Recent advances in smart responsive hydrogel microspheres for tissue regeneration: preparation, characteristics and applications
Zijian
Wu†
,
Guang
Shi†
,
Luhao
Li†
,
Zhenzhou
Piao†
,
Junwu
Wang
,
Renxin
Chen
,
Zhuowen
Hao
,
Zheyuan
Zhang
,
Zouwei
Li
,
Yilong
Huang
and
Jingfeng
Li
*
Department of Orthopedics, Zhongnan Hospital of Wuhan University, Wuhan, 430071, China. E-mail: jingfengli@whu.edu.cn
First published on 12th August 2025
Abstract
The tissue engineering strategy represents a balanced and dynamic approach that utilizes various tissue cells, bio-scaffolds and bio-agents (such as cells and drugs) to facilitate tissue repair and regeneration. Hydrogel microspheres (HMS), distinguished by their unique micrometer-scale spherical architecture featuring high specific surface area, high porosity, and low invasiveness, coupled with advantageous biochemical properties including biocompatibility, biodegradability, hydrophilicity, and lubricity, are regarded as promising bio-scaffolds in tissue engineering. Notably, smart responsive hydrogel microspheres (SRHMS) can intelligently release bio-agents to actively promote tissue regeneration in response to internal or external stimuli, in a spatiotemporally controlled manner. As a result, SRHMS have attracted escalating attention for their application potential in tissue engineering. This review summarizes recent advancements in SRHMS applications within tissue engineering, emphasizing their fabrication processes, functional characteristics, and therapeutic applications across diverse domains such as bone, cartilage, skin, nerve, and cardiac tissue engineering.
Wider impactRecently, smart responsive hydrogel microspheres (SRHMS) have emerged as a novel class of intelligent materials, demonstrating transformative potential in the tissue engineering strategy due to their unique structures and multifunctional properties. Advances in key techniques such as microfluidics, electrospray, and three dimensional (3D) bio-printing have enabled precise control over microsphere properties and bio-agent delivery. The dynamic responsiveness of SRHMS to endogenous and exogenous stimuli enhances their spatiotemporal precision in multifunctional applications, including bone, cartilage, skin, nerve and cardiac regeneration. The integration of smart responsiveness with rolling lubricity further highlights the distinctive functional advantages of SRHMS. As a result, their superior controllability and smart application capabilities make them more attractive to researchers compared to other biomaterials. The interdisciplinary convergence of SRHMS with artificial intelligence may pioneer an innovative pathway for regenerative medicine and minimally invasive therapies. Future directions will involve multidimensional (4D/5D) structuring, multi-stimuli integration, and self-healing capabilities, further bridging the gap between laboratory research and clinical translation. This review provides a comprehensive framework to guide the design of next-generation SRHMS systems, fostering innovation in smart therapeutics and advancing materials science toward scalable, patient-specific applications. |
1. Introduction
The tissue engineering strategy represents an integrated process that combines scaffolds, cells, and bio-agents to assemble three-dimensional (3D) complexes for tissue repair and functional regeneration.1,2 As essential biomaterials, scaffolds must possess biocompatibility, biodegradability, and bioactivity to facilitate their diverse applications in bone, cartilage, skin, nerve and cardiac tissue engineering.3 Among these, hydrogels with 3D networks have emerged as highly promising scaffold materials due to their exceptional biocompatibility, biodegradability, hydrophilicity, tunable mechanical properties, and ability to mimic the extracellular matrix (ECM) microenvironment.4 Consequently, hydrogel delivery systems have not only significantly expanded the range of available biomaterials but have also provided a highly adaptable design space for tissue engineering, laying a critical foundation for precision medicine and personalized therapeutic strategies.5,6 However, their limitations are becoming increasingly apparent:7,8 bulk hydrogels suffer from poor injectability, hindering precise filling of irregular tissue defects; their homogeneous structure impedes spatial cell distribution and gradient diffusion of nutrients; moreover, their static physicochemical properties limit their ability to dynamically respond to complex pathological microenvironments.To overcome these limitations, HMS have emerged as a promising solution. Compared to traditional hydrogels, the micron-sized spherical structure of HMS offers distinct advantages: the uniform size and smooth surface of HMS significantly enhance injectability and tissue lubrication, reducing mechanical damage during the implantation process and improving uniform distribution within targeted tissues.9,10 The natural gaps between the microspheres not only facilitate cell migration and proliferation but also create pathways for nutrient and metabolic waste transport.3,10,11 More importantly, the high specific surface area and tunable porosity of HMS greatly increase the loading and release efficiency of drugs, cells, and bio-agents.12 These characteristics enable HMS to precisely conform to the complex morphology of tissue defects and dynamically regulate the microenvironment in both spatial and temporal dimensions, thus laying the foundation for the development of smart responsive delivery systems.
