toughness recovery was 67% after 2 min. The biodegradability was investigated in vitro in PBS and evaluated via the loss of mechanical properties (close to zero after 24 h) and mass (27–47% after 24 d).

Cai et al. [135] designed soft hydrogels with maximum G' values of ≈ 100 Pa with two physical cross-linking mechanisms: weak heterodimeric *ex vivo* molecular recognition and in vivo cross-linking due to the thermo-responsive formation of a reinforcing network. The first network was composed of a star-shaped proline-rich peptide-PEG copolymer assembled with a recombinant linear protein bearing CC43 WW and RGD cell-binding domains connected by hydrophilic spacers. The second was based on PNIPAM chains conjugated to the PEG copolymer, and the erosion of the material was ~70% after 14 d in vitro.

## 6.3. Dual Self-Healing via a Combination Of Dynamic Chemical Bonds and Physical Interactions

### 6.3.1. Combination of Schiff-base and Physical Interactions

Unsurprisingly, due to the numerous self-healing systems relying on Schiff-base reactions, this dynamic bond is the most common one in dual self-healing systems combining dynamic chemical bonds and physical interactions. Several examples of self-healing hydrogels based on the combination of dynamic Schiff-base bonds as the chemical bonds and copolymer micelle cross-linking as the physical interactions have been reported (Table 10). Guo et al. [136] prepared hydrogels by mixing quaternized chitosan with micelles of benzaldehyde-terminated Pluronic F-127 (PF-127) under physiological conditions. Similarly, hydrogels based on CEC and benzaldehyde-terminated PF-127/CNTs were prepared.[137] Self-healing was due to the Schiff-base and physical cross-linking interactions of the PF-127 micelles and CNTs. In both examples, the self-healing times were 2–3 h and the G' values were in the range of 10–30 kPa.

Other dual self-healing hydrogels combine Schiff-base bonds and catechol-Fe coordination. This is the case for catechol-modified poly(l-lysine) and A-Dex [138] or catechol-modified oxidized HA and aminated gelatin,[139] which are associated with Fe³⁺ ions. The first system yielded soft gels with G' values limited to 7.6 kPa compared to the G' of the second system of 535 kPa. The hydrogels rapidly self-healed with SHEs of >91% and could degrade in PBS or in the presence of iron-chelating competitors. The cytocompatibility of the catechol-modified poly(l-lysine)/A-Dex hydrogel was confirmed toward L929 and NIH 3T3 murine cells.

Finally, Schiff-base linkages were also combined with host-guest interactions [140] or hydrogen bonds.[141] This is exemplified by hydrogels composed of the guest polymer, phenolphthalein-grafted CEC, the host molecule, hexamethylenediamine-modified β-CD, and A-Alg,[140] or chitosan hydrogels cross-linked with vanillin.[141] In the latter case, the amino group of chitosan could react with the aldehyde group of vanillin via a Schiff-base reaction, and its hydroxyl groups could form hydrogen bonds with the hydroxyl or amino groups of another chitosan chain. For these two types of hydrogels, repair times of a few hours were required to reach SHEs of ≥80%.

### 6.3.2. Dual-self-healing systems: other combinations

Two recent examples based on hydrogen bonds and boronate esters are reported in the literature (Table 10). The first was composed of PVA, borax, and carboxymethyl cellulose.[142] In this case, $\sigma_{max}$ and E could reach 720 and 340 kPa, respectively, and self-healing was realized after 16 h at 20 °C. In the second example, hydrogels were obtained using dynamic benzoxaborole-sugar interactions and quadruple hydrogen bonds of UPy moieties in poly(DMA-st-META-st-LAEMA) and poly(MPC-st-UPyHEMA-st-MAABO) copolymers.[143] Depending on the compositions of the different gels, G' ranged from 0.3 to 1.1 kPa. Self-healing times were rapid, with shape recoveries in only 20 s for the samples evaluated via the cutting method.

Disulfide bonds have also been used in systems with hydrogen bonds. Wen et al. proposed a degradable PU hydrogel composed of PCL, PEG, and IPDI, containing disulfide bonds and hydrogen bonds.[144] The SHE, as quantified via tensile test reached 97%, and the materials exhibited high mechanical strengths, with $\sigma_{max}$ values of up to 3.3 MPa. At all PU/PCL ratios, the mass loss in PBS was limited to 8% after 30 d.

A combination of DA adducts and electrostatic interactions was proposed by Li et al., using pectin/chitosan hydrogels. Pectin was modified with furfural to form the conjugated diene partner, and 6-maleidocaproic acid was grafted onto chitosan to yield the dienophile partner.[145] The hydrogels healed in 5 h at 37 °C and could bear a mass of 5 N without damage.

Finally, Jiang et al. imparted self-healing properties on elastomers due to oxime-urethane and hydrogen bonds.[146] PU copolymers based on dimethylglyoxime, IPDI, PTHF, and glycerol were prepared. Depending on the glycerol content, $\sigma_{max}$, E, and $\epsilon_{max}$ values of 33 kPa–4.4 MPa, 172 kPa–3.7 MPa, and 500–3300%, respectively, were obtained. The self-healing capacity was also a function of the glycerol content. After healing at room temperature for 5 min, the SHEs in terms of $\sigma_{max}$ were >80% for elastomers but limited to 37% at low and high glycerol contents. Under in vitro enzymatic conditions, the elastomers lost 7.5% of their masses after nine days.

## 6.4. Multi-mechanism self-healing systems

Finally, although uncommon, the self-healing properties may be due to the combination of more than two mechanisms. To the best of our knowledge, only two examples have combined two types of chemical linkages with one physical interaction. In both cases, boronate ester bonds are associated with Schiff-base linkages and hydrogen bonds [147] or electrostatic interactions.[148] In the first system, hydrogels were prepared using PVA and borax as the matrix, which was reinforced with cellulose nanofibers and dopamine-grafted oxidized carboxymethyl cellulose. In the second study, gels were prepared using borax-functionalized oxidized chondroitin sulfate (BOC), gelatin, and BOC-doped polypyrrole. Both hydrogels were cytocompatible in vitro and exhibited similar healing times of ≈ 5 min and G' values in the range of 0.4–10 kPa.

A combination of two physical interactions and one chemical reversible bond is also possible. This is the case in a study by Tran et al. [149] that reported hydrogels based on dynamic
<table><caption>**Table 10.** Self-healing hydrogels and elastomers based on combination of dynamic chemical bonds and physical interactions (materials, mechanisms, mechanical properties, SH evaluation, SH conditions, SHE, and degradation).</caption><thead><tr><th>Nature</th><th>Mechanism</th><th>Materials</th><th>Mechanical properties</th><th>SH evaluation</th><th>SH conditions</th><th>SHE</th><th>Degradation</th><th>Ref.</th></tr></thead><tbody><tr><td>Hydrogel</td><td>Schiff base<br>Copolymer micelle<br>cross-linking</td><td>Quaternized chitosan<br>Benzaldehyde-terminated<br>Pluronic micelles</td><td>E = 21 to 37 kPa<br>G' = 22 to 53 kPa</td><td>Cut</td><td>2 h, 25 °C</td><td>100% (G')</td><td>In vitro</td><td>[136]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Cross-linking inter-<br>actions of micelles<br>and CNTs</td><td>CEC<br>Benzaldehyde Pluronic<br>micelles<br>CNTs</td><td>G' = 6 to 16 kPa</td><td>Cut</td><td>3 h</td><td>100%</td><td>In vitro, 49 to 74%<br>mass loss in 13d</td><td>[137]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Catechol-Fe<br>coordination</td><td>Catechol modified<br>poly(L-lysine)<br>A-Dex<br>Fe</td><td>G' = 0.7 to 7.6 kPa<br>Compression stress = 25<br>to 80 kPa</td><td>Cut</td><td>Water</td><td>91%</td><td>In vitro, with<br>deferoxamine<br>mesylate</td><td>[138]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Catechol-Fe<br>coordination</td><td>Catechol-modified<br>oxidized HA<br>Aminated gelatin<br>Fe<sup>3+</sup></td><td>G' up to 535 kPa</td><td>Cut<br>Syringe injection</td><td>15 min</td><td>97% (G')</td><td>In vitro, 12d</td><td>[139]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Host-guest<br>interactions</td><td>Phenolphthalein-grafted<br>CEC<br>Hexamethylenediamine<br>modified β-CD<br>A-Alg</td><td>Compression stress up to<br>12.5 kPa</td><td>Cut</td><td>4 h</td><td>80%</td><td>In vitro, 54 to 66%<br>mass loss in 168h</td><td>[140]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Hydrogen bonds</td><td>Chitosan<br>Vanillin</td><td>G' up to 80 kPa</td><td>Cut</td><td>5 h</td><td>93% (σ<sub>max</sub>)<br>99% (ε<sub>max</sub>)</td><td>–</td><td>[141]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Hydrogen bonds</td><td>PVA<br>Borax<br>Carboxymethylcellulose</td><td>E up to 340 kPa<br>σ<sub>max</sub> up to 760 kPa</td><td>Cut</td><td>16 h, 20 °C</td><td>80% (σ<sub>max</sub>)</td><td>–</td><td>[142]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Hydrogen bonds</td><td>Poly(DMA-st-META-st-<br>LAEMA)<br>Poly (MPC-st-UPyHEMA-<br>st-MABBO)</td><td>G' = 0.3 to 1.1 kPa</td><td>Cut</td><td>20 s</td><td>100% (G')</td><td>In vitro, 120 min<br>with fructose</td><td>[143]</td></tr><tr><td>Hydrogel</td><td>Disulfide bonds<br>Hydrogen bonds</td><td>PU (IPDI) — PCL—PEG</td><td>σ<sub>max</sub> = 3.3 MPa</td><td>Cut</td><td>10 min at 50 °C,<br>then 1 h at RT and<br>24 h in water</td><td>97% (σ<sub>max</sub>)</td><td>In vitro, 4 to 8%<br>mass loss in 30 d</td><td>[144]</td></tr><tr><td>Hydrogel</td><td>DA bonds<br>Electrostatic<br>interactions</td><td>Furfural-pectin<br>6-maleidocaproic acid<br>chitosan</td><td>–</td><td>Cut</td><td>5 h, 37 °C</td><td>–</td><td>–</td><td>[145]</td></tr><tr><td>Elastomer</td><td>Oxime-urethane<br>bonds<br>Hydrogen bonds</td><td>PU based on dimethylg-<br>lyoxime, IPDI, PTMEG,<br>and glycol</td><td>E = 172 kPa to 3.7 MPa<br>σ<sub>max</sub> = 33 kPa to 4.4 MPa<br>ε<sub>max</sub> = 500 to 3300%</td><td>Cut</td><td>5 min, RT</td><td>80% (σ<sub>max</sub>)</td><td>In vitro, 7.5% mass<br>loss with enzyme<br>after 9d</td><td>[146]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Schiff base<br>Hydrogen bonds</td><td>PVA<br>Borax<br>Cellulose nanofibers<br>Dopamine-grafted<br>oxidized carbomethyl<br>cellulose</td><td>G' = 2 to 10 kPa</td><td>Cut</td><td>5 min</td><td>100% (G')</td><td>In vitro, pH-depen-<br>dent, up to 60%<br>mass loss in 3.5 h<br>at pH 5</td><td>[147]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Schiff base<br>Electrostatic<br>interactions</td><td>BOC<br>Gelatin<br>BOC-doped polypyrrole</td><td>G' = 0.4 to 1.6 kPa</td><td>Cut</td><td>5min</td><td>–</td><td>In vivo, 14 to 21d</td><td>[148]</td></tr><tr><td>Hydrogel</td><td>Disulfide bonds<br>Hydrogen bonds<br>Ionic interactions</td><td>2,3-dimercapto-<br>1-propanol<br>Meso-2,3-<br>dimercaptosuccinic acid<br>Ca<sup>2+</sup></td><td>E = 13.4 kPa<br>σ<sub>max</sub> = 12.1 kPa<br>ε<sub>max</sub> = 4334%</td><td>Cut</td><td>1 min, 25 °C, air or<br>water</td><td>93 to 100% (E,<br>σ<sub>max</sub>, ε<sub>max</sub>)</td><td>In vitro, 5 h with<br>glutathione</td><td>[149]</td></tr></tbody></table>
<table><caption>**Table 10.** Self-healing hydrogels and elastomers based on combination of dynamic chemical bonds and physical interactions (materials, mechanisms, mechanical properties, SH evaluation, SH conditions, SHE, and degradation).</caption><thead><tr><th>Nature</th><th>Mechanism</th><th>Materials</th><th>Mechanical properties</th><th>SH evaluation</th><th>SH conditions</th><th>SHE</th><th>Degradation</th><th>Ref.</th></tr></thead><tbody><tr><td>Hydrogel</td><td>Schiff base<br>Copolymer micelle<br>cross-linking</td><td>Quaternized chitosan<br>Benzaldehyde-terminated<br>Pluronic micelles</td><td>E = 21 to 37 kPa<br>G' = 22 to 53 kPa</td><td>Cut</td><td>2 h, 25 °C</td><td>100% (G')</td><td>In vitro</td><td>[136]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Cross-linking inter-<br>actions of micelles<br>and CNTs</td><td>CEC<br>Benzaldehyde Pluronic<br>micelles<br>CNTs</td><td>G' = 6 to 16 kPa</td><td>Cut</td><td>3 h</td><td>100%</td><td>In vitro, 49 to 74%<br>mass loss in 13d</td><td>[137]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Catechol-Fe<br>coordination</td><td>Catechol modified<br>poly(L-lysine)<br>A-Dex<br>Fe</td><td>G' = 0.7 to 7.6 kPa<br>Compression stress = 25<br>to 80 kPa</td><td>Cut</td><td>Water</td><td>91%</td><td>In vitro, with<br>deferoxamine<br>mesylate</td><td>[138]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Catechol-Fe<br>coordination</td><td>Catechol-modified<br>oxidized HA<br>Aminated gelatin<br>Fe<sup>3+</sup></td><td>G' up to 535 kPa</td><td>Cut<br>Syringe injection</td><td>15 min</td><td>97% (G')</td><td>In vitro, 12d</td><td>[139]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Host-guest<br>interactions</td><td>Phenolphthalein-grafted<br>CEC<br>Hexamethylenediamine<br>modified β-CD<br>A-Alg</td><td>Compression stress up to<br>12.5 kPa</td><td>Cut</td><td>4 h</td><td>80%</td><td>In vitro, 54 to 66%<br>mass loss in 168h</td><td>[140]</td></tr><tr><td>Hydrogel</td><td>Schiff base<br>Hydrogen bonds</td><td>Chitosan<br>Vanillin</td><td>G' up to 80 kPa</td><td>Cut</td><td>5 h</td><td>93% (σ<sub>max</sub>)<br>99% (ε<sub>max</sub>)</td><td>—</td><td>[141]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Hydrogen bonds</td><td>PVA<br>Borax<br>Carboxymethylcellulose</td><td>E up to 340 kPa<br>σ<sub>max</sub> up to 760 kPa</td><td>Cut</td><td>16 h, 20 °C</td><td>80% (σ<sub>max</sub>)</td><td>—</td><td>[142]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Hydrogen bonds</td><td>Poly(DMA-st-META-st-<br>LAEMA)<br>Poly (MPC-st-UPyHEMA-<br>st-MABBO)</td><td>G' = 0.3 to 1.1 kPa</td><td>Cut</td><td>20 s</td><td>100% (G')</td><td>In vitro, 120 min<br>with fructose</td><td>[143]</td></tr><tr><td>Hydrogel</td><td>Disulfide bonds<br>Hydrogen bonds</td><td>PU (IPDI) — PCL—PEG</td><td>σ<sub>max</sub> = 3.3 MPa</td><td>Cut</td><td>10 min at 50 °C,<br>then 1 h at RT and<br>24 h in water</td><td>97% (σ<sub>max</sub>)</td><td>In vitro, 4 to 8%<br>mass loss in 30 d</td><td>[144]</td></tr><tr><td>Hydrogel</td><td>DA bonds<br>Electrostatic<br>interactions</td><td>Furfural-pectin<br>6-maleidocaproic acid<br>chitosan</td><td>—</td><td>Cut</td><td>5 h, 37 °C</td><td>—</td><td>—</td><td>[145]</td></tr><tr><td>Elastomer</td><td>Oxime-urethane<br>bonds<br>Hydrogen bonds</td><td>PU based on dimethylg-<br>lyoxime, IPDI, PTMEG,<br>and glycol</td><td>E = 172 kPa to 3.7 MPa<br>σ<sub>max</sub> = 33 kPa to 4.4 MPa<br>ε<sub>max</sub> = 500 to 3300%</td><td>Cut</td><td>5 min, RT</td><td>80% (σ<sub>max</sub>)</td><td>In vitro, 7.5% mass<br>loss with enzyme<br>after 9d</td><td>[146]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Schiff base<br>Hydrogen bonds</td><td>PVA<br>Borax<br>Cellulose nanofibers<br>Dopamine-grafted<br>oxidized carbomethyl<br>cellulose</td><td>G' = 2 to 10 kPa</td><td>Cut</td><td>5 min</td><td>100% (G')</td><td>In vitro, pH-depen-<br>dent, up to 60%<br>mass loss in 3.5 h<br>at pH 5</td><td>[147]</td></tr><tr><td>Hydrogel</td><td>Boronate ester<br>bonds<br>Schiff base<br>Electrostatic<br>interactions</td><td>BOC<br>Gelatin<br>BOC-doped polypyrrole</td><td>G' = 0.4 to 1.6 kPa</td><td>Cut</td><td>5min</td><td>—</td><td>In vivo, 14 to 21d</td><td>[148]</td></tr><tr><td>Hydrogel</td><td>Disulfide bonds<br>Hydrogen bonds<br>Ionic interactions</td><td>2,3-dimercapto-<br>1-propanol<br>Meso-2,3-<br>dimercaptosuccinic acid<br>Ca<sup>2+</sup></td><td>E = 13.4 kPa<br>σ<sub>max</sub> = 12.1 kPa<br>ε<sub>max</sub> = 4334%</td><td>Cut</td><td>1 min, 25 °C, air or<br>water</td><td>93 to 100% (E,<br>σ<sub>max</sub>, ε<sub>max</sub>)</td><td>In vitro, 5 h with<br>glutathione</td><td>[149]</td></tr></tbody></table>
<table><caption>**Table 10. Continued.**</caption><thead><tr><th>Nature</th><th>Mechanism</th><th>Materials</th><th>Mechanical properties</th><th>SH evaluation</th><th>SH conditions</th><th>SHE</th><th>Degradation</th><th>Ref.</th></tr></thead><tbody><tr><td>Hydrogel</td><td>Boronate ester<br>Hydrogens bonds<br>Hydration</td><td>Guar gum-g-poly(acrylamidoglycolic acid)<br>Silver nitrate Sodium borohydride</td><td>Compression (module 0.043 to 0.112 MPa; strain = 80 to 89%)</td><td>Cut<br>Syringe injection</td><td>25 °C</td><td>–</td><td>–</td><td>[150]</td></tr></tbody></table>

