Microribbon-hydrogel composite scaffold accelerates cartilage regeneration in vivo with enhanced mechanical properties using mixed stem cells and chondrocytes

Biomaterials 228 (2020) 119579

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Microribbon-hydrogel composite scaffold accelerates cartilage regeneration in vivo with enhanced mechanical properties using mixed stem cells and chondrocytes

T

Heather Rogana, Francisco Ilaganb, Xinming Tongc, Constance R Chuc, Fan Yanga,c,∗ a

Department of Bioengineering, Stanford University, Stanford, CA, 94305, USA Department of Biology, Stanford University, Stanford, CA, 94305, USA c Department of Orthopaedic Surgery, Stanford University, Stanford, CA, 94305, USA b

A R T I C LE I N FO

A B S T R A C T

Keywords: Cartilage tissue engineering ADSCs Chondrocytes Co-culture Microribbons Mechanical properties

Juvenile chondrocytes are robust in regenerating articular cartilage, but their clinical application is hindered by donor scarcity. Stem cells offer an abundant autologous cell source but are limited by slow cartilage deposition with poor mechanical properties. Using 3D co-culture models, mixing stem cells and chondrocytes can induce synergistic cartilage regeneration. However, the resulting cartilage tissue still suffers from poor mechanical properties after prolonged culture. Here we report a microribbon/hydrogel composite scaffold that supports synergistic interactions using co-culture of adipose-derived stem cells (ADSCs) and neonatal chondrocytes (NChons). The composite scaffold is comprised of a macroporous, gelatin microribbon (μRB) scaffolds filled with degradable nanoporous chondroitin sulfate (CS) hydrogel. We identified an optimal CS concentration (6%) that best supported co-culture synergy in vitro. Furthermore, 7 days of TGF-β3 exposure was sufficient to induce catalyzed cartilage formation. When implanted in vivo, μRB/CS composite scaffold supported over a 40-fold increase in compressive moduli of cartilage produced by mixed ADSCs/NChons to ~330 kPa, which surpassed even the quality of cartilage produced by 100% NChons. Together, these results validate μRB/CS composite as a promising scaffold for cartilage regeneration using mixed populations of stem cells and chondrocytes.

1. Introduction Articular cartilage has limited capacity to self-repair upon injury due to low cellularity and vascularity. Cartilage defects can lead to early onset of the painful and debilitating disease osteoarthritis (OA), a leading disability among adults [1–3]. Driven by significant need, cartilage tissue was one of the first target tissue types for tissue engineering, yet achieving effective articular cartilage regeneration remains difficult after decades of research efforts [4–8]. A critical component to consider for articular cartilage regeneration is the appropriate choice of cell source. An ideal cell source would allow easy harvest in a minimally invasive manner and expandability in vitro while maintaining the right phenotype for producing high quality articular cartilage. Chondrocytes are the resident cell type in articular cartilage and possess the appropriate phenotype for new cartilage formation. The ability of chondrocytes to produce quality cartilage decreases with age and disease state [9–11], with juvenile chondrocytes being the most robust cell sources for producing cartilage. However,

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obtaining sufficient numbers of chondrocytes is difficult due to donor site morbidity, scarcity of donor tissues, and tendency to dedifferentiate during in vitro expansion [3,8]. Unlike chondrocytes, mesenchymal stem cells are much more abundant and can be isolated from multiple adult tissues in a minimally invasive manner. A variety of mesenchymal stem cell sources have been used for cartilage tissue engineering including bone marrow-derived mesenchymal stem cell (BM-MSCs) [12–15], synovium-derived mesenchymal stem cells (S-MSCs) [16,17], and adipose-derived stem cells (ADSCs) [12,18,19]. While stem cells offer a more abundant autologous cell source, the quality of resulting cartilage is generally inferior than cartilage produced by chondrocytes and is characterized by fibrocartilage with inferior mechanical properties [4]. This leaves a scientific challenge to minimize the number of chondrocytes, due to their scarcity, and still produce robust cartilage with an abundant cell source, such as stem cells. To overcome the limitations of using stem cells or chondrocytes alone, here we aim to utilize mixed populations of chondrocytes and stem cells as cell sources for cartilage regeneration. By optimizing the mixed cell ratio and

Corresponding author. Department of Bioengineering, Stanford University, Stanford, CA, 94305, USA. E-mail address: [email protected] (F. Yang).

https://doi.org/10.1016/j.biomaterials.2019.119579 Received 17 May 2019; Received in revised form 26 September 2019; Accepted 24 October 2019 Available online 31 October 2019 0142-9612/ © 2019 Elsevier Ltd. All rights reserved.