Building upon the advantages of traditional HMS, researchers have further developed SRHMS, which can dynamically sense and respond to specific endogenous and exogenous stimuli, thereby enabling precise regulation during tissue repair.13–15 These materials can trigger drug release, mechanical property changes, or cell behavior modulation in response to endogenous stimuli (such as pH, enzymes, reactive oxygen species (ROS)) or exogenous signals (such as light, temperature, magnetic fields, ultrasound (US)), greatly enhancing the spatiotemporal controllability of tissue engineering.15,16 Compared to traditional HMS, SRHMS offer more controllable application performance, showcasing superior drug release kinetics and establishing themselves as an emerging frontier in the tissue engineering strategy.
Recent research has demonstrated the advantages and significance of SRHMS in biomedical applications, driven by continuous innovations in the tissue engineering strategy. However, there is a lack of comprehensive reviews in this field. Therefore, this review systematically summarizes the latest research advancements in SRHMS for tissue engineering (Scheme 1). We first examine the properties, advantages, and disadvantages of commonly used materials for fabricating SRHMS. We then provide an in-depth discussion of the manufacturing processes involved in producing SRHMS, including microfluidics, emulsion technique, electrospray, mechanical fragmentation, lithography, and 3D bio-printing, as well as the strengths and weaknesses of each method. Next, we describe how SRHMS achieve smart responsiveness through chemical or physical crosslinking and explain the mechanisms behind each type of smart response. The review also highlights how the SRHMS delivery systems integrate smart responsive therapy with microsphere lubrication, demonstrating the functional advantages of SRHMS across various tissue engineering applications. Finally, we present a great deal of examples of SRHMS applications in diverse tissue engineering fields. We hope sincerely that by summarizing and looking forward to the progress of SRHMS research, we can cultivate new perspectives and expand the applications of SRHMS.
 | ||
| Scheme 1 SRHMS respond to endogenous or exogenous stimuli to release bio-agents in a spatiotemporally controlled manner, thereby promoting tissue regeneration. Created in https://BioRender.com. | ||
2. Preparation of smart responsive hydrogel microspheres
2.1. Biomaterials for hydrogel microspheres
HMS can be engineered from natural polymers, synthetic polymers (Fig. 1), or a hybrid of both, depending on the material source. These microspheres must fulfill essential characteristics such as hydrophilicity, biocompatibility, biosafety, biodegradability, mechanical properties, sol–gel transition capability, and tailored functionalities. Polymers with high hydrophilicity possess surface functional groups that form stable hydrogen bonds with water molecules, contributing to the preservation of structural integrity. The critical role of biocompatibility and biosafety in the application of HMS within tissue engineering is undeniable. Optimal biodegradability ensures that the degradation rate aligns with tissue regeneration, facilitating the sustained release of bio-agents throughout the healing process, thereby improving therapeutic efficacy and patient adherence to the therapy without the need for additional surgical interventions. Matching the mechanical properties of the microspheres to those of the surrounding tissues is crucial for effective tissue regeneration. SRHMS are capable of undergoing sol–gel phase transitions in response to internal or external stimuli, allowing for the on-demand release of bio-agents. Tailored functionalities of HMS are evident in their targeting ability, efficient loading and delivery mechanisms, and inherent lubricity. In the following section, we will provide a concise overview of some natural and synthetic polymers used in their fabrication. | ||
| Fig. 1 Natural and synthetic polymers employed in the fabrication of SRHMS. Created in https://BioRender.com. | ||
2.1.1.1. Hyaluronic acid. HA, a natural and negatively charged polysaccharide composed of N-acetylglucosamine and D-glucuronic acid, is ubiquitously present in the ECM of various tissues, including cartilage, skin, and synovial joint fluid.19 HA is renowned for its exceptional hydrophilicity, biocompatibility, and biodegradability. Beyond these intrinsic properties, HA plays a pivotal role in modulating cellular activities and functions through its molecular interactions with specific receptors on cell membranes. For instance, HA facilitates collagen deposition, epithelialization, and wound vascularization, thereby promoting the migration and proliferation of fibroblasts and keratinocytes, which significantly aids in wound healing.20 Despite these advantages, HA-based HMS face limitations in load-bearing tissue engineering applications due to their poor mechanical properties and susceptibility to degradation by endogenous enzymes resulting in their fragility under mechanical stress and a short in vivo half-life.21 The HA chain is rich in reactive groups, such as amino, hydroxyl, and carboxyl groups, which can be chemically modified through various methods—including carbodiimide reactions, enzymatic reactions, protein reactions, photo-reactions, Schiff base reactions, Michael addition reactions, and click chemistry reactions.21,22 Such modifications can significantly alter the mechanical properties, chemical characteristics, and bioactivity of HA-based HMS. A particularly noteworthy approach is dynamic covalent coupling (DCC) chemistry, which enables the creation of adaptable and reversible polymer networks within HA-based HMS.23 This technique imparts advanced functionalities such as injectability, shear-thinning, and self-healing properties, allowing the hydrogels to respond intelligently to environmental stimuli. In a groundbreaking study, Lei et al. enhanced the lubricity of the system by complexing HA molecules with liposomes through non-covalent interactions.24 Concurrently, they introduced methacrylic anhydride to the HA chain to generate HA methacrylate (HAMA), thereby endowing HA with photo-curability. Leveraging microfluidic technology and photopolymerization processes, they developed HA-based HMS encapsulating rapamycin-loaded liposomes (RAPA@Lipo@HMs), offering a novel therapeutic approach for the alleviation of osteoarthritis (OA).
