Acta Biomaterialia 12 (2015) 30–41

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Visible-light-initiated hydrogels preserving cartilage extracellular signaling for inducing chondrogenesis of mesenchymal stem cells Bogyu Choi a, Soyon Kim b, Brian Lin a, Kevin Li c, Olga Bezouglaia d, Jinku Kim e, Denis Evseenko f, Tara Aghaloo d, Min Lee a,b,⇑ a

Division of Advanced Prosthodontics, University of California, Los Angeles, CA 90095, USA Department of Bioengineering, University of California, Los Angeles, CA 90095, USA Department of Chemical and Biomolecular Engineering, University of California, Los Angeles, CA 90095, USA d Division of Diagnostic and Surgical Sciences, University of California, Los Angeles, CA 90095, USA e Department of Bio and Chemical Engineering, Hongik University, Sejong 339-701, Republic of Korea f Department of Orthopaedic Surgery, University of California, Los Angeles, CA 90095, USA b c

a r t i c l e

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Article history: Received 11 June 2014 Received in revised form 10 October 2014 Accepted 13 October 2014 Available online 20 October 2014 Keywords: Photocrosslinkable chitosan hydrogels Collagen Transforming growth factor Mesenchymal stem cells Chondrogenesis

a b s t r a c t Hydrogels have a unique opportunity to regenerate damaged cartilage tissues by introducing mesenchymal stem cells (MSCs) in a highly swollen environment similar to articular cartilage. During cartilage development, collagen–cell interactions play an important role in mediating early mesenchymal condensation and chondrogenesis with transforming growth factor-b1 (TGF-b1) stimulation. Here, a hydrogel environment that can enhance cell–matrix interactions and chondrogenesis by stabilizing type-II collagen (Col II) and TGF-b1 into photopolymerizable (methacrylated) chitosan (MeGC) with simple entrapment and affinity binding is demonstrated. The MeGC hydrogel was designed to gel upon initiation by exposure to visible blue light in the presence of riboflavin, an aqueous initiator from natural vitamin. The incorporation of Col II into MeGC hydrogels increased cellular condensation and deposition of cartilaginous extracellular matrix by encapsulated chondrocytes. MeGC hydrogels containing Col II supported the release of TGF-b1 in a controlled manner over time in chondrogenic medium and the incorporated TGF-b1 further enhanced chondrogenesis of encapsulated chondrocytes and MSCs, especially synovial MSCs. Subcutaneous implantation of hydrogel cultures showed greatly improved neocartilage formation in constructs loaded with TGF-b1 compared with controls. These findings suggest that cartilage mimetic hydrogels have a high potential for cartilage repair. Ó 2014 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction The regeneration of damaged articular cartilage is clinically challenging because of its poor self-healing capacity [1]. Although autologous chondrocyte implantation (ACI) is an attractive treatment option for articular cartilage lesions [2], it is limited by the need for multiple surgical procedures, cell expansion that often induces rapid dedifferentiation of isolated chondrocytes in monolayer culture [3,4] and hypertrophic differentiation of implanted cells with subsequent ossification in vivo. In addition, variations in the chondrogenic potential of the harvested chondrocytes from patients with osteoarthritis make ACI less than ideal [5]. Alternative cell-based therapies using mesenchymal stem cells (MSCs) ⇑ Corresponding author at: UCLA School of Dentistry, 10833 Le Conte Avenue, CHS 23-088F, Los Angeles, CA 90095-1668, USA. Tel.: +1 310 825 6674; fax: +1 310 825 6345. E-mail address: [email protected] (M. Lee). http://dx.doi.org/10.1016/j.actbio.2014.10.013 1742-7061/Ó 2014 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

are attractive options to improve cartilage repair. MSCs can be obtained from various sources, including bone marrow, adipose tissue, periosteum, muscle and synovium, and have the ability to be expanded to large cell numbers and differentiated into various cell lineages [6–8]. In particular, synovium-derived MSCs (SMSCs) are promising candidates for cartilage tissue engineering due to their high chondrogenic potential and the ease with which they can be expanded in vitro compared to MSCs from other sources [9,10]. Furthermore, SMSCs can be easily isolated from patients by minimally invasive harvesting procedures. Direct application of MSCs alone to the defect site has not yet been proven to be efficient, partly due to leaching of the precursors away from the site and the lack of an appropriate microenvironment for directing chondrogenesis [11,12]. Therefore, there is increasing interest in developing a scaffold system with which to localize MSCs in the defect area, promote and maintain their chondrogenic differentiation, and improve their integration with the

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host tissue. During cartilage development, chondrogenic differentiation is regulated by cell–cell and cell–matrix interactions in the presence of various soluble factors. In particular, collagen (Col)– cell interactions with Transforming growth factor-b (TGF-b) stimulation play an important role in mediating an early mesenchymal condensation process during chondrogenesis [13–17]. These physical and soluble cues can be engineered into biomaterial scaffolds. Type-II collagen (Col II) is a main structural protein in articular cartilage and has been proven to play a crucial role in chondrogenesis. Although Col can form a gel spontaneously, the use of Col alone is limited in tissue engineering applications due to its significant shrinkage and rapid degradation [18,19]. Chemical modification can enhance the physical stability of Col, but this can reduce the bioactivity of natural Col [20,21]. TGF-b1 is known to be important in inducing the chondrogenesis of MSCs and the formation of cartilage tissue. However, the therapeutic efficacy of TGF-b1 is affected by delivery kinetics. Extracellular matrix (ECM) is a natural reservoir for many growth factors and potentiates their bioactivities. Recent studies have implicated the binding of TGF-b1 to various ECM molecules, such as Col and heparin [22–24]. Here, we seek to take advantage of this affinity binding found in natural systems for the controlled delivery of TGF-b1. We have previously developed injectable hydrogels using methacrylated chitosan (MeGC), a derivative of natural chitin, and riboflavin (RF), a visible blue light (VBL) initiator in the vitamin B group, which supported proliferation and deposition of cartilaginous ECM of encapsulated chondrocytes [25]. This gelation system can be rapidly prepared under mild physiological conditions without toxic chemicals. In this study, we report a simple and fast method to stabilize natural cues of Col II binding and TGF-b1 signaling present in the cartilage extracellular environment into MeGC hydrogels without affecting the bioactivities of the signals. Specifically, a hydrogel was formed by mixing solutions of unpolymerized MeGC with Col II and TGF-b1, followed by VBL-crosslinking in the presence of RF (Fig. 1). The feasibility of this hydrogel system to support chondrogenic differentiation was investigated by encapsulating SMSCs and further compared with bone-marrow-derived MSCs (BMSCs) and chondrocytes. We further investigated the ability of this system to promote chondrogenesis in an in vivo subcutaneous model.