A-Alg, aldehyde-modified alginate; A-Dex, aldehyde-modified dextran; BOC, borax-functionalized oxidized chondroitin sulfate; CD, cyclodextrin; CEC, carboxyethyl chitosan; CNTs, carbon nanotubes; DMA, N-N-dimethylacrylamide; E, Young's modulus; G', storage modulus; HA, hyaluronic acid; HEMA, 2-hydroxyethyl methacrylate; IPDI, isophorone diisocyanate; LAEMA, 2-lactobioamidoethyl methacrylamide; MAABO, 5-methacrylamido-1,2-benzoxaborole; META, [2-(methacryloyloxy)ethyl] trimethylylammonium chloride; MPC, 2-methacryloyloxyethyl phosphorylcholine; PCL, polycaprolactone; PEG, poly(ethylene glycol); PTMEG, poly(tetramethylene ether)glycol; PU, polyurethane; PVA, poly(vinyl alcohol); RT, room temperature; ε<sub>max</sub>, strain at break; σ<sub>max</sub>, stress at break.

poly(disulfide) backbones and rapidly reversible physical cross-links, such as hydrogen bonds and ionic interactions (Figure 9). They copolymerized 2,3-dimercapto-1-propanol and meso-2,3-dimercaptosuccinic acid before generating hydrogels in aqueous Ca²⁺ solutions. These hydrogels immediately reconnected after cutting and exhibited rapid self-healing in air at 25 °C or water. The quantitative SHE was confirmed via mechanical property studies, and degradation in a solution of glutathione was complete after 5 h.

Finally, Palem et al. [150] prepared silver nanocomposite hydrogels via in situ addition of guar gum-g-poly(acrylamidoglycolic acid), silver nitrate, and sodium borohydride. The hydrogels were apparently self-healing owing to the dynamic covalent interactions between the borax and cis-diol groups, but also due to different interactions, such as intermolecular hydrogen

bonding and strong hydration during the process. An MTT assay toward the CCD-986sk cell lines confirmed the compatibility of the hydrogels.

# 7. Degradable hydrogels and elastomers with self-healing properties in biomedical applications

## 7.1. Self-Healable Hydrogels in Biomedical Applications

Self-healing hydrogels are attracting increasing interest in the biomedical field owing to their similarities to biological systems (Table 11). Among the self-healing hydrogels, injectable hydrogels are critical. These hydrogels may be extruded through a syringe as liquids due to their shear-thinning behaviors and

<figure><img src="image_11.png" /><figcaption>**Figure 9.** Schematic structure of degradable hydrogel based on 2,3-dimercapto-1-propanol and meso-2,3-dimercaptosuccinic acid copolymers – P(BAL-co-DMSA) – cross-linked with divalent ion Ca²⁺ and its self-healing mechanism. Adapted with permission.[149] Copyright 2020, Elsevier.</figcaption></figure>
<table><caption>**Table 10. Continued.**</caption><thead><tr><th>Nature</th><th>Mechanism</th><th>Materials</th><th>Mechanical properties</th><th>SH evaluation</th><th>SH conditions</th><th>SHE</th><th>Degradation</th><th>Ref.</th></tr></thead><tbody><tr><td>Hydrogel</td><td>Boronate ester<br>Hydrogens bonds<br>Hydration</td><td>Guar gum-g-poly(acrylamidoglycolic acid)<br>Silver nitrate Sodium borohydride</td><td>Compression (module 0.043 to 0.112 MPa; strain = 80 to 89%)</td><td>Cut<br>Syringe injection</td><td>25 °C</td><td>–</td><td>–</td><td>[150]</td></tr></tbody></table>

A-Alg, aldehyde-modified alginate; A-Dex, aldehyde-modified dextran; BOC, borax-functionalized oxidized chondroitin sulfate; CD, cyclodextrin; CEC, carboxyethyl chitosan; CNTs, carbon nanotubes; DMA, N-N-dimethylacrylamide; E, Young's modulus; G', storage modulus; HA, hyaluronic acid; HEMA, 2-hydroxyethyl methacrylate; IPDI, isophorone diisocyanate; LAEMA, 2-lactobioamidoethyl methacrylamide; MAABO, 5-methacrylamido-1,2-benzoxaborole; META, [2-(methacryloyloxy)ethyl] trimethylammonium chloride; MPC, 2-methacryloyloxyethyl phosphorylcholine; PCL, polycaprolactone; PEG, poly(ethylene glycol); PTMEG, poly(tetramethylene ether)glycol; PU, polyurethane; PVA, poly(vinyl alcohol); RT, room temperature; ε<sub>max</sub>, strain at break; σ<sub>max</sub>, stress at break.

poly(disulfide) backbones and rapidly reversible physical cross-links, such as hydrogen bonds and ionic interactions (Figure 9). They copolymerized 2,3-dimercapto-1-propanol and meso-2,3-dimercaptosuccinic acid before generating hydrogels in aqueous Ca²⁺ solutions. These hydrogels immediately reconnected after cutting and exhibited rapid self-healing in air at 25 °C or water. The quantitative SHE was confirmed via mechanical property studies, and degradation in a solution of glutathione was complete after 5 h.

Finally, Palem et al. [150] prepared silver nanocomposite hydrogels via in situ addition of guar gum-g-poly(acrylamidoglycolic acid), silver nitrate, and sodium borohydride. The hydrogels were apparently self-healing owing to the dynamic covalent interactions between the borax and cis-diol groups, but also due to different interactions, such as intermolecular hydrogen

bonding and strong hydration during the process. An MTT assay toward the CCD-986sk cell lines confirmed the compatibility of the hydrogels.

# 7. Degradable hydrogels and elastomers with self-healing properties in biomedical applications

## 7.1. Self-Healable Hydrogels in Biomedical Applications

Self-healing hydrogels are attracting increasing interest in the biomedical field owing to their similarities to biological systems (Table 11). Among the self-healing hydrogels, injectable hydrogels are critical. These hydrogels may be extruded through a syringe as liquids due to their shear-thinning behaviors and