2.1.1.2. Chitosan. CS, a naturally derived linear aminopolysaccharide, is obtained through the deacetylation of chitin and consists of glucosamine and N-acetylglucosamine units.25 As the sole cationic polysaccharide known to date, CS is highly valued for its cost-effectiveness, exceptional bioactivity, biocompatibility, and notably its potent antibacterial properties as well as its adjustable biodegradability by varying the degree of deacetylation, making it a prime candidate for the fabrication of HMS.26 The electrostatic interactions between the cationic amino groups of CS and negatively charged molecules or anions are pivotal in the formation of CS-based HMS, as well as in drug delivery and tissue engineering applications. Studies have shown that the biodegradability of CS is contingent upon the degree of deacetylation of chitin; a deacetylation degree below 70% facilitates easy degradation, whereas a degree above 70% renders it more resistant.25 Furthermore, CS's sensitivity to pH and temperature makes it an ideal material for crafting pH-responsive and temperature-responsive injectable HMS.27 CS-based HMS can be formed through either physical crosslinking (involving ionic, hydrogen bonding, hydrophobic, and electrostatic interactions) or chemical crosslinking (including Schiff base reactions, Diels–Alder reactions, Michael addition reactions, thiol–ene reactions, and nucleophilic ring-opening reactions).28 The CS backbone is rich in hydroxyl and amino groups, which can be chemically modified to engineer multifunctional CS-based HMS for tissue engineering.22 Nonetheless, CS-based HMS are not without their limitations, such as the constrained compressive strength and elastic modulus observed in hydrogen-bonded variants.27 Lei et al. developed a multifunctional HM system by using magnetic CS microspheres (MCS) embedded with Fe3O4 nanoparticles and integrated with Zn2+, which significantly promoted wound healing.29 The biohybrid MCS, by immobilizing Zn2+ ions, efficiently captures histidine-tagged vascular endothelial growth factor (VEGF) produced by Escherichia coli (E. coli), thereby accelerating wound repair. More importantly, the combination of Zn2+ and CS demonstrates synergistic antibacterial effects during the wound healing process, effectively mitigating infection and inflammation.
2.1.1.3. Collagen. Col, one of the primary structural proteins in the ECM, is predominantly composed of RGD sequences and is rich in glycine, aspartic acid, and arginine sequences.30 It is ubiquitously found in the connective tissues of animals, including cartilage, intervertebral discs, skin, and tendons, where it exerts a profound influence on cellular activities in vivo. Under physiological conditions, Col has the remarkable ability to self-assemble into a triple-helical fibrous structure, which confers exceptional tensile strength and durability.31 Traditionally, Col is sourced from animals, but this approach has several drawbacks. Animal-derived Col is often insoluble in water, has limited availability, requires time-consuming extraction processes, and, most concerningly, may harbor harmful contaminants such as viruses and provoke immune responses.30,32 Recombinant humanized Col has been developed by cloning Col cDNA fragments into appropriate vectors, expressing them in cellular systems, and purifying the resultant protein.33 This recombinant Col not only preserves the biological properties inherent to natural Col but also boasts reduced immunogenicity, enhanced water solubility, and superior processability.32 The array of outstanding biological properties that Col possesses renders it an ideal candidate for use in HMS. Nonetheless, HMS composed solely of Col typically suffer from inadequate physical strength and mechanical performance. To ameliorate these deficiencies, physical crosslinking methods (such as ultraviolet (UV) irradiation or heat treatment) and chemical crosslinking agents (including formaldehyde, glutaraldehyde, and 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC)) can be utilized.26 Moreover, the integration of polymers like elastin, CS, glycosaminoglycans, polyvinyl alcohol, or polyacrylic acid (PAC) during the crosslinking process can significantly enhance biological functionalities, including cell proliferation, differentiation, and migration, as well as bolster the mechanical strength and functional efficacy of the HM system. In a study, researchers developed a type I Col recombinant peptide HM (RCP-MS) loaded with bone morphogenetic protein 2 (BMP-2).34 These microspheres were then dispersed within three distinct hydrogels: high mannuronic acid (SLM) alginate, high guluronic acid (SLG) alginate, and a thermoresponsive HA derivative (HApN), thereby creating three unique smart responsive systems aimed at fostering new bone formation. The experimental outcomes revealed that the mechanically responsive SLG + RCP-MS system outperformed its counterparts—the mechanically responsive SLM + RCP-MS system and the thermally responsive HApN + RCP-MS system—in terms of therapeutic efficacy in bone formation.