2. Materials and methods 2.1. Preparation of photocrosslinkable hydrogels MeGC was prepared as previously described [25,26]. Briefly, glycidyl methacrylate (Sigma–Aldrich, St. Louis, MO) was added to a 2% (w/v) aqueous solution (pH 9.0) of glycol chitosan (GC; Sigma–Aldrich, 500 kDa) with a 1.1 M ratio of glycidyl methacrylate to the primary amine groups in chitosan and allowed to react for 36 h with gentle shaking at room temperature. The reaction mixture was then neutralized, dialyzed against water using a membrane with a cutoff molecular weight of 50 kDa for 15 h and lyophilized. The MeGC solution was sterile-filtered through a 0.45 lm polyethersulfone membrane. The degree of methacrylation to the GC was 26%, as determined by 1H nuclear magnetic resonance imaging. An MeGC (2% w/v) composite solution containing 0.4% w/v of Col II (MeGC/Col II) was prepared by gently mixing stock solutions of MeGC (4% w/v, in phosphate-buffered saline (PBS)) and Col II (from chicken sternal cartilage, Sigma–Aldrich, 1.0% w/v in 0.05% acetic acid), and neutralized with 0.05 N NaOH and 10 PBS on ice. No precipitation was observed after mixing. MeGC solutions (2% w/v) without additives were prepared by diluting the 4% w/v MeGC in PBS. The hydrogel was formed by

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exposing 40 ll of the solution to VBL (400–500 nm, 500– 600 mW cm 2, Bisco Inc., Schaumburg, IL) in the presence of RF (final concentration 6 lM) as a photoinitiator. After irradiation, the MeGC/Col II hydrogel was placed in a 37 °C incubator for 20 min. 2.2. Hydrogel characterization The gelation time of the MeGC and MeGC/Col II composite solutions under VBL irradiation with different concentrations of RF (1.5, 3, 6, 12 and 24 lM) was investigated by a test-tube tilting method as previously described (n = 3) [27]. To observe the distribution of Col II in hydrogels, disk-shaped hydrogels (diameter = 5 mm, height = 2 mm) were embedded in optimal cutting temperature medium (Fisher Scientific) and frozen at 80 °C. The frozen blocks were sectioned along the plane of the construct at 20 lm thickness and mounted on glass slides. The cryosections collected from the top, middle and bottom of the construct were stained with 0.1% Picrosirius red solution (Polysciences, Inc., Warrington, PA) and specimens were observed under a polarized light microscope. The compressive modulus of hydrogels (6 lM RF, 40 s VBL irradiation) was measured via an indentation experiment with an Instron electromechanical testing machine (Instron, Model 5564, Norwood, MA). The indentation tests were conducted using cylindrical hydrogel disks (diameter = 9.8 mm, height = 5 mm) with a load cell capacity of 2.5 N and a flat-ended indenter (3 mm in diameter) at a strain rate of 1.67% s 1. The Poisson’s ratio (m) of the hydrogels was 0.25, which was measured using the optical method described previously [28–32]. 2.3. Scanning electron microscopy (SEM) The interior morphology of MeGC and MeGC/Col II hydrogels was observed using a Nova NanoSEM 230 microscope (FEI, Hillsboro, OR). The hydrogels were fixed with 2.5% glutaraldehyde for 2 h at room temperature and interior morphology was imaged in low vacuum mode (magnification 10,000) and wet condition without sample drying and surface coating to minimize deformation of the hydrogel structure. 2.4. In vitro degradation Hydrogel degradation was observed by incubating the hydrogels in 3 ml of PBS with or without lysozyme (2 mg ml 1) at 37 °C for up to 21 days. The crosslinked hydrogels were washed three times in deionized water prior to the degradation test to remove uncrosslinked MeGC from the hydrogels. The lysozyme solution was replaced twice a week. At various time points, the incubating medium was removed and the wet mass was obtained after gentle blotting without drying the hydrogels. The degradation ratios were expressed as the ratio of the remaining weight to the initial weight of the hydrogels. All of the experiments were performed in triplicate. 2.5. In vitro release study and evaluation of TGF-b1 bioactivity For protein encapsulation in hydrogels, 400 ng of Texas RedÒ-conjugated avidin (Invitrogen, Carlsbad, CA) or TGF-b1 (PeproTech, Rocky Hill, NJ) was mixed with 40 ll of 2% MeGC solution in the presence or absence of Col II to give a final protein concentration of 10 lg ml 1. The protein suspension was irradiated with VBL for 40 s to form a hydrogel at a final RF initiator concentration of 6 lM and incubated at 37 °C for 20 min before using. Since both MeGC and MeGC/Col II gelation occurred within 20 s of irradiation at the given RF concentration, most of proteins initially used were expected to be entrapped in the hydrogel network formed.

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Fig. 1. Design of cartilage mimetic hydrogel. MeGC, a mixed solution of MeGC and Col II, and a mixed solution of TGF-b1, MeGC and Col II were irradiated with VBL in the presence of RF (photoinitiator) to form MeGC, MeGC/Col II and MeGC/Col II/TGF-b1 hydrogels, respectively.