<figure><img src="image_11.png" /><figcaption>**Figure 9.** Schematic structure of degradable hydrogel based on 2,3-dimercapto-1-propanol and meso-2,3-dimercaptosuccinic acid copolymers – P(BAL-co-DMSA) – cross-linked with divalent ion Ca²⁺ and its self-healing mechanism. Adapted with permission.[149] Copyright 2020, Elsevier.</figcaption></figure>
<table><caption>**Table 11.** Degradable self-healing hydrogels and elastomers and their potential applications in the biomedical field.</caption><thead><tr><th>Mechanism</th><th>Nature</th><th>Materials</th><th>Potential biomedical applications</th><th>Ref.</th></tr></thead><tbody><tr><td rowspan="20">Schiff base</td><td rowspan="19">Hydrogel</td><td>Phenol-functionalized chitosan/PEG-BA</td><td>3D Bioprinting</td><td>[40]</td></tr><tr><td>Glycolated chitosan/PEG-BA</td><td>Tissue engineering, vascular repair</td><td>[42]</td></tr><tr><td>Chitosan-g-aniline/PEG-BA</td><td>Cell delivery</td><td>[33]</td></tr><tr><td>CMC/PEG-BA</td><td>Hemostasis</td><td>[43]</td></tr><tr><td>Quaternized methacryloyl chitosan/PEG-BA</td><td>Wound healing</td><td>[44]</td></tr><tr><td>Zwitterionic L-GA acid-functionalized chitosan/PEG-BA</td><td>Delivery</td><td>[45]</td></tr><tr><td>CEC/PEG-BA/Polyacrylamide</td><td>Tissue engineering, actuators, wearable devices</td><td>[46]</td></tr><tr><td>CEC/A-HA</td><td>Drug delivery for tumor therapy</td><td>[36]</td></tr><tr><td>N-succinyl-chitosan/Chondroitin sulfate multiple aldehydes</td><td>Cell delivery and tissue engineering</td><td>[31]</td></tr><tr><td>CEC/A-Alg</td><td>Cell delivery</td><td>[32]</td></tr><tr><td>CMC/Aldehyde-functionalized carboxymethyl cellulose</td><td>Cell delivery</td><td>[37]</td></tr><tr><td>Glycolated chitosan/A-Dex</td><td>Protein release</td><td>[35]</td></tr><tr><td>CMC/A-Dex</td><td>Prevention and treatment of postoperative adhesion</td><td>[48]</td></tr><tr><td>CEC/Dextran-g-aniline tetramer-g-4-formylbenzoic acid</td><td>Tissue engineering and cell delivery</td><td>[34]</td></tr><tr><td>Acrylamide modified chitin/A-Alg</td><td>Drug delivery and tissue engineering</td><td>[38]</td></tr><tr><td>CMC/A-Alg/Hap/CaCO3</td><td>Tissue engineering and drug delivery</td><td>[50]</td></tr><tr><td>CMC/A-Alg/MGMs</td><td>Tissue engineering and drug delivery</td><td>[51]</td></tr><tr><td>L-arginine conjugated chitosan/PEG-BA/pDA-NPs</td><td>Wound dressing</td><td>[54]</td></tr><tr><td>CMC/Poly(dextran-g-4-formylbenzoic acid)/Peptide nanofibers</td><td>Wound healing</td><td>[55]</td></tr><tr><td>Ethylenediamine functional gelatin/Dialdehyde carboxymethylcellulose</td><td>Drug and cell delivery and tissue engineering</td><td>[57]</td></tr><tr><td>Aldehyde-modified xanthan gum/Phosphatidylethanolamine liposomes</td><td>Cell delivery and tissue engineering</td><td>[58]</td></tr><tr><td rowspan="12">Boronate ester bonds</td><td rowspan="12">Hydrogel</td><td>Poly(methacryloyloxyethyl phosphorylcholine) functionalized with benzoxaborole or catechol pendant groups</td><td>Drug delivery and tissue engineering</td><td>[61]</td></tr><tr><td>5-methacrylamido-1,2-benzoxaborole/3-gluconamidoprpyl methacrylamide/Acrylamide</td><td>Bioprinting and cell encapsulation and delivery</td><td>[62]</td></tr><tr><td>Nopoldiol/2-gluconamidoethylmethacrylamide/PEG methyl ether methacrylate/MAABO</td><td>Gene delivery, cell therapy, and tissue engineering</td><td>[63]</td></tr><tr><td>Poly(DMA-st-MAABO)/2-lactobionamidoethyl methacrylamide, exhibiting galactose residues on the surface</td><td>Tissue engineering and drug delivery</td><td>[64]</td></tr><tr><td>Dopamine functionalized 4-armed PEG/Phenylboronic acid modified 4-armed PEG</td><td>Bioadhesion, drug delivery, and tissue engineering</td><td>[65]</td></tr><tr><td>PVA/Poly(aspartic acid) derivatives with boronic acid groups</td><td>Wound dressing, tissue repairing</td><td>[66]</td></tr><tr><td>PVA/Carboxyethyl cellulose grafted with phenylboronic acid</td><td>Tissue engineering, wound healing, and drug delivery</td><td>[68]</td></tr><tr><td>Methacrylated hyaluronic acid/3-aminophenylboronic acid-modified sodium alginate</td><td>Cell delivery and tissue regeneration</td><td>[69]</td></tr><tr><td>PVA/Poly(acrylamide-co-dopamine methacrylamide) bis(phenylboronic acid carbamoyl) cystamine</td><td>Drug delivery in cancer treatment</td><td>[70]</td></tr><tr><td>CTL/Boric acid/Mannitol</td><td>Tissue engineering cartilage regeneration</td><td>[71]</td></tr><tr><td rowspan="7">Hydrazone bonds</td><td rowspan="7">Hydrogel</td><td>Tetraphenylethylene-poly(DMA-stat-diacetone acrylamide)2/Acylhydrazide functionalized pectin</td><td>Drug delivery and tissue engineering</td><td>[72]</td></tr><tr><td>Pectin aldehyde/Poly(NIPAM-stat-acylhydrazide)</td><td>Anticancer drug carrier</td><td>[73]</td></tr><tr><td>A-HA/Hydrazide-HA</td><td>RNA delivery</td><td>[74]</td></tr><tr><td>Adipic dihydrazide-grafted carboxyethyl chitin/PEG-BA Phe-NH2</td><td>Drug and cell delivery for tissue regenerative medicine</td><td>[75]</td></tr><tr><td>Oxidized xanthan/8-arm PEG-hydrazine</td><td>Chemotherapeutic agent delivery, cell therapy, and tissue engineering</td><td>[76]</td></tr><tr><td>Multi-hydrazide PEG/A-HA/Gelatin</td><td>Improving cell engraftment and retention</td><td>[77]</td></tr><tr><td>Pectin achlydrazide/Poly(DMA-stat-4-FPA)</td><td>Tissue engineering, drug delivery, and biosensors</td><td>[78]</td></tr></tbody></table>
<table><caption>**Table 11.** Degradable self-healing hydrogels and elastomers and their potential applications in the biomedical field.</caption><thead><tr><th>Mechanism</th><th>Nature</th><th>Materials</th><th>Potential biomedical applications</th><th>Ref.</th></tr></thead><tbody><tr><td rowspan="20">Schiff base</td><td rowspan="20">Hydrogel</td><td>Phenol-functionalized chitosan/PEG-BA</td><td>3D Bioprinting</td><td>[40]</td></tr><tr><td>Glycolated chitosan/PEG-BA</td><td>Tissue engineering, vascular repair</td><td>[42]</td></tr><tr><td>Chitosan-g-aniline/PEG-BA</td><td>Cell delivery</td><td>[33]</td></tr><tr><td>CMC/PEG-BA</td><td>Hemostasis</td><td>[43]</td></tr><tr><td>Quaternized methacryloyl chitosan/PEG-BA</td><td>Wound healing</td><td>[44]</td></tr><tr><td>Zwitterionic L-GA acid-functionalized chitosan/PEG-BA</td><td>Delivery</td><td>[45]</td></tr><tr><td>CEC/PEG-BA/Polyacrylamide</td><td>Tissue engineering, actuators, wearable devices</td><td>[46]</td></tr><tr><td>CEC/A-HA</td><td>Drug delivery for tumor therapy</td><td>[36]</td></tr><tr><td>N-succinyl-chitosan/Chondroitin sulfate multiple aldehydes</td><td>Cell delivery and tissue engineering</td><td>[31]</td></tr><tr><td>CEC/A-Alg</td><td>Cell delivery</td><td>[32]</td></tr><tr><td>CMC/Aldehyde-functionalized carboxymethyl cellulose</td><td>Cell delivery</td><td>[37]</td></tr><tr><td>Glycolated chitosan/A-Dex</td><td>Protein release</td><td>[35]</td></tr><tr><td>CMC/A-Dex</td><td>Prevention and treatment of postoperative adhesion</td><td>[48]</td></tr><tr><td>CEC/Dextran-g-aniline tetramer-g-4-formylbenzoic acid</td><td>Tissue engineering and cell delivery</td><td>[34]</td></tr><tr><td>Acrylamide modified chitin/A-Alg</td><td>Drug delivery and tissue engineering</td><td>[38]</td></tr><tr><td>CMC/A-Alg/Hap/CaCO3</td><td>Tissue engineering and drug delivery</td><td>[50]</td></tr><tr><td>CMC/A-Alg/MGMs</td><td>Tissue engineering and drug delivery</td><td>[51]</td></tr><tr><td>L-arginine conjugated chitosan/PEG-BA/pDA-NPs</td><td>Wound dressing</td><td>[54]</td></tr><tr><td>CMC/Poly(dextran-g-4-formylbenzoic acid)/Peptide nanofibers</td><td>Wound healing</td><td>[55]</td></tr><tr><td>Ethylenediamine functional gelatin/Dialdehyde carboxymethylcellulose</td><td>Drug and cell delivery and tissue engineering</td><td>[57]</td></tr><tr><td>Aldehyde-modified xanthan gum/Phosphatidylethanolamine liposomes</td><td>Cell delivery and tissue engineering</td><td>[58]</td></tr><tr><td rowspan="12">Boronate ester bonds</td><td rowspan="12">Hydrogel</td><td>Poly(methacryloyloxyethyl phosphorylcholine) functionalized with benzoxaborole or catechol pendant groups</td><td>Drug delivery and tissue engineering</td><td>[61]</td></tr><tr><td>5-methacrylamido-1,2-benzoxaborole/3-gluconamidoprpyl methacrylamide/Acrylamide</td><td>Bioprinting and cell encapsulation and delivery</td><td>[62]</td></tr><tr><td>Nopoldiol/2-gluconamidoethylmethacrylamide/PEG methyl ether methacrylate/MAABO</td><td>Gene delivery, cell therapy, and tissue engineering</td><td>[63]</td></tr><tr><td>Poly(DMA-st-MAABO)/2-lactobionamidoethyl methacrylamide, exhibiting galactose residues on the surface</td><td>Tissue engineering and drug delivery</td><td>[64]</td></tr><tr><td>Dopamine functionalized 4-armed PEG/Phenylboronic acid modified 4-armed PEG</td><td>Bioadhesion, drug delivery, and tissue engineering</td><td>[65]</td></tr><tr><td>PVA/Poly(aspartic acid) derivatives with boronic acid groups</td><td>Wound dressing, tissue repairing</td><td>[66]</td></tr><tr><td>PVA/Carboxyethyl cellulose grafted with phenylboronic acid</td><td>Tissue engineering, wound healing, and drug delivery</td><td>[68]</td></tr><tr><td>Methacrylated hyaluronic acid/3-aminophenylboronic acid-modified sodium alginate</td><td>Cell delivery and tissue regeneration</td><td>[69]</td></tr><tr><td>PVA/Poly(acrylamide-co-dopamine methacrylamide) bis(phenylboronic acid carbamoyl) cystamine</td><td>Drug delivery in cancer treatment</td><td>[70]</td></tr><tr><td>CTL/Boric acid/Mannitol</td><td>Tissue engineering cartilage regeneration</td><td>[71]</td></tr><tr><td rowspan="7">Hydrazone bonds</td><td rowspan="7">Hydrogel</td><td>Tetraphenylethylene-poly(DMA-stat-diacetone acrylamide)2/Acylhydrazide functionalized pectin</td><td>Drug delivery and tissue engineering</td><td>[72]</td></tr><tr><td>Pectin aldehyde/Poly(NIPAM-stat-acylhydrazide)</td><td>Anticancer drug carrier</td><td>[73]</td></tr><tr><td>A-HA/Hydrazide-HA</td><td>RNA delivery</td><td>[74]</td></tr><tr><td>Adipic dihydrazide-grafted carboxyethyl chitin/PEG-BA Phe-NH2</td><td>Drug and cell delivery for tissue regenerative medicine</td><td>[75]</td></tr><tr><td>Oxidized xanthan/8-arm PEG-hydrazine</td><td>Chemotherapeutic agent delivery, cell therapy, and tissue engineering</td><td>[76]</td></tr><tr><td>Multi-hydrazide PEG/A-HA/Gelatin</td><td>Improving cell engraftment and retention</td><td>[77]</td></tr><tr><td>Pectin achlydrazide/Poly(DMA-stat-4-FPA)</td><td>Tissue engineering, drug delivery, and biosensors</td><td>[78]</td></tr></tbody></table>
<table><caption>Table 11. Continued.</caption><thead><tr><th>Mechanism</th><th>Nature</th><th>Materials</th><th>Potential biomedical applications</th><th>Ref.</th></tr></thead><tbody><tr><td rowspan="2">Disulfide bonds</td><td rowspan="2">Hydrogel</td><td>A-HA/3,3'-dithiobis(propionic hydrazide)</td><td>Tissue adhesive</td><td>[79]</td></tr><tr><td>P(NIPAM-FPA-DMA)/3,3'-dithiobis (propionohydrazide)<br>PEO dihydrazide</td><td>Drug delivery, tissue engineering, cell culture</td><td>[6]</td></tr><tr><td rowspan="2">Hydrogen bonds</td><td rowspan="2">Hydrogel</td><td>PAEH/Dialdehyde-PEG</td><td>Tissue repairing and drug release</td><td>[80]</td></tr><tr><td>BSA</td><td>Tissue engineering and 3D bioprinting</td><td>[81]</td></tr><tr><td rowspan="5">Hydrophobic interactions</td><td rowspan="5">Hydrogel</td><td>Cytosine and guanosine-modified HA/1,6-hexamethylenediamine</td><td>Drug delivery, tissue engineering, cell scaffold materials, and regenerative medicine</td><td>[91]</td></tr><tr><td>BSA/Epychlorhydrin</td><td>Drug delivery</td><td>[92]</td></tr><tr><td>PGS with UPy</td><td>Tissue engineering and drug delivery</td><td>[98]</td></tr><tr><td>PSeD with UPy</td><td>Tissue engineering</td><td>[99]</td></tr><tr><td>Ferrocene-modified chitosan</td><td>Drug delivery</td><td>[102]</td></tr><tr><td rowspan="3">Host-guest interactions</td><td rowspan="3">Hydrogel</td><td>Alginate/Micelles of regenerated silk fibroin incorporating stearyl methacrylate</td><td>Drug delivery and tissue engineering</td><td>[103]</td></tr><tr><td>4-arm poly(ethylene glycol)-b-poly(γ-c-nitrobenzyl-L-glutamate)</td><td>Drug delivery</td><td>[101]</td></tr><tr><td>HA-Ad/HA-CD</td><td>Drug and cell delivery</td><td>[104]</td></tr><tr><td rowspan="4">Electrostatic/Ionic interactions</td><td rowspan="4">Hydrogel</td><td>β-cyclodextrin-modified PGA/Cholesterol-modified triblock<br>PGA-b-PEG-b-PGA</td><td>Tissue engineering and drug delivery</td><td>[106]</td></tr><tr><td>β-CD-modified alginate/Adamantine-modified graphene oxide</td><td>Tissue engineering</td><td>[105]</td></tr><tr><td>Poly(CD) Acrylamide/N-vinyl-pyrrolidinone</td><td>Drug delivery</td><td>[107]</td></tr><tr><td>Chitosan/Carboxymethyl cellulose</td><td>Drug delivery</td><td>[108]</td></tr><tr><td rowspan="2">Metal-coordination</td><td rowspan="2">Hydrogel</td><td>PCL diol modified with IPDI before chain extension with anionic oligo-alginate and cationic N-methyldiethanolamine</td><td>Tissue engineering</td><td>[109]</td></tr><tr><td>CMC/EDTA:Fe³⁺/Silver nanoparticles</td><td>Wound healing</td><td>[112]</td></tr><tr><td rowspan="5">Combination of dynamic chemical bonds</td><td rowspan="5">Hydrogel</td><td>CMC/(Fe³⁺/Al³⁺)</td><td>Tissue engineering</td><td>[113]</td></tr><tr><td>CMC/PCAD/Al³⁺</td><td>Tissue sealant</td><td>[114]</td></tr><tr><td>Polyaspartamide/histamine conjugate with/Cu2+</td><td>Adhesive</td><td>[115]</td></tr><tr><td>Thiolated BSA/Silver nitrate</td><td>Tissue engineering</td><td>[116]</td></tr><tr><td>Gelatin/A-Alg/Adipic acid dihydrazide</td><td>Tissue engineering and drug delivery</td><td>[118]</td></tr><tr><td rowspan="2">Combination of physical interactions</td><td rowspan="2">Hydrogel</td><td>Gelatin/A-Dex/Adipic acid dihydrazide/bFGF@PLGA/Chlorhexidine acetate</td><td>Drug delivery and wound healing</td><td>[119]</td></tr><tr><td>Polyglycerol sebacate/PEG methyl ether methacrylate/α-CD</td><td>Drug delivery</td><td>[124]</td></tr><tr><td rowspan="5">Combination of dynamic chemical bonds and physical interactions</td><td rowspan="5">Elastomer</td><td>Denatured BSA</td><td>Drug delivery</td><td>[125]</td></tr><tr><td>Gelatin methacrylate/2-(2-methoxyethoxy)ethyl methacrylate</td><td>Cell culture, cell delivery, and tissue engineering</td><td>[126]</td></tr><tr><td>Gelatin/PEDOT:PSS/MWCNTs-COOH</td><td>Wound healing</td><td>[127]</td></tr><tr><td>Poly(glycerol sebacate)-co-poly(ethylene glycol)-g-catechol/Fe³⁺/<br>UPy-functional gelatin</td><td>Wound healing and wound closure</td><td>[132]</td></tr><tr><td>κ-carrageenan/PAA/Zr⁴⁺/D-galactose/Micelles of n-octadecyl acrylate dissolved in sodium dodecyl sulfate</td><td>Tissue engineering</td><td>[134]</td></tr><tr><td rowspan="2">Hydrogel</td><td rowspan="2">Hydrogel</td><td>Polyurethane (PCL/HDI)/Sulfated alginate</td><td>Tissue engineering</td><td>[130]</td></tr><tr><td>Quaternized chitosan/Benzaldehyde-terminated Pluronic micelles</td><td>Wound healing</td><td>[136]</td></tr><tr><td></td><td></td><td></td><td>Wound dressing</td><td>[137]</td></tr></tbody></table>
<table><caption>Table 11. Continued.</caption><thead><tr><th>Mechanism</th><th>Nature</th><th>Materials</th><th>Potential biomedical applications</th><th>Ref.</th></tr></thead><tbody><tr><td rowspan="2">Disulfide bonds</td><td rowspan="2">Hydrogel</td><td>A-HA/3,3'-dithiobis(propionic hydrazide)</td><td>Tissue adhesive</td><td>[79]</td></tr><tr><td>P(NIPAM-FPA-DMA)/3,3'-dithiobis (propionohydrazide)<br>PEO dihydrazide</td><td>Drug delivery, tissue engineering, cell culture</td><td>[6]</td></tr><tr><td rowspan="2">Hydrogen bonds</td><td rowspan="2">Hydrogel</td><td>PAEH/Dialdehyde-PEG</td><td>Tissue repairing and drug release</td><td>[80]</td></tr><tr><td>BSA</td><td>Tissue engineering and 3D bioprinting</td><td>[81]</td></tr><tr><td rowspan="4">Hydrophobic<br>interactions</td><td rowspan="4">Hydrogel</td><td>Cytosine and guanosine-modified HA/1,6-hexamethylenediamine</td><td>Drug delivery, tissue engineering, cell scaffold materials, and regenerative medicine</td><td>[91]</td></tr><tr><td>BSA/Epychlorhydrin</td><td>Drug delivery</td><td>[92]</td></tr><tr><td>PGS with UPy</td><td>Tissue engineering and drug delivery</td><td>[98]</td></tr><tr><td>PSeD with UPy</td><td>Tissue engineering</td><td>[99]</td></tr><tr><td rowspan="2">Hydrophobic<br>interactions</td><td rowspan="2">Hydrogel</td><td>Ferrocene-modified chitosan</td><td>Drug delivery</td><td>[102]</td></tr><tr><td>Alginate/Micelles of regenerated silk fibroin incorporating stearyl methacrylate</td><td>Drug delivery and tissue engineering</td><td>[103]</td></tr><tr><td rowspan="2">Host-guest<br>interactions</td><td rowspan="2">Hydrogel</td><td>4-arm poly(ethylene glycol)-b-poly(γ-c-nitrobenzyl-L-glutamate)</td><td>Drug delivery</td><td>[101]</td></tr><tr><td>HA-Ad/HA-CD</td><td>Drug and cell delivery</td><td>[104]</td></tr><tr><td rowspan="3">Electrostatic/Ionic<br>interactions</td><td rowspan="3">Hydrogel</td><td>β-cyclodextrin-modified PGA/Cholesterol-modified triblock<br>PGA-b-PEG-b-PGA</td><td>Tissue engineering and drug delivery</td><td>[106]</td></tr><tr><td>β-CD-modified alginate/Adamantine-modified graphene oxide</td><td>Tissue engineering</td><td>[105]</td></tr><tr><td>Poly(CD) Acrylamide/N-vinyl-pyrrolidinone</td><td>Drug delivery</td><td>[107]</td></tr><tr><td rowspan="2">Hydrogel</td><td rowspan="2">Elastomer</td><td>Chitosan/Carboxymethyl cellulose</td><td>Drug delivery</td><td>[108]</td></tr><tr><td>PCL diol modified with IPDI before chain extension with anionic oligo-alginate and cationic N-methyldiethanolamine</td><td>Tissue engineering</td><td>[109]</td></tr><tr><td rowspan="5">Metal-coordination</td><td rowspan="5">Hydrogel</td><td>CMC/EDTA:Fe³⁺/Silver nanoparticles</td><td>Wound healing</td><td>[112]</td></tr><tr><td>CMC/(Fe³⁺/Al³⁺)</td><td>Tissue engineering</td><td>[113]</td></tr><tr><td>CMC/PCAD/Al³⁺</td><td>Tissue sealant</td><td>[114]</td></tr><tr><td>Polyaspartamide/histamine conjugate with/Cu2+</td><td>Adhesive</td><td>[115]</td></tr><tr><td>Thiolated BSA/Silver nitrate</td><td>Tissue engineering</td><td>[116]</td></tr><tr><td rowspan="2">Hydrogel</td><td rowspan="2">Elastomer</td><td>Gelatin/A-Alg/Adipic acid dihydrazide</td><td>Tissue engineering and drug delivery</td><td>[118]</td></tr><tr><td>Gelatin/A-Dex/Adipic acid dihydrazide/bFGF@PLGA/Chlorhexidine<br>acetate</td><td>Drug delivery and wound healing</td><td>[119]</td></tr><tr><td rowspan="4">Hydrogel</td><td rowspan="4">Elastomer</td><td>Polyglycerol sebacate/PEG methyl ether methacrylate/α-CD</td><td>Drug delivery</td><td>[124]</td></tr><tr><td>Denatured BSA</td><td>Drug delivery</td><td>[125]</td></tr><tr><td>Gelatin methacrylate/2-(2-methoxyethoxy)ethyl methacrylate</td><td>Cell culture, cell delivery, and tissue engineering</td><td>[126]</td></tr><tr><td>Gelatin/PEDOT:PSS/MWCNTs-COOH</td><td>Wound healing</td><td>[127]</td></tr><tr><td rowspan="2">Hydrogel</td><td rowspan="2">Elastomer</td><td>Poly(glycerol sebacate)-co-poly(ethylene glycol)-g-catechol/Fe³⁺/<br>UPy-functional gelatin</td><td>Wound healing and wound closure</td><td>[132]</td></tr><tr><td>κ-carrageenan/PAA/Zr⁴⁺/D-galactose/Micelles of n-octadecyl acrylate<br>dissolved in sodium dodecyl sulfate</td><td>Tissue engineering</td><td>[134]</td></tr><tr><td rowspan="2">Hydrogel</td><td rowspan="2">Elastomer</td><td>Polyurethane (PCL/HDI)/Sulfated alginate</td><td>Tissue engineering</td><td>[130]</td></tr><tr><td>Quaternized chitosan/Benzaldehyde-terminated Pluronic micelles</td><td>Wound healing</td><td>[136]</td></tr><tr><td rowspan="2">Hydrogel</td><td rowspan="2">Elastomer</td><td>Quaternized chitosan/Benzaldehyde-terminated Pluronic micelles</td><td>Wound dressing</td><td>[137]</td></tr><tr><td>Quaternized chitosan/Benzaldehyde-terminated Pluronic micelles</td><td>Wound dressing</td><td>[137]</td></tr></tbody></table>
<table><caption>**Table 11. Continued.**</caption><thead><tr><th>Mechanism</th><th>Nature</th><th>Materials</th><th>Potential biomedical applications</th><th>Ref.</th></tr></thead><tbody><tr><td></td><td></td><td>Catechol modified poly(L-lysine)/A-Dex/Fe</td><td>Wound healing</td><td>[138]</td></tr><tr><td></td><td></td><td>Catechol-modified oxidized HA/Aminated gelatin/Fe<sup>3+</sup></td><td>Wound dressing</td><td>[139]</td></tr><tr><td></td><td></td><td>Chitosan/Vanillin</td><td>Tissue engineering and wound repairing</td><td>[141]</td></tr><tr><td></td><td></td><td>Poly(DMA-st-META-st-LAEMA)/Poly (MPC-st-UPyHEMA-st-MABBO)</td><td>Cell and tissue engineering</td><td>[143]</td></tr><tr><td></td><td></td><td>Furfural-pectin/6-maleidocaproic acid chitosan</td><td>Drug delivery</td><td>[145]</td></tr><tr><td></td><td></td><td>PVA/Borax/Cellulose nanofibers/Dopamine-grafted oxidized carbomethyl cellulose</td><td>Wound healing</td><td>[147]</td></tr><tr><td></td><td></td><td>BOC/Gelatin/BOC-doped polypyrrole</td><td>Treatment of traumatic spinal cord injury</td><td>[148]</td></tr><tr><td></td><td></td><td>Guar gum-g-poly(acrylamidoglycolic acid)/Silver nitrate/Sodium borohydride</td><td>Wound healing</td><td>[150]</td></tr></tbody></table>

