Multiscale mechanical-adapted hydrogels for the repair of intervertebral disc degeneration

Abstract

Nucleus pulposus (NP) tissue engineering brings new hope in the repair of intervertebral disc degeneration (IVDD). IVDD is often accompanied by multiscale changes in the mechanical microenvironment, including the changes of mechanical property of collagen fibril, NP tissue, and mechanical instability of spine. In this study, a multiscale mechanically-adapted strategy is proposed to promote NP repair. To achieve this goal, a viscoelastic-adapted dual-network hydrogel (PVA-DN) is constructed. The hydrogel with multiscale tunable viscoelasticity and dynamic compression condition is used to meet the multiscale mechanical requirements of NP regeneration. The results show that the viscoelastic hydrogel promotes the proliferation, migration and adhesion of nucleus pulposus cell (NPC) as well as the secretion of NP-specific extracellular matrix. RNA-seq results show that it attenuates the inflammatory microenvironment by inhibiting the IL-17 signaling pathway. Appropriate dynamic compression applied to the viscoelastic scaffold further promotes the physiological function of NPC, and the viscoelasticity of hydrogel protects against NPC damage induced by excessive compression. Animal experiments demonstrate that the viscoelastic hydrogel effectively restores disc mechanical function and delays disc degeneration in rats. Findings from this study highlight that the multiscale mechanically-adapted strategy is effective in the repair of IVDD.

Graphical abstract

Highlights

• •A multiscale mechanically-adapted hydrogel is constructed for the repair of IVDD.
• •Viscoelastic hydrogel with fast stress relaxation times at multiscale can attenuate the inflammatory microenvironment by blocking the IL-17 signaling pathway.
• •Viscoelastic hydrogel with fast stress relaxation times at multiscale promotes the formation of stress fibers, earlier adhesion and ECM secretion of NPCs.

1. Introduction

Low back pain (LBP) is a very common musculoskeletal disorder that places a heavy burden on the quality of life of patients and economic development of the society [1,2]. Intervertebral disc degeneration (IVDD) is the predominant cause of LBP in humans and the degree of degeneration is closely correlated with the level of pain. Although the pathological mechanism is not fully understood, the reduction in nucleus pulposus cell (NPC) number, the disruption of extracellular matrix (ECM), and changes in the local biomechanical microenvironment are involved in its progression [3,4]. Therefore, the repair of degenerated NP is the key treatment to solve this common health problem. Current clinical strategies for IVDD including non-surgical treatments such as medication combined with physiotherapy, and surgical treatments that decompress the nerve roots (discectomy) and remove the degenerated discs (disc replacement and fusion), all focused on symptom relief and functional improvement. However, the long-term effects are limited because the progression of degeneration fails to be reversed, and it has high recurrence rate after surgery [5,6]. Indeed, some patients experience recurrence or deterioration since most surgeries alter the original structural integrity, tissue homeostasis, or the mechanical properties of the spine [7]. Tissue engineering has provided a new approach for NP repair in recent years [8].

IVDD is often accompanied by multiscale changes. This disease originates at the molecular scale and progresses to higher cellular and organizational scales, ultimately leading to impaired organ function. At the macroscopic level, nucleus pulposus (NP) is a highly hydrated tissue located at the center of IVD, maintaining the mechanical stability of the disc [9]. Notochordal cells resident in the NP region are replaced by chondrocyte-like cells during IVDD and the proteoglycans are progressively lost, resulting in a decrease in NP water content and mechanical instability [[10], [11], [12]]. In turn, altered biomechanics lead to a variety of spinal disorders like disc herniation and spinal stenosis [13]. At the microscopic level, NPCs reside in the ECM consisting of a network of collagen and glycosaminoglycan that is constantly remodeled in both normal and pathological conditions [14]. The degenerative process leads to a loss of proteoglycans and disorganization of ECM architecture, resulting in an increase in stiffness of NP tissue [15,16]. NPCs could respond to the change in the stiffness or viscoelasticity of the ECM by altering the cytoskeleton, leading to a range of changes in biological functions, such as cell spreading, motility, and proliferation [[17], [18], [19]]. At the nanoscopic level, the collagen fibril become thicker and stiffer during IVDD. This mechanical signal can also be transduced into cells by cell surface receptors, regulating several cellular functions vital for maintaining normal homeostasis [14]. Mechanical properties at the micro- and nano-scale affect the mechanical microenvironment around the NPC, consequently affecting the physiological function. Therefore, multiscale mechanical properties are critical for both cellular function and disc function (see Scheme 1).

Scheme 1.

A multiscale mechanical-adapted hydrogel for the repair of IVDD. At the nanoscopic level, the fast stress relaxation hydrogel modulates the distribution of the integrin-binding domain/actin-binding region of vinculin, promotes earlier adherence of NPCs and has more pseudopods that can grip more firmly. At the microscopic level, the fast stress relaxation hydrogel further modulates cytoskeletal rearrangement, promotes NPC spreading, attenuates the inflammatory microenvironment and inhibits the degradation of ECM by blocking the IL-17 signaling pathway. Ultimately, the fast stress relaxation hydrogel contributes to NP repair and restoration of mechanical function of the IVD after implantation at the macroscopic level.

Although some mechanical properties of NP correlate with the degree of degeneration, research on the multiscale mechanical properties under dynamic conditions is still scarce. Therefore, it is necessary to further understand the biomechanics of tissues at the molecular level (nanoscopic mechanical properties), cellular level (microscopic mechanical properties) and organ level (macroscopic static and dynamic mechanical properties) and the corresponding mechanobiological process. Currently, various hydrogels for tissue engineering are being developed to reconstruct a healthy mechanical microenvironment and promote the repair of NP [20,21]. Among them, the viscoelasticity of hydrogel is increasingly studied because most matrices surrounding NPCs are viscoelastic materials which exhibit both solid- and liquid-like mechanical behavior. It has been reported that hydrogels with NP-like viscoelasticity could maintain the viability, or promote the ECM secretion of stem cells, are proved to be effective on IVD repair [22,23]. Long-term cell survival and NP-specific ECM secretion of adipose-derived stem cells are significantly influenced by matrix viscoelasticity, where the deposition of aggrecan and type-II collagen is significantly enhanced in the fast-relaxing hydrogel [23]. In addition, the microenvironment of the IVD is not only influenced by the stiffness or the viscoelasticity of matrix, but it is also regulated by external mechanical stimuli [22,24,25]. The IVD is located in a unique microenvironment that is biomechanically constrained, and it is a load-bearing organ and its metabolic and cellular activity is closely associated to the biomechanical microenvironment [26]. However, most current studies are based on the macro-mechanical properties of substrates, while there is a lack of research on the design of the mechanical properties of materials at the micro- and nano-scale level, and it is still unclear how dynamic loading acts synergistically with multiscale mechanical properties, particularly viscoelastic properties.

IVDD can be accelerated by the release of inflammatory mediators, such as IL‐1α, IL‐1β, IL‐6, IL‐8, IL‐17 and TNF‐α, and the recruitment of immune cells for infiltration [[27], [28], [29]]. Activation of the MAPK/ERK signaling pathway and the nuclear factor-κB (NF-κB) pathway leads to an imbalance in ECM metabolism, cell loss, increased cellular senescence and oxidative stress in NPCs, and inhibition of NPC proliferation, resulting in IVDD. Activation of these signaling pathways can trigger the release of pro-inflammatory cytokines from IVD cells, such as IL-1β and TNF-α [30]. IL-1β and TNF-α were found to be significantly upregulated in the NP of patients with IVDD. And TNF-α promotes the progression of IVDD and induces pain exacerbation, and its concentration is positively correlated with the degree of IVDD [31]. In addition, NF-κB can be silenced in the IVD by injecting specific siRNA in the NP degeneration model, thereby reducing injury-induced NP degeneration [32].