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In our previous paper, we have demonstrated that paracrine signal exchange between ADSCs and NChons are responsible for catalyzed cartilage formation [20]. To enable retention of paracrine signals, while also leveraging the advantage of μRB scaffolds to accelerate cartilage regeneration, we designed a novel μRB/hydrogel composite scaffold containing macroporous gelatin μRBs filled with degradable nanoporous CS hydrogel. This would allow paracrine signals to be retained initially to induce synergistic interactions between ADSCs and NChons, and subsequent degradation of CS hydrogel would allow the released cells to leverage the benefit of macroporous μRB scaffolds. We first determine an optimal CS concentration that best supports co-culture synergy, and compared the effect of 7-days or 21-days of TGF-β exposure on the cartilage production by mixed cell populations in vitro. The optimized μRB/CS scaffold was subsequently examined for its ability to accelerate restoration of mechanical properties of resulting cartilage in vivo using a mouse subcutaneous model. Nanoporous PEG/ CS hydrogels, the hydrogel formulation used in our previous study, was included as a control. Outcomes were analyzed by quantifying cell proliferation, mechanical testing (unconfined compression), histology, and immunohistochemistry. The broader impacts of this work could improve not only cartilage tissue regeneration, but also other applications such as bone or muscle. The concept of forming a composite macroporous μRB scaffold with a rapidly degrading nanoporous hydrogel could aid in paracrine signal exchange for other mixed cell populations and improve the mechanical properties of many regenerated tissues.

culture conditions, using a mixed cell source has been shown to result in better articular cartilage formation with substantially fewer chondrocytes [20–28]. While previous work has shown promise in using a mixed cell population of stem cells and chondrocytes for cartilage regeneration, one remaining challenge is that resulting cartilage tissue still suffers from poor mechanical properties after prolonged in vitro culture. Many previous strategies using mixed cell populations for cartilage repair have used hydrogels (HGs) as 3D scaffolds given their injectability and wide use in the cartilage tissue engineering field [20–22]. Both synthetic and natural polymers have been used to fabricate hydrogels. Among the most commonly used ones include polyethylene glycol (PEG), chondroitin sulfate (CS), hyaluronic acid (HA), heparan sulfate (HS), and gelatin (Gel). These conventional hydrogel networks are generally nanoporous, which is beneficial for retaining paracrine signals. Recent studies from our group and others have shown that synergy during coculture is driven by soluble factor exchange between the stem cells and chondrocytes [20,29]. Specifically, when ADSCs and juvenile chondrocytes were mixed co-cultured in PEG/CS hydrogels, the newly formed cartilage was produced solely by juvenile chondrocytes, whereas the ADSCs catalyzed chondrocytes via paracrine signaling, increasing the proliferation and total cartilage ECM production by chondrocytes. However, the nanoporosity of hydrogel networks also pose physical constraints on the encapsulated cells, leading to delayed new matrix deposition. This is due to the need for degradation, which further decreases the compressive modulus of engineered tissues. Furthermore, cells encapsulated in nanoporous hydrogels generally resulted in newly formed cartilage restricted to pericellular regions, with a minimal increase in the overall mechanical properties of engineered cartilage [5,30]. One approach to overcome this challenge would be to vary the hydrogel composition. To assess whether varying hydrogel stiffness or biochemical compositions could accelerate the increase in the mechanical properties of cartilage produced by mixed cell populations, our group recently reported a combinatorial study by encapsulating mixed chondrocytes and ADSCs in 42 hydrogel compositions with tunable biochemical cues or stiffnesses [31]. However, regardless of the hydrogel compositions, the compressive moduli of resulting cartilage either stayed constant or even decreased over time due to degradation outpacing the speed of new matrix deposition [31]. There remains a critical need to design 3D scaffolds that can support synergistic interactions using mixed cell populations, while accelerating cartilage deposition with enhanced mechanical properties. We aim to achieve this goal by utilizing a novel composite scaffold of gelatinbased microribbons (μRBs) and chondroitin sulfate hydrogel. Our lab has previously reported a 3D cell niche of macroporous gelatin-based μRB scaffolds, which are formed via crosslinking of gelatin-based, μRB shaped hydrogel building blocks. This scaffold structure is particularly attractive for cartilage regeneration due to its great shock-absorbing mechanical properties. When subject to cyclic compression testing, μRB scaffolds can deform transiently then immediately return to original shape, whereas conventional hydrogels would easily fragment or deform [32]. Compared to nanoporous gelatin hydrogels, the macroporous μRB hydrogels also accelerated MSC-based cartilage deposition with significantly enhanced compressive moduli and interconnectivity of deposited matrix [32,33]. In a previous study, μRBs led to over a 20-fold increase in Young's modulus after 21 days of in vitro culture with improved interconnectivity of newly deposited cartilage throughout the scaffold [33]. In contrast, only a 2.8-fold increase in Young's modulus was observed in the stem cell-laden hydrogel control [33]. One limitation of this prior study is the undesirable fibrocartilage phenotype, which is typical for cartilage produced using MSCs as single cell sources. To enhance the quality of tissue engineered cartilage, one possible strategy is to explore the use of mixed cell populations, yet the potential of μRB scaffolds for supporting cartilage regeneration using a mixed population of stem cells and chondrocytes has never been explored before.

