were one order of magnitude stiffer than cartilage on be a combined effect of the osmotic pressure of fixed charges on initial fast loading at 9% strain. However, the modulus of proteoglycans and an organized extracellular matrix. reinforced hydrogels displayed a steep decrease to an equilibrium Etching the high-porosity scaffolds did not result in a further value roughly one-third the equilibrium modulus of cartilage, yet increase in stiffness of the composites. This approach was on the still one order of magnitude larger than the equilibrium modulus basis of our previous finding that covalent attachment of the of hydrogel alone. hydrogel component to a scaffold that was fabricated from The effect of reinforcement can be explained by the highly modified, methacrylated PCL, resulted in increased construct organized fibre architecture of the scaffolds. The mathematical strength . It should be noted that etching results in an increase model we developed demonstrated that the hydrogel places the in van de Waals forces but does not establish covalent bonds PCL fibres under tension on axial compression, and predicts a between the hydrogel and the reinforcing scaffold. theoretical upper bound on the attainable stiffness. However, the Infiltrating hydrogel into scaffolds that had been fabricated theoretical stiffness of the reinforced hydrogels is one order of with traditional 3D melt-printing techniques with thick fibres magnitude larger than that observed experimentally. This is (Z88 mm) did not show a significant mechanical effect. Axial reasonable since an idealized construct was considered, in which loading of these constructs requires compression of the PCL the fibres are initially taut at zero strain, the hydrogel is purely through a strong vertical column of fibre crossings, which is not elastic and there is no slip between the fibres and the hydrogel. In easily deformed. Video imaging of the compression cycle fact, in our experimental set-up, we showed that very little water, confirmed that thick fibres were not stretched and water flowed or hydrogel, was compressed out of the constructs; therefore, the out of these scaffolds without providing a synergistic effect. In fibres did not completely lock the hydrogel. The stress relaxation contrast, the columns of fine fibre crossings are easily deformed on isostrain (Fig. 3) reflects the exudation of the aqueous under axial loading; however, they are supported by the hydrogel component (water or gel) from the reinforcing scaffold. For component in the composites. Supplementary Movie 2 shows cartilage we observed a higher equilibrium modulus, which may some local distortion of the columns, but no large-scale buckling 6 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. Stress (kPa) Stress (kPa) Stress (kPa) Stress (kPa) Stress (kPa) NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE 0.06 0.0012 100% NC NC NC 0.05 C 0.0010 C 25 C 80% 0.04 0.0008 20 60% 0.03 0.0006 15 40% 0.02 0.0004 10 20% 0.01 0.0002 5 0% 0 0 0 Day1 Day7 GeIMA 93% GeIMA 93% GeIMA 93% 0.030 0.05 1,000 NC NC NC C C 0.025 800 0.04 0.020 600 0.03 0.015 400 0.02 0.010 200 0.01 0.005 0 0 GeIMA 93% GeIMA 93% GeIMA 93% Figure 6 | Differentiation of chondrocytes embedded in reinforced GelMA-HA hydrogels. (a) Stereomicroscopy image of a GelMA-HA gel and (b)ofa GelMA-HA gel reinforced with a PCL scaffold (93% porosity) on day 1 (scale bars, 2 mm). (c) Haematoxylin/eosin staining after 7 days of culture shows that chondrocytes remained within the GelMA component and retained a round morphology (scale bar, 500 mm, inlay, 200 mm). (d) DAPI staining confirms a homogenous distribution of the cells throughout the construct (scale bar, 200 mm). (e) Chondrocytes remained viable over 7 days of culture in the gels. Gene expression analysis for (f) ACAN, (g) COL2A1, (h) COL1A1 and (i) COL10A1 on day 14 for compressed (C) and non-compressed (NC) groups (n ¼ 4, mean s.d.; *Po0.05, two-way ANOVA). All gene expression levels were normalized to the housekeeping gene b2 microglobulin (B2M). (j) Stiffness of the constructs following long-term culture (n ¼ 4). (k) GAG/ww values on day 28 (n ¼ 5). of the columns; therefore, the microfibres are put under tension. to