In this study, a multiscale mechanically-adapted strategy was proposed to promote NP repair. For this purpose, a viscoelastic dual-network hydrogel with multiscale tunable mechanical properties was constructed. The mechanical properties of the viscoelastic hydrogel were investigated to adapt to the mechanical properties of the NP tissue and to allow proliferation, adhesion and migration of NPCs on the surface of the hydrogels. Next, dynamic compressive strain was applied to the cell-loaded viscoelastic scaffolds to assess the expression of NP-specific metabolic markers and to investigate the potential mechanisms of how the viscoelastic hydrogel affects ECM secretion and regeneration in 3D culture. Finally, the viscoelastic hydrogels were implanted into the rat tail disc and the regeneration of the degenerated disc was evaluated. This study might provide a new strategy to repair disc degeneration.

2. Results

2.1. Preparation and characterization of the viscoelastic dual-network hydrogel

The preparation of methacryloylated polyglutamic acid grafted with phenylboronic acid (PBA-m-PGA) and gelatin methacryloyl (GelMA) was cross-linked with each other to form a dual-network structure, and the introduction of polyvinyl alcohol (PVA) formed a boric acid bond with PBA-m-PGA, which was used as a bridge molecule to connect the PVA chains into the dual network, finally cross-linking to form a hydrogel under UV light (Fig. 1A). The spectrum of the proton nuclear magnetic resonance (¹H NMR) of m-PGA showed two new peaks at 5.66 and 6.09 ppm, corresponding to the chemical shifts of protons of the vinyl groups (C=C), while 4.13–4.41 ppm was the chemical shift of methine protons in γ-PGA. The proton peak at a chemical shift of 7.61 ppm in PBA-m-PGA indicated the successful grafting of the phenylboronic acid moiety (Fig. 1B). The formation of m-PGA and PBA-m-PGA was further confirmed by Fourier transform infrared spectrometer (FTIR). FTIR spectra showed that the relative intensity of the C-O stretching vibrational peak at 1165 cm⁻¹ became stronger in m-PGA as compared to PGA molecules, while the spectra of the PBA-m-PGA molecules showed a characteristic peak at 1687 cm⁻¹ of the C=C structure of the benzene ring, suggesting the successful grafting of the phenylboronic acid. Moreover, the O-H absorption peak near 3269 cm⁻¹ significantly increased after the addition of PVA, indicating the formation of a stronger borate dynamic covalent bond in the hydrogel (Fig. 1C).

Fig. 1.

Preparation and characterization of the viscoelastic dual-network (PVA-DN) hydrogel. (A) Schematic diagram showing the structure and cross-linking mechanism of the hydrogels. (B) ¹H-NMR spectra. (C) FTIR spectra. (D) Swelling properties of the different PVA-DNs. (E) SEM images of the different PVA-DNs. (F) Rheological analysis of the different PVA-DNs. (G) Live-dead staining images of NPCs on the hydrogel. Green: live cells; red: dead cells. (H) CCK-8 assay at 1, 3, 5 and 7 days. ∗, P < 0.05; ∗∗, P < 0.01; ∗∗∗, P < 0.001; ∗∗∗∗, P < 0.0001.

The viscoelastic dual-network (PVA-DN) hydrogels was prepared. The injection property and the self-healing visualization property of the hydrogels were shown in Fig. S1. The swelling capacity of hydrogels was assessed by calculating the mass swelling ratio. All of them almost reached the swelling equilibrium at 3 h, with the highest swelling rate in the 1 % PVA-DN hydrogel with increasing time, which could be related to the density and structural characteristic of the hydrogels (Fig. 1D). The SEM images showed that all groups had an interconnected porous network structure, which was conducive to internal nutrient transport and 3D cell culture. The 1 % PVA-DN hydrogel had the largest pore size, while the 5 % PVA-DN hydrogel had the smallest pore size (Fig. 1E). Degradation experiments revealed that 5 % PVA-DN hydrogel degraded only approximately 30 % after 14 days, whereas the other two hydrogels degraded at a slower rate (Fig. S2A). Importantly, the mechanical properties (elastic modulus and stress relaxation time) of all hydrogels were essentially stable for 14 days under tissue culture conditions (Figs. S2B and C). The dynamic covalent borate bonds possessed a viscoelastic behavior as demonstrated by rheological experiments in the frequency range from 0.1 to 10 Hz that revealed a gradual decrease in the loss factor (G′′/G′) with increasing frequency. The largest loss factor occurred in the 5 % PVA-DN hydrogel, further confirming its viscoelastic properties (Fig. 1F). The cytocompatibility of the viscoelastic hydrogel was further assessed by live/dead staining after seeding NPCs on hydrogels for 48 h. The results showed that cells survived when cultured on the surface of all hydrogels, with an almost negligible number of dead cells, indicating that the viscoelastic hydrogels were not cytotoxic when the PVA concentration did not exceed 5 wt% (Fig. 1G). In addition, cell proliferation increased with the increase of PVA concentration (Fig. 1H).

2.2. Multiscale mechanical modulation adapts to the microenvironment of NP tissue

At the macroscopic level, the dual-network hydrogels are mechanically adapted to NP tissue by adjusting the concentration of GelMA, PBA-m-PGA, and PVA (Fig. 2A). The elastic modulus, representing the stiffness of the material, was determined by fitting linearly the stress-strain curve from the compression test. In this study, three different relative concentrations of dual-network hydrogels were obtained (1 % PVA-DN - “slow stress relaxation”, 3 % PVA-DN - “medium stress relaxation” and 5 % PVA-DN - “fast stress relaxation”) by adjusting the PVA cross-linking density. The initial elastic modulus was determined at 8 kPa between the different groups, which was not statistically significantly different (Fig. 2A1). Furthermore, all hydrogels resisted compression to at least 70 % strain before fracture, this highlighting their high ductility (Fig. S3A). Viscoelasticity, on the other hand, was regulated by the density of borate bonds formed by the coupling of PVA to the PBA-m-PGA chains. τ_1/2 is defined as the time taken for the stress on hydrogel dropping to half of its initial value, which ranged from 2711 ± 496.4 s in slow stress relaxation hydrogels, to 1274 ± 218.4 s in medium stress relaxation hydrogels and 444 ± 62.6 s in fast stress relaxation hydrogels, as revealed by stress relaxation tests in compression (Fig. 2A2, S3B). Among them, the elastic moduli of the hydrogels were similar to fresh NP at macro-scale, and the relaxation time of the fast stress relaxation hydrogel was closest to those of fresh NP tissue, although there were still differences between them. The increase in proliferation was attributed to an isolated change in stress relaxation, since the initial elastic modulus was constant, due to the very low cell adhesion of the PVA hydrogel [33].

Fig. 2.

Multiscale mechanical adaptation of hydrogels to NP tissue. (A) Macro-mechanical adaptation: (A1) Statistical results of the elastic modulus at the macro-scale, (A2) Statistical results of the relaxation times at the macro-scale. (B) Dynamic mechanical adaptation: Stress–strain curves of (B1) fast stress relaxation hydrogels and (B2) NP tissue at different compression cycles up to 100 cycles. Tests were performed at a strain of 20 % applied at a displacement rate of 1 mm/min. (C) Micro-mechanical adaptation: (C1) Statistical results of the elastic modulus at the micro-scale, (C2) Relaxation times obtained at a surface residence time of 4 s at the micro-scale. (D) Nano-mechanical adaptation: (D1) Statistical results of the elastic modulus at the nano-scale, (D2) Relaxation times obtained at a surface residence time of 4 s at the nano-scale. ∗, P < 0.05; ∗∗, P < 0.01; ∗∗∗, P < 0.001; ∗∗∗∗, P < 0.0001.