To investigate the release of proteins from hydrogels, the obtained hydrogels were transferred to 24-well plate and incubated in PBS or chondrogenic medium (CM) consisting of Dulbecco’s modified Eagle’s medium (DMEM) with 10% fetal bovine serum (FBS), ITS+ Premix supplement (BD Biosciences, Bedford, MA), 100 nM dexamethasone, 40 lg ml 1 L-proline, 1 mM sodium pyruvate and 50 lg ml 1 L-ascorbic acid 2-phosphate (all Sigma– Aldrich). The final concentration of each avidin and TGF-b1 in hydrogels was 10 lg ml 1. Triplicate experiments were carried out. The release medium was replaced at the designated time interval. Released avidin was detected by fluorescence. The amount of TGF-b1 released into the release medium was analyzed by ELISA (R&D systems). The bioactivity of the TGF-b1 released from the hydrogels was assessed by measuring its ability to inhibit proliferation of human lung epithelia cells (CCL-185) as previously described [33]. TGFb1-loaded hydrogels were incubated in PBS for 24 h and the supernatant was collected for the assay. Freshly reconstituted TGF-b1 was used as a control. CCL-185 (ATCC, Manassas, VA) cells were cultured in DMEM supplemented with 10% FBS, 100 U ml 1 penicillin and 100 lg ml 1 streptomycin at 37 °C in a 5% CO2 humidified atmosphere. At the beginning of the assay, cultures were

washed once and replaced with fresh DMEM, then the hydrogel supernatant was added to the culture medium at final TGF-b1 concentrations ranging from 0 to 1.25, 2.5 and 5 ng ml 1. After incubation for 24 h, the proliferation of cells was assessed using a Cell Counting Kit-8 (CCK-8, Dojindo, Kumamoto, Japan) according to the manufacturer’s protocol.

2.6. In vitro 3-D culture of cells in hydrogels Rabbit chondrocytes were isolated from the knee of 3-monthold New Zealand white rabbits and expanded in basal culture medium (BM; DMEM with 10% FBS, 100 U ml 1 penicillin, and 100 lg ml 1 streptomycin) at 37 °C in a 5% CO2 humidified atmosphere. To evaluate the effect of Col II on chondrogenesis in MeGC hydrogels, rabbit chondrocytes were photoencapsulated as previously described [27]. Briefly, 4 ll of cell suspension (passage 3) in BM was mixed with 40 ll of MeGC or MeGC/Col II solution prepared in PBS containing 6 lM RF at a final concentration of 2  106 cells ml 1). The cell suspension was exposed to VBL for 40 s to form hydrogels and cultured in BM. The medium was replaced twice per week for up to 42 days.

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Dr. Denis Evseenko kindly provided human fetal chondrocytes, BMSCs and SMSCs. The human fetal chondrocytes were expanded in DMEM/F12 with 10% FBS, 100 U ml 1 penicillin and 100 lg ml 1 streptomycin. The BMSCs and SMSCs were expanded in DMEM with 20% FBS, 100 U ml 1 penicillin and 100 lg ml 1 streptomycin. To evaluate the multilineage differentiation capacity of both BMSCs and SMSCs, the cells were cultured in specific differentiation medium (n = 3). Adipogenesis was induced using adipogenic medium consisting of BM, 5 lg ml 1 insulin, 0.5 mM 3-isobutyl1-methylxanthine, 0.1 mM indomethacin and l lM dexamethasone for 14 days, and lipid accumulation in cells was detected with oil red O staining. For chondrogenesis, cell-pellets (1  106 cells per pellet) were cultured for 21 days in CM supplemented with TGF-b1 (10 ng ml 1), and sulfated glycosaminoglycan (sGAG) secretion from the cells was observed with Safranin-O staining. To investigate osteogenesis, cells were cultured for 3 days in osteogenic medium consisting of BM, 50 lg ml 1 L-ascorbic acid, 10 mM glycerophosphate and 100 nM dexamethasone, and alkaline phosphatase (ALP) expression from cells was detected with ALP staining. Both SMSCs and BMSCs showed the ability to differentiate into fat, cartilage and bone cells (Supplementary data Fig. S1). To further evaluate the effect of TGF-b1 on chondrogenesis in MeGC/Col II hydrogels, human fetal chondrocytes (passages 3–4), BMSCs (passages 4–5) or SMSCs (passages 4–5) were encapsulated in MeGC/Col II or MeGC/Col II/TGF-b1 (the final concentration of TGF-b1 was 10 lg ml 1) hydrogels at a density of 10  106 cells ml 1. Hydrogels were cultured in CM and the medium was replaced twice per week for up to 21 days. The 3-D cultured samples were analyzed for cellular aggregation, proliferation and viability. Cellular aggregation was determined using bright-field microscope images and the size distribution of the aggregates was quantified using NIH-ImageJ software (http:/rsb.info.nih.gov/ ij/) (n = 3). The proliferation of cells in the hydrogels was assessed using CCK-8 according to the manufacturer’s protocol (n = 3). The viability of the cells in the hydrogels was determined using a Live/Dead assay kit (Invitrogen). Stained cells were observed using an Olympus IX71 microscope (Olympus, Tokyo, Japan). 2.7. In vivo subcutaneous implantation For in vivo subcutaneous implantation, we prepared 100 ll of VBL-crosslinked hydrogels of MeGC/Col II ( TGF-b1) or MeGC/Col II/TGF-b1 (the final TGF-b1 concentration was 10 lg ml 1; +TGFb1) encapsulating various cells (SMSCs, BMSCs, or chondrocytes) at a density of 10  106 cells ml 1. Four-week-old CD-1 male mice (Charles River Laboratories, Cambridge, MA) were randomly divided into six groups: (1) SMSCs TGF-b1, (2) SMSCs + TGF-b1, (3) BMSCs TGF-b1, (4) BMSCs + TGF-b1, (5) chondrocytes TGFb1 and (6) chondrocytes + TGF-b1. After a week of in vitro chondrogenic induction, the constructs were implanted subcutaneously on the backs of mice (two hydrogels per mouse; three mice per group). The mice were sacrificed 3 weeks after the implantation. All animal handling and studies were performed in accordance with the guidelines of the University of California, Los Angeles (UCLA) Animal Research Committee. 2.8. Histological analysis For histology, cultured constructs were fixed in 10% neutral buffered formalin, embedded in paraffin and sectioned at 5 lm. Sections were deparaffinized and stained with Alcian blue or Safranin-O to assess sGAG secretion, and Col II expression was determined by immunohistochemistry (IHC). For IHC, sections were incubated with primary antibody against Col II (anti-human Col II; EMD Millipore, Billerica, MA) and the antibody was detected