2. Materials and methods 2.1. Materials Gelatin type-A from porcine skin, dimethyl sulfoxide (DMSO, > 97% purity), methacrylic acid N-hydroxysuccinimide ester, glutaraldehyde, L-lysine hydrochloride, dexamethasone (> 97% purity), ascorbic-2-phosphate, proline, paraformaldehyde, Safranin O, Weigert's Hematoxylin solution, Alizarin Red stain, and Fluoromount were purchased from Sigma-Aldrich (St. Louis, MO). Penicillin and streptomycin, sodium pyruvate, Dulbecco's Modified Eagle Medium (DMEM), AlexaFluor 700, Bovine serum albumin (BSA), Pierce BCA Protein Assay, LIVE/DEAD™ cell viability assay, Fast Green stain, Quant-iT™ PicoGreen™ dsDNA assay, Alexa Fluor 488 goat anti-rabbit secondary antibody, Hoechst, optimal cutting temperature (OCT) compound, and Masson's Trichrome staining kit were purchased from Thermo Fisher Scientific (Pittsburgh, PA). All other sources of materials have been indicated within methods. All materials were used as directed by product information.

2.2. Gelatin microribbon synthesis Gelatin microribbons (μRBs) were synthesized with a modified protocol as previously described [32,33]. Gelatin was dissolved in dimethyl sulfoxide (DMSO, 19 wt%), before wet spinning. Spinning was performed into a bath of ethanol stirred at 500 rpm and 1.8 m from gelatin extrusion point. A syringe pump was used to extrude gelatin at a constant rate of at 5 ml/h. Precipitated gelatin was dried in 100% acetone for 3 h to flatten gelatin fibers and form ribbon structures with an average width of approximately 40 μm μRBs were then chopped in 100% ethanol to achieve an average length of 3 mm. Chopped μRBs were methacrylated by treating with methacrylic acid N-hydroxysuccinimide ester (1 mol% of gelatin) in methanol. μRBs were intracrosslinked by glutaraldehyde (0.08 mol% of gelatin) in methanol overnight at room temperature. Solution was quenched with 1% L-lysine in PBS for 2 h μRBs were washed thoroughly with PBS followed by water prior to freeze drying and storage at −20 °C. 2

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100 μg/mL sodium pyruvate, 100 units/mL penicillin and 100 μg/mL streptomycin [20,21]. For HG samples the precursor solution of 5% PEGDA, 3% CS, and 0.1% photoinitiator LAP in PBS was mixed homogeneously with cells, transferred to cylindrical molds (5 mm in diameter) and exposed to ultraviolet (UV) light (365 nm, 2 mW/cm2) for 4 min for polymerization. HG scaffolds were then cultured in chondrogenic media for 21 days prior to analysis or until implantation. For μRB and μRB/CS samples, lyophilized μRBs were first rehydrated with either PBS or CS (3%, 6%, or 9%), then mixed with cell suspension and LAP photoinitiator (0.1%). The final μRB concentration was 8% (w/v). Cells in μRB mixture were sandwiched between glass slides with a 2 mm spacer and exposed to UV light (365 nm, 2 mW/cm2) for 4 min for crosslinking. After 24 h of incubation in chondrogenic media, a biopsy punch (4.5 mm diameter) was used to punch out cylindrical scaffolds. Individual scaffolds were then cultured in chondrogenic media for 21 days prior to analysis or until implantation. In vitro analyses included cell viability assessment, cell proliferation quantification, mechanical testing, and histology.

2.3. Scanning electron microscopy (SEM) Acellular scaffolds were formed using 8% (w/v) with varying methacrylated chondroitin sulfate (CS) concentrations (3%, 6%, or 9%) or PBS and 0.1% photoinitiator (lithium phenyl-2,4,6-trimethylbenzoylphosphinate, LAP). CS was synthesized as previously reported [19]. Scaffolds were crosslinked under UV light (365 nm, 2 mW/cm2) for 4 min and allowed to swell for 24 h in PBS. Environmental SEM was performed on hydrated samples using a Hitachi S-3400 N variable pressure microscope (Schaumburg, IL). Sample stage was cooled gradually to −15 °C during pressurization to 50 Pa to avoid freeze-drying scaffolds. Images were captured with an electron beam of 15 kV and a distance around 5 mm. 2.4. Confocal microscopy Fluorescently labeled μRBs and CS were used for confocal microscopy of composite scaffolds. μRBs were labeled with a 5 μg/mL solution of AlexaFluor 700 in glutaraldehyde. FITC conjugated CRGDS peptide (Biomatik, Wilmington, DE) was mixed with CS gel proportionately to CS concentration prior to crosslinking: 3% CS/100 μM FITCRGD, 6% CS/200 μM FITC-RGD, and 9% CS/300 μM FITC-RGD. Acellular scaffolds were formed using 8% (w/v) labeled μRBs and either PBS or CS/FITC-RGD with 0.1% LAP. HG control was made using 5% PEGDA and 3% CS/100 μM FITC-RGD. All samples were equilibrated in PBS for 24 h prior to imaging. Leica SP8 microscope (Buffalo Grove, IL) was used to take images. Fiji/ImageJ software (NIH) with the 3D viewer plugin was used to reconstruct 3D images and take 360° movies.