The mechanical properties of hydrogels under dynamic compression were evaluated by cyclic compression tests to achieve the macro-dynamical adaptation to the NP (Fig. 2B). Faster rates of stress relaxation corresponded to greater rates of viscous energy dissipation in the material per unit time. The mechanical properties of the hydrogel under dynamic compression were evaluated by cyclic compression tests. The energy dissipation of the hydrogel was observed during continuous cyclic compression for up to 100 cycles, with the dispersion of the stress-strain curves being most pronounced in the fast stress relaxation hydrogel (Fig. 2B1,S3D and S3E). For fast stress relaxation hydrogel, the area enclosed by the stress-strain curve decreases most rapidly, indicating the strongest capacity of energy absorption (Table S1). Among them, the fast stress relaxation hydrogel had the closest energy absorption capacity to NP tissue (Fig. 2B2). More importantly, all hydrogels fully recovered from the cyclic compression test, demonstrating the good self-recovery property of the material, ensuring the maintenance of its properties and providing a basis for a subsequent mechanical dynamic loading (Fig. S3C). The relatively weak cross-linking of dynamic covalent bonds provided an effective energy dissipation mechanism, resulting in the formation of hydrogels with good capacity of energy absorption and self-recovery properties.

The mechanical property of hydrogels was determined by atomic force microscopy (AFM) to achieve the micro- and nano-mechanical adaptation to the NP (Fig. 2C and D). AFM uses a tip mounted at the end of a microcantilever to indent the hydrogels or cells, allowing an active characterization of their mechanical properties at the micro- and nano-scale, including elasticity and viscoelasticity [34,35]. Force curve during the approach-reside-retract movement of the probe reflected the contact of the tip with the hydrogel or tissue (Figs. S3F and G). The elastic modulus measured at the micro-scale was the same as that measured at the macro-scale, with no significant difference among them (Fig. 2C1). The elastic moduli of slow, medium and fast stress relaxation hydrogels at the nano-scale were 22.6, 45.3 and 73.6 kPa, respectively, which were not as same as those at the micro-scale (Fig. 2D1). The stiffness of hydrogels at the micro- and nano-scale was closely related to the density of their internal scaffolds with dual network structure. The oscilloscope recorded the deflection signal and height signal during the approach-reside-retract movement of the AFM probe, representing a typical stress relaxation curve (red) and a corresponding probe height change curve (grey) (Fig. S3H). Stress relaxation curves were also obtained on the substrate as contrast. The force remains constant, indicating that no relaxation appeared when performing the movement of the AFM probe on the substrate (Fig. S3I). The stress relaxation of hydrogels at the micro-scale was quite different from that at the macro-scale, although showing the same trend. The fast stress relaxation group had the fastest stress relaxation time of approximately 4 s (Fig. 2C2). On the other hand, the stress relaxation of hydrogels at the nano-scale showed the same trend as that at the micro-scale (Fig. 2D2). In particular, the elastic moduli of the fast stress relaxation hydrogel were similar to fresh NP at micro- and nano-scale, and the relaxation times of the fast stress relaxation hydrogel were closest to those of fresh NP tissue at different scales, although there were still differences between them.

2.3. Cell adhesion, spreading and migration on the viscoelastic hydrogel

Cell adhesion leads to changes in cell morphology and skeleton, which is the first stage of cell-ECM interaction. In turn, the cytoskeleton is remodeled mainly by F-actin polymerization and depolymerization, which are important for cell motility, mechanical properties and function. Cytoskeletal F-actin was stained with Alexan Fluor 488-labeled rhodamine-phalloidin to reveal the effect of the viscoelastic substrate on the arrangement of intracellular stress fibers. At the early stage of adhesion after seeding for 2 h, NPCs cultured on the slow and medium stress relaxation hydrogel substrates showed no formation of stress fibers in the cytoplasm, but only punctate distributions of F-actin, whereas NPCs cultured on fast stress relaxation hydrogel substrate showed more short bundles of stress fibers, indicating a higher aggregation and turnover of F-actin in the cytoplasm. Notably, at the late stage of adhesion after seeding for 8 h, stress fibers were formed in NPCs of all groups, with a diffuse distribution of most of the cellular stress fibers on fast stress relaxation hydrogel substrate, whereas the stress fibers on slow stress relaxation substrate were mostly polar. These results suggested that fast stress relaxation substrate might be more favorable to the formation to intracellular stress fibers, as well as to the polymerization and turnover of the microfilament backbone during adhesion (Fig. 3A). The nascent stress fibers were more distributed at the membrane edge when NPCs were cultured for 48 h. However, the cytoskeletal fibers were rearranged and the original network was gradually transformed into a parallel structure under increasing ECM stress relaxation time, with increased density of stress fibers, while the spreading area of the cells was gradually reduced (Fig. 3C). This suggested that different viscoelastic hydrogels had significant effects on cytoskeletal assembly, and that NPCs sensed the different mechanical microenvironments and regulated their polarization and spreading according to the viscoelasticity of the substrate. AFM was performed on cells at the micro-scale to better understand these variations. The 2D and 3D plots showed that NPCs were flat on the fast stress relaxation hydrogel substrate, but prominent on the slow stress relaxation substrate (Fig. 3B). The height and stiffness of NPCs increased as their spreading morphology changed. NPCs growing on the fast stress relaxation substrate had more and longer filopodia and lamellipodia at their edges (Fig. 3D). The formation of these pseudopods facilitated the development of the migratory phenotype and greatly increased the motility of NPCs. Vinculin protein normally accumulates at focal adhesions and participates in the cytoplasmic adaptation through integrin-binding domains/actin-binding regions, linking the cytoskeleton to the ECM [36,37]. Talin works with vinculin to attach cells and adhere to viscoelastic substrates. The integrin binding domain/actin-binding region of talin and vinculin was more widely distributed and larger in the fast stress relaxation hydrogel, as shown by immunofluorescence results (Fig. 3E). Cell migration is a highly complex multi-step process requiring the formation of adhesion patches followed by the formation of cell membrane protrusions, including filopodia and lamellipodia, closely linked to dynamic F-actin assembly and depolymerization [38]. No significant migration of NPCs on the slow stress relaxation substrate was observed at 24 h after the removal of the scratch inserts. Scratches partially healed on the medium stress relaxation substrate, and they healed most significantly on the fast stress relaxation substrate, with the highest rate of cell migration (Fig. 3F and G).

Fig. 3.

The effect of hydrogel viscoelasticity on cell spreading, adhesion and migration. (A) Cytoskeletal F-actin labeled with Actin-Tracker Green-488 staining. Fluorescent images were taken at 2, 8 and 24 h. (B) The contours, the sub-microscopic structure of membrane protrusions and the 3D reconstruction of living cells using AFM. (C) Average spreading area of NPCs grown on hydrogels. (D) Quantification of the elastic modulus, the height, number of filopodia and lamellipodia in the NPCs by AFM. (E) Immunofluorescence images of talin and vinculin after 48 h. (F) Scratch test at 0 and 24 h. (F) Quantification of relative migration rates. ∗, P < 0.05; ∗∗, P < 0.01; ∗∗∗, P < 0.001; ∗∗∗∗, P < 0.0001.

2.4. Effects of multiscale mechanical adapted hydrogels on ECM anabolism/catabolism in NPCs under static condition

Type II collagen and aggrecan are the main components of ECM in NP tissues and are associated with the physiological functions of normal IVD. Therefore, the biological function of the NPCs in the hydrogels was investigated after 7 days of static 3D culture using real-time quantitative polymerase chain reaction and Western blot analysis. The fast stress relaxation hydrogel promoted the expression of the NP-specific ECM anabolism-related genes Acan and Coll2a1, and significantly suppressed the expression of ECM catabolism-related genes Mmp13 and Mmp3 (Fig. 4A). The results on the expression of the associated proteins were similar to those of gene expression (Fig. 4B and C). The ratio of aggrecan to collagen in the tissues grown in the proposed hydrogel was found to be about 15:1, more like the IVD tissues (Fig. S4). The NP-specific ECM markers aggrecan and collagen II had relatively high expression in the fast stress relaxation hydrogel, suggesting that viscoelasticity promoted ECM secretion and synthesis.

Fig. 4.