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using the SuperPicture™ polymer detection kit with DAB substrate (Invitrogen) per the manufacturer’s instructions. Images were obtained using the Olympus IX71 microscope. The areas where sGAG and Col II were expressed were relatively quantified using NIH-ImageJ software (n = 3). Image analysis was performed on the positively stained cellular areas as well as on the ECM surrounding the cells since the synthesized sGAG or Col II could be localized in the intracellular area of cells. 2.9. Statistical analysis Statistical analysis was performed using one- or two-way analysis of variance followed by Tukey’s post hoc test. A value of p < 0.05 was considered as significant. 3. Results and discussion 3.1. Hydrogel characterization VBL activates and decomposes RF (vitamin B2) to form free radicals [34]. The methacrylate groups in MeGC polymerize via free radical polymerization and form a crosslinked network [26]. Both MeGC and MeGC/Col II solutions showed sol-to-gel transition by VBL irradiation in the presence of RF (Fig. 2a). The sol-to-gel phase transition time was investigated as a function of RF concentration using a vial-tilting method. The MeGC solution with 1.5 lM of RF showed a sol-to-gel transition after 80 s of VBL irradiation. The addition of Col II significantly decreased the gelation time from 80 to 24 s at 1.5 lM of RF. As the RF concentration increased, the gelation time decreased for both MeGC and MeGC/Col II solutions. The observed decreases in gelation time caused by the addition of Col II and the increase in RF concentration can be attributed to the formation of an interpenetrating network by Col II in the crosslinked MeGC network and higher amounts of free radicals produced by RF, respectively. The sol content of the MeGC hydrogels rapidly decreased from 72.6 to 16.4% with an increase in irradiation time from 10 to 40 s as measured using the technique previously described (Supplementary data Fig. S2) [26]. There was no significant change in sol content with irradiation beyond 40 s. The addition of Col II did not result in significant differences in sol content. In further studies, 6 lM RF and 40 s of irradiation time were selected, since gelation occurred within 20 s above 3 lM RF and the gelation time did not change significantly. The interior morphology of the hydrogels was determined using SEM in low vacuum mode. A cross-sectional SEM image of MeGC showed a porous microstructure with smooth MeGC struts. In contrast, MeGC/Col II hydrogels exhibited nanofibrous structures within the MeGC pores, indicating the formation of characteristic Col fibers by the incorporation of Col II (Fig. 2b). The homogeneous distribution of Col II was verified throughout the hydrogels, as shown by Picrosirius red staining (Fig. 2c). The compressive modulus of the hydrogels was characterized in the presence or absence of Col II (Fig. 2d). The addition of Col II significantly increased the modulus of MeGC, from 4.6 to 7.0 kPa. These results suggest that Col II incorporation into MeGC induced the formation of nanofibrous structures, which increased the stiffness of the hydrogel. Although the observed modulus of the hydrogels are orders of magnitudes lower than those of human articular cartilage (4–19 MPa) [35], our hydrogel system was designed to deliver progenitor cells and growth factors in the defect site and to provide excellent microenvironments for cell growth and differentiation. It is expected that implanted cells will generate new cartilage tissue as the hydrogel degrades, which in turn will provide ultimate mechanical support. Although the gel modulus can be increased with longer irradiation up to 300 s without compromising cell viability [25,27], increasing the hydrogel modulus over

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Fig. 2. Characterization of hydrogel systems. (a) Gelation time of chitosan (MeGC) and MeGC/Col II solutions as a function of RF concentration determined by vial tilting method (n = 3; ⁄⁄p < 0.01 compared with MeGC). (b) Interior morphology of MeGC and MeGC/Col II determined by SEM. The scale bar is 10 lm. (c) Polarized light microscopic images of MeGC and MeGC/Col II hydrogels stained with Picrosirius red (gel surface). Scale bars are 100 lm. Picrosirius red staining was also performed on the thin cryosections (20 lm) collected from the top, middle and bottom of disk-shaped hydrogels (diameter = 5 mm, height = 2 mm). Scale bars are 200 lm. (d) Mechanical strength of hydrogels (n = 3; ⁄⁄p < 0.01 compared with MeGC). (e) In vitro gel degradation in PBS with (filled markers) or without (empty markers) lysozyme.

13 kPa significantly inhibited sGAG secretion from chondrocytes encapsulated in the hydrogel (data not shown), indicating the importance of a good balance between initial mechanical strength and cell differentiation for successful chondrogenesis. Scaffolds for tissue engineering should be biodegradable in vivo to promote tissue ingrowth and release incorporated growth factors. Chitosan is enzymatically degraded in vivo by lysozyme present in a variety of tissues and body fluids, including cartilage [36,37]. The degradation behavior of VBL-crosslinked chitosan hydrogels was evaluated in a physiologically relevant lysozyme solution (0.2% w/v) as the lysozyme level of human hyaline cartilage can range from 0.80 to 3.03 mg g 1 cartilage [38] and lysozyme can degrade chitosan by binding to the sugar chains present in the chitosan backbone. Approximately 14% mass loss was observed in MeGC hydrogels during the first day in the presence of lysozyme, with a cumulative mass loss of 32.5% after 21 days, whereas only 4.3% mass loss was observed after 21 days without lysozyme (Fig. 2e). The observed large initial mass loss is likely due to the degradation of loosely crosslinked MeGC from the hydrogel surface. The addition of Col II to the MeGC hydrogels did not significantly affect their degradation behavior.