A-Alg, aldehyde modified alginate; A-Dex, aldehyde modified dextran; A-HA, aldehyde modified hyaluronic acid; Ad, adamantane; bFGF@PLGA, PLGA microspheres integrated with basic fibroblast growth factor; BOC, borax-functionalized oxidized chondroitin sulfate; BSA, bovine serum albumin; CD, cyclodextrin; CEC, carboxyethyl chitosan; CMC, carboxymethyl chitosan; CTL, lactose modified chitosan; DMA, N,N-dimethylacrylamide; FPA, formylphenyl acrylate; HA, hyaluronic acid; HAp, hydroxyapatite; HDI, hexamethylene diisocyanate; IPDI, isophorone diisocyanate; MAABO, 5-methacrylamido-1,2-benzoxaborole; MGMs, magnetic gelatin microspheres: MWCNTs, multiwalled carbon nanotubes; NIPAM, N-isopropylacrylamide; PAA, poly(acrylic acid); PAEH, poly(aspartic acid) derivatives with hydrazide functional groups; PCAD, photoluminescent citric acid derivatives; PCL, polycaprolactone; pDa-NPs, polydopamine nanoparticles; PEG, poly(ethylene glycol), PEG-BA, benzaldehyde-poly(ethylene glycol); PEO, poly(ethylene oxide); PGS, poly(glycerol-co-sebacate); PSeD, poly(sebacoyl diglyceride); PVA, poly(vinyl alcohol); UPy, 2-ureido-4-[1H]-pyrimidinone.

then reform into solid gels in situ. In most studies discussed in this review, these hydrogels are presented as potential candidates for use in biomedical applications, such as drug delivery, encapsulation of cells or proteins, tissue engineering, and wound dressing (Figure 10). However, these assertions have not been demonstrated and are only described as perspectives. In this section, we focus on the systems discussed in this review, for which proofs of concept for specific medical uses were clearly reported.

### 7.1.1. Self-Healable Hydrogels for Use in Wound Dressing

Numerous self-healing hydrogels have been applied in wound healing.[44,48,54,55,66,69,113,119,127,132,136–138,147,150] The skin acts as a protective barrier, and rapidly repairing skin damage that may be caused by trauma or burns or various diseases is crucial. Numerous hydrogels exhibit antibacterial properties, e.g., Ling et al. produced a self-healing hydrogel with antibacterial and angiogenic activities to accelerate skin wound healing.[54] The addition of pDA-NPs to chitosan/PEG-BA gels improved the inhibition of bacterial growth and enhanced angiogenesis. When evaluated in a rat model, wound healing was accelerated compared to that in the control group (99% versus 79%). In another example, Shen et al. reported that their poly(aspartic acid)-based system exhibited accelerated wound repair and improved skin cell proliferation owing to the prevention of bacterial infections in mice in vivo.[66] Hydrogels composed of N,O-CMC and A-Dex were fabricated as wound dressings and for use in preventing and treating postoperative adhesions.[48] The occurrence of postoperative adhesions in rat models with damaged sidewalls was significantly reduced (10%) in vivo, and furthermore, these gels could prevent infection and reduce inflammation. Qiu et al. developed chitosan-based hydrogels containing peptide nanofibers to heal chronic wounds.[55] This system exhibited antibacterial and hemostatic properties, which enabled the elimination of bacterial biofilms and induced blood-

cell and platelet aggregation. Antibacterial adhesive injectable hydrogels were also designed as wound dressings for use in joint skin wound healing,[136] and the quaternized chitosan/benzaldehyde-terminated Pluronic system was evaluated in vivo. Accelerated wound healing due to curcumin-loaded hydrogels was revealed by higher granulation tissue thickness, collagen deposition, and upregulated vascular endothelial growth factor. Wang et al. developed a self-healing antibacterial hydrogel composed of quaternized methacryloyl chitosan and PEG-BA for use in wound treatment.[44] The gel displayed anti-infection properties and promoted the healing of infected wounds in vivo. Antimicrobial hydrogels reinforced with cellulose were

<figure><img src="image_12.png" /><figcaption>**Figure 10.** Potential biomedical applications of degradable self-healing hydrogels and elastomers.</figcaption></figure>
<table><caption>**Table 11.** Continued.</caption><thead><tr><th>Mechanism</th><th>Nature</th><th>Materials</th><th>Potential biomedical applications</th><th>Ref.</th></tr></thead><tbody><tr><td></td><td></td><td>Catechol modified poly(L-lysine)/A-Dex/Fe</td><td>Wound healing</td><td>[138]</td></tr><tr><td></td><td></td><td>Catechol-modified oxidized HA/Aminated gelatin/Fe<sup>3+</sup></td><td>Wound dressing</td><td>[139]</td></tr><tr><td></td><td></td><td>Chitosan/Vanillin</td><td>Tissue engineering and wound repairing</td><td>[141]</td></tr><tr><td></td><td></td><td>Poly(DMA-st-META-st-LAEMA)/Poly (MPC-st-UPyHEMA-st-MABBO)</td><td>Cell and tissue engineering</td><td>[143]</td></tr><tr><td></td><td></td><td>Furfural-pectin/6-maleidocaproic acid chitosan</td><td>Drug delivery</td><td>[145]</td></tr><tr><td></td><td></td><td>PVA/Borax/Cellulose nanofibers/Dopamine-grafted oxidized carbomethyl cellulose</td><td>Wound healing</td><td>[147]</td></tr><tr><td></td><td></td><td>BOC/Gelatin/BOC-doped polypyrrole</td><td>Treatment of traumatic spinal cord injury</td><td>[148]</td></tr><tr><td></td><td></td><td>Guar gum-g-poly(acrylamidoglycolic acid)/Silver nitrate/Sodium borohydride</td><td>Wound healing</td><td>[150]</td></tr></tbody></table>

A-Alg, aldehyde modified alginate; A-Dex, aldehyde modified dextran; A-HA, aldehyde modified hyaluronic acid; Ad, adamantane; bFGF@PLGA, PLGA microspheres integrated with basic fibroblast growth factor; BOC, borax-functionalized oxidized chondroitin sulfate; BSA, bovine serum albumin; CD, cyclodextrin; CEC, carboxyethyl chitosan; CMC, carboxymethyl chitosan; CTL, lactose modified chitosan; DMA, N,N-dimethylacrylamide; FPA, formylphenyl acrylate; HA, hyaluronic acid; HAp, hydroxyapatite; HDI, hexamethylene diisocyanate; IPDI, isophorone diisocyanate; MAABO, 5-methacrylamido-1,2-benzoxaborole; MGMs, magnetic gelatin microspheres: MWCNTs, multiwalled carbon nanotubes; NIPAM, N-isopropylacrylamide; PAA, poly(acrylic acid); PAEH, poly(aspartic acid) derivatives with hydrazide functional groups; PCAD, photoluminescent citric acid derivatives; PCL, polycaprolactone; pDa-NPs, polydopamine nanoparticles; PEG, poly(ethylene glycol), PEG-BA, benzaldehyde-poly(ethylene glycol); PEO, poly(ethylene oxide); PGS, poly(glycerol-co-sebacate); PSeD, poly(sebacoyl diglyceride); PVA, poly(vinyl alcohol); UPy, 2-ureido-4[1H]-pyrimidinone.

then reform into solid gels in situ. In most studies discussed in this review, these hydrogels are presented as potential candidates for use in biomedical applications, such as drug delivery, encapsulation of cells or proteins, tissue engineering, and wound dressing (Figure 10). However, these assertions have not been demonstrated and are only described as perspectives. In this section, we focus on the systems discussed in this review, for which proofs of concept for specific medical uses were clearly reported.