Effects of multiscale mechanical adapted hydrogels on ECM anabolism/catabolism in NPCs under static and dynamic conditions. The expression of ECM anabolism- and catabolism-related (A) genes and (B) proteins in NPCs cultured in static 3D hydrogels with different viscoelasticity for 7 days. (C) Quantification of the grey bands of (B). The expression of ECM anabolism- and catabolism-related (D) genes and (E) proteins in NPCs cultured in the different viscoelastic hydrogels after dynamic compression of 5 % or 20 % for 3 days, respectively. (F) Quantification of the grey bands of (E). ∗, P < 0.05; ∗∗, P < 0.01; ∗∗∗, P < 0.001; ∗∗∗∗, P < 0.0001.

2.5. Effects of multiscale mechanical adapted hydrogels on ECM anabolism/catabolism in NPCs under dynamic condition

The imbalance in the anabolism/catabolism metabolism of ECM in NP is closely related to the local mechanical microenvironment that initiates and promotes IVDD. The effect of dynamic compression at different amplitudes on the biosynthesis of NPCs in the 3D culture system was investigated setting the compression amplitude to 0 (static), 5 % (low strain) and 20 % (high strain) at a frequency of 1 Hz. ECM anabolism-related genes (Acan and Coll2a1) were significantly up-regulated after 3 days of loading at 5 % strain. However, the expression of ECM anabolism-related genes showed a significant downward trend as the strain increased to 20 % (Fig. 4D). The results on the expression of the associated proteins were similar to those of the gene expression. The highest expression of ECM-related molecules in NPCs in the 3D culture system was observed after applying low strain stress to fast stress relaxation hydrogel, while their expression was reduced when subjected to high compressive stress loading, suggesting that low strain stress synergistically acted with the fast stress relaxation hydrogel to better promote the secretion of ECM. More interestingly, no significant difference was observed in the expression of ECM between high strain stress and static combination of fast stress relaxation hydrogels, suggesting that viscoelastic hydrogel protected NPCs under high stress (Fig. 4E and F).

2.6. Regulation mechanism of the multiscale mechanical adapted hydrogels on NPCs

RNA-seq was used to investigate the differentially expressed genes in NPCs cultured in the slow and fast stress relaxation hydrogels for 3 days to understand the underlying molecular mechanisms of how the viscoelastic hydrogel affects ECM secretion and regeneration. The volcano plot showed that significant differences in genes expressions between the fast stress relaxation and the slow stress relaxation group (Fig. 5A). The heat map also showed that the expression of the inflammatory genes IL6, IL17 and TNF and the ECM catabolism-related genes Mmp3, Mmp9, Mmp10 and Mmp13 in NPCs in the fast stress relaxation group were significantly downregulated (Fig. 5B). GO analysis showed that these differences were mainly related to catabolism/anabolism of the ECM and a negative regulation of inflammatory response with the fast stress relaxation group was observed (Fig. 5C). In addition, Gene Set Enrichment Analysis (GSEA) demonstrated that “cytokine activity” and “collagen catabolic process” were significantly inhibited in the fast stress relaxation hydrogel (Fig. 5D). Kyoto Encyclopedia of Genes and Genomes (KEGG) pathway enrichment analysis was performed to identify potential signaling pathways involved in these processes. The results showed that the IL-17 signaling pathway was downregulated in the fast stress relaxation hydrogel (Fig. 5E). Consistent with this, protein blotting of IL-17A, IL-6, TNF-α and IL-1β confirmed the RNA-seq results (Fig. 5F). IL-17 is located upstream the NF-κB and MAPK signaling, and its activation leads to the amplification of the inflammatory cascade, loss of cartilage matrix and remodeling [39,40]. Thus, our hypothesis was that fast stress relaxation hydrogel partly mitigated the damage of over-activated IL-17 signaling on NPCs.

Fig. 5.

Transcriptomic analysis. (A) Volcano plot of DEGs. Red dots indicated upregulated genes; blue dots indicate the downregulated genes. (B) Heatmap showing a series of DEGs in NPCs. (C) GO analysis showing the top 20 biological processes involved in DEGs. (D) KEGG enrichment bubble plot showing the top 20 pathways involved in DEGs. (E) GSEA showing that the “cytokine activity” and “collagen catabolic process” were inhibited in the fast stress relaxation group. (F) Western blot of IL-17A, IL-6, TNF-α and IL-1β.

2.7. In vivo repair effect of the multiscale mechanical adapted hydrogels in rats

2.7.1. Imaging evaluation

The therapeutic effect of the slow and fast stress relaxation hydrogel in the rat tail disc was investigated by X-ray and MRI performed after 4 and 8 weeks of implantation. The Defect group showed a significant decrease in disc height at 4 weeks compared to the Sham group, while the disc height in the Slow and Fast groups was well maintained. The vertebrae in the Defect group had blurred cone margins and continued disc height loss at 8 weeks. The disc height was also reduced in the Slow group. The cone rim in the Fast group remained smooth and the disc height did not change significantly (Fig. 6A and B). The signal intensity of the MRI images in the Fast group, which reflected the water content of the IVD, was higher than that of the Defect and Slow groups at 4 and 8 weeks, except for the Sham group. And the Pfirrmann degenerative score of IVD in the Fast group was lowest (Fig. 6C and D). The progression of IVDD was delayed by the implantation of the fast stress relaxation hydrogel as demonstrated by the images.

Fig. 6.

Imaging evaluation of in vivo repair performance. (A) MRI images (right) and X-ray images (left) of rat caudal vertebra at 4 and 8 weeks. Quantitative analyses of (B) DHI%, (C) the relative water content and (D) Pfirrmann degenerative score of IVD based on the MRI image. ∗, P < 0.05; ∗∗, P < 0.01; ∗∗∗, P < 0.001; ∗∗∗∗, P < 0.0001.

2.7.2. Histological evaluation and biomechanical analysis

Further analysis by H&E and SO/FG staining showed significant structural repair of IVD after the implantation of the fast stress relaxation hydrogel (Fig. 7A). The NP tissue in the Defect group was almost disappeared at 4 weeks and other structures of IVD were normal. However, the normal structure of IVD was replaced by acellular, flocculent tissue at 8 weeks, and the annulus fibrosus, the cartilage endplate, and the even bony endplates were very severely damaged. The structure of IVD in the Slow group was similar to that of the Defect group, with no significant repair of the NP tissue. The boundaries between the NP tissue and the annulus fibrosus in the IVD in the Fast group were well defined and some degree of regenerative repair of the NP tissue with the endplate intact was observed at 4 and 8 weeks. Interestingly, differences in degeneration and repair among different groups were still evident as revealed by the histological scores at 4 and 8 weeks (Fig. 7B). The results showed that viscoelastic hydrogel promoted ECM synthesis, protected the NP structural integrity, and thus promoted the regeneration and repair of the NP tissue.

Fig. 7.

Histological evaluation and biomechanical analysis. (A) H&E staining (first row) and SO/FG staining (second row) of IVD at 4 and 8 weeks. (B) Histological grade of IVD in different groups at 4 and 8 weeks based on H&E staining. (C) Axial tension-compression biomechanical testing. (D) Representative force-displacement curve of the Sham group. (E) Results of range of motion (ROM), torsional neutral zone (NZ) length, compressive stiffness and tensile stiffness. ∗, P < 0.05; ∗∗, P < 0.01; ∗∗∗, P < 0.001; ∗∗∗∗, P < 0.0001.

The biomechanical test was used to study the restoration of the biomechanical function after IVD repair using viscoelastic hydrogel (Fig. 7C and D). The neutral zone (NZ) and range of motion (ROM) are two important biomechanical factors that are commonly used to assess the instability of motion segments and biomechanics after disc repair strategies in the cyclic mechanical studies. ROM length, NZ length, compressive and tensile modulus were significantly lower in the Defect group compared to the Sham group as revealed by the axial tension-compression test, suggesting the loss of mechanical function during disc degeneration. The ROM length, NZ length, compressive and tensile modulus in the Slow group were similar to those in the Defect group, suggesting that mechanical function was not restored after slow stress relaxation hydrogel implantation. However, the decrease in ROM and NZ was not significant and the recovery of the compressive and tensile modulus was also the most pronounced in the Fast group, indicating the biomechanical function of IVD recovered well after implantation of the fast stress relaxation hydrogel (Fig. 7E).