2.8. Subcutaneous mouse implantation All guidelines for the care and use of laboratory animals were followed and protocols were approved by the Stanford institutional animal care and use committee. Scaffolds (HG and μRB/6%CS) containing co-culture (3:1 ratio of ADSCs:NChons) or monoculture controls were cultured for 7 days in vitro prior to subcutaneous implantation. 6-Week old athymic, female, nude mice were used (Charles River). Scaffolds were harvested after 21 and 56 days in vivo. Outcomes were analyzed by mechanical testing (n = 5) and histology (n = 3).

2.5. Protein elution BSA was used as a model protein to determine release profiles for all scaffold compositions following modified methods as previously reported [34]. Briefly, a final concentration of 2% (w/v) BSA was encapsulated in each scaffold: μRB, μRB+3% CS, μRB+6% CS, μRB+9% CS, and 5% PEGDA+3% CS. Following polymerization, scaffolds (n = 5 per group) were placed in 500 μl PBS. Elution was collected and scaffolds were moved to new wells with fresh PBS at each time point (5 min, 30 min, 1 h, 2 h, 4 h, 6 h, 8 h, and 24 h). BSA concentration of each collected elution was quantified using the Pierce BCA Protein Assay. Cumulative BSA released was calculated by summing the amount of BSA measured at each time point with all prior time points.

2.9. Cell viability and proliferation assay Cell viability was assessed at day 1 and day 21 of in vitro culture using the LIVE/DEAD™ cell viability assay. Images were taken using a Zeiss fluorescence microscope to visualize the live cells (green) and dead cells (red). Samples for cell proliferation quantification were lyophilized and digested in papainase solution (Worthington Biochemical, Lakewood, NJ) at 60 °C for 16 h. The Quant-iT™ PicoGreen™ dsDNA assay was used to quantify DNA. Ratio of day 21 DNA to day 1 DNA was used to calculate fold of cell proliferation.

2.6. Cell isolation and culture 2.10. Mechanical testing ADSC isolation: hADSCs were isolated from human adipose tissue as previously described [18]. ADSCs were expanded to passage 5 before use. Growth medium was composed of high glucose DMEM supplemented with 10 ng/mL basic fibroblast growth factor (bFGF) (Peprotech, Rocky Hill, NJ), 100 U/ml penicillin, and 0.1 mg/mL streptomycin. NChon isolation: Chondrocytes were harvested from the femoropatellar groove of calf stifles (3 days old) (Research 87, Boylston, MA). Articular cartilage was removed and digested as previously described [20]. NChons were stored in liquid nitrogen and did not undergo any expansion in vitro.

In vitro (n = 3) and in vivo (n = 5) samples were harvested for unconfined compression tests (Instron 5944 materials testing system (Instron Corporation, Norwood, MA), 10 N load cell (Interface, Inc., Scottsdale, AZ)). Tests were performed as previously described [35,36]. Briefly samples were compressed at a rate of 1% strain/sec to a maximum strain of 30% in PBS. Young's modulus was determined between 10 and 20% strain, a range that is physiologically relevant and has been widely used in previous literature by our group and others [35,36]. This allows direct comparison of elastic modulus of resulting cartilage across different studies.

2.7. 3D encapsulation and in vitro culture

2.11. Histology and immunohistochemistry

For co-culture ADSCs and NChon were encapsulated at a 3:1 ratio [20,23]. Cell density was kept constant at 15 million cells/ml. All scaffolds were cultured in chondrogenic medium composed of high glucose DMEM, 10 ng/mL recombinant human transforming growth factor beta 3 (TGF-β3) (Peprotech, Rocky Hill, NJ), 100 nM dexamethasone, 50 μg/mL ascorbic-2-phosphate, 40 μg/mL proline, 5 μg/ mL insulin-transferrin-selenium premix (BD Biosciences, San Jose, CA),