In our previous study, we showed that MeGC hydrogels could feasibly be used as a chondrogenic matrix [25]. To further investigate the effect of Col II on chondrogenesis, rabbit chondrocytes were encapsulated in MeGC in the presence or absence of Col II and cultured for 6 weeks (Fig. 3). Bright-field observations showed that Col II significantly increased the formation of chondrocyte aggregation. MeGC hydrogels showed a high level of cell viability (over 93%), and the addition of Col II did not have any significant effect on that cell viability. Larger cellular aggregations were found in MeGC/Col II compared to MeGC hydrogels. After 6 weeks in culture, sGAG, one of main cartilaginous ECM components, was expressed more intensely in MeGC/Col II hydrogels compared to in MeGC hydrogels, as assessed by Alcian blue and Safranin-O staining. These results suggest that Col II enhances cartilaginous ECM production in VBL-crosslinked hydrogel systems. 3.2. In vitro release of TGF-b1 from MeGC/Col II hydrogel TGF-b1 is a well-documented potential chondrogenic factor. However, direct use of TGF-b1 is restricted due to its instability and rapid enzymatic degradation, requiring high therapeutic doses

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Fig. 3. Chondrogenic ability of MeGC and MeGC/Col II hydrogel systems were compared through 3-D culture of chondrocytes in each hydrogel. Cell cluster formation, cell viability and sGAG accumulation were investigated by optical images, Live/Dead staining, Alcian blue staining and Safranin-O staining, respectively. The scale bar is 200 lm. The relative area of cell clusters and sGAG expression were analyzed by ImageJ (n = 3; ⁄⁄p < 0.01 compared with MeGC).

that can lead to inefficient cartilage formation and unwanted osteophyte formation [39,40]. To investigate the feasibility of using MeGC hydrogels as sustained protein release carriers, avidin or TGF-b1 was encapsulated in the hydrogels by mixing proteins with solutions of unpolymerized MeGC and Col II followed by VBLcrosslinking in the presence of RF (Fig. 4a). Avidin was selected as a model protein as its net positive charge (pI = 10.5) is similar to that of TGF-b1 (pI = 9.5) at physiological pH. The encapsulated protein was homogeneously distributed throughout the hydrogels, as shown by using fluorescently labeled avidin (Supplementary data Fig. S3). The release of fluorescently labeled avidin from hydrogels was measured by incubating the hydrogels in PBS for 21 days (Fig. 4b). When encapsulated in MeGC hydrogel, approximately 81% of the initially loaded avidin was released from the hydrogels during the first day, followed by a gradual release up to day 21. The incorporation of Col II into the MeGC hydrogels reduced the initial burst of protein to 18%, with a cumulative release of 42% after 21 days. This may be due to the affinity binding of cationic avidin to Col II in the MeGC/Col II hydrogel, which significantly suppressed the initial burst release of avidin. These results indicate that MeGC/Col II has a high potential as a sustained delivery system for positively charged proteins such as TGF-b1. ELISA assay showed that a significantly higher initial burst release of TGF-b1 was observed from MeGC hydrogels compared with MeGC/Col II hydrogels in PBS, suggesting affinity binding of TGFb1 to Col II in the MeGC/Col II hydrogels (Fig. 4c). We further investigated the feasibility of the MeGC/Col II hydrogels to support sustained release of TGF-b1 in both PBS and in CM containing serum and other chondrogenic molecules, which may better represent physiological conditions (Fig. 4d). In CM, 16% of the initially loaded TGF-b1 was released from MeGC/Col II at day 1, followed by slow release up to day 21, with a cumulative release of 26%. Compared to the release in PBS, the release of TGF-b1 was higher in CM. This is possibly due to the reduced electrostatic interaction between TGF-b1 and Col II by competition between proteins in CM and TGF-b1 for the same binding sites on Col II. Similar observations of increased release in the presence of serum proteins have been reported using alginate beads loaded with vascular endothelial growth factor [41], and Col coatings loaded with bovine serum albumin or histone [42].

Free radicals produced during photopolymerization can cause incomplete protein release from hydrogels by covalently immobilizing encapsulated TGF-b1 onto the crosslinked MeGC hydrogels [43]. Moreover, the produced free radicals can have detrimental effects on encapsulated TGF-b1 by protein denaturation, and the released TGF-b1 may not be as active as native protein. Therefore, we assessed the bioactivity of released TGF-b1 by a CCL-185 growth-inhibition assay as previously described [33]. The supernatant collected from TGF-b1-loaded MeGC or MeGC/Col II hydrogels decreased the proliferation of CCL-185 cells in a dose-dependent manner, indicating that the bioactivity of TGF-b1 was preserved during the photoencapsulation process in the hydrogels. The supernatant from TGF-b1-loaded MeGC or MeGC/Col II hydrogels decreased cell proliferation to a similar extent to that of freshly reconstituted TGF-b1 (Fig. 5). 3.3. In vitro culture of SMSCs in MeGC/Col II hydrogels releasing TGFb1 SMSCs were encapsulated in MeGC/Col II hydrogels with or without TGF-b1. After 3 weeks in culture, bright-field images showed that most of the SMSCs were present as single cells, with few cell aggregations, in MeGC/Col II hydrogels (Fig. 6a). In contrast, TGF-b1 increased the number of cell aggregations in the hydrogels, and the relative proportion of large cellular aggregations over 600 lm2 was increased significantly, from 6.3 to 33.1% (Fig. 6b). The culture of SMSCs in the hydrogels was compared with that of chondrocytes or BMSCs. The seeded chondrocytes and BMSCs remained as single cells after 3 weeks in MeGC/Col II hydrogels. Abundant chondrocyte aggregations were observed, with some spindle cells, in hydrogels loaded with TGF-b1. However, BMSCs formed only a few aggregations in the MeGC/Col II hydrogel loaded with TGF-b1. SMSCs encapsulated in MeGC/Col II hydrogels showed a significant increase in cell proliferation compared to chondrocytes or BMSCs (Fig. 6c) in the presence and absence of TGF-b1. Live/Dead staining showed a high level of viability (>95%) of encapsulated cells over the 3 week culture period in both MeGC/Col II and MeGC/Col II/TGF-b1 hydrogels (Fig. 7a). sGAG secretion were determined by Safranin-O staining (Fig. 7b). The SMSC aggregation density in MeGC/Col II/TGF-b1 hydrogels was

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Fig. 4. Protein release from hydrogels. (a) Avidin or TGF-b1 was encapsulated in MeGC or MeGC/Col II hydrogel and then incubated in PBS. TGF-b1 was encapsulated in MeGC/ Col II hydrogel and then incubated in PBS or CM. Release profile of avidin (b) or TGF-b1 (c) from MeGC and MeGC/Col II hydrogels in PBS. (d) TGF-b1 release profile from MeGC/Col II in PBS or CM.

