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Cartilage-like protein hydrogels engineered via entanglement

Nature volume 618, pages 740–747 (2023)Cite this article

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Abstract

Load-bearing tissues, such as muscle and cartilage, exhibit high elasticity, high toughness and fast recovery, but have different stiffness (with cartilage being significantly stiffer than muscle)1,2,3,4,5,6,7,8. Muscle achieves its toughness through finely controlled forced domain unfolding–refolding in the muscle protein titin, whereas articular cartilage achieves its high stiffness and toughness through an entangled network comprising collagen and proteoglycans. Advancements in protein mechanics and engineering have made it possible to engineer titin-mimetic elastomeric proteins and soft protein biomaterials thereof to mimic the passive elasticity of muscle9,10,11. However, it is more challenging to engineer highly stiff and tough protein biomaterials to mimic stiff tissues such as cartilage, or develop stiff synthetic matrices for cartilage stem and progenitor cell differentiation12. Here we report the use of chain entanglements to significantly stiffen protein-based hydrogels without compromising their toughness. By introducing chain entanglements13 into the hydrogel network made of folded elastomeric proteins, we are able to engineer highly stiff and tough protein hydrogels, which seamlessly combine mutually incompatible mechanical properties, including high stiffness, high toughness, fast recovery and ultrahigh compressive strength, effectively converting soft protein biomaterials into stiff and tough materials exhibiting mechanical properties close to those of cartilage. Our study provides a general route towards engineering protein-based, stiff and tough biomaterials, which will find applications in biomedical engineering, such as osteochondral defect repair, and material sciences and engineering.

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Fig. 1: Schematics of the chain-entangled network structure of the N-DC (FL)8 hydrogel.
Fig. 2: Chain entanglements enhance the stiffness of (FL)8 hydrogels.
Fig. 3: N-DC (FL)8 hydrogels showed superb compressive mechanical properties.
Fig. 4: Stiff hydrogel group shows good repair efficacy of subchondral bone and cartilage at 12 weeks post-implantation.
Fig. 5: Histology analysis shows significantly improved subchondral bone and cartilage repair in the stiff hydrogel group at 12 weeks post-implantation.

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Acknowledgements

This work is supported by the Natural Sciences and Engineering Research Council of Canada (H.L.), the National Key R&D Program of China (grant no. 2020YFA0908100) (Y.C.), the National Natural Science Foundation of China (81991514 (Q.J.) and 32271408 (L.L.)), and the Natural Science Foundation of Jiangsu Province (BK202200121 (L.L.) and BK20202013 (Q.J.)). We thank R. Shadwick for assistance with Instron tests.

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  1. These authors contributed equally: Linglan Fu, Lan Li

Authors and Affiliations

  1. Department of Chemistry, The University of British Columbia, Vancouver, British Columbia, Canada

    Linglan Fu, Qingyuan Bian, Jiayu Li & Hongbin Li

  2. State Key Laboratory of Pharmaceutical Biotechnology, Division of Sports Medicine and Adult Reconstructive Surgery, Department of Orthopedic Surgery, Branch of National Clinical Research Center for Orthopedics, Drum Tower Hospital affiliated to Medical School of Nanjing University, Nanjing, People’s Republic of China

    Lan Li, Jing Jin & Qing Jiang

  3. National Laboratory of Solid State Microstructures, Department of Physics, Nanjing University, Nanjing, People’s Republic of China

    Bin Xue & Yi Cao

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Contributions

H.L. conceived the project. L.F. conceptualized the DC hydrogelation strategy and performed experiments on the engineering and characterization of hydrogels and analysed the data. Q.B. and B.X. performed some of the engineering and characterization work on hydrogels and prepared hydrogel scaffolds for in vitro and in vivo experiments. J.L. performed part of the protein engineering work. L.L., Y.C., Q.J. and H.L. designed the in vivo experiments. L.L. and J.J. carried out the in vitro and in vivo experiments. L.F., L.L. and H.L. wrote the initial manuscript. All authors discussed the results and contributed to the final manuscript.

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Correspondence to Hongbin Li.

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Competing interests

L.F. and H.L. are co-inventors on a provisional patent application (application number PCT/CA2021/051622, submitted by The University of British Columbia) covering the N-DC hydrogelation method described in this article. The other authors declare no competing interests.