native cartilage. The porosity of these woven scaffolds was We further note that this axial loading reduces the horizontal 70–74%, compared with 93–98% in the present study. The pore size, improving gel confinement. To our knowledge, we are stiffness of the woven composites was reported up to 0.2 Mpa (at the first to show the synergistic effect of a well-designed scaffold/ equilibrium stress), which was twice as stiff as the scaffold hydrogel system; the work shown in this paper could therefore without the gel . Both the woven and the melt-electrospun lead to a paradigm shift in the field. composites showed axial recovery after compression for 10% and Reinforced GelMA hydrogels possess stress–strain curves that 20%, respectively. The water that was compressed out of the closely resemble those of healthy articular cartilage. In addition, scaffolds was likely reabsorbed during the relaxation phase, in absolute terms, the stiffness of the biodegradable composites comparable to fluid dynamics in the articular cartilage . Our was comparable to the stiffness of articular cartilage, which has reinforced hydrogels showed decreased resistance if compressed been reported to range from 400 to 800 kPa (refs 43–45). On the by 35% or more, which may be because of fractures or other hand, the stiffness of scaffolds fabricated with traditional delamination in the PCL scaffold junctions. For translational 3D-printing was one order of magnitude larger than that of the purposes, it is important to realize that native cartilage is 30,33 articular cartilage, which is consistent with previous reports . exposed to 15–45% axial deformation under long-term static The porosity of these scaffolds ranged between 72 and 89%, compression . whereas the porosities reported in the literature range from 28.9 Incorporation of organized fibrous PCL scaffolds within a well- and 91.2% (refs 39,46,47). Nevertheless, when aiming to fabricate characterized hydrogel system makes it possible to culture cells in high-porosity scaffolds (480%) from thick fibres, large fibre a customizable, mechanically diversified environment. The spacing (41.5 mm) is required, which causes scaffold hydrogel component of the composite constructs will degrade 14,29 disintegration under strains of B10%. In addition, the thicker within months , allowing the regeneration of new tissue, while fibres occupy a relatively high volume that is inaccessible for the PCL component degrades within years forming a temporary tissue formation until the scaffold has degraded. reinforcing network to the new tissue. Under physiological Hydrogels reinforced with woven scaffolds, composed of either compressive loading of 20% strain and 1 Hz, our gene expression 8 21,22 polyglycolic acid or PCL , have previously been reported to data suggest that chondrocyte expression of matrix mRNAs is possess tensile, compressive and shear properties comparable significantly upregulated in composite hydrogels compared with NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications 7 & 2015 Macmillan Publishers Limited. All rights reserved. Cell viability COL10A1 mRNA (normalized to B2M) Stiffness (kPa) ACAN mRNA (day 28) (normalized to B2M) COL2A1 mRNA –1 GAG/ww (µg µg ) (normalized to B2M) (day 28) COL1A1 mRNA (normalized to B2M) ARTICLE NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 non-reinforced, weak hydrogels. These results highlight the Directly after mixing in the TEMED, the GelMA solution as described above was injected in the mold untill all PCL scaffolds were fully infused. Crosslinking of the importance of developing cell-culture platforms and tissue- GelMA was allowed in this mold at 37 C for 1 h. Cell-free samples were then engineered constructs that better mimic the in vivo stored in PBS at 37 C. Chondrocyte-laden samples were kept in a chondrocyte mechanoenvironment of natural articular cartilage. Still, the expansion medium. In order to evaluate the effect of etching time of the scaffolds exact mechanisms of stress transfer to the cells in fibre-reinforced on construct stiffness, GelMA was reinforced with scaffolds etched for 0–16 h (all n ¼ 5) and crosslinked with 25 mM APS/TEMED. In addition, 93–98% porosity hydrogels remain to