Fig. 5. Bioactivity of TGF-b1 released from hydrogels. CCL-185 cells were exposed to the supernatant collected from TGF-b1-loaded MeGC or MeGC/Col II hydrogels at final TGF-b1 concentrations ranging from 0 to 5 ng ml 1. Cell proliferation was assessed by the CCK-8 assay, compared with freshly reconstituted TGF-b1.

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Fig. 6. In vitro growth of cells cultured in hydrogels. (a) Bright-field images of SMSCs, BMSCs and chondrocytes after 3 weeks of culture in vitro. The scale bar is 200 lm. (b) Image analysis of cell cluster formation over 3 weeks of culture in MeGC/Col II hydrogel (n = 3; ⁄p < 0.05, ⁄⁄p < 0.01 compared with MeGC/Col II without TGF-b1). Areas of cell clusters were normalized to the area of BMSC clusters. (c) Cell proliferation analyzed using CCK-8 assay (n = 3).

higher than in TGF-b1-free MeGC/Col II hydrogels, consistent with the bright-field observations. Safranin-O staining was more intense in the cell aggregations of TGF-b1-loaded hydrogels compared with the SMSCs cultured in TGF-b1-free hydrogels, indicating increased GAG accumulation. Quantitation of Safranin-O staining by image analysis showed that positively stained area was increased 4.6-fold with the addition of TGF-b1 in MeGC/Col II hydrogels (Fig. 7c). BMSCs did not display noticeable staining for sGAG in MeGC/Col II hydrogels. Although chondrocytes were present as single cells

or small clusters in TGF-b1-free MeGC/Col II hydrogels, the region of Safranin-O-positive ECM extended into the hydrogel matrix surrounding the cell clusters. Delivery of TGF-b1 greatly increased Safranin-O staining of chondrocytes cultured in MeGC/Col II hydrogels. In the hydrogels, SMSCs appeared to proliferate faster with increased GAG deposition compared to BMSCs, indicating that SMSCs may be a desirable mesenchymal cell source for cartilage tissue engineering. Previous studies demonstrated that SMSCs had superior proliferative and chondrogenic abilities compared to

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Fig. 7. In vitro viability and differentiation of cells cultured in hydrogels. (a) Viability of cells in hydrogels after 3 weeks in culture. Cells were stained using a Live (green)/ Dead (red) kit. The scale bar is 200 lm. (b) Histological analysis of 3-D cultured cells in hydrogel after 3 weeks in culture. Sections were stained with Safranin-O to visualize sGAG expression. The scale bar is 200 lm. (c) Relative sGAG expression analysis using ImageJ base on the stained area and the intensity of the stain (n = 3; ⁄p < 0.05, ⁄⁄p < 0.01 compared with MeGC/Col II without TGF-b1).

Fig. 8. In vivo neocartilage formation. (a) Cells were encapsulated in MeGC/Col II or MeGC/Col II/TGF-b1 hydrogel and, after 1 week of in vitro chondrogenic induction, the constructs were implanted into subcutaneous pockets of nude mice. (b) Hydrogels after 3 weeks of implantation in vivo.

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Fig. 9. (a) Histological analysis via Safranin-O staining to determine the sGAG accumulation after 3 weeks of implantation. The scale bar is 200 lm. (b) sGAG expression was analyzed by ImageJ (n = 3; ⁄⁄p < 0.01 compared with compared with MeGC/Col II without TGF-b1).

other MSCs derived from bone marrow, adipose tissue, periosteum, and skeletal muscle [9,44]. 3.4. In vivo evaluation of TGF-b1-loaded MeGC/Col II hydrogels in a subcutaneous model We further evaluated the chondrogenic ability of MeGC/Col II/ TGF-b1 by implanting SMSC-laden hydrogels in a subcutaneous

pocket of nude mice (Fig. 8a). All animals survived with no adverse effects during the 3-week implantation. Gross histology of the constructs after sacrifice showed no significant difference in their final sizes (Fig. 8b). To visualize cartilaginous ECM expression, samples were sectioned and stained with Safranin-O for sGAG (Fig. 9a) and immunohistochemical staining for Col II (Fig. 10a). As observed in vitro, TGF-b1-delivered groups exhibited chondrocytic lacunae with significantly enhanced Safranin-O positive staining compared

Fig. 10. (a) Immunostaining to visualize Col II production. The scale bar is 200 lm. (b) The relative Col II expression was quantified using ImageJ based on the stained area and the intensity of the stain (n = 3; ⁄p < 0.05, ⁄⁄p < 0.01 compared with MeGC/Col II without TGF-b1).