In vitro samples were harvested at day 21 and in vivo samples were harvested at day 0 (time of implantation, following 7 days in vitro), day 21, and day 56. Sample fixation was performed in 4% paraformaldehyde for 30 min prior to washing in PBS and transferring to OCT compound for 48 h and freezing in liquid nitrogen. Samples were sectioned with a thickness of 12 μm per slice. sGAG deposition was evaluated by staining with Safranin O and counterstaining with Fast Green. 3

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To test our hypothesis that increasing CS concentration would decrease protein release (loss of paracrine signals), we measured protein release using bovine serum albumin (BSA) as a model protein. μRBs alone resulted in the fastest protein release, and HG resulted in the slowest elution profile (Fig. 1C). For μRB + CS composite scaffolds, increasing CS concentration led to slower protein release, supporting our hypothesis. Adding 3% CS delayed the burst release observed in μRB alone control (Fig. 1C), which is likely due to delayed protein release from the nanoporous CS portion, and potential protein interactions with the CS.

Cell nuclei were counterstained with Weigert's Hematoxylin solution. Total collagen and tissue morphology were visualized with Masson's Trichrome staining. Alizarin Red stain was used evaluate mineralization. Samples were mounted with VectaMount (Vector Laboratories, Burlingame, CA). Immunostaining for type I, type II, and type X collagens was performed. Rabbit polyclonal antibodies type I, type II, and type X collagens (Abcam, Cambridge, MA) were diluted at 1:100 and incubated overnight at 4 °C. Secondary antibody (Alexa Fluor 488 goat anti-rabbit) was diluted at 1:200 with Hoechst (4 μg/mL) and incubated for 1 h at room temperature. Samples were mounted with the aqueous mounding medium, Fluoromount. Fluorescent imaging was performed on a Zeiss Observer. Z1 microscope (Dublin, CA) with Zen 2 software.

3.2. Composite scaffold composed of μRB with 6% CS supports highest synergy

2.12. Statistical analysis Following material characterization, we then evaluated the potential of composite scaffolds containing μRBs and CS to support catalyzed cell proliferation and subsequent cartilage production using mixed populations of ADSCs and NChons (3:1 ratio) (Fig. 2). Twenty-four hours after encapsulation, all composite scaffold groups showed high cell viability with a round cell morphology similar to that in the nanoporous HG control (Fig. 2A). In contrast, rapid cell spreading was observed in μRB alone control. These results confirmed that cells were initially encapsulated in the nanoporous CS gel portion within the composite scaffolds. After 21 days of culture, some cell spreading was observed in the μRB + CS groups, as the CS gel was degraded by the cells and mixed cell population were no longer physically trapped within the nanoporous CS hydrogels. We also observed large cell nodules in μRB +6% CS and 9% CS groups, but not in the 3% CS composite group (Fig. 2A). Prior reports evaluating co-culture synergy have linked cell nodule size with an increase in proliferation and catalyzed cartilage formation [20,22,31]. Quantitative evaluation of cell proliferation correlated with nodule size (Fig. 2B). As nodules began decreasing in size with μRBs containing 9% CS and the HG group, there was also a statically significant drop in fold of cell proliferation. After 21 days of culture, both μRBs alone and μRB+3% CS scaffolds failed to maintain their shape and decreased substantially in size due to extensive cellular contraction forces exerted by spreading cells (Fig. 2C). Taken together, we chose μRB +6% CS as the optimal formulation for further evaluation in supporting catalyzed cartilage formation by ADSCs and NChons. Since CS stains positive for Safranin O staining, it allows the use of histology to monitor the degradation of CS component in our composite scaffold (Fig. S2). By comparing the staining of day 1 and day 7 cellseeded samples, we can observe rapid degradation of the CS hydrogel component after 7 days.