3. Discussion

IVDD is a chronic condition characterized by the progressive loss of mechanical stability and shock absorption, which lead to osteophyte formation and restricted motion of spinal segments. Many adverse events are involved in the progression of IVDD, including an overactive inflammatory microenvironment, ECM anabolism and catabolism imbalance, and abnormal mechanical loading [41]. These factors are interrelated and reinforce each other, resulting in a vicious circle [42]. The IVDD process is accompanied by a multiscale alteration of mechanical properties that begins at the molecular scale, progresses to the cellular scale, and ultimately leads to organ-scale damage. Therefore, the design of engineered materials with specific mechanical properties for NP tissue to treat IVDD should not only be considered at the macro-scale, but the micro- and nano-scale should be also considered to obtain a mechanical adaptation to the healthy NP tissue.

Hydrogels, which are usually composed of natural polymers or synthetic polymers, are good candidates to realize multiscale mechanical-adaption to NP tissue [21]. In this study, a dual-network hydrogel was developed, in which one network is chemically crosslinked and the other physically crosslinked and included the violet-crosslinked PBA-m-PGA/GelMA hydrogel as the rigid first network and a dynamic boronic acid-bonded PVA hydrogel as the flexible second network. The dynamic PVA-borate network provides a unique viscoelasticity to the hydrogel, in addition to the injectable and self-healing properties, allowing the hydrogel to effectively mimic the dynamic mechanical microenvironment of ECM [43]. The change of the cross-linking agent or polymer concentration as well as the control of the elastic modulus and stress relaxation time constant by an order of magnitude or more can alter the viscoelastic property of hydrogels, which is an important mechanical property of biomaterial systems that can be tailored to promote the growth of specific cell types [37,44].

At the macroscopic level, the NP tissue when IVD is loaded transmits the loading and energy to the surrounding tissues, dispersing the energy and pressure on the disc. The macroscopic changes in the overall mechanical properties during IVDD lead to a mechanical imbalance in the disc, resulting in the inability to load the disc. On the other hand, growing nerves and the increased number of nociceptors and mechanoreceptors in the degenerate disc are mechanically stimulated in the biomechanically incompetent and abnormally loaded degenerate IVD, leading to increased LBP generation [45]. Therefore, the mechanical properties of implanted biomaterial must not only be sufficiently stiff to support the disc, but also viscoelastic to absorb the externally applied energy to meet the mechanical functional requirements of IVD or spine. Thus, in this study, the viscoelasticity of PVA-DN hydrogels is tunable by adjusting the concentrations of GelMA, PBA-m-PGA and PVA. The results show that different PVA-DN hydrogels could possess the initial elastic modulus similar to that of fresh NP tissue. In addition, the viscoelasticity of fast stress relaxation PVA-DN hydrogel at the macro-scale was closet to fresh NP tissue. The cyclic compressive test demonstrated that the hydrogel material as a whole had good self-recovery property, while the fast stress relaxation hydrogel had the strongest energy absorption capacity, closest to NP. Furthermore, the result of axial tension-compression test showed that the mechanical function of IVD recovered well after 4 and 8 weeks of implantation of the fast stress relaxation hydrogel into the disc, with results similar to those of the Sham group.

At the microscopic level, the gelatinous NP tissue of degenerated IVD becomes less hydrated and the stiffness of ECM increases, causing the formation of more fibrotic NP [15,46]. Once the ECM mechanical microenvironment is altered, NPCs behaviors would be affected in a mechanosensitive manner. Transmembrane ECM receptors act as mechanoreceptors that transfer mechanical signals to the cytoskeleton, thereby controlling cell behavior through a process known as mechanotransduction [47]. Cytoskeletal remodeling and receptor-mediated signaling are important mechanotransduction events that affect the phenotype and biological activity [48]. Thus, in the degenerative pathological microenvironment, NPCs are not only structurally altered, but also functionally impaired, leading to a decrease in ECM secretion. In this work, the fast stress relaxation hydrogels showing similar mechanical properties at a micro-scale with that of NP tissues, was used to provide a suitable mechanical microenvironment for cell survival. According to our experimental results, NPCs in the fast stress relaxation group were the first to sense the mechanical signal and respond by skeleton spreading. The morphology of NPCs gradually changed from polygonal to elongated with increasing matrix stress relaxation time, and the mechanical properties of NPCs also varied with shape. When the stress relaxation of ECM was faster, stress fibers in the cytoskeleton were reduced and the elastic modulus also decreased, meaning that NPCs were easily deformed. In addition, NPCs had more and longer filopodia and 1–2 wider lamellipodia, which greatly improved the motility of cells, as confirmed in the cell migration experiments. Moreover, the spreading area of NPCs was larger when grown on the fast stress relaxation ECM. The shape of the NPCs revealed that they deteriorated as the stress relaxation time of the ECM increased. Gene and protein expression of the NP-specific ECM markers COL2 and ACAN was relatively high in the fast stress relaxation hydrogel, suggesting that viscoelastic hydrogels promoted ECM secretion and synthesis. The low stress microenvironment promoted the biological activity and function of tissue-engineered NP and enhanced ECM synthesis. Local mechanical conditions can be fully exploited using surgical techniques to provide an optimal low-pressure stress microenvironment for the treatment of IVDD. The development of technology allows the precise control of disc loading.

At the nanoscopic level, the integrin-mediated adhesions, which link the cell's actin machinery to ECM molecules presented on a substrate, are intrinsic to the state of cytoskeletal organization and tension in a cell [37]. Integrins on the cell surface are considered mechanosensory receptors and cluster at focal adhesions, acting as mechano-transducers [49]. Vinculin, paxillin and talin are involved in the cytoplasmic adaptation through the integrin binding domain/actin-binding region of these proteins [36,37]. Cells attach to a substrate by binding to cell adhesion ligands on the ECM through integrin receptors, and a cascade of signals occur, such as adherens junction signaling, to connect the cytoskeleton and the ECM [50]. The mechanical strength of the actin-to-ECM linkage increases in response to forces applied at this site. This is achieved by the local accumulation of actin filaments and linker proteins connecting actins to integrins, at force-bearing adhesion sites, leading to an increase in the number of molecular bonds between the actin cytoskeleton- and ECM-bound integrins [51]. Our study showed that changes in mechanical properties at the nano-scale affected the distribution and size of integrin-binding domains/actin-binding regions, resulting in a significant impact on the cellular response to morphogenetic, migratory, and proliferative functions. The mechanical properties of materials must therefore be tuned at multiple scales to adapt to the mechanical environment of cell regeneration.

The elastic moduli of the viscoelastic hydrogels and NP tissues at the macroscopic scale were similar to those at the microscopic scale, which is a few kilopascals and not statistically different. However, at the nanoscopic scale, the moduli were all in the tens of kilopascals and there was a gradual increase in modulus from slow to fast relaxation hydrogels. The stress relaxation times of the viscoelastic hydrogels and NP tissues were very different at the macroscopic and microscopic scales, but the stress relaxation curves were essentially similar at the micrometer and nanometer scales. The relaxation times of the fast stress relaxation hydrogels were closest to those of the fresh NP tissues at all scales, although there are still differences between them. In gels with ionic crosslinks, stress relaxation is mainly achieved by the breaking and subsequent reformation of the ionic cross-links. In contrast, in gels with only covalent cross-links, stress relaxation is mainly achieved by the water migration and the relaxation rate slows down with increasing sample size [52,53]. The dual-network hydrogel we constructed is formed by the formation of the first flexible network by borate bonding, a dynamic covalent bond, and the second rigid network by covalent cross-linking by violet light irradiation. Thus, stress relaxation is attributed mainly to viscoelastic behavior, and the water migration also has an effect. Multiscale extracellular mechanical or structural environments are very important to promoting cellular function and regeneration in a variety of tissues. Through the development of multiscale mechanically tunable decellularised extracellular matrix (dECM), the changes in the micromechanical properties lead to macrophage-mediated differential host immune responses to the material, thus playing a key role in determining the outcome of tissue regeneration and wound healing [54]. The cellular infiltration and migration capacity within the electrospun scaffold is also influenced by controlling the multiscale mechanical properties of the electrostatically spun membrane [55]. The multiscale hierarchical structures and mechanical environments created by using resorbable electrostatically spun poly (L-lactic acid) (PLLA) nanofibres are also promising for in vivo tendon and ligament regeneration [56]. Therefore, the multiscale mechanical adaptation might be an effective strategy for cellular function and tissue regeneration in the nucleus pulposus.