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to TGF-b1-free groups, thus confirming increased proteoglycan accumulation. Similar to in vitro findings, delivery of TGF-b1 significantly increased the number of SMSC aggregations with intense sGAG secretion. Quantitation of Safranin-O staining showed that sGAG secretion was increased 2.7-fold following TGF-b1 delivery (Fig. 9b). In addition, SMSC groups expressed sGAG 3.8-fold higher than that detected in BMSC groups in the presence of TGF-b1. The sGAG expression in SMSC groups was comparable to that detected in chondrocytes groups. Immunochemical staining showed Col II expression in a similar manner (Fig. 10). Highly intense Col II staining was observed in SMSC-laden MeGC/Col II/TGF-b1 groups, which was comparable to chondrocyte groups, while little staining was observed in BMSC groups. These results showed that the MeGC hydrogels designed to confer the unique natural extracellular molecular signaling enhanced the chondrogenesis of MSCs. Future work will elucidate the molecular mechanisms involved in Col II binding and TGF-b1-mediated chondrogenesis by investigating gene expression profiling in the MeGC/Col II/TGF-b1 hydrogel system. Although we chose the subcutaneous implantation model for the initial testing of our hydrogel systems, mechanical stimuli would be crucial to the cartilage repair process. Therefore, additional studies in articular cartilage defects under mechanical loading will be performed for the future clinical translation. Moreover, further assessment of the resulting systems in cartilage defects including a chronic osteoarthritis lesion would establish molecular strategies recapitulating the natural extracellular signaling to promote articular cartilage regeneration. A chronic osteoarthritis lesion frequently includes large articular cartilage defects where the containment of therapeutic agents is challenging in these defects. Our hydrogel system may serve as appropriate delivery modalities for stem cells and growth factors to the site of extensive cartilage damage. 4. Conclusion We have developed a cartilage mimetic hydrogel microenvironment by combining important physical and soluble cues for cartilage development, including Col II and TGF-b1, into a VBLcrosslinkable MeGC hydrogel. Our in vitro studies showed that incorporation of Col II in the chitosan matrix enhanced cellular condensation and chondrogenesis of encapsulated chondrocytes and supported sustained release of TGF-b1. Col II-mediated chondrogenesis was significantly enhanced in the presence of TGF-b1 both in vitro and in a subcutaneous model. Our findings suggest that TGF-b1-loaded MeGC/Col II hydrogels provide a favorable microenvironment to promote chondrogenesis and have a great potential as a biomaterial for cartilage regeneration. Acknowledgements This work was supported by the UCLA Academic Senate Research Award and a UCLA School of Dentistry Faculty seed grant. Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figures 1–4 and 6–10 are difficult to interpret in black and white. The full colour images can be found in the on-line version, at http://dx.doi.org/10.1016/ j.actbio.2014.10.013. Appendix B. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.actbio.2014.10. 013.

References [1] Hunziker EB. Articular cartilage repair: basic science and clinical progress. A review of the current status and prospects. Osteoarthr Cartilage 2002;10:432–63. [2] Brittberg M, Lindahl A, Nilsson A, Ohlsson C, Isaksson O, Peterson L. Treatment of deep cartilage defects in the knee with autologous chondrocyte transplantation. N Engl J Med 1994;331:889–95. [3] Benya PD, Padilla SR, Nimni ME. Independent regulation of collagen types by chondrocytes during the loss of differentiated function in culture. Cell 1978;15:1313–21. [4] Diaz-Romero J, Gaillard JP, Grogan SP, Nesic D, Trub T, Mainil-Varlet P. Immunophenotypic analysis of human articular chondrocytes: changes in surface markers associated with cell expansion in monolayer culture. J Cell Physiol 2005;202:731–42. [5] Tallheden T, Bengtsson C, Brantsing C, Sjogren-Jansson E, Carlsson L, Peterson L, et al. Proliferation and differentiation potential of chondrocytes from osteoarthritic patients. Arthritis Res Ther 2005;7:R560–8. [6] Pittenger MF, Mackay AM, Beck SC, Jaiswal RK, Douglas R, Mosca JD, et al. Multilineage potential of adult human mesenchymal stem cells. Science 1999;284:143–7. [7] Richardson SM, Hoyland JA, Mobasheri R, Csaki C, Shakibaei M, Mobasheri A. Mesenchymal stem cells in regenerative medicine: opportunities and challenges for articular cartilage and intervertebral disc tissue engineering. J Cell Physiol 2010;222:23–32. [8] Al-Nbaheen M, Vishnubalaji R, Ali D, Bouslimi A, Al-Jassir F, Megges M, et al. Human stromal (mesenchymal) stem cells from bone marrow, adipose tissue and skin exhibit differences in molecular phenotype and differentiation potential. Stem Cell Rev Rep 2013;9:32–43. [9] Sakaguchi Y, Sekiya I, Yagishita K, Muneta T. Comparison of human stem cells derived from various mesenchymal tissues – superiority of synovium as a cell source. Arthritis Rheum 2005;52:2521–9. [10] Yoshimura H, Muneta T, Nimura A, Yokoyama A, Koga H, Sekiya I. Comparison of rat mesenchymal stem cells derived from bone marrow, synovium, periosteum, adipose tissue, and muscle. Cell Tissue Res 2007;327:449–62. [11] Chaipinyo K, Oakes BW, Van Damme MPI. The use of debrided human articular cartilage for autologous chondrocyte implantation: maintenance of chondrocyte differentiation and proliferation in type I collagen gels. J Orthop Res 2004;22:446–55. [12] Vinatier C, Mrugala D, Jorgensen C, Guicheux J, Noel D. Cartilage engineering: a crucial combination of cells, biomaterials and biofactors. Trends Biotechnol 2009;27:307–14. [13] Ahmed TAE, Hincke MT. Strategies for articular cartilage lesion repair and functional restoration. Tissue Eng Part B 2010;16:305–29. [14] Zhang L, Yuan T, Guo LK, Zhang XD. An in vitro study of collagen hydrogel to induce the chondrogenic differentiation of mesenchymal stem cells. J Biomed Mater Res A 2012;100A:2717–25. [15] Pabbruwe MB, Kafienah W, Tarlton JF, Mistry S, Fox DJ, Hollander AP. Repair of meniscal cartilage white zone tears using a stem cell/collagen-scaffold implant. Biomaterials 2010;31:2583–91. [16] Lu Z, Doulabi BZ, Huang C, Bank RA, Helder MN. Collagen type II enhances chondrogenesis in adipose tissue-derived stem cells by affecting cell shape. Tissue Eng Part A 2010;16:81–90. [17] Kopesky PW, Byun S, Vanderploeg EJ, Kisiday JD, Frisbie DD, Grodzinsky AJ. Sustained delivery of bioactive TGF-beta1 from self-assembling peptide hydrogels induces chondrogenesis of encapsulated bone marrow stromal cells. J Biomed Mater Res A 2013. [18] Chang KY, Hung LH, Chu IM, Ko CS, Lee YD. The application of type II collagen and chondroitin sulfate grafted PCL porous scaffold in cartilage tissue engineering. J Biomed Mater Res A 2010;92A:712–23. [19] Vernon RB, Sage EH. Contraction of fibrillar type I collagen by endothelial cells: a study in vitro. J Cell Biochem 1996;60:185–97. [20] Wang AY, Leong S, Liang YC, Huang RC, Chen CS, Yu SM. Immobilization of growth factors on collagen scaffolds mediated by polyanionic collagen mimetic peptides and its effect on endothelial cell morphogenesis. Biomacromolecules 2008;9:2929–36. [21] Dong CM, Wu X, Caves J, Rele SS, Thomas BS, Chaikof EL. Photomediated crosslinking of C6-cinnamate derivatized type I collagen. Biomaterials 2005;26:4041–9. [22] Jung Y, Chung YI, Kim SH, Tae G, Kim YH, Rhie JW. In situ chondrogenic differentiation of human adipose tissue-derived stem cells in a TGF-beta1 loaded fibrin–poly(lactide-caprolactone) nanoparticulate complex. Biomaterials 2009;30:4657–64. [23] Holland TA, Tabata Y, Mikos AG. In vitro release of transforming growth factorbeta 1 from gelatin microparticles encapsulated in biodegradable, injectable oligo(poly(ethylene glycol) fumarate) hydrogels. J Control Release 2003;91:299–313. [24] Diederichs S, Baral K, Tanner M, Richter W. Interplay between local versus soluble transforming growth factor-beta and fibrin scaffolds: role of cells and impact on human mesenchymal stem cell chondrogenesis. Tissue Eng Part A 2012;18:1140–50. [25] Hu JL, Hou YP, Park H, Choi B, Hou SY, Chung A, et al. Visible light crosslinkable chitosan hydrogels for tissue engineering. Acta Biomater 2012;8:1730–8. [26] Amsden BG, Sukarto A, Knight DK, Shapka SN. Methacrylated glycol chitosan as a photopolymerizable biomaterial. Biomacromolecules 2007;8:3758–66.