GraphPad Prism (Graphpad Software, San Diego, CA) was used to perform statistical analysis and create graphs. Two-way analysis of variance and multiple pairwise comparisons with the Dunnett test were used to determine statistical significance (p < 0.05). All data are represented as mean ± standard deviation with at least three replicates per group for in vitro analysis and five replicates for in vivo analysis. 3. Results and discussion 3.1. Increasing CS concentration reduces pore size and protein diffusion in μRB/HG composite Most prior reports of synergistic interactions between stem cells and chondrocytes were conducted in 3D nanoporous HG culture [20–22,31]. While synergy was observed in conventional HG, the mechanical properties of resulting cartilage remain poor, with Young's modulus only reaching ~10% of the value of native cartilage (500–600 kPa) [31,37–39]. Using MSCs as a model cell type, our group has recently demonstrated macroporous gelatin μRB-based scaffolds enable a rapid increase in compressive modulus of tissue engineered cartilage, reaching over 200 kPa after only three weeks of culture [33]. To leverage the advantages of μRB-based scaffolds while still enabling paracrine signal induced synergistic interactions using mixed cell populations, here we developed a novel composite scaffold by incorporating fast degrading CS nanoporous hydrogels into macroporous μRB scaffolds. To determine the optimal CS concentration needed for supporting synergy, we first varied CS concentration (3%, 6% and 9%) and characterized compressive moduli and protein elution. As expected, increasing CS concentration increased compressive moduli of composite scaffolds from 5 kPa (for 3% CS) to 75 kPa (for 9% CS) (Fig. S1). Scanning electron microscopy (SEM) revealed there were minimal changes in porosity between μRB and μRB + 3% CS (Fig. 1A). μRB +6% CS showed a decrease in macroporosity, which was further reduced in μRB +9% CS. In comparison to the PEG/CS nanoporous HG control, μRB +9% CS still showed detectable ribbon structures and pores. To help visualize how CS was incorporated into the gelatin μRB composite scaffolds, gelatin μRB and CS were fluorescently labeled separately (red for μRBs, green for CS), and scaffolds were imaged using confocal microscopy. μRBs alone, showed intense red fluorescence of individual μRBs (Fig. 1B). In all composite scaffolds, individual μRBs were red covered by green coating, indicating part of the CS solution was also absorbed by the μRBs (Fig. 1B). Increasing haze of green fluorescence was observed in μRB composite scaffolds containing 3%–9% CS, indicating CS formed nanoporous gels, filling up the macropores of μRB scaffolds. PEG/CS HG control showed uniform green fluorescence, with no noticeable macropores (Fig. 1B). We also included reconstructed 3D movies of the scaffold morphology, which showed consistent results with confocal images (Supplemental videos). Supplementary video related to this article can be found at http:// dx.doi.org/10.1016/j.biomaterials.2019.119579.

3.3. Short term TGF-β exposure was sufficient to support synergistic cartilage formation using μRB/CS composite scaffold Our previous protocol used TGF-β throughout the culture time in vitro, which is costly, and prolonged in vitro culture is not desirable for clinical translation. We then assessed whether short TGF-β exposure (7 days) would be sufficient to support co-culture synergy, comparable to full TGF-β supplementation throughout the 21 days of in vitro culture. We used μRB +6% CS scaffolds containing a mixed population of ADSCs and NChons. While 21-day TGF-β exposure led to about 40% higher cell proliferation than 7-day TGF-β exposure in the co-culture group (Fig. 3B), both groups showed comparable Young's moduli by the end of 21 days (Fig. 3C). The only statically significant difference observed in Young's modulus was from ADSC mono-culture (Fig. 3C). This was expected because it is well established that ADSCs require TGF-β for chondrogenesis. Only 7-days of TGF-β was insufficient to fully induce and maintain chondrogenesis of ADSCs. Additionally, there was no significant difference in cell proliferation between 7 and 21-day TGF-β exposures for ADSC and NChon mono-culture controls (Fig. 3B). Interestingly, NChon mono-culture appeared to perform slightly better 4

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Fig. 1. Characterization of the effects of varying chondroitin sulfate (CS) concentration on morphology, porosity and protein elution profiles from μRB/CS composite scaffolds. Composite scaffold was fabricated by mixing gelatin μRBs with varying concentrations of methacrylated CS hydrogel precursor solution. Nanoporous hydrogel (HG) made of 5% PEGDA and 3% CS was included as a control. (A) Scanning electron microscopy images of all scaffolds; Scale bar = 250 μm; (B) Confocal microscopy images of all scaffolds. Gelatin μRBs were fluorescently labeled red and CS was labeled green; (C) Protein elution profile of BSA from all scaffolds over 24 h (n = 5/group). (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)

exposure was sufficient to induce maximal synergy between ADSCs and NChons for optimal cartilage deposition.

with shorter TGF-β exposure. Compared with 21 days of TGF-β treatment, the group treated with only 7 days of TGF-β led to cartilage with improved mechanical strength (Fig. 3C) and greater amount of cartilage matrix deposition (Fig. 3D and Fig. S3), although the difference is not statistically significant. These results suggest that prolonged exposure of chondrocytes to TGB-β is not necessary and may inhibit hyaline cartilage deposition. This observation is consistent with recent reports that extended exposure to active TGF-β is only found in OA joints, not healthy joints [40,41]. However, due to the lack of significance observed here, future investigations would need to be performed to verify this finding. Immunostaining of cartilage markers further validated 7-day TGF-β was sufficient to induce cartilage deposition by mixed cell populations (Fig. 3D). Co-culture showed the most intense and extensive type II collagen staining compared to mono-culture controls. Regardless of the TGF-β exposure time, both co-culture and ADSC mono-culture showed extensive type I collagen staining, suggesting fibrocartilage phenotype (Fig. 3D). In contrast, minimal type I collagen was expressed by NChon mono-cultures. Together, these results suggest both NChon and ADSCs contribute to type II collagen deposition during co-culture, and type I collagen deposition in the co-culture group was mostly contributed by ADSCs. Total collagen and sGAG deposition also showed comparable levels of total matrix deposition between 7-day and 21-day TGF-β exposures (Fig. S3). Together, our results validate that 7 days of TGF-β