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Extended data figures and tables

Extended Data Fig. 1 Viscosity of native and denatured (FL)8 protein solutions in PBS and 7M GdHCl.

Denatured (FL)8 (red triangles) displays higher viscosity than native (FL)8 (blue diamonds).

Extended Data Fig. 2 Schematics of the NC-(FL)8 (native crosslinked) hydrogels.

The elastomeric protein (FL)8 was first dissolved in PBS to a high concentration (~200 mg/mL) (a). Upon photo-crosslinking, (FL)8 were crosslinked into a hydrogel network lacking any chain entanglements, owing to the short length of folded (FL)8, resulting in the N-NC (native NC) hydrogel (b). When denatured in GdHCl, the (FL)8 in the hydrogel network unfold, and behave as random coils. The resultant D-NC (denatured NC) hydrogel also remains free of chain entanglements (c).

Extended Data Fig. 3 Photographs of scanning electron microscopy imaging of N-DC and N-NC (FL)8 hydrogels.

Both hydrogels have porous network structures; however, the pore size of the N-DC hydrogel (~2 μm) is significantly smaller than that of the N-NC hydrogel (~20 μm).

Extended Data Fig. 4 Mechanical properties of N-DC and D-DC (FL)8 hydrogels in tensile testing.

(a) Young’s modulus and breaking strain of 20% N-DC and D-DC (FL)8 hydrogels. (b) Toughness and swelling ratio of 20% N-DC and D-DC (FL)8 hydrogels. It is evident that the N-DC (FL)8 hydrogels exhibited much higher Young’s modulus and higher toughness than the N-NC hydrogel. (c-d) Young’s modulus/breaking strain (c) and toughness/swelling ratio (d) of N-DC (FL)8 hydrogels at different protein concentrations. (e-f) Young’s modulus/breaking strain (e) and toughness/swelling ratio (f) of 20% N-DC hydrogels based on (FL)4, (FL)8 and (FL)16, which consist of 4, 8 and 16 tandem repeats of FL. 20% N-DC (FL)4 hydrogels showed similar mechanical properties as those of (FL)8 and (FL)16, suggesting that the length of unfolded (FL)4 is sufficient for chain entanglements. The data is presented as average ± standard deviation (ave. ± S.D.). The number of events n ranges from 3–5 for different experiments, with the exception that n = 18 for the Young’s modulus of 20% (FL)8 N-DC hydrogel.

Extended Data Fig. 5 Compression properties of N-DC (FL)8 hydrogels.

(a) Stress-strain curves of a N-DC (FL)8 hydrogel compressed to failure. Inset shows the photographs of the hydrogel right after failure (1st cycle) and after three more consecutive compression-unloading cycles (4th cycle). Cracks were observed right after failure. Subsequent compression led to propagation of the crack, but the failure is not brittle. (b) Photographs of the N-DC (FL)8 hydrogel during various stages of compression. After unloading, the hydrogel recovered its shape rapidly. (c) Schematics of the hydrogel network structure during compression-unloading. The structure is drawn in a highly schematic manner, and for illustration purpose only. Red circle indicates the chain entanglement, red arrows point to two folded FL domains that may undergo forced-unfolding at high strain, and then refolding after unloading. When compressed to 50% strain and then unloaded, the hydrogel can quickly recover. When compressed to 80% strain, a small number of FL domains unfold. Upon unloading, a large fraction of unfolded FL domains refolds quickly, while a small fraction refolds following a slower kinetics. For simplicity, hydrophobically collapsed FL aggregates are not shown. The force-induced unfolding of the folded FL domains help dissipate energy during loading, and upon unloading, the refolding of FL domains help the hydrogel to regain its shape and mechanical properties. (d—e) Consecutive compression-unloading curves of a N-DC (FL)8 hydrogel at a frequency of 0.08 Hz (d) and 0.67 Hz (e). The loading rate was 20 mm/min in (d) and 200 mm/min in (e). The N-DC hydrogels showed limited fatigue (insets). The percentage is the peak stress of each cycle compared to the initial stress when loaded to strain of 60%.