B. Choi et al. / Acta Biomaterialia 12 (2015) 30–41 [27] Park H, Choi B, Hu J, Lee M. Injectable chitosan hyaluronic acid hydrogels for cartilage tissue engineering. Acta Biomater 2013;9:4779–86. [28] Jurvelin JS, Buschmann MD, Hunziker EB. Optical and mechanical determination of Poisson’s ratio of adult bovine humeral articular cartilage. J Biomech 1997;30:235–41. [29] Wang CC, Chahine NO, Hung CT, Ateshian GA. Optical determination of anisotropic material properties of bovine articular cartilage in compression. J Biomech 2003;36:339–53. [30] Kelly TA, Ng KW, Wang CC, Ateshian GA, Hung CT. Spatial and temporal development of chondrocyte-seeded agarose constructs in free-swelling and dynamically loaded cultures. J Biomech 2006;39:1489–97. [31] Elliott DM, Narmoneva DA, Setton LA. Direct measurement of the Poisson’s ratio of human patella cartilage in tension. J Biomech Eng 2002;124:223–8. [32] Lee SY, Pereira BP, Yusof N, Selvaratnam L, Yu Z, Abbas AA, et al. Unconfined compression properties of a porous poly(vinyl alcohol)–chitosan-based hydrogel after hydration. Acta Biomater 2009;5:1919–25. [33] Zheng X, De Paiva CS, Rao K, Li DQ, Farley WJ, Stern M, et al. Evaluation of the transforming growth factor-beta activity in normal and dry eye human tears by CCL-185 cell bioassay. Cornea 2010;29:1048–54. [34] Liang JY, Yuann JM, Cheng CW, Jian HL, Lin CC, Chen LY. Blue light induced free radicals from riboflavin on E. coli DNA damage. J Photochem Photobiol, B 2013;119:60–4. [35] Shepherd DE, Seedhom BB. The ‘instantaneous’ compressive modulus of human articular cartilage in joints of the lower limb. Rheumatology 1999;38:124–32. [36] Pangburn SH, Trescony PV, Heller J. Lysozyme degradation of partially deacetylated chitin, its films and hydrogels. Biomaterials 1982;3:105–8.

41

[37] Varum KM, Myhr MM, Hjerde RJ, Smidsrod O. In vitro degradation rates of partially N-acetylated chitosans in human serum. Carbohydr Res 1997;299:99–101. [38] Greenwald RA, Josephson AS, Diamond HS, Tsang A. Human cartilage lysozyme. J Clin Investig 1972;51:2264–70. [39] van Lent PL, Blom AB, van der Kraan P, Holthuysen AE, Vitters E, van Rooijen N, et al. Crucial role of synovial lining macrophages in the promotion of transforming growth factor beta-mediated osteophyte formation. Arthritis Rheum 2004;50:103–11. [40] van Beuningen HM, Glansbeek HL, van der Kraan PM, van den Berg WB. Osteoarthritis-like changes in the murine knee joint resulting from intraarticular transforming growth factor-beta injections. Osteoarthritis Cartilage 2000;8:25–33. [41] Gu F, Amsden B, Neufeld R. Sustained delivery of vascular endothelial growth factor with alginate beads. J Control Release 2004;96:463–72. [42] Lee M, Chen TT, Iruela-Arispe ML, Wu BM, Dunn JC. Modulation of protein delivery from modular polymer scaffolds. Biomaterials 2007;28:1862–70. [43] Lin CC, Sawicki SM, Metters AT. Free-radical-mediated protein inactivation and recovery during protein photoencapsulation. Biomacromolecules 2008;9:75–83. [44] De Bari C, Dell’Accio F, Karystinou A, Guillot PV, Fisk NM, Jones EA, et al. A biomarker-based mathematical model to predict bone-forming potency of human synovial and periosteal mesenchymal stem cells. Arthritis Rheum 2008;58:240–50.