3.4. μRB/CS composite, but not HG, accelerated restoration of cartilagemimicking compressive moduli in vivo While our previous work has demonstrated the ability of PEG/CS nanoporous hydrogels to support catalyzed cartilage formation using mixed ADSC/NChons, one key remaining limitation is that the mechanical strength of the resulting cartilage did not increase over time. This is because hydrogel degradation is necessary to enable new cartilage nodule deposition. Hydrogel degradation leads to a decrease in compressive moduli, which is compensated by the newly deposited cartilage nodules, but the overall compressive moduli stays the same. Our motivation in designing the μRB/CS composite scaffold is to enable faster restoration of cartilage mechanical strength by leveraging the advantages of macroporous μRB scaffolds. To assess whether the optimized μRB/CS composite scaffold (μRB + 6% CS) could support catalyzed cartilage regeneration in vivo with enhanced mechanical properties, we chose a mouse subcutaneous model which has been widely used for initial in vivo screening due to relative low cost and high sample size [21]. All samples were first precultured in vitro to induce co-culture synergy, then implanted subcutaneously in vivo for up to 56 days (Fig. 4A). NChons or ADSCs alone 5

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Fig. 2. Effect of varying scaffold composition on cell proliferation, morphology and cartilage formation in vitro. Mixed populations of 75% adipose-derived stem cells and 25% neonatal chondrocytes were encapsulated for all groups. (A) Live/dead staining at day 1 after encapsulation. Green: live cells, Red: dead cells, scale bars = 250 μm; (B) Quantification of cell proliferation by day 21 (n = 5/group); *p < 0.05 compared to μRB group. (C) Masson's Trichrome staining of all groups at day 21. μRB and μRB/3% CS composite both shrank over time (blue: collagen, black: nuclei, scale bar = 1 mm). (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)

μRB +6% CS groups showed higher cell proliferation than HG control (Fig. S4B). All groups showed an increase in opacity over the time of in vivo culture, suggesting new matrix deposition (Fig. S5). Over 56 days in

encapsulated in μRB +6% CS scaffolds were included as controls. A coculture group was also encapsulated in PEGDA/CS hydrogel (HG), the same composition as our previous publication, as a control. At day 0 of implantation, all groups showed high cell viability (Fig. S4A). All three

Fig. 3. Short TGF-β3 exposure (7 days) was sufficient to induce synergy with enhanced cell proliferation and cartilage regeneration in vitro. Mixed ADSC/NChons (75%:25%) were encapsulated in a μRB composite scaffold containing 6% CS. (A) Experimental design. Samples were cultured in chondrogenic medium supplemented with TGF-β3 for 7 or 21 days; (B) Fold of cell proliferation by day 21; (C) Young's modulus at day 21; (D) Immunostaining for Type II and Type I Collagen. Green: collagen, blue = nuclei, scale bars = 200 μm. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.) 6

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Fig. 4. μRB/CS composite, but not HG, accelerates cartilage regeneration and increase in compressive moduli of engineered cartilage using co-culture in a mouse subcutaneous model. (A) Schematic of experimental design. Mixed co-culture of ADSC/NChons were encapsulated either in μRB +6% CS or nanoporous PEG/CS hydrogels (HG). 100% NChon or 100% ADSCs were also encapsulated in μRB/CS composite as controls. (B) Young's moduli over time (n = 5/group), #: p < 0.05 compared to HG co-culture at day 0, ^:p < 0.05 compared to μRB/CS co-culture at day 21, *p < 0.05 (C, D) Safranin O staining for sGAG and Masson's Trichrome staining for collagen (red: sGAG, blue: collagen); Scale bars = 250 μm. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)

deposition was highest with μRB/CS co-culture, but the extent of deposition was less than that of sGAG. Histology results trends correlated with mechanical testing results (Fig. 4B). μRB/CS co-culture led to the highest Young's modulus and most extensive sGAG and collagen deposition, followed by μRB/CS-NChons. In contrast, HG co-culture showed the least increase in newly deposited cartilage and increase in compressive moduli.