Our work further found that fast stress relaxation PVA-DN hydrogels had anti-inflammatory effects, as revealed by sequencing. A relationship exists between viscoelasticity of hydrogel and inflammation, and cells cultured in viscoelastic hydrogel block their inflammatory process by inhibiting the myeloid-specific isoform of phosphatidylinositol PI3K-γ, compared with monocytes cultured in a rigid matrix, with a positive correlation between myelofibrotic grades and viscoelasticity [57]. Our results revealed that the PI3K-Akt signaling pathway in the fast stress relaxation hydrogel was inhibited, as revealed by sequencing, and our speculation was that viscoelasticity of hydrogel inhibited IL-17 expression by blocking the PI3K-Akt signaling pathway. IL-17 activates downstream NF-κB, MAPK and ERK signaling pathways by binding to the IL-17 receptor. These downstream signaling pathways further activate the expression of factors involved in inflammation and ECM catabolism. The pro-inflammatory cytokine IL-17 functions in a complex cytokine network and is considered a key component of autoimmune and inflammatory processes, which contribute to the pathogenesis of various inflammatory diseases such as psoriasis, psoriatic arthritis and ankylosing spondylitis [[58], [59], [60]]. IL-17 increases the expression of the receptor activator of NF-κB ligand (RANKL) in the bone destruction segment in osteoblasts and is abundantly expressed in synovial tissue. This is the mechanism used by IL-17 to destroy bones in rheumatoid arthritis [61]. Moreover, Due to its potency, IL-17 has become a clinical target for the treatment of diseases, and three IL-17 inhibitors have already been approved for the treatment of psoriasis, psoriatic arthritis, and ankylosing spondylitis [62]. IL-17 binds to TNF and acts on epithelial cells (e.g., keratinocytes in psoriasis) to produce chemokines that recruit neutrophils and macrophages and promote inflammation. Recent studies have shown that IL-17 regulates the proliferation, migration, and apoptosis of vascular endothelial cells and vascular smooth muscle cells through multiple pathways and promotes the secretion of a variety of cytokines, which contribute to the onset and progression of atherosclerosis [63]. Although only few study on IL-17 in NP are available, a multifunctional and microenvironment-responsive metal-phenolic network release platform, termed TMP@Alg-PBA/PVA, attenuates NPC apoptosis by inhibiting the IL-17/ERK signaling pathway [64]. According to previous studies and our findings, increased IL-17 expression may be one of the causes of excessive ECM catabolism in IVDD and is a potential therapeutic target for IVDD.

NP is a gelatinous tissue with a network of randomly organized collagen fibrils and elastin fibrils [65]. The NP in degenerated IVD tissue is replaced by fibrotic tissue, along with clustering of cells and loss of normal matrix. In addition, degenerated endplates are characterized by irregularity and sclerosis, with microfractures and reduced thickness, that compromise the nutrient supply to NP and internal annulus fibrosus [66]. X-ray and MRI results confirmed that the fast stress relaxation hydrogel injection group maximally maintained the height of the intervertebral space and the water content of the NP compared to the other groups. Partially restored NP ECM tissue was observed in the tissue sections of the hydrogel group tissue sections at 4 and 8 weeks post-operatively, which was structurally more complete. All these in vivo experimental results confirmed the viscoelastic performance of our hydrogels that mechanically adapted to medullary tissues effectively reversing multiple adverse factors leading to IVDD and promoting the regeneration of the ECM portion of the IVD.

In this study, we propose a multiscale mechanically-adapted strategy for intervertebral disc repair. Translating research findings into clinical applications would be a huge step forward, but there is still a long way to go. Production scale-up, stringent regulatory requirements, and different patients' specific conditions have become major challenges for clinical translation. Therefore, we will further improve the theoretical basis of the multiscale mechanically-adapted strategy and verify the safety and effectiveness of this strategy through more animal experiments and clinical trials. Although the viscoelasticity of biomaterial is currently able to promote specific ECM secretion from NPCs, the regulatory mechanism still needs further investigation. In addition, there is a lack of long-term data or specific mechanobiological parameters in multiscale tissue engineering, and ideal mechanical conditions need to be further established. In future studies, we will extend the experimental period to collect longer-term data and explore more relevant specific mechanobiological parameters.

4. Conclusion

This study developed a dual-network hydrogel with tunable viscoelasticity for the multiscale mechanically-adapted strategy aiming to promote NP repair. At the macroscopic level, the fast stress relaxation PVA-DN hydrogels could possess initial elastic modulus and viscoelasticity closest to that of fresh NP tissue. The cyclic compressive test demonstrated that the hydrogel material as a whole had good self-recovery property, while the fast stress relaxation hydrogel had the strongest energy absorption capacity, closest to NP. The results confirmed that changes in the viscoelasticity of the PVA-DN hydrogels induced the reorganization of the cytoskeleton in NPCs at the nanoscopic level, and affecting cell spreading area and cell morphology at the microscopic level. The viscoelastic hydrogel also promoted the secretion of NP-specific ECM and protected against damage to NPCs induced by excessive compressive strain. Moreover, the dynamic compressive strain of 5 % applied to the viscoelastic scaffold further promoted the physiological function of NPCs. RNA-seq results further showed that the fast stress relaxation hydrogels attenuated the inflammatory microenvironment by inhibiting the IL-17 signaling pathway. The in vivo implantation of the fast stress relaxation hydrogels in the degenerated disc could maintain the disc height, water content, intact structure and biomechanical property, thereby promoting the regeneration of IVD. In conclusion, the mechanical design of materials at multiple scales brings new insight for the repair of IVD.

5. Materials and methods

5.1. Preparation of methacryloylated polyglutamic acid grafted with phenylboronic acid (PBA-m-PGA)

At room temperature, γ-Polyglutamic acid (γ-PGA, 5 g, Macklin, China) was completely dissolved in distilled water (50 mL). The γ-PGA solution was activated by adding 4-dimethylpyridine (DMAP, 125 mg, Energy Chemical, China) for 1 h. Methacryloylated polyglutamic acid (m-PGA) was obtained by adding glycidyl methacrylate (GMA, 1.25 mL, Macklin, China) and reacting for 24 h. The prepared m-PGA (100 mg) was completely dissolved in distilled water (100 mL). 1-ethyl-(3-dimethylaminopropyl) carbodiimide hydrochloride (150 mg,Macklin, China) and N-hydroxy-5-norbornene-2, 3-diimide (200 mg,Aladdin, China) were added the m-PGA solution and activated for 3 h. Subsequently, 3-aminobenzeneboronic acid (100 mg, Macklin, China) was added and the reaction was kept away from light for 12 h. The product was dialyzed and lyophilized into PBA-m-PGA.

5.2. Preparation of the dual-network hydrogels

PVA (degree of polymerization 1750, China) was dissolved in distilled water at 95 °C with stirring to prepare 2, 6, and 10 wt% PVA solutions. A 10 wt% PBA-m-PGA solution was prepared. And a 10 wt% GelMA with 60 % substitution (EFL, China) containing 0.05 wt% lithium phenyl (2, 4, 6-trimethylbenzoyl) phosphonate (LAP) (EFL, China) was prepared. A hydrogel prepolymer solution was prepared by mixing equal volumes of the PBA-m-PGA solution and GelMA solution. The above hydrogel prepolymer solution and PVA solutions were mixed at a ratio of 1:1 and stirred continuously at 50 °C to prepare 1 % PVA-DN group, 3 % PVA-DN group and 5 % PVA-DN group. Then they were cross-linked with ultraviolet light (with a wavelength of 405 nm) for 30 s to form the hydrogels.