be investigated. PCL scaffolds (all n ¼ 5) were infused with alginate in the mold. The alginate In conclusion, the current work represents a significant step composite constructs were then crosslinked by immersion in a 102-mmol towards developing biomechanically functional tissue constructs. calciumchloride solution for 30 min. The stiffness of the constructs was significantly enhanced, achieving values similar to those of native articular cartilage, Harvest of equine cartilage. Full-thickness cartilage was harvested from the knee by combining hydrogels with 3D-printed, high-porosity melt- joint from one equine donor (10 years old) with macroscopically healthy cartilage. electrospun PCL scaffolds. This synergistic effect could be This was performed after consent of the horse owner and according to the ethical guidelines of the Utrecht University. Equine cartilage was used because of the modulated by altering the porosity of the reinforcing scaffolds. availability and its similarities to human cartilage . Cylindrical samples with a The composite constructs have a strong elastic component diameter of 5 mm and a height of 2 mm (n ¼ 8) were taken with a biopsy punch and recover after physiological axial strains. Finally, human from the medial patellofemoral groove and stored in PBS for up to 3 h. chondrocytes encapsulated in the GelMA/PCL composites were found to be more responsive to mechanical loading, which led to Mechanical analyses. The stiffness (or compressive modulus) of GelMA and significant changes in gene expression in vitro. alginate hydrogels, PCL scaffolds, the composite constructs and articular cartilage (all n ¼ 5) was measured by single uniaxial unconfined compression in air at room temperature, after 1–3 days of submersion in PBS from their preparation. Methods We confirmed that for single-cycle compression testing, the absence of a PBS Fabrication of PCL microfibre scaffolds. Several scaffolds were 3D-printed from immersion bath does not influence the stress responses of hydrogel samples. The medical-grade PCL (PURAC, Gorinchem, the Netherlands) with a custom-made effect of crosslinking with 12.5 or 25 mM APS/TEMED was tested. The stress–strain melt-electrospinning device . PCL was heated to 103 C in a 2 ml polypropylene curves of (reinforced) GelMA that was crosslinked with 12.5 mM APS/TEMED syringe and extruded at a rate of 18 mlh using a syringe pump. In combination were compared with cartilage samples. Over a period of 2 min, a force ramp (axial with an electrostatic field of 8–10 kV between the syringe needle tip (23G) and the  1 strain rate ca. 25% min ) was applied to the samples employing a microtester collector plate (3 cm distance), a stable PCL jet was obtained. Defined scaffold (Instron, Melbourne, Australia) or a Dynamic Mechanical Analyser (DMA 2980, architectures, with dimensions up to 120  120  1 mm, were realized through TA Instruments, New Castle, DE, USA). The stiffness was calculated from the computer-aided movement of the aluminium collector plate at a speed of linear derivative of the stress–strain curve at 12–15% strain. In the low-porosity 1,400 mm min , using the Mach3 motion control software (Artsoft, USA). scaffolds that were fabricated with traditional 3D-printing, the 6–9% strain region A predefined 0–90 architecture was imposed, with a fibre spacing of 0.2, 0.4, 0.6, was taken, as these constructs often disintegrated when strained beyond 10%. The 0.8 or 1.0 mm. The melt-electrospinning writing was terminated when scaffolds compression cycle of GelMA reinforced with 97 and 82% porosity scaffolds was reached a height of 1.0 mm, to ensure maximum quality of the architecture. To captured from the side (hand-held digital microscope 1.3 MP, Meade instruments, investigate the effect of the PCL flow rate on fibre diameter and scaffold porosity, Europe GmbH & Co, Rhede, Germany) in order to analyse the lateral expansion of we printed constructs at 4  and 10  flow rate, with a fibre spacing of 0.4 and both components of the composite construct. In addition, images from the top of 1.0 mm, respectively. Cylindrical samples were extracted from the scaffolds with a these constructs were taken with a stereomicroscope, when uncompressed