Extended Data Fig. 6 N-DC Fn-(Cys-FL)4-RGD-(Cys-FL)4-RDG (FLRGD) hydrogels support cell adhesion, spreading, proliferation and differentiation of mouse osteoprogenitor MC3T3-E1 cells.

(a) Epi-fluorescence images of stained cells on N-DC FLRGD hydrogels (On gel) and cell-culture dish (Control). Cells were simultaneously stained with Calcein AM dye and propidium iodide (PI) for staining live and dead cells, respectively. (b) Cell viability determined by live/dead cell staining on N-DC FLRGD hydrogels and cell-culture dish (Control). Error bars correspond to the standard deviation. (c–d) Immunostaining of MC3T3-E1 cells after cultured for 5 days (c) and 10 days (d) in blank, N-DC FLRGD hydrogel and PS culturing dish control groups. The Col I and Runx2 are specific markers for the differentiation in the direction of osteoblast. Clearly, Col I and Runx2 were up-regulated in MC3T3-E1 cells cultured on N-DC-FLRGD hydrogel (Gel) and on coated cell culture dish (control) groups, while Col I and Runx2 were hardly observed in blank (uncoated cell culture dish) groups. (e and f) Expression of osteoblast differentiation-related genes (Col I, RUNX2, ALP and OPN) for MC3T3-E1 cells after cultured for 5 and 10 days. The data represent the relative mRNA levels of the target genes normalized to the levels of the reference genes and are expressed with the levels in the blank group as 1. The values represent the mean ± s.d. Differences between groups were assessed by student’s t-test; NS: p > 0.05; *: p < 0.05; **: p < 0.01; ***: p < 0.001. Cells were cultured on uncoated cell culture dish (Blank), N-DC-FLRGD hydrogel (Gel), and coated cell culture dish (Control).

Extended Data Fig. 7 Tissue remodelling and regeneration process at different time points.

(a) Imaging results and histological staining at week 4 and week 8. Scale bars: 100 μm. Arrows indicate the remaining hydrogel implants after 4 weeks of implantation. After 8 weeks of implantation, all hydrogel implants were degraded. (b) The quantitative analysis at week 4 and week 8, including Micro CT analysis, ICRS scoring system, and O’Driscoll scoring system. Data are presented as mean ± s.d. (n = 6). No obvious difference was observed at week 4 among the three groups. However, the subchondral bone and cartilage in stiff hydrogel treated group exhibited better regeneration phenomenon at week 8. The statistical difference was observed in trabecular parameters and O’Driscoll scoring system. *: p < 0.05; **: p < 0.01; ***: p < 0.001; N.S.: not significant.

Extended Data Fig. 8 Immunological response at different time points evaluated using histology staining.

(a) Histological analysis (H&E staining) of major organs and skins at time point of 48 h and 7 days. Scale bars: 100 μm. (b) Liver function test results at two time points. Data are presented as mean ± s.d. (n = 3). (c) Histological analysis (H&E staining) of the major organs from the New Zealand rabbits after implantation for 12 weeks. Scale bars: 100 μm.

Extended Data Fig. 9 The repairing effect of highly stiff (FL-M23C)8-Fn hydrogel at three time points.

(a) General view of the cartilage and subchondral defects regions at three time points post implantation of the highly stiff hydrogel. Irregular tissue remained in the defect region after 12 weeks implantation. Micro MRI analysis (b) and Micro CT analysis (c) indicated that the regeneration of cartilage regeneration was scarce, and the remolding of subchondral bone was also limited. (d) Histological analysis (H&E staining) of cartilage and subchondral bone regeneration. The hydrogel implants (indicated by the arrows) can be clearly observed even after implantation for 12 weeks. Scale bars: 100 μm.

Extended Data Table 1 Enhancement of mechanical properties via the DC hydrogelation method (E: tensile modulus; Y: compressive modulus)
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Supplementary Video 1

Consecutive compression–unloading cycles of the N-DC (FL)8 hydrogel. The vdeo shows the deformation and recovery of the N-DC (FL)8 hydrogel during a consecutive loading–unloading test. The final strain is 65%.

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Fu, L., Li, L., Bian, Q. et al. Cartilage-like protein hydrogels engineered via entanglement. Nature 618, 740–747 (2023). https://doi.org/10.1038/s41586-023-06037-0

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