5.2.1. Uniaxial compressive test

The hydrogel sample of approximately 3.8 mm in height and 11.8 mm in diameter was obtained using a mold. The hydrogel sample was placed on the platform of the mechanical testing machine, the height of the upper indenter was carefully adjusted to make contact with the hydrogel, and then all parameters were set to 0. To obtain the stress-strain curve, the load was applied at a compressive strain rate of approximately 0.004/s. The Young's modulus was calculated from the slope of the first 10 % of stress-strain curve. The stress relaxation test was used to test the viscoelastic properties of hydrogels. After compressing the hydrogel to strain of 15 % and maintaining the strain constant, the load is gradually reduced to 50 % of the initial load, which is the stress relaxation time [67]. It is classified into Slow stress relaxation group (1 % PVA-DN), Medium stress relaxation group (3 % PVA-DN) and Fast stress relaxation group (5 % PVA-DN) according to the length of stress relaxation time.

5.2.2. Cyclic compressive test

The hydrogel sample was placed on the platform of the mechanical testing machine, the height of the upper indenter was carefully adjusted to make contact with the hydrogel, and then all parameters were set to 0. For cyclic compression testing, the hydrogel was compressed to strain of 20 % and then unloaded to stress of 0. The maximum strain of each cycle range was 20 %, the minimum stress was 0, and the cycle was repeated for 100 times.

5.2.3. Rheological test

The rheological properties of the hydrogels (20 mm diameter and 1 mm thickness) were measured using a Discovery HR 2 hybrid rheometer (TA Instrument, USA). Measurement starts when the gap between the rotor and the sample table reaches the set value of 1 mm. Frequency sweep tests were performed from 0.1 Hz to 10 Hz at 37 °C and strain of 1 %. Three key parameters were determined: energy storage modulus (G′), loss modulus (G″), and loss factor (tanδ = G′′/G′).

5.3. Hydrogel characterization

The prepared samples were fixed in 2.5 % neutral glutaraldehyde (Macklin, China) for 1 h, rinsed with PBS solution and frozen for 24 h. After freeze-drying for 3 days, the frozen samples were cut with a scalpel blade to expose the internal structure. The dried samples were mounted on the SEM stage with the cut side up using conductive adhesive, and then the samples were gold-sprayed using an ion sputterer (gold-spraying parameters: 20 mA, 45 s). Microscopic morphology (excitation voltage: 15–20 kV) was observed using a scanning electron microscope (SEM, FEI, USA). The identification of m-PGA and PBA-m-PGA were explored with proton nuclear magnetic resonance spectroscopy (¹ H NMR, Varian FT-500 MHz, America). The proton peaks at chemical shift 6.08 ppm and chemical shift 5.61 ppm in the plots indicates the successful grafting of methacrylate group, while the proton peak at chemical shift 7.61 ppm indicates the successful grafting of phenylboronic acid group [68]. Hydrogels were characterized and analyzed using a fourier-transform infrared spectroscope (FTIR, Thermo Fisher Scientific, USA). Briefly, the flakes were prepared with a potassium bromide grinding tablet and scanned 128 times in the range of 400–4000 cm⁻¹ at a resolution of 4 cm⁻¹. The classic gravimetric method was used to evaluate the swelling properties of hydrogels. Briefly, hydrogels were lyophilized to determine their dry weight (W_d). At various time points, the swollen samples (W_s) were weighed during immersion in PBS at 37 °C. The mass swelling ratio was defined as:

SR_m = [(W_s-W_d)/W_d] × 100% (1)

The hydrogel was decomposed by keeping it in PBS and weighing it at different times. The mass remaining ratio was defined as:

RRm = W1/W0 × 100% (2)

W₀ is the initial weight after dissolution equilibrium, W₁ is the weight weighed during degradation.

5.4. Atomic force microscopy (AFM)

To measure the mechanical properties of hydrogels and NP tissue at micro- and nano-scales, experiments were performed by using an AFM scanner (Dimension ICON, Bruker, USA) in force-volume mode in a 0.15 M PBS (pH = 7.4) liquid cell. For microscale experiments, a spherical borosilicate glass tip (Bruker, USA) with a spring constant of 0.06 N/m and a diameter of 5000 nm was bonded to the V-shaped silicon nitride cantilever. The following equation was used to calculate the elastic modulus:

$$E=\frac{\sqrt{\pi }}{2}\left(1-v^2\right)\frac{S}{\sqrt{A}}$$ (3)

where E is the elastic modulus, υ is Poisson's ratio, S is contact stiffness, and A is the contact area of the spherical indenter. The detailed calculation referred to previous studies [69,70].

While for nanoscale experiments, elastic modulus was obtained using the tip (Bruker, USA) with a diameter of 30 nm and a spring constant of 0.06–0.18 N/m. The detail determination of modulus referred to a previous study [71]:

$$F=\frac{4E}{3(1-v^2)}\sqrt{R}{\delta }^{3/2}$$ (4)

where F is indentation force, E is Young's modulus, υ is Poisson's ratio, R is radius of the indenter and δ is indentation.

Microscale stress relaxation curves were recorded by an oscilloscope (LeCroy, New York, USA) connected to the AFM with a voltage of 1 V and surface residence time of 4 s. All experiments were performed in the central region of the hydrogel and cells.

For cell imaging, cells were fixed with 4 % paraformaldehyde for 30 min, and then scanned using a spherical-tip probe (tip radius <10 nm). The scan size was 60 × 60 μm with a resolution of 128 × 128.

5.5. Proliferation, viability and migration of NPCs

To assess the viability of NPCs in different groups, cell viability assays were performed using a live/dead staining kit (Invitrogen, USA). NPCs on the hydrogel were incubated for 48 h and then stained for 30 min. The condition of cells was observed under a Zeiss Axiovert 40CFL microscope (Zeiss, Oberkochen, Germany). Cell Counting Kit-8 assay (CCK-8) (Invitrogen, USA).was used to evaluate the effect of viscoelasticity on cell proliferation. NPCs were plated on the surface of hydrogels at an initial density of 1000 cells/cm². The CCK-8 solution was added on days 1, 3, 5, and 7, respectively, and NPCs were incubated for 1 h at 37 °C in the absence of light. Absorbance values at 450 nm (OD450nm) were measured using a microplate spectrophotometer (BioTek, VT, USA). A scratch test was used to measure the migratory ability of NPCs. A stainless steel square column 1 mm wide was placed on the surface of hydrogel. After inoculation and complete spreading of the cells, the square columns were removed and replaced with serum-free medium, and photographs were taken at 0 and 24 h. Relative mobility was calculated using Image J software (NIH, MD, USA).

5.6. Immunofluorescence staining

After incubating NPCs on different stress relaxation substrate surfaces for 48 h, we fixed the cells with 4 % paraformaldehyde for 30 min. Cells were permeabilized with 0.3 % Triton X-100 in PBS for 10 min and then soaked in Immunol Staining Blocking Buffer for 1 h to block nonspecific binding. For cytoskeletal staining, cells were incubated in Actin-Tracker Green-488 staining solution for 50 min at 37 °C. For focal adhesion staining, cells were incubated with 1:200 diluted anti-vinculin antibodies at 4 °C overnight, Alexa Fluor 555 antibodies (1:1000) were used for fluorescent labeling. Cell nuclei were stained with DAPI. Immunofluorescence images were visualized and captured using a Zeiss Axiovert 40CFL microscope (Zeiss, Oberkochen, Germany).