### 7.1.1. Self-Healable Hydrogels for Use in Wound Dressing

Numerous self-healing hydrogels have been applied in wound healing.[44,48,54,55,66,69,113,119,127,132,136–138,147,150] The skin acts as a protective barrier, and rapidly repairing skin damage that may be caused by trauma or burns or various diseases is crucial. Numerous hydrogels exhibit antibacterial properties, e.g., Ling et al. produced a self-healing hydrogel with antibacterial and angiogenic activities to accelerate skin wound healing.[54] The addition of pDA-NPs to chitosan/PEG-BA gels improved the inhibition of bacterial growth and enhanced angiogenesis. When evaluated in a rat model, wound healing was accelerated compared to that in the control group (99% versus 79%). In another example, Shen et al. reported that their poly(aspartic acid)-based system exhibited accelerated wound repair and improved skin cell proliferation owing to the prevention of bacterial infections in mice in vivo.[66] Hydrogels composed of N,O-CMC and A-Dex were fabricated as wound dressings and for use in preventing and treating postoperative adhesions.[48] The occurrence of postoperative adhesions in rat models with damaged sidewalls was significantly reduced (10%) in vivo, and furthermore, these gels could prevent infection and reduce inflammation. Qiu et al. developed chitosan-based hydrogels containing peptide nanofibers to heal chronic wounds.[55] This system exhibited antibacterial and hemostatic properties, which enabled the elimination of bacterial biofilms and induced blood-

cell and platelet aggregation. Antibacterial adhesive injectable hydrogels were also designed as wound dressings for use in joint skin wound healing,[136] and the quaternized chitosan/benzaldehyde-terminated Pluronic system was evaluated in vivo. Accelerated wound healing due to curcumin-loaded hydrogels was revealed by higher granulation tissue thickness, collagen deposition, and upregulated vascular endothelial growth factor. Wang et al. developed a self-healing antibacterial hydrogel composed of quaternized methacryloyl chitosan and PEG-BA for use in wound treatment.[44] The gel displayed anti-infection properties and promoted the healing of infected wounds in vivo. Antimicrobial hydrogels reinforced with cellulose were

<figure><img src="image_12.png" /><figcaption>**Figure 10.** Potential biomedical applications of degradable self-healing hydrogels and elastomers.</figcaption></figure>
prepared using PVA and borax.[147] They were impregnated with neomycin and appeared to be effective against numerous bacteria occurring in wounds. Chen et al. produced self-healing hydrogels for use in the sequential delivery of growth factors and antibacterial agents to heal wounds.[119] The gels were composed of aminated gelatin, adipic acid dihydrazide, and oxidized dextran and embedded with bFGF@PLGA microspheres and CHA. In vivo studies of the wounded skins of rats confirmed that the gels could prevent infection due to the delivery of CHA. Moreover, cell proliferation and wound healing could be accelerated via the release of bFGF. Another study reported conductive adhesive hydrogel wound dressings for use in the photothermal therapy of infected skin wounds.[137] The hydrogels were based on chitosan, benzaldehyde-terminated Pluronic, and CNTs and exhibited photothermal antibacterial properties in vivo. They could significantly improve wound healing, collagen deposition, and angiogenesis. To continue with photothermal antibacterial activity, Guo et al. developed self-healing hydrogel adhesives for use in wound closure or healing.[132] In vivo studies revealed that these gels could effectively kill methicillin-resistant S. aureus, exhibited good hemostatic properties, and enabled superior wound closure and healing of skin incisions compared to that observed using medical glue or surgical sutures. Self-healable, electroactive hydrogels based on gelatin were designed by Zheng et al. for use in motion sensing and acceleration of skin wound recovery via electrical stimulation.[127] An in vivo study performed using a full-thickness rat skin defect model highlighted a faster wound recovery with wound area reduction, granulation tissue formation, collagen deposition improvement, vascularization, and re-epithelialization. Finally, Palem et al. reported guar gum-grafted-poly(acrylamidoglycolic acid)-based hydrogels containing nano-silver, which appeared to accelerate wound healing.[150]

### 7.1.2. Self-Healable Hydrogels for Use in Drug Delivery

Drug delivery also benefits from the design of self-healing hydrogels.[36,50,51,73,74,76,78,80,92,100,101,108,118,124] Several methods may be used to control drug release, and two systems that involve magnetic fields have been reported. A self-healing chitosan-alginate hydrogel incorporating MGMs was designed for use in drug delivery.[51] When the gels were exposed to an external magnetic field, the drug release exhibited no clear difference in the early stage but significantly increased after the fifth day to reach 75% (compared to 60% without a magnetic field). In the second example, hydrogels were prepared using chitosan and cellulose and impregnated with magnetic chitosan microspheres.[108] Under an external magnetic field, the release increased to 91% compared to 62% without stimulation.

Self-healing hydrogels have been studied for use in anticancer treatment, and drug release is often controlled by pH. Bilalis et al. used a polypeptide system that could release gemcitabine to treat pancreatic cancer.[100] The injectable hydrogel was responsive to pH, and thus, only melted close to the pancreatic cancer cells, which enabled a targeted release of the drug toward cancerous tissues. An injectable BSA hydrogel was developed by Upadhyay et al. for use in controlled drug release toward cancer cells.[92] Once loaded with doxorubicin (Dox), the

hydrogels could kill 70-80% of cancer cells. Sharma et al. prepared xanthan-PEG hydrogels loaded with Dox that exhibited pH-responsive drug release.[76] Drug delivery was accelerated at tumoral pH compared to that at physiological pH. Polysaccharide-based hydrogels designed by Qian et al. also demonstrated pH-responsive Dox release in PBS at various pH values.[36] Dox was released more rapidly under acidic conditions than in an alkaline medium, and the drug-loaded gels induced the deaths of HeLa cancer cells. In another example, pectin aldehyde was used in combination with an acylhydrazide-functionalized polymer to prepare an injectable system with anticancer drug-releasing properties.[73] Gels loaded with Dox and combretastatin A4 disodium phosphate affected CT26 tumors in vivo. Zhao et al. reported photodegradable self-healing hydrogels composed of PEG and polypeptides.[101] This system exhibited Dox release under UV irradiation due to the presence of photolabile o-nitrobenzyl ester groups. Increasing the amount of Dox released improved the apoptosis ratio of the HeLa cells. Another example of a Dox-loaded hydrogel is a poly(glycerol sebacate)-based system, which was reported as an injectable matrix.[124] Dox release was studied in vitro and exhibited a biphasic profile. The first phase was driven by diffusion, whereas the second phase was driven by hydrogel erosion.

The last two systems were designed to release proteins or small interfering ribonucleic acids (siRNAs). Cho et al. reported an injectable system prepared using glycol chitosan and A-Dex, and revealed that the release kinetics of BSA could be tuned by adjusting the needle size upon injection.[35] This could be explained by the modulation of the sizes and shapes of the fragmented hydrogels after passing through the needles. Wang et al. designed hydrogels for use in siRNA sequestration and initiated delivery to the heart.[74] Their system released siRNA on demand in response to proteolytic activity after myocardial infarction.

### 7.1.3. Self-Healable Hydrogels for Use in Cell Encapsulation and Delivery

Various self-healable hydrogels have been used in cell encapsulation and delivery,[31-33,58,61,69,75-77,126,135] e.g., a self-healing conductive injectable hydrogel based on chitosan and PEG-BA was prepared as a cell-delivery carrier for use in cardiac cell therapy.[33] The conductivities of the hydrogels were similar to that of native cardiac tissue. C2C12 myoblasts proliferated in the gels, and their viabilities were maintained after injection. Moreover, C2C12 myoblasts and H9c2 cardiac cells could be released at tunable rates. Yang et al. studied chitin hydrogels that could support the proliferation and multipotent differentiation of rat bone marrow-derived stem cells, and these hydrogels could be used as 3D scaffolds in stem-cell encapsulation and release.[75] Ma et al. developed an injectable polymer-liposome hydrogel as a cell carrier, wherein encapsulated cells were viable for an extended duration.[58] Several systems for use in cell retention have also been identified. Recently, stem cell transplantation emerged as a promising method of treating numerous diseases and injuries. However, although cell transplantation via injection is the preferred method of targeted cell delivery, this therapy results in low cell retention and engraftment. Cai et al.
prepared using PVA and borax.[147] They were impregnated with neomycin and appeared to be effective against numerous bacteria occurring in wounds. Chen et al. produced self-healing hydrogels for use in the sequential delivery of growth factors and antibacterial agents to heal wounds.[119] The gels were composed of aminated gelatin, adipic acid dihydrazide, and oxidized dextran and embedded with bFGF@PLGA microspheres and CHA. In vivo studies of the wounded skins of rats confirmed that the gels could prevent infection due to the delivery of CHA. Moreover, cell proliferation and wound healing could be accelerated via the release of bFGF. Another study reported conductive adhesive hydrogel wound dressings for use in the photothermal therapy of infected skin wounds.[137] The hydrogels were based on chitosan, benzaldehyde-terminated Pluronic, and CNTs and exhibited photothermal antibacterial properties in vivo. They could significantly improve wound healing, collagen deposition, and angiogenesis. To continue with photothermal antibacterial activity, Guo et al. developed self-healing hydrogel adhesives for use in wound closure or healing.[132] In vivo studies revealed that these gels could effectively kill methicillin-resistant S. aureus, exhibited good hemostatic properties, and enabled superior wound closure and healing of skin incisions compared to that observed using medical glue or surgical sutures. Self-healable, electroactive hydrogels based on gelatin were designed by Zheng et al. for use in motion sensing and acceleration of skin wound recovery via electrical stimulation.[127] An in vivo study performed using a full-thickness rat skin defect model highlighted a faster wound recovery with wound area reduction, granulation tissue formation, collagen deposition improvement, vascularization, and re-epithelialization. Finally, Palem et al. reported guar gum-grafted-poly(acrylamidoglycolic acid)-based hydrogels containing nano-silver, which appeared to accelerate wound healing.[150]

### 7.1.2. Self-Healable Hydrogels for Use in Drug Delivery

Drug delivery also benefits from the design of self-healing hydrogels.[36,50,51,73,74,76,78,80,92,100,101,108,118,124] Several methods may be used to control drug release, and two systems that involve magnetic fields have been reported. A self-healing chitosan-alginate hydrogel incorporating MGMs was designed for use in drug delivery.[51] When the gels were exposed to an external magnetic field, the drug release exhibited no clear difference in the early stage but significantly increased after the fifth day to reach 75% (compared to 60% without a magnetic field). In the second example, hydrogels were prepared using chitosan and cellulose and impregnated with magnetic chitosan microspheres.[108] Under an external magnetic field, the release increased to 91% compared to 62% without stimulation.

Self-healing hydrogels have been studied for use in anticancer treatment, and drug release is often controlled by pH. Bilalis et al. used a polypeptide system that could release gemcitabine to treat pancreatic cancer.[100] The injectable hydrogel was responsive to pH, and thus, only melted close to the pancreatic cancer cells, which enabled a targeted release of the drug toward cancerous tissues. An injectable BSA hydrogel was developed by Upadhyay et al. for use in controlled drug release toward cancer cells.[92] Once loaded with doxorubicin (Dox), the

hydrogels could kill 70-80% of cancer cells. Sharma et al. prepared xanthan-PEG hydrogels loaded with Dox that exhibited pH-responsive drug release.[76] Drug delivery was accelerated at tumoral pH compared to that at physiological pH. Polysaccharide-based hydrogels designed by Qian et al. also demonstrated pH-responsive Dox release in PBS at various pH values.[36] Dox was released more rapidly under acidic conditions than in an alkaline medium, and the drug-loaded gels induced the deaths of HeLa cancer cells. In another example, pectin aldehyde was used in combination with an acylhydrazide-functionalized polymer to prepare an injectable system with anticancer drug-releasing properties.[73] Gels loaded with Dox and combretastatin A4 disodium phosphate affected CT26 tumors in vivo. Zhao et al. reported photodegradable self-healing hydrogels composed of PEG and polypeptides.[101] This system exhibited Dox release under UV irradiation due to the presence of photolabile o-nitrobenzyl ester groups. Increasing the amount of Dox released improved the apoptosis ratio of the HeLa cells. Another example of a Dox-loaded hydrogel is a poly(glycerol sebacate)-based system, which was reported as an injectable matrix.[124] Dox release was studied in vitro and exhibited a biphasic profile. The first phase was driven by diffusion, whereas the second phase was driven by hydrogel erosion.

The last two systems were designed to release proteins or small interfering ribonucleic acids (siRNAs). Cho et al. reported an injectable system prepared using glycol chitosan and A-Dex, and revealed that the release kinetics of BSA could be tuned by adjusting the needle size upon injection.[35] This could be explained by the modulation of the sizes and shapes of the fragmented hydrogels after passing through the needles. Wang et al. designed hydrogels for use in siRNA sequestration and initiated delivery to the heart.[74] Their system released siRNA on demand in response to proteolytic activity after myocardial infarction.

### 7.1.3. Self-Healable Hydrogels for Use in Cell Encapsulation and Delivery

Various self-healable hydrogels have been used in cell encapsulation and delivery,[31–33,58,61,69,75–77,126,135] e.g., a self-healing conductive injectable hydrogel based on chitosan and PEG-BA was prepared as a cell-delivery carrier for use in cardiac cell therapy.[33] The conductivities of the hydrogels were similar to that of native cardiac tissue. C2C12 myoblasts proliferated in the gels, and their viabilities were maintained after injection. Moreover, C2C12 myoblasts and H9c2 cardiac cells could be released at tunable rates. Yang et al. studied chitin hydrogels that could support the proliferation and multipotent differentiation of rat bone marrow-derived stem cells, and these hydrogels could be used as 3D scaffolds in stem-cell encapsulation and release.[75] Ma et al. developed an injectable polymer-liposome hydrogel as a cell carrier, wherein encapsulated cells were viable for an extended duration.[58] Several systems for use in cell retention have also been identified. Recently, stem cell transplantation emerged as a promising method of treating numerous diseases and injuries. However, although cell transplantation via injection is the preferred method of targeted cell delivery, this therapy results in low cell retention and engraftment. Cai et al.
studied injectable hydrogels that could resolve the problem of cell death after transplantation.[135] Their system, which was composed of a peptide-PEG copolymer assembled with a recombinant protein and PNIPAM chains conjugated to PEG, could protect the cells from mechanically disruptive forces during needle injection and provide a suitable 3D microenvironment for cell survival and retention. Moreover, in another system, hydrogels could enhance the survival, attachment, and engraftment of cells due to rapid sol-gel transitions.[77] Another hydrogel was prepared by Ren et al. to modulate the retention of transplanted cells.[126] As in the previous example, stem cells delivered within the gel were well protected from mechanical damage, which led to improved cell retention for three weeks in vivo.