and at 5-mm diameter biopsy punch. Two scaffolds were stacked in order to achieve a 30% strain between two glass slides. height of 2 mm, which is comparable to the thickness of human cartilage in the In order to test the time-dependent stress response, GelMA, GelMA reinforced knee joint . The porosity and mechanical properties of the melt-electrospun with 93% porosity scaffolds and articular cartilage samples were subjected to a scaffolds were compared with scaffolds fabricated with traditional 3D-printing series of isostrain steps . Samples were 2% pre-strained for 5 min, followed by technologies. To this end, scaffolds were 3D-printed from PCL with a BioScaffolder strains of 6, 9 and 13% that were consecutively applied for 15 min each. The system (SYS þ ENG, Salzgitter-Bad, Germany) as described previously . Briefly, experiments were performed in PBS and the stress response of all samples was PCL was heated till 70 C and extruded with an Auger screw through a 27-gauge recorded. The modulus on fast initial loading and the equilibrium modulus were needle. Scaffolds measuring 40  40  2 mm were fabricated with a fibre spacing of extracted, and the stress decay rate was estimated by fitting to a bi-exponential 0.75, 1.0, 1.25, 1.5 and 1.75 mm. The quality of the scaffolds was imaged both with function. stereomicroscopy and Scanning Electron Microscopy (SEM, Hitachi TM3000, The resistance to axial deformation of GelMA and GelMA reinforced with PCL Japan and Quanta 200, FEI, Milton, Australia). The fibre diameter was measured constructs (porosity 93 and 97%) was measured after a cyclic (20  ) axial strain of with the ImageJ software (National Institutes of Health, USA). The porosity of the 20% (Allround-line Z020, Zwick Roell, Germany). In addition, the resistance after PCL scaffolds was determined gravimetrically. 20 cycles of axial deformation of 50% was measured for GelMA constructs. The recovery measurements were performed in PBS and constructs were allowed to recover for 1 min after every cycle. GelMA constructs reinforced with a 93% porous Etching of PCL microfibre scaffolds. PCL scaffolds were etched in order to scaffold underwent compression with incremental maximum strains of 20–40% increase its hydrophilicity and surface area, which could potentially contribute to (with 5% increments) in order to analyse the maximal strain that could be exerted the stiffness of reinforced GelMA constructs. Cylindrical PCL scaffolds with a before irreversible damage would occur. porosity of 94% were treated with 70% ethanol and subsequently etched with 5 M sodium hydroxide for 0, 1, 2, 4, 8 and 16 h. After etching, the scaffolds were washed in deionized H O until the pH reached 7.4, and then air-dried. The effect of etching Modelling the fibre reinforcement of hydrogels. A mathematical model was was evaluated by measuring the fibre diameter (ImageJ) from SEM images, and by constructed to investigate further the mechanisms by which the 3D-printed scaf- assessing the weight from 2  5 mm diameter scaffolds (expressed in relative weight folds reinforce the hydrogels. There is an extensive literature on the modelling of 41,42 loss). Since mild etching will increase hydrophilicity and may facilitate perfusion of fibre-reinforced biological materials, with recent cartilage-focussed examples . GelMA through the PCL scaffold, 2 h of etching was performed for all other Fibre-reinforced materials are often modelled by assuming that it is reasonable to experiments presented. define a continuously varying fibre density and orientation and then making the material properties a function of these. In this instance, given that we know the arrangement and approximate number of fibres in each plane of the material, we Preparation of GelMA and alginate gels. GelMA was synthesized by reaction of take the more direct approach of considering how each fibre stretches as the type A gelatin (Sigma-Aldrich, St Louis, MO, USA) with methacrylic anhydride at scaffold deforms. Our model takes into account the number of fibres in the scaffold 50 C for 1 h as previously described . GelMA was dissolved in PBS at 60 Cata and the fibre diameter, the Young’s modulus of PCL and the construct