5.7. Cyclic compression on hydrogels laden with NPCs

A mechanical driven bioreactor unit was used to study to cell response to dynamic compression, which consisted of a computer system, a control unit, a drive motor, and a cell culture chamber, mounted in a cell culture vessel. The mechanical cell culture system (LinMot-Talk, China) provided vertical cyclic compression with adjustable frequency and strain to apply mechanical stimulation to NPCs (initial density of 1 × 10⁶ cells/cm²). After sterilization of the device, the cells were encapsulated in a hydrogel for 24 h in the 3D static culture. The hydrogel laden with cells were compressed with frequency of 1 Hz, strain amplitude of 5 % or 20 %, and daily duration of 2 h for 3 days.

5.8. Reverse transcription-quantitative polymerase chain reaction (RT-qPCR)

To collect NPCs in 3D culture, the configured GelMA lysis solution (6 μL/mL) was added to the hydrogel and placed in the incubator to lyse for 2 h, then filtered and centrifuged to obtain the cell precipitate. Total RNA was extracted from NPCs using TRIzol reagent, and the RNA concentration was measured using a NanoDrop 2000 spectrophotometer (Thermo Fisher Scientific, Waltham, USA). RNA was converted to first-strand cDNA for quantitative polymerase chain reaction. RT-PCR was performed on a CFX96™ Real-Time PCR System using the iTap™ Universal SYBR® Green Supermix Kit (Bio-Rad, CA, USA). The transcript levels of NP-specific ECM synthesis and degradation genes, including Acan, Coll2a1, Mmp13, and Mmp3, were detected; the primer sequences of the target genes are listed in the Table, and the housekeeping gene GAPDH was used as an internal standard. The relative transcript levels of the target genes were normalized to GAPDH and analyzed using the 2^–ΔΔCt method.

5.9. Western blot assay

NPCs were harvested as described above, and total cell proteins were extracted with RIPA lysis buffer (Solarbio, China) containing protease inhibitors on ice for 30 min. Protein concentration was measured using the BCA protein assay kit (Beyotime, China). Equal amounts of proteins from each extract were denatured and then transferred to nitrocellulose membranes by electrophoresis. The membranes were blocked with 5 % wt/v skim milk powder for 30 min and then incubated with the appropriate primary antibody solution overnight at 4 °C. The membranes were then incubated with the appropriate horseradish peroxidase-conjugated secondary antibodies (goat anti-rabbit or goat anti-mouse, 1:1000) for 1 h at room temperature, and the bands were visualized using developing solutions and instruments. The intensity of the bands was quantified using Image J software.

5.10. RNA-seq and bioinformatic analyses

NPCs were grown in slow and fast stress relaxation hydrogels for 3D culture for 3 days, with 3 sets of parallel samples. RNA was extracted from NPCs using TRIzol reagent (Invitrogen, CA, USA). Sequencing of RNA libraries was performed on the illumine Novaseq™ 6000 platform from LC Bio Technology CO, Ltd (China). 2 × 150bp paired-end sequencing (PE150) was performed on Illumina NovaseqTM6000 (LC-Bio Technology Co., Ltd., China) according to the supplier's recommended method. Analysis of the differential expression of genes between the two groups was performed using DESeq2 software. Genes were considered to be differentially expressed if they had a false discovery rate <0.05 and an absolute fold change ≥2. GO terms that met the conditions with p < 0.05 were defined as significantly enriched GO terms in the differentially expressed genes (DEGs). Gene set enrichment analyses were performed using GSEA (v4.1.0) and MSigDB software to determine whether a set of genes for a given GO term and KEGG pathway showed significant differences between the two groups. The KEGG pathways with |NES| > 1, p-val <0.05 and FDR q-val <0.25 that met this condition under GO conditions were considered to be different between the two groups. The National Genebank of China was consulted for all sequencing reads.

5.11. Animal models and surgical procedures

All procedures followed the NIH Guide for the Care and Use of Laboratory Animals and were approved by the Institutional Animal Care and Use Committee of Soochow University (SUDA20231113A03). A total of 30 male Sprague-Dawley rats (8–10 weeks old) were used in the study and randomized into three groups: the Sham group, the Defect group (removal plus PBS injection), the Slow and Fast group (removal plus slow or fast stress relaxation hydrogel injection). The disc removal model was established by using IDD postoperative aspiration for NP removal at the same segment Co7/8. For the Sham group, only the skin was incised. The rats were anaesthetised by intraperitoneal injection of isoflurane and the central part of the NP tissue was accessed by percutaneous puncture with a 21G needle. After puncturing the annulus, the needle was rotated and held in the disc for 30 s at a depth of 5 mm. NP tissue was visible on the needle after withdrawal. For the Slow and Fast group, the hydrogel precursor was then injected using a micropump at 10 μL with violet cross-linking after needle puncture. For the Defect group, the procedure was performed as described above and finally 10 μL PBS was injected. After surgery, the rats were transferred to a warm, ventilated environment. Disc samples were performed after 4 and 8 weeks of incubation.

5.12. Radiology evaluation

All rats were anaesthetised and placed in the supine position with the tail upright. The rat caudal spine was photographed using an X-ray machine (exposure time: 10 s, 26 kV) and a 1.5T magnetic resonance imaging (MRI) scanner (scanning sequence: FRFSE-XL, layer thickness: 1.4 m) at 4 and 8 weeks after surgery to obtain the disc height index and T2-weighted images, which were processed with Image J software. The Pfirrmann degenerative score of the discs surrounding the implant was based on the MRI image. The disc height index (DHI) was defined as:

$$\text{DHI}=\frac{2\times \left(D+E+F\right)}{(\mathrm{A}+B+C+G+H+I)}$$ (5)

where D, E, F are the anterior, middle, and posterior edge heights of the intervertebral space, A, B, C, G, H, I are the anterior, middle, and posterior heights of adjacent vertebral bodies.

5.13. Histological analysis

The target rat spinal samples were fixed with 4 % paraformaldehyde (Biosharp, China) for 48 h and decalcified with 10 % ethylenediaminetetraacetic acid (EDTA; pH = 7.4) (Shanghai yuanye Bio-Technology Co. LTD, China) for 6 weeks. The decalcified samples were dehydrated and embedded in paraffin. The midsagittal sections (6 μm) were made using a paraffin slicer (Leica, Germany) and used for histologic staining. Hematoxylin and eosin (H&E) staining and Safranin O-fast (S-O) green staining were performed with a staining kit in accordance with standard protocols (Solarbio, China). Histologic scores were assessed as previously described by Masuda [72].

5.14. Biomechanical tests on IVD

The changes in the biomechanical properties of IVD before and after surgery were evaluated by the axial tension-compression tests performed using a universal material testing machine (Hengyi, China). The specimens were loaded for 20 cycles with force amplitude of ±8 N and frequency of 0.5 Hz in a force control mode. The compressive stiffness, tensile stiffness, range of motion (ROM), and axial neutral zone (NZ) length were obtained using the method described in the literature [71,73].

5.15. Statistical analysis

The results were given as means ± standard deviation (SD) of at least three independent experiments. Statistical significance was evaluated with One-way ANOVA or Two-way ANOVA using GraphPad Prism 9.5.0 software (CA, USA). If the P value ≤ 0.05, the difference was considered statistically significant.

CRediT authorship contribution statement

Wenbin Cai: Writing – original draft, Formal analysis, Data curation, Conceptualization. Fanlei Yang: Software, Formal analysis, Data curation. Chao Yang: Validation, Methodology, Data curation. Yu Liu: Validation, Software, Methodology, Data curation. Hao Xu: Supervision, Methodology. Wen Zhang: Supervision, Project administration, Methodology. Bin Li: Investigation, Conceptualization. Fengxuan Han: Writing – review & editing, Supervision, Resources, Funding acquisition. Zongping Luo: Writing – review & editing, Supervision, Funding acquisition. Ting Liang: Writing – review & editing, Supervision, Funding acquisition, Conceptualization.

Ethics approval and consent to participate

All procedures followed the NIH Guide for the Care and Use of Laboratory Animals and were approved by the Institutional Animal Care and Use Committee of Soochow University (SUDA20231113A03).

Declaration of competing interest

The authors declare that they have no competing interests.

Appendix A. Supplementary data

The following is the Supplementary data to this article.