vivo, all three μRB/CS groups showed a steady increase in compressive moduli. The greatest increase was observed in the co-culture group, with a Young's modulus reaching over 330 kPa (Fig. 4B), a 40-fold increase from time of implantation, and reaching ~60% of the value of native cartilage (500–600 kPa) [38,39]. This modulus was even higher than that obtained with 100% NChons (230 kPa). The lowest Young's modulus was observed with μRB/CS ADSCs (33 kPa). Consistent with our previous report, the HG co-culture group showed no significant increase in compressive moduli over 56 days (Fig. 4B). Together, these results validated our hypothesis that the μRB/CS composite scaffold enabled faster restoration of cartilage-mimicking mechanical properties using mixed cell populations. Here, we utilized unconfined compression to determine the Young's modulus of each sample. This is a standard mechanical testing method used for measuring cartilage stiffness and allowed accurate relative comparisons between groups for this study [36]. However, other more precise methods that take into account the hyperelasticity of the scaffold could also be employed to allow comparison to other studies [42,43]. To visualize the intensity and distribution of newly deposited cartilage, safranin O staining (sGAG) and Masson's Trichrome staining (total collagen) were performed. By day 21, μRB/CS co-culture resulted in the most intense sGAG and collagen deposition, distributed in an interconnected manner (Fig. 4C and D). Almost the entire scaffold was remodeled and subsequently filled with sGAG at day 21. Collagen

3.5. Characterization of cartilage phenotype using immunostaining To evaluate the phenotype of newly formed cartilage in vivo, immunostaining was performed for markers of articular cartilage (type II collagen), fibrocartilage (type I collagen), and hypertrophic cartilage (type X collagen). While the μRB/CS co-culture group showed substantially higher amounts of type II collagen deposition than HG coculture at both day 21 and day 56, it also induced more extensive collagen I deposition (Fig. 5A and B). Type I collagen production was highest in ADSC control, and lowest in NChon control using μRB/CS composite scaffolds. These results suggest type I collagen observed in the μRB/CS co-culture group was most likely contributed by ADSCs. Overall, there was minimal type X collagen deposition from all groups. Similar to our observation, previous studies using stem cells alone in subcutaneous models have reported high expression of collagen I and undesirable endochondral ossification [44]. Unlike prior reports using 7

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Fig. 5. Immunostaining of cartilage markers (Type II, I, and X collagens) at day 21 and day 56 in vivo. Green: collagen, blue: nuclei, scale bars = 200 μm. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)

component can be broadly applied to enhance tissue regeneration using other mixed cell populations where paracrine signal exchange drives synergistic interactions and catalyzed tissue formation.

stem cells alone, using mixed cell populations did not lead to bone formation or mineralization, as shown by the negative Alizarin red staining results (Fig. S6). A subcutaneous model was used for this study as it is a widely used model for initial in vivo investigation, and it enables testing larger sample sizes at different points in a cost-effective manner. Now that our results have validated the potential of μRB/CS hydrogels as a promising scaffold for supporting robust cartilage regeneration, they provide solid ground to support future work utilizing a more clinically relevant disease models such as a rat osteochondral defect model [45,46]. While we focus on ADSC/NChon mixed cell populations and cartilage regeneration in this study, the concept of combining a macroporous μRB scaffold with a fast degrading nanoporous hydrogel

4. Conclusions In summary, here we report a novel scaffold composed of macroporous μRBs filled with degradable CS hydrogel, for supporting cartilage regeneration using a mixed population of ADSCs and juvenile chondrocytes. We identified 6% CS as the optimal formulation to support synergistic interactions between ADSCs and NChons while maintaining scaffold size and shape. Furthermore, 7 days of TGF-β3 exposure was sufficient to induce catalyzed cartilage formation. When 8

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implanted in a mouse subcutaneous model, μRB/CS scaffolds containing mixed co-culture enabled a rapid increase in compressive modulus of engineered cartilage reaching ~60% of the modulus of native cartilage. Impressively, the mechanical properties of cartilage engineered using a mixed cell population even surpassed the modulus of cartilage produced using 100% chondrocytes. Together, our results validate a μRB/CS composite scaffold as a promising scaffold for cartilage regeneration using mixed populations of stem cells and chondrocytes. Future studies will further validate the potential of μRB/CS scaffolds for cartilage regeneration using more disease-relevant models such as rat osteochondral defect models [45,46].

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Acknowledgements [20]

The authors acknowledge NIH R01DE024772 (F.Y.), NSF CAREER award CBET-1351289 (F.Y.), California Institute for Regenerative Medicine Tools and Technologies Award RT3-07804 (F.Y.), the Stanford Bio-X Interdisciplinary Initiative Seed grant (F.Y.), the Stanford Child Health Research Institute Faculty Scholar Award (F.Y.), NSF Graduate Research Fellowship Program (H.R.), and the Stanford Interdisciplinary Graduate Fellowship from the Stanford Bio-X program (H. R.) for support.

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Appendix A. Supplementary data Supplementary data to this article can be found online at https:// doi.org/10.1016/j.biomaterials.2019.119579.

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Data availability

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The raw and processed data required to reproduce these findings are available to download from data.mendeley.com.

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