dimensions. concentration of 10% (w/v) containing 12.5 or 25 mM APS. TEMED (12.5 or In the model, the composite construct was viewed as an elastic solid, in which the 25 mM) was added followed by 5 s vortexing in order to initiate crosslinking of the PCL fibres are placed under tension by the hydrogel on axial compression. GelMA. Sodium alginate (IMCD, Amersfoort, the Netherlands) was dissolved in PBS at 3% (w/v), and used as a control gel to GelMA for the mechanical analyses. Harvest of human chondrocytes. Macroscopically healthy cartilage was harvested either from a discarded talus bone that was resected from a 7-year-old patient Preparation of reinforced hydrogel constructs. The scaffolds were placed in an undergoing an orthopaedic intervention, or from the femoral condyles of knee injection mold that was custom-made from polymethylmethacrylate in order to fit replacement surgery patients (age: 71.0 4.1; n ¼ 6) with consent. This was in 10 cylinders with a diameter of 5 mm and a height of 2 mm. All cylindrical voids in concordance with the institutional code of conduct regarding the use of discarded the mold were interconnected so that gel could be serially perfused. The mold was tissues in the University Medical Center Utrecht, and ethics approval was also sealed with a sheet of polyethylene that was fixed between two glass histology slides. obtained from Queensland University of Technology and Prince Charles hospital 8 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved. NATURE COMMUNICATIONS | DOI: 10.1038/ncomms7933 ARTICLE before sample collection. Cartilage was cut into small slices and washed with PBS Immunohistochemistry. After 28 days, the reinforced GelMA-HA gels were supplemented with penicillin and streptomycin. The cartilage was digested over- paraffin-embedded and sectioned at 5 mm. For antigen retrieval, ready-to-use night in 0.15% type II collagenase at 37 C. The resulting cell suspension was proteinase K solution (Dako) was used for 10 min at 37 C. Sections were filtered (100 mm cell strainer) and washed three times with PBS. Then, cells were blocked with 2% FBS solution before exposure to primary antibodies: I-8H5 resuspended in chondrocyte expansion medium (DMEM supplemented with 10% (MP Biomedicals), dilution 1:300 for collagen type I; II-II6B3 (DSHB), dilution 1  1 fetal bovine serum (FBS), 100 units ml penicillin and 100 mgml streptomycin, 1:200 for collagen type II. Following incubation in fluorescence-labelled goat and 10 ng ml FGF-2) and expanded for 10 days in monolayer cultures (seeding anti-mouse secondary antibody (Alexa Fluor 488, Invitrogen), sections were density 5,000 cells cm ). For the short-term study (7 day), chondrocytes from the mounted with Prolong Gold (Invitrogen) and visualized using a confocal talus bone were used at passage 2, and for the long-term study (28 day) chon- fluorescence microscope (A1R Confocal, Nikon). drocytes obtained from the femoral condyles of the different donors were pooled and used at passage 1. Statistics. All data were expressed as mean s.d. of five independent samples, unless stated otherwise. One-way analysis of variance (ANOVA) with Bonferroni Cell viability. Chondrocytes were mixed in the GelMA/APS solution at a correction was performed to compare the stiffness of the reinforced hydrogels; the concentration of 10 million ml . Crosslinking was performed with different effect of scaffold porosity was compared for every hydrogel and the effect of the proportions of APS/TEMED (12.5/12.5 and 25/25 mM) in order to evaluate hydrogel was compared for every scaffold porosity. Two-way ANOVA was used to cytotoxicity of the redox crosslinking process. The viability of the chondrocytes was evaluate the effect of compression and hydrogel types on gene expression, GAG/ measured for 4 h (day 1) and 7 days after embedding and crosslinking of the ww and stiffness. An independent samples t-test, not assuming equal variances, was GelMA constructs. To visualize cell viability, a LIVE/DEAD Viability Assay performed to evaluate cell viability (SPSS, IBM software, USA). Differences were (Molecular Probes MP03224, Eugene, USA) was performed according to the considered significant when Po0.05. manufacturer’s instructions. 