2.1.1.4. Gelatin. Gel is a water-soluble polypeptide, derived from the irreversible degradation of Col and inheriting many of its advantageous properties such as exceptional hydrophilicity, biocompatibility, and biodegradability.22 These characteristics make Gel highly effective in promoting cellular activities, including migration, adhesion, growth, proliferation, and differentiation. However, the preparation process disrupts Col's triple-helical structure, significantly reducing Gel's mechanical strength compared to its precursor.35 Moreover, Gel's crosslinked network is susceptible to rapid hydrolysis by proteinase K within an hour, leading to swift degradation rates in vivo and thereby constraining its applications.36 Like Col, Gel can be crosslinked through various techniques to produce bio-scaffolds with improved mechanical strength and functionality. These techniques encompass physical crosslinking (such as thermal or radiation methods), chemical crosslinking (utilizing agents like glutaraldehyde, genipin, or methacrylic anhydride), or enzymatic crosslinking (e.g., with transglutaminase).22,26 Additionally, the integration of Gel with inorganic particles can enhance the system's bioactivity and mechanical properties.37 In the realm of tissue engineering, Gel surpasses Col in several aspects, including lower immunogenicity, enhanced hydrophilicity, and thermoresponsiveness.26 A notable feature of Gel is its reversible sol–gel transition at temperatures between 25 and 35 °C (the critical solution temperature), which underpins the theoretical basis for developing thermoresponsive HMS in tissue engineering.38 Li and his team employed microfluidic technology to fabricate matrix metalloproteinase (MMP)-responsive Gel methacrylate (GelMA) HMS under UV light.39 They developed G5-AHP/miR-140 nanoparticles by complexing miR-140 with a fifth-generation polyamidoamine (G5-AHP) that was modified with arginine, histidine, and phenylalanine. These nanoparticles were then encapsulated within the HMS, resulting in “nano-micro” combined gene hydrogels. This innovative strategy offers a novel cell-free solution for mitigating OA, showcasing the immense potential of localized injectable gene delivery systems in therapeutic applications.
2.1.1.5. Alginate. Alg is a natural anionic polysaccharide polymer rich in carboxyl groups, and it is derived from the cytoplasm and cell walls of brown algae. It consists of repeating blocks of α-L-guluronic acid (G blocks) and β-D-mannuronic acid (M blocks), which are linked by 1,4-glycosidic bonds.22,38,40 Due to its exceptional biocompatibility, biodegradability, hydrophilicity, low toxicity, cost-effectiveness, and ease of availability, Alg hydrogels are widely employed in tissue engineering research for the repair and regeneration of various tissues, including bone, cartilage, cardiac, and nerve tissue.41 However, the limited cell adhesion properties and mechanical strength of Alg-based HMS present challenges for their broader application in tissue engineering.38 Alg exhibits polycation behavior in aqueous solutions, and its G blocks can crosslink with divalent cations such as Ca2+, Ba2+, and Zn2+ to form HM systems with optimal mechanical strength.40 Additionally, by covalently crosslinking Alg with multifunctional molecules (e.g., polyacrylamide-co-acylhydrazone, polyethylene glycol-diamine), composite Alg hydrogels with enhanced mechanical strength can be achieved.26,41 A study has shown that incorporating inorganic particles into a hydrogel system composed of oxidized Alg and Gel can significantly enhance the system's ability to promote cell activity.42 Qu et al. successfully developed a novel pH-responsive HM system (E7-Lipo@Alg/Cs) using gas microfluidics, ionic crosslinking, and electrostatic adsorption techniques for targeted delivery via the intestinal-bone axis.43 This system is designed to regulate mitochondrial function and prevent bone loss, thereby enhancing the therapeutic efficacy for osteoporosis (OP). BMSC-affinitive peptides were conjugated to liposomes encapsulating resveratrol, which were then incorporated into alginate-calcium HMS. CS, electrostatically adsorbed onto the microsphere surface, provides resistance to gastric acid and facilitates targeted delivery to BMSCs through the intestinal-bone axis. Furthermore, mammals lack enzymes capable of degrading Alg, but Alg can be partially oxidized through crosslinking with divalent cations to enhance its degradability.26,44 Partial oxidation induces a conformational change in the