### 7.1.4. Self-Healable Hydrogels for Use in Tissue Engineering and Tissue Adhesion

Self-healing hydrogels may be useful in bone tissue engineering. An injectable polysaccharide hydrogel loaded with HAp and calcium carbonate was developed for use in the regeneration of irregular bone defects.[50] Liu et al. prepared a protein-based hydrogel that was injected into large cranial defects of rabbits as a scaffold with osteoinductive and osteoconductive activities.[116] These in vivo studies revealed improved and accelerated bone repair. Another system was designed to repair the central nervous system.[41] The growth of neurosphere-like progenitors was accelerated within the hydrogels, and they could differentiate into neuron-like cells. The injection of this hydrogel enabled an 81% recovery of neural development. Hsieh et al. designed a biodegradable self-healing chitosan-fibrin hydrogel to induce blood-capillary formation,[42] and vascular endothelial cells seeded in this hydrogel could form capillary-like structures. Luo et al. reported the development of an injectable, electroconductive hydrogel to enhance tissue repair after traumatic spinal cord injury by bridging cavity spaces.[148] The gels exhibited mechanical and conductive properties similar to those of natural spinal cord tissues and could promote neuronal differentiation, improve axon outgrowth, and inhibit astrocyte differentiation in vitro. Significant locomotor functions were restored in rats with spinal cord injuries in vivo.

Finally, the development of self-healing hydrogels as tissue adhesives remains in progress. Identifying an acceptable compromise between the adhesive strength, mechanical properties, and degradation rate is often challenging. Xu et al. fabricated an injectable, self-healing, multi-responsive hydrogel as an on-demand dissolution tissue adhesive.[79] The system based on A-HA and a disulfide-containing cross-linker was compared to the commercial tissue adhesive BioGlue and exhibited an increased lap shear strength on porcine skin (up to 120 kPa, which is 65.8% higher than that of BioGlue). Li et al. developed another bioadhesive hydrogel that could efficiently close an open wound and enable post-wound-closure care.[138] In this case, the gel was composed of catechol-modified poly(lysine) and oxidized dextran and it enabled post-closure care after efficiently closing the skin incisions, as the bioadhesive could be easily removed and reapplied to the reopened wounds.

## 7.2. Self-Healable Elastomers in Biomedical Applications

Although less common than hydrogels, degradable, biocompatible elastomers are useful in the field of biomedical engineering because they mimic the mechanical properties of numerous soft tissues while resisting multiple mechanical stresses (Table 11). They are accepted by host tissues and resorb over time, thus avoiding possible inflammation and complications. However, once the chemically cross-linked elastomer is damaged, it cannot be repaired and thus may lead to complications in the healing of the host tissue. Therefore, the development of self-healing elastomers is necessary for use in biomedical applications, particularly in complex, dynamic in vivo environments.

Wu et al. synthesized PGS and grafted a UPy unit onto PGS (PGS-U) via HDI groups.[98] They focused on the elaboration of multifunctional 3D materials to deliver drugs, yield an antibacterial surface, and enable a cellular tri-culture. For drug delivery, 5-aminosalicylic acid (5-ASA) was incorporated into a PGS-4U film (ratio of 4/10 between the UPy and PGS units). A sandwich material consisting of the superposition of two layers without embedded 5-ASA and one layer with 5-ASA displayed linear release for 18 d. In parallel, antibacterial poly-L-lysine-UPy was physically associated with the PGS-U films via hydrogen bonding using drop casting. The antibacterial activity of the modified PGS-U against Gram-negative *E. coli* was confirmed, with a bactericidal ratio of up to 100% and no observable colony-forming units. Finally, to reconstruct tissues at the interface between the ligament, cartilage, and bone, PGS-U was used as a co-culture system for L929 fibroblasts, chondrocytes, and MC3T3 osteoblasts. Cells were cultured on independent films before exploiting the self-healing properties of PGS-U in pressing these films together at 37 °C for 1 min to generate a single culture surface.

Daemi et al. used a natural polymer based on alginate with cationic PU to achieve a tensile strength of 10 MPa for vascular applications.[109] Cytocompatibility was confirmed in vitro using L929 fibroblasts and human umbilical vein endothelial cells, whereas a mild inflammatory response 20 d post-implantation was observed in vivo in rat tissues, with the formation of new vessels. In addition, the same group modified its initial cationic polymer to decrease its toughness (61 MJ m⁻³).[130] With the highest amount of sulfated alginate, it displayed superior anticoagulation activity and blood compatibility and no cytotoxicity or chronic inflammation after in vivo implantation into Wistar rats for 8 weeks. In vivo degradation following subcutaneous implantation resulted in the remaining mass decreasing to 85% after 8 weeks.

The last example concerns tissue reconstruction (soft or hard). Jiang et al. designed self-healable elastomers based on PUs with dual-dynamic bonds (oxime-urethane and hydrogen bonds).[146] After demonstrating their viabilities, they were applied to 1) limit the progression of aortic aneurysms, 2) heal the sciatic nerve, and 3) immobilize the sternum (Figure 11). The assays were performed in vivo in mice using soft or hard elastomers, depending on the ratio of glycerol used in their preparation. First, to limit the progression of aortic aneurysms, arteries were wrapped with elastomers, and the optimal in vivo results were obtained using elastomers with *E* values similar to that of the vascular tissue (1 MPa). Second, in nerve reconstruction, a softer elastomer with a lower glycerol content
studied injectable hydrogels that could resolve the problem of cell death after transplantation.[135] Their system, which was composed of a peptide-PEG copolymer assembled with a recombinant protein and PNIPAM chains conjugated to PEG, could protect the cells from mechanically disruptive forces during needle injection and provide a suitable 3D microenvironment for cell survival and retention. Moreover, in another system, hydrogels could enhance the survival, attachment, and engraftment of cells due to rapid sol-gel transitions.[77] Another hydrogel was prepared by Ren et al. to modulate the retention of transplanted cells.[126] As in the previous example, stem cells delivered within the gel were well protected from mechanical damage, which led to improved cell retention for three weeks in vivo.

### 7.1.4. Self-Healable Hydrogels for Use in Tissue Engineering and Tissue Adhesion

Self-healing hydrogels may be useful in bone tissue engineering. An injectable polysaccharide hydrogel loaded with HAp and calcium carbonate was developed for use in the regeneration of irregular bone defects.[50] Liu et al. prepared a protein-based hydrogel that was injected into large cranial defects of rabbits as a scaffold with osteoinductive and osteoconductive activities.[116] These in vivo studies revealed improved and accelerated bone repair. Another system was designed to repair the central nervous system.[41] The growth of neurosphere-like progenitors was accelerated within the hydrogels, and they could differentiate into neuron-like cells. The injection of this hydrogel enabled an 81% recovery of neural development. Hsieh et al. designed a biodegradable self-healing chitosan-fibrin hydrogel to induce blood-capillary formation,[42] and vascular endothelial cells seeded in this hydrogel could form capillary-like structures. Luo et al. reported the development of an injectable, electroconductive hydrogel to enhance tissue repair after traumatic spinal cord injury by bridging cavity spaces.[148] The gels exhibited mechanical and conductive properties similar to those of natural spinal cord tissues and could promote neuronal differentiation, improve axon outgrowth, and inhibit astrocyte differentiation in vitro. Significant locomotor functions were restored in rats with spinal cord injuries in vivo.

Finally, the development of self-healing hydrogels as tissue adhesives remains in progress. Identifying an acceptable compromise between the adhesive strength, mechanical properties, and degradation rate is often challenging. Xu et al. fabricated an injectable, self-healing, multi-responsive hydrogel as an on-demand dissolution tissue adhesive.[79] The system based on A-HA and a disulfide-containing cross-linker was compared to the commercial tissue adhesive BioGlue and exhibited an increased lap shear strength on porcine skin (up to 120 kPa, which is 65.8% higher than that of BioGlue). Li et al. developed another bioadhesive hydrogel that could efficiently close an open wound and enable post-wound-closure care.[138] In this case, the gel was composed of catechol-modified poly(lysine) and oxidized dextran and it enabled post-closure care after efficiently closing the skin incisions, as the bioadhesive could be easily removed and reapplied to the reopened wounds.

## 7.2. Self-Healable Elastomers in Biomedical Applications

Although less common than hydrogels, degradable, biocompatible elastomers are useful in the field of biomedical engineering because they mimic the mechanical properties of numerous soft tissues while resisting multiple mechanical stresses (Table 11). They are accepted by host tissues and resorb over time, thus avoiding possible inflammation and complications. However, once the chemically cross-linked elastomer is damaged, it cannot be repaired and thus may lead to complications in the healing of the host tissue. Therefore, the development of self-healing elastomers is necessary for use in biomedical applications, particularly in complex, dynamic in vivo environments.

Wu et al. synthesized PGS and grafted a UPy unit onto PGS (PGS-U) via HDI groups.[98] They focused on the elaboration of multifunctional 3D materials to deliver drugs, yield an antibacterial surface, and enable a cellular tri-culture. For drug delivery, 5-aminosalicylic acid (5-ASA) was incorporated into a PGS-4U film (ratio of 4/10 between the UPy and PGS units). A sandwich material consisting of the superposition of two layers without embedded 5-ASA and one layer with 5-ASA displayed linear release for 18 d. In parallel, antibacterial poly-L-lysine-UPy was physically associated with the PGS-U films via hydrogen bonding using drop casting. The antibacterial activity of the modified PGS-U against Gram-negative *E. coli* was confirmed, with a bactericidal ratio of up to 100% and no observable colony-forming units. Finally, to reconstruct tissues at the interface between the ligament, cartilage, and bone, PGS-U was used as a co-culture system for L929 fibroblasts, chondrocytes, and MC3T3 osteoblasts. Cells were cultured on independent films before exploiting the self-healing properties of PGS-U in pressing these films together at 37 °C for 1 min to generate a single culture surface.

Daemi et al. used a natural polymer based on alginate with cationic PU to achieve a tensile strength of 10 MPa for vascular applications.[109] Cytocompatibility was confirmed in vitro using L929 fibroblasts and human umbilical vein endothelial cells, whereas a mild inflammatory response 20 d post-implantation was observed in vivo in rat tissues, with the formation of new vessels. In addition, the same group modified its initial cationic polymer to decrease its toughness (61 MJ m⁻³).[130] With the highest amount of sulfated alginate, it displayed superior anticoagulation activity and blood compatibility and no cytotoxicity or chronic inflammation after in vivo implantation into Wistar rats for 8 weeks. In vivo degradation following subcutaneous implantation resulted in the remaining mass decreasing to 85% after 8 weeks.

The last example concerns tissue reconstruction (soft or hard). Jiang et al. designed self-healable elastomers based on PUs with dual-dynamic bonds (oxime-urethane and hydrogen bonds).[146] After demonstrating their viabilities, they were applied to 1) limit the progression of aortic aneurysms, 2) heal the sciatic nerve, and 3) immobilize the sternum (Figure 11). The assays were performed in vivo in mice using soft or hard elastomers, depending on the ratio of glycerol used in their preparation. First, to limit the progression of aortic aneurysms, arteries were wrapped with elastomers, and the optimal in vivo results were obtained using elastomers with *E* values similar to that of the vascular tissue (1 MPa). Second, in nerve reconstruction, a softer elastomer with a lower glycerol content
<figure><img src="image_13.png" /><figcaption>**Figure 11.** 1) Schematic diagram of chemical structure formula, tunable mechanical and self-healing properties, and in vivo biomedical applications. 2) Sternum immobilization by elastomer in rat. 3) Gross observations of aorta with cross-linked self-healing elastomer (SHE). The blood vessel pointed by the black arrow is the abdominal aorta, scale bar = 1 mm. 4. Paper-cut simulation (clip) of sciatic nerve coaptation. Adapted with permission.[146] Copyright 2021, Nature.</figcaption></figure>

was used. A film of this elastomer was cut in half to yield ends with castle-like shapes, and each nerve end was wrapped in the film. Once the ends were joined, they displayed a weld that was three- to four-fold faster than sutures. This material exhibited superior repair results than those obtained using sutures, fibrin glue, and PCL conduit groups. Finally, a hard elastomer (with a high glycerol content) was used in bone repair after fracture. After 6 weeks, the fracture was repaired without a pneumothorax, and the results were identical to those obtained using the steel wire control.

**8. Concluding Remarks, Opportunities, and Challenges**

In this review, we analyzed the latest examples of self-healable degradable hydrogels and elastomers for use in biomedical applications, including tissue engineering, drug delivery, and wound repair. The nature of the dynamic cross-linking, cova-lent or physical, was described, in addition to the conditions/stimuli required to induce self-healing (e.g., heat and pH). We focused on the most recent degradable networks that combine at least two self-healing mechanisms, in particular. To provide a useful compilation for readers and aid them in selecting potential strategies for their requirements, we also analyzed the mechanical properties, repair time, and SHE. These properties are summarized in the diagrams shown in **Figure 12** and pre-sented as functions of the natures of the self-healing mecha-nisms of the degradable networks analyzed. These diagrams

confirm that most of the networks exhibit high SHEs of >75%, regardless of the type of self-healing. The time required to obtain a high SHE highlights the superiority of networks with multiple self-healing mechanisms. Their repair times, although similar to those observed with dynamic chemical bonds, are clearly superior to those of self-healing networks with only physical interactions, which display a wide range of repair times (Figure 12.1). Furthermore, a large range of mechanical properties, from a few pascals to a few tens of megapascals, may be realized using degradable self-healing networks based on physical interactions (Figure 12.2), as this class of biomate-rials includes most elastomers listed in this review. In compar-ison, the networks with self-healing based on chemical bonds discussed in this review are mostly characterized by moduli of ≈ 10 kPa, whereas the moduli of the networks with dual self-healing mechanisms lie in between. Based on the systems ana-lyzed here, a contradiction between mechanical strength and self-healing behavior is apparent. This may be due to the nature of the polymeric chains used in the various systems. Most self-healing systems are designed with hydrophilic polymers, which leads to a majority of self-healable hydrogels and a minority of self-healable elastomers. Consequently, most of the analyzed self-healing networks display poor mechanical properties. However, when considering only the elastomers, self-healing properties may also be imparted to stronger materials (e.g., E > 150 MPa for systems based on PCL and DA reactions [86,87]).