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University of Twente for facilitating mechanical testing on their Zwick universal 34. Brown, T. D., Dalton, P. D. & Hutmacher, D. W. Direct writing by way of melt mechanical tester. We thank professor Harrie Weinans for fruitful discussions on car- electrospinning. Adv. Mater. 23, 5651–5657 (2011). tilage biomechanics. The II-II6B3 monoclonal antibody developed by T.F. Linsenmayer 35. Farrugia, B. L. et al. Dermal fibroblast infiltration of poly(epsilon-caprolactone) was obtained from the Developmental Studies Hybridoma Bank developed under the scaffolds fabricated by melt electrospinning in a direct writing mode. auspices of the NICHD and maintained by The University of Iowa, Department of Biofabrication 5, 025001 (2013). Biology, Iowa City, IA 52242. J.V. was supported by a grant from the Dutch government 36. Schuurman, W. et al. Gelatin-methacrylamide hydrogels as potential to the Netherlands Institute for Regenerative Medicine (NIRM, grant nFES0908); biomaterials for fabrication of tissue-engineered cartilage constructs. Macromol. F.P.W.M. was supported by a Marie Curie grant from the European Commission (PIOF- Biosci. 13, 551–561 (2013). GA-272286) and J.M. was supported by the Dutch Arthritis Foundation. This work was 37. Chen, Y. C. et al. Functional human vascular network generated in also supported by the Australia National Health and Medical Research Council and the photocrosslinkable gelatin methacrylate hydrogels. Adv. Funct. Mater. 22, European Community’s Seventh Framework Programme (FP7/2007-2013) under grant 2027–2039 (2012). agreement n309962 (HydroZONES). 38. Hong, W., Zhao, X. H., Zhou, J. X. & Suo, Z. G. A theory of coupled diffusion and large deformation in polymeric gels. J. Mech. Phys. Solids 56, 1779–1793 (2008). Author contributions 39. Hutmacher, D. W. Scaffolds in tissue engineering bone and cartilage. J.V., F.P.W.M., D.W.H. and J.M. designed the study. F.P.W.M., P.D.D. and D.W.H. Biomaterials 21, 2529–2543 (2000). developed the melt electrospinning set-up. J.V., F.P.W.M. and E.M.v.B. fabricated the 40. Pham, Q. P., Sharma, U. & Mikos, A. G. Electrospun poly(epsilon- scaffolds and performed the mechanical experiments. J.E.J. and E.M.v.B. conducted the caprolactone) microfiber and multilayer nanofiber/microfiber scaffolds: dynamic loading experiment. L.S.K. and H.M.B. did the mathematical modelling. characterization of scaffolds and measurement of cellular infiltration. W.J.A.D., P.D.D., D.W.H. and J.M. directed the study. J.V. prepared the manuscript Biomacromolecules 7, 2796–2805 (2006). together with all co-authors. 41. Pierce, D. M., Ricken, T. & Holzapfel, G. A. A hyperelastic biphasic fibre- reinforced model of articular cartilage considering distributed collagen fibre orientations: continuum basis, computational aspects and applications. Comput. Methods Biomech. Biomed. Eng. 16, 1344–1361 (2013). Additional information Supplementary Information accompanies this paper at http://www.nature.com/ 42. Hosseini, S. M., Wilson, W., Ito, K. & van Donkelaar, C. C. How naturecommunications preconditioning affects the measurement of poro-viscoelastic mechanical properties in biological tissues. Biomech. Model Mechanobiol. 13, 503–513 Competing financial interests: The authors declare no competing financial interests. (2014). 43. Athanasiou, K. A., Agarwal, A. & Dzida, F. J. Comparative study of the intrinsic Reprints and permission information is available online at http://npg.nature.com/ mechanical properties of the human acetabular and femoral head cartilage. reprintsandpermissions/ J. Orthopaed. Res. 12, 340–349 (1994). 44. Jurvelin, J. S., Buschmann, M. D. & Hunziker, E. B. Optical and mechanical How to cite this article: Visser, J. et al. Reinforcement of hydrogels using determination of Poisson’s ratio of adult bovine humeral articular cartilage. three-dimensionally printed microfibres. Nat. Commun. 6:6933 J. Biomech. 30, 235–241 (1997). doi: 10.1038/ncomms7933 (2015). 10 NATURE COMMUNICATIONS | 6:6933 | DOI: 10.1038/ncomms7933 | www.nature.com/naturecommunications & 2015 Macmillan Publishers Limited. All rights reserved.

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