As shown during this review, self-healable degradable/bio-eliminable networks open large avenues in designing smart materials for use in biomedical applications and should
<figure><img src="image_13.png" /><figcaption>**Figure 11.** 1) Schematic diagram of chemical structure formula, tunable mechanical and self-healing properties, and in vivo biomedical applications. 2) Sternum immobilization by elastomer in rat. 3) Gross observations of aorta with cross-linked self-healing elastomer (SHE). The blood vessel pointed by the black arrow is the abdominal aorta, scale bar = 1 mm. 4. Paper-cut simulation (clip) of sciatic nerve coaptation. Adapted with permission.[146] Copyright 2021, Nature.</figcaption></figure>

was used. A film of this elastomer was cut in half to yield ends with castle-like shapes, and each nerve end was wrapped in the film. Once the ends were joined, they displayed a weld that was three- to four-fold faster than sutures. This material exhibited superior repair results than those obtained using sutures, fibrin glue, and PCL conduit groups. Finally, a hard elastomer (with a high glycerol content) was used in bone repair after fracture. After 6 weeks, the fracture was repaired without a pneumothorax, and the results were identical to those obtained using the steel wire control.

**8. Concluding Remarks, Opportunities, and Challenges**

In this review, we analyzed the latest examples of self-healable degradable hydrogels and elastomers for use in biomedical applications, including tissue engineering, drug delivery, and wound repair. The nature of the dynamic cross-linking, cova-lent or physical, was described, in addition to the conditions/stimuli required to induce self-healing (e.g., heat and pH). We focused on the most recent degradable networks that combine at least two self-healing mechanisms, in particular. To provide a useful compilation for readers and aid them in selecting potential strategies for their requirements, we also analyzed the mechanical properties, repair time, and SHE. These properties are summarized in the diagrams shown in **Figure 12** and pre-sented as functions of the natures of the self-healing mecha-nisms of the degradable networks analyzed. These diagrams

confirm that most of the networks exhibit high SHEs of >75%, regardless of the type of self-healing. The time required to obtain a high SHE highlights the superiority of networks with multiple self-healing mechanisms. Their repair times, although similar to those observed with dynamic chemical bonds, are clearly superior to those of self-healing networks with only physical interactions, which display a wide range of repair times (Figure 12.1). Furthermore, a large range of mechanical properties, from a few pascals to a few tens of megapascals, may be realized using degradable self-healing networks based on physical interactions (Figure 12.2), as this class of biomate-rials includes most elastomers listed in this review. In compar-ison, the networks with self-healing based on chemical bonds discussed in this review are mostly characterized by moduli of ≈ 10 kPa, whereas the moduli of the networks with dual self-healing mechanisms lie in between. Based on the systems ana-lyzed here, a contradiction between mechanical strength and self-healing behavior is apparent. This may be due to the nature of the polymeric chains used in the various systems. Most self-healing systems are designed with hydrophilic polymers, which leads to a majority of self-healable hydrogels and a minority of self-healable elastomers. Consequently, most of the analyzed self-healing networks display poor mechanical properties. However, when considering only the elastomers, self-healing properties may also be imparted to stronger materials (e.g., E > 150 MPa for systems based on PCL and DA reactions [86,87]).

As shown during this review, self-healable degradable/bio-eliminable networks open large avenues in designing smart materials for use in biomedical applications and should
<figure><img src="image_14.png" /><figcaption>**Figure 12.** Compared properties of the self-healable degradable networks presented in the review as a function of the nature of the self-healing mechanism.</figcaption></figure>

continue to attract attention due to their unique features, which may be summarized as follows (Figure 13). First, the range of mechanical properties that may be reached covers most of the requirements in soft-tissue engineering and repair, with moduli ranging from a few pascals for the softest hydrogels to a few tens of megapascals for elastomers. This opens opportunities to develop networks mimicking the extracellular matrix for use in cell culturing and delivery or engineer vessels, muscles, ligaments, etc., as only bones and their mechanical properties are currently out of reach for the reported materials. Second, the healing times may be as fast as a few seconds, depending on the reversible bonds and the mobility of the polymer chains. Clearly, this time also depends on the parameter that is retained to evaluate self-healing, as the time required for visual repair may be drastically shorter than the time required for the recovery of the mechanical properties. A third advantage with respect to their biomedical applications is their degradability, which may exhibit large variations in degradation time, from hours to months, depending on the individual components used: hydrophilic or hydrophobic chains, biopolymers or synthetic chains, and hydrolytic or enzymatic degradation. Finally, the intrinsic multifunctionality of these systems renders them particularly attractive for use in drug delivery because the interactions responsible for the dynamic network may be exploited to increase the loading and retention of drugs and biologics (e.g., ionic interactions and hydrogen bonding).

However, these clear opportunities should not conceal the challenges that remain (Figure 13). The first challenge is the design of stronger networks. As shown in this review, hydrogels reign supreme in the field, but still lack resistance to mechanical stress and do not cover the entire spectrum of soft tissue reconstruction. Self-healable degradable elastomers are thus natural candidates to address this limitation, but very few studies have been conducted to date, as most efforts in the field focus on bio-based, recyclable, reprocessable elastomers that are not intended for biomedical use. This is particularly true for vitrimers, which are permanent chemical networks with dynamic covalent bonds that enable the network to change its topology while maintaining a constant number of chemical bonds. Although self-healable by nature, and in several cases based on the same dynamic bonds (e.g., imines and boronic esters), vitrimers are still rarely designed for use in biomedical

applications.[151] In addition, combining covalent bonds and dynamic/reversible chemical bonds or noncovalent interactions/physical bonds may be investigated to provide self-healable, strong, degradable networks. Although the replacement of covalent bonds with dynamic bonds in networks diminishes their inherent mechanical strengths,[152] a sound ratio between the two types of bonds should ensure the mechanical properties while guaranteeing a degree of self-healing. Another challenge is the generation of additional characteristics along with self-healing. Microbial resistance, conductivity, and anisotropic and adhesive properties are examples of the characteristics necessary for developing a novel generation of smart biomaterials. Although the examples analyzed in this review illustrate these aspects, they remain rare and require more complex formulations that may hinder their translation to products or the use of potentially toxic precursors, threatening their biocompatibility. Another challenge lies in the definition of self-healing networks. Although dynamic bonds are the key prerequisite in self-healing networks, several composite materials based on non-dynamic cross-linked networks and thermoplastic chains

<figure><img src="image_15.png" /><figcaption>**Figure 13.** Summary of the opportunities and challenges associated with self-healable degradable networks for biomedical applications.</figcaption></figure>
<figure><img src="image_14.png" /><figcaption>**Figure 12.** Compared properties of the self-healable degradable networks presented in the review as a function of the nature of the self-healing mechanism.</figcaption></figure>

continue to attract attention due to their unique features, which may be summarized as follows (Figure 13). First, the range of mechanical properties that may be reached covers most of the requirements in soft-tissue engineering and repair, with moduli ranging from a few pascals for the softest hydrogels to a few tens of megapascals for elastomers. This opens opportunities to develop networks mimicking the extracellular matrix for use in cell culturing and delivery or engineer vessels, muscles, ligaments, etc., as only bones and their mechanical properties are currently out of reach for the reported materials. Second, the healing times may be as fast as a few seconds, depending on the reversible bonds and the mobility of the polymer chains. Clearly, this time also depends on the parameter that is retained to evaluate self-healing, as the time required for visual repair may be drastically shorter than the time required for the recovery of the mechanical properties. A third advantage with respect to their biomedical applications is their degradability, which may exhibit large variations in degradation time, from hours to months, depending on the individual components used: hydrophilic or hydrophobic chains, biopolymers or synthetic chains, and hydrolytic or enzymatic degradation. Finally, the intrinsic multifunctionality of these systems renders them particularly attractive for use in drug delivery because the interactions responsible for the dynamic network may be exploited to increase the loading and retention of drugs and biologics (e.g., ionic interactions and hydrogen bonding).

However, these clear opportunities should not conceal the challenges that remain (Figure 13). The first challenge is the design of stronger networks. As shown in this review, hydrogels reign supreme in the field, but still lack resistance to mechanical stress and do not cover the entire spectrum of soft tissue reconstruction. Self-healable degradable elastomers are thus natural candidates to address this limitation, but very few studies have been conducted to date, as most efforts in the field focus on bio-based, recyclable, reprocessable elastomers that are not intended for biomedical use. This is particularly true for vitrimers, which are permanent chemical networks with dynamic covalent bonds that enable the network to change its topology while maintaining a constant number of chemical bonds. Although self-healable by nature, and in several cases based on the same dynamic bonds (e.g., imines and boronic esters), vitrimers are still rarely designed for use in biomedical

applications.[151] In addition, combining covalent bonds and dynamic/reversible chemical bonds or noncovalent interactions/physical bonds may be investigated to provide self-healable, strong, degradable networks. Although the replacement of covalent bonds with dynamic bonds in networks diminishes their inherent mechanical strengths,[152] a sound ratio between the two types of bonds should ensure the mechanical properties while guaranteeing a degree of self-healing. Another challenge is the generation of additional characteristics along with self-healing. Microbial resistance, conductivity, and anisotropic and adhesive properties are examples of the characteristics necessary for developing a novel generation of smart biomaterials. Although the examples analyzed in this review illustrate these aspects, they remain rare and require more complex formulations that may hinder their translation to products or the use of potentially toxic precursors, threatening their biocompatibility. Another challenge lies in the definition of self-healing networks. Although dynamic bonds are the key prerequisite in self-healing networks, several composite materials based on non-dynamic cross-linked networks and thermoplastic chains

<figure><img src="image_15.png" /><figcaption>**Figure 13.** Summary of the opportunities and challenges associated with self-healable degradable networks for biomedical applications.</figcaption></figure>
have been claimed as self-healable biomaterials. The use of free PCL chains that may flow at temperatures above their melting temperatures through the damaged network and promote the partial recovery of the mechanical and structural properties upon cooling is, for example, reported for degradable self-healable elastomers.[153] The extension of self-healing to hydrogels obtained via classical electrostatic or hydrophobic interactions is also questionable, as any shear-thinning hydrogel based on such mechanisms (e.g., alginate/Ca²⁺ gels or PLA-PEG-PLA triblock copolymer gels) could be considered as a self-healable biomaterial. Finally, the evaluation of self-healing properties is crucial in terms of two aspects. First, the self-healing capacities of the final devices or scaffolds are generally not evaluated. The possibility of processing 3D self-healable biomaterials has been investigated in previous studies using techniques such as 3D printing/bioprinting,[154] particle leaching,[155] freeze-drying,[156] thermally induced phase separation,[157] gas foaming,[158] solvent casting,[159] nanofiber self-assembly,[160] and fiber-based techniques.[161] However, the self-healing properties of 3D biomaterials have not been studied and should be investigated. Second, as shown in this review, various methods are used to estimate the repair time and SHE (Figure 3). Wide discrepancies remain in practice, rendering comparisons of the performances of materials challenging. This also renders the performance of any trend analysis, which is required for the future sound design of self-healable biomaterials dedicated to specific, demanding biomedical applications, challenging. More standardized methods (e.g., via a normative framework) to evaluate and quantify the self-healing properties to categorize self-healable materials are necessary to ensure further development and potential transfer to the market of this class of innovative biomaterials. In this sense, efforts to model the mechanics of self-healing polymer networks cross-linked via dynamic bonds may enable the more rational design of these systems. This may also enable comparison between them by offering tools to predict healing times and SHEs based on macromolecular and reactivity considerations.[27] Therefore, this aspect should be considered in optimizing self-healing degradable networks.

## Acknowledgements

The authors thank the French Agency for Research (ANR) for support of LG through ANR2016-BIOSCAFF (ANR-16-CE09-0024) and of MG through ANR2019-OPENN (ANR-19-CE19-0022-02).

## Conflict of Interest

The authors declare no conflict of interest.

## Keywords

biomedical, degradable networks, elastomers, hydrogels, self-healing

Received: May 10, 2022

Revised: December 11, 2022

Published online: January 25, 2023

[1] S. Utrera-Barrios, R. Verdejo, M. A. López-Manchado, M. Hernández Santana, *Mater. Horiz.* **2020**, *7*, 2882.

[2] S. R. White, N. R. Sottos, P. H. Geubelle, J. S. Moore, M. R. Kessler, S. R. Sriram, E. N. Brown, S. Viswanathan, *Nature* **2001**, *409*, 794.

[3] M. Deng, F. Guo, D. Liao, Z. Hou, Y. Li, *Polym. Chem.* **2018**, *9*, 98.

[4] P. Wang, L. Yang, B. Dai, Z. Yang, S. Guo, G. Gao, L. Xu, M. Sun, K. Yao, J. Zhu, *Eur. Polym. J.* **2020**, *123*, 109382.

[5] S.-M. Kim, H. Jeon, S.-H. Shin, S.-A. Park, J. Jegal, S. Y. Hwang, D. X. Oh, J. Park, *Adv. Mater.* **2018**, *30*, 1705145.

[6] H. An, K. Xu, L. Chang, Y. Wang, J. Qin, W. Li, *Polymer* **2018**, *147*, 38.

[7] N. Kuhl, S. Bode, R. K. Bose, J. Vitz, A. Seifert, S. Hoeppener, S. J. Garcia, S. Spange, S. van der Zwaag, M. D. Hager, U. S. Schubert, *Adv. Funct. Mater.* **2015**, *25*, 3295.

[8] Y. Song, Y. Liu, T. Qi, G. L. Li, *Angew. Chem.* **2018**, *130*, 14034.

[9] A. J. R. Amaral, G. Pasparakis, *Polym. Chem.* **2017**, *8*, 6464.

[10] X. Wu, J. Li, G. Li, L. Ling, G. Zhang, R. Sun, C.-P. Wong, *J. Appl. Polym. Sci.* **2018**, *135*, 46532.

[11] Y. Peng, Y. Yang, Q. Wu, S. Wang, G. Huang, J. Wu, *Polymer* **2018**, *157*, 172.

[12] D. K. Smith, *Chem. Soc. Rev.* **2009**, *38*, 684.

[13] D. Ma, Y. Wang, W. Dai, *Mater. Sci. Eng. C* **2018**, *89*, 456.

[14] S. Doppalapudi, A. Jain, W. Khan, A. J. Domb, *Polym. Adv. Technol.* **2014**, *25*, 427.

[15] R. Muthuraj, M. Misra, A. K. Mohanty, *J. Appl. Polym. Sci.* **2018**, *135*, 45726.

[16] Medical Polymers Market Growth & Trends, https://www.grand-viewresearch.com/press-release/global-medical-polymer-market, accessed: March, 2022.

[17] Q. Chen, S. Liang, G. A. Thomas, *Prog. Polym. Sci.* **2013**, *38*, 584.

[18] C. Mangeon, E. Renard, F. Thevenieau, V. Langlois, *Mater. Sci. Eng. C* **2017**, *80*, 760.

[19] Y. Tu, N. Chen, C. Li, H. Liu, R. Zhu, S. Chen, Q. Xiao, J. Liu, S. Ramakrishna, L. He, *Acta Biomater.* **2019**, *90*, 1.

[20] N. Roy, B. Bruchmann, J.-M. Lehn, *Chem. Soc. Rev.* **2015**, *44*, 3786.

[21] L. Zhai, A. Narkar, K. Ahn, *Nano Today* **2020**, *30*, 100826.

[22] D. G. Bekas, K. Tsirka, D. Baltzis, A. S. Paipetis, *Compos. Part B Eng.* **2016**, *87*, 92.

[23] R. V. S. P. Sanka, B. Krishnakumar, Y. Leterrier, S. Pandey, S. Rana, V. Michaud, *Front. Mater* **2019**, *6*, 137.

[24] D. L. Taylor, M. in het Panhuis, *Adv. Mater.* **2016**, *28*, 9060.

[25] X. Yan, F. Wang, B. Zheng, F. Huang, *Chem. Soc. Rev.* **2012**, *41*, 6042.

[26] B. Cheng, X. Lu, J. Zhou, R. Qin, Y. Yang, *ACS Sustainable Chem. Eng.* **2019**, *7*, 4443.

[27] K. Yu, A. Xin, Q. Wang, *J Mech Phys Solids* **2018**, *121*, 409.

[28] S. Lyu, D. Untereker, *Int. J. Mol. Sci.* **2009**, *10*, 4033.

[29] E. Yousif, R. Haddad, *SpringerPlus* **2013**, *2*, 398.

[30] A. Banerjee, K. Chatterjee, G. Madras, *Mater. Sci. Technol.* **2014**, *30*, 567.

[31] S. Lü, C. Gao, X. Xu, X. Bai, H. Duan, N. Gao, C. Feng, Y. Xiong, M. Liu, *ACS Appl. Mater. Interfaces* **2015**, *7*, 13029.

[32] Z. Wei, J. Zhao, Y. M. Chen, P. Zhang, Q. Zhang, *Sci. Rep.* **2016**, *6*, 1.

[33] R. Dong, X. Zhao, B. Guo, P. X. Ma, *ACS Appl. Mater. Interfaces* **2016**, *8*, 17138.

[34] B. Guo, J. Qu, X. Zhao, M. Zhang, *Acta Biomater.* **2019**, *84*, 180.

[35] I. S. Cho, T. Ooya, *J. Biomater. Sci., Polym. Ed.* **2018**, *29*, 145.

[36] C. Qian, T. Zhang, J. Gravesande, C. Baysah, X. Song, J. Xing, *Int. J. Biol. Macromol.* **2019**, *123*, 140.

[37] J. Thomas, A. Sharma, V. Panwar, V. Chopra, D. Ghosh, *ACS Appl. Bio Mater* **2019**, *2*, 2013.

[38] F. Ding, S. Wu, S. Wang, Y. Xiong, Y. Li, B. Li, H. Deng, Y. Du, L. Xiao, X. Shi, *Soft Matter* **2015**, *11*, 3971.

[39] S. Maity, A. Chatterjee, N. Chakraborty, J. Ganguly, *New J. Chem.* **2018**, *42*, 5946.

[40] Y. Liu, C. W. Wong, S. W. Chang, S. Hui Hsu, *Acta Biomater.* **2021**, *122*, 211.
have been claimed as self-healable biomaterials. The use of free PCL chains that may flow at temperatures above their melting temperatures through the damaged network and promote the partial recovery of the mechanical and structural properties upon cooling is, for example, reported for degradable self-healable elastomers.[153] The extension of self-healing to hydrogels obtained via classical electrostatic or hydrophobic interactions is also questionable, as any shear-thinning hydrogel based on such mechanisms (e.g., alginate/Ca²⁺ gels or PLA-PEG-PLA triblock copolymer gels) could be considered as a self-healable biomaterial. Finally, the evaluation of self-healing properties is crucial in terms of two aspects. First, the self-healing capacities of the final devices or scaffolds are generally not evaluated. The possibility of processing 3D self-healable biomaterials has been investigated in previous studies using techniques such as 3D printing/bioprinting,[154] particle leaching,[155] freeze-drying,[156] thermally induced phase separation,[157] gas foaming,[158] solvent casting,[159] nanofiber self-assembly,[160] and fiber-based techniques.[161] However, the self-healing properties of 3D biomaterials have not been studied and should be investigated. Second, as shown in this review, various methods are used to estimate the repair time and SHE (Figure 3). Wide discrepancies remain in practice, rendering comparisons of the performances of materials challenging. This also renders the performance of any trend analysis, which is required for the future sound design of self-healable biomaterials dedicated to specific, demanding biomedical applications, challenging. More standardized methods (e.g., via a normative framework) to evaluate and quantify the self-healing properties to categorize self-healable materials are necessary to ensure further development and potential transfer to the market of this class of innovative biomaterials. In this sense, efforts to model the mechanics of self-healing polymer networks cross-linked via dynamic bonds may enable the more rational design of these systems. This may also enable comparison between them by offering tools to predict healing times and SHEs based on macromolecular and reactivity considerations.[27] Therefore, this aspect should be considered in optimizing self-healing degradable networks.

## Acknowledgements

The authors thank the French Agency for Research (ANR) for support of LG through ANR2016-BIOSCAFF (ANR-16-CE09-0024) and of MG through ANR2019-OPENN (ANR-19-CE19-0022-02).

## Conflict of Interest

The authors declare no conflict of interest.

## Keywords

biomedical, degradable networks, elastomers, hydrogels, self-healing

Received: May 10, 2022

Revised: December 11, 2022

Published online: January 25, 2023

[1] S. Utrera-Barrios, R. Verdejo, M. A. López-Manchado, M. Hernández Santana, *Mater. Horiz.* **2020**, *7*, 2882.

[2] S. R. White, N. R. Sottos, P. H. Geubelle, J. S. Moore, M. R. Kessler, S. R. Sriram, E. N. Brown, S. Viswanathan, *Nature* **2001**, *409*, 794.

[3] M. Deng, F. Guo, D. Liao, Z. Hou, Y. Li, *Polym. Chem.* **2018**, *9*, 98.

[4] P. Wang, L. Yang, B. Dai, Z. Yang, S. Guo, G. Gao, L. Xu, M. Sun, K. Yao, J. Zhu, *Eur. Polym. J.* **2020**, *123*, 109382.

[5] S.-M. Kim, H. Jeon, S.-H. Shin, S.-A. Park, J. Jegal, S. Y. Hwang, D. X. Oh, J. Park, *Adv. Mater.* **2018**, *30*, 1705145.

[6] H. An, K. Xu, L. Chang, Y. Wang, J. Qin, W. Li, *Polymer* **2018**, *147*, 38.

[7] N. Kuhl, S. Bode, R. K. Bose, J. Vitz, A. Seifert, S. Hoeppener, S. J. Garcia, S. Spange, S. van der Zwaag, M. D. Hager, U. S. Schubert, *Adv. Funct. Mater.* **2015**, *25*, 3295.

[8] Y. Song, Y. Liu, T. Qi, G. L. Li, *Angew. Chem.* **2018**, *130*, 14034.

[9] A. J. R. Amaral, G. Pasparakis, *Polym. Chem.* **2017**, *8*, 6464.

[10] X. Wu, J. Li, G. Li, L. Ling, G. Zhang, R. Sun, C.-P. Wong, *J. Appl. Polym. Sci.* **2018**, *135*, 46532.

[11] Y. Peng, Y. Yang, Q. Wu, S. Wang, G. Huang, J. Wu, *Polymer* **2018**, *157*, 172.

[12] D. K. Smith, *Chem. Soc. Rev.* **2009**, *38*, 684.

[13] D. Ma, Y. Wang, W. Dai, *Mater. Sci. Eng. C* **2018**, *89*, 456.

[14] S. Doppalapudi, A. Jain, W. Khan, A. J. Domb, *Polym. Adv. Technol.* **2014**, *25*, 427.

[15] R. Muthuraj, M. Misra, A. K. Mohanty, *J. Appl. Polym. Sci.* **2018**, *135*, 45726.

[16] Medical Polymers Market Growth & Trends, https://www.grand-viewresearch.com/press-release/global-medical-polymer-market, accessed: March, 2022.

[17] Q. Chen, S. Liang, G. A. Thomas, *Prog. Polym. Sci.* **2013**, *38*, 584.

[18] C. Mangeon, E. Renard, F. Thevenieau, V. Langlois, *Mater. Sci. Eng. C* **2017**, *80*, 760.

[19] Y. Tu, N. Chen, C. Li, H. Liu, R. Zhu, S. Chen, Q. Xiao, J. Liu, S. Ramakrishna, L. He, *Acta Biomater.* **2019**, *90*, 1.

[20] N. Roy, B. Bruchmann, J.-M. Lehn, *Chem. Soc. Rev.* **2015**, *44*, 3786.

[21] L. Zhai, A. Narkar, K. Ahn, *Nano Today* **2020**, *30*, 100826.

[22] D. G. Bekas, K. Tsirka, D. Baltzis, A. S. Paipetis, *Compos. Part B Eng.* **2016**, *87*, 92.

[23] R. V. S. P. Sanka, B. Krishnakumar, Y. Leterrier, S. Pandey, S. Rana, V. Michaud, *Front. Mater* **2019**, *6*, 137.

[24] D. L. Taylor, M. in het Panhuis, *Adv. Mater.* **2016**, *28*, 9060.

[25] X. Yan, F. Wang, B. Zheng, F. Huang, *Chem. Soc. Rev.* **2012**, *41*, 6042.

[26] B. Cheng, X. Lu, J. Zhou, R. Qin, Y. Yang, *ACS Sustainable Chem. Eng.* **2019**, *7*, 4443.

[27] K. Yu, A. Xin, Q. Wang, *J Mech Phys Solids* **2018**, *121*, 409.

[28] S. Lyu, D. Untereker, *Int. J. Mol. Sci.* **2009**, *10*, 4033.

[29] E. Yousif, R. Haddad, *SpringerPlus* **2013**, *2*, 398.

[30] A. Banerjee, K. Chatterjee, G. Madras, *Mater. Sci. Technol.* **2014**, *30*, 567.

[31] S. Lü, C. Gao, X. Xu, X. Bai, H. Duan, N. Gao, C. Feng, Y. Xiong, M. Liu, *ACS Appl. Mater. Interfaces* **2015**, *7*, 13029.

[32] Z. Wei, J. Zhao, Y. M. Chen, P. Zhang, Q. Zhang, *Sci. Rep.* **2016**, *6*, 1.

[33] R. Dong, X. Zhao, B. Guo, P. X. Ma, *ACS Appl. Mater. Interfaces* **2016**, *8*, 17138.

[34] B. Guo, J. Qu, X. Zhao, M. Zhang, *Acta Biomater.* **2019**, *84*, 180.

[35] I. S. Cho, T. Ooya, *J. Biomater. Sci., Polym. Ed.* **2018**, *29*, 145.

[36] C. Qian, T. Zhang, J. Gravesande, C. Baysah, X. Song, J. Xing, *Int. J. Biol. Macromol.* **2019**, *123*, 140.

[37] J. Thomas, A. Sharma, V. Panwar, V. Chopra, D. Ghosh, *ACS Appl. Bio Mater* **2019**, *2*, 2013.

[38] F. Ding, S. Wu, S. Wang, Y. Xiong, Y. Li, B. Li, H. Deng, Y. Du, L. Xiao, X. Shi, *Soft Matter* **2015**, *11*, 3971.

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<figure><img src="image_16.png" /></figure>

**Mathilde Grosjean** graduated as a chemical engineer from the National Graduate School of Chemistry of Montpellier in 2019. In 2020, she started a Ph.D. in the Department of Polymers for Health and Biomaterials of the Max Mousseron Biomolecules Institute (IBMM) in Montpellier, under the supervision of Prof. Benjamin Nottelet. Her Ph.D. project, funded by the French National Research Agency (ANR), aims to develop degradable biomaterials with actuation, bioadhesive, or self-healing properties for use in biomedical applications.

<figure><img src="image_17.png" /></figure>

**Louis Gangolphe** graduated as Chemical Engineer from CPE Lyon in 2016. In 2020, he received his Ph.D. in Biomaterials at the University of Montpellier and the University of Grenoble Alpes. After working on the design of photocross-linked degradable elastomers for tissue reconstruction, he gained experience in biotechnology with the use of natural polymers and microfluidic technology to develop liver organoids for drug metabolism evaluation. He recently started a new adventure by joining Linxens as R&D Materials Engineer. His research interests focus on the use of functional polymer materials and their processing for biotechnology and medical devices sectors.

<figure><img src="image_18.png" /></figure>

**Benjamin Nottelet** received his Ph.D. in Materials Chemistry from the University of Montpellier, France (2005). He was a post-doctoral fellow at ENSCM, France (2006) and in the Department of Pharmaceutics and Biopharmaceutics at the University of Geneva, Switzerland (2006–2008). In 2008, he joined the University of Montpellier and the Department of Artificial Biopolymers as an Associate Professor. He was appointed Full Professor in 2018 and is currently co-leader of the Department of Polymers for Health and Biomaterials of the IBMM. His research focuses on degradable and multifunctional polymers for biomedical applications in drug delivery, diagnostic, medical devices, and tissue engineering.
<figure><img src="image_16.png" /></figure>

**Mathilde Grosjean** graduated as a chemical engineer from the National Graduate School of Chemistry of Montpellier in 2019. In 2020, she started a Ph.D. in the Department of Polymers for Health and Biomaterials of the Max Mousseron Biomolecules Institute (IBMM) in Montpellier, under the supervision of Prof. Benjamin Nottelet. Her Ph.D. project, funded by the French National Research Agency (ANR), aims to develop degradable biomaterials with actuation, bioadhesive, or self-healing properties for use in biomedical applications.

<figure><img src="image_17.png" /></figure>

**Louis Gangolphe** graduated as Chemical Engineer from CPE Lyon in 2016. In 2020, he received his Ph.D. in Biomaterials at the University of Montpellier and the University of Grenoble Alpes. After working on the design of photocross-linked degradable elastomers for tissue reconstruction, he gained experience in biotechnology with the use of natural polymers and microfluidic technology to develop liver organoids for drug metabolism evaluation. He recently started a new adventure by joining Linxens as R&D Materials Engineer. His research interests focus on the use of functional polymer materials and their processing for biotechnology and medical devices sectors.

<figure><img src="image_18.png" /></figure>

**Benjamin Nottelet** received his Ph.D. in Materials Chemistry from the University of Montpellier, France (2005). He was a post-doctoral fellow at ENSCM, France (2006) and in the Department of Pharmaceutics and Biopharmaceutics at the University of Geneva, Switzerland (2006–2008). In 2008, he joined the University of Montpellier and the Department of Artificial Biopolymers as an Associate Professor. He was appointed Full Professor in 2018 and is currently co-leader of the Department of Polymers for Health and Biomaterials of the IBMM. His research focuses on degradable and multifunctional polymers for biomedical applications in drug delivery, diagnostic, medical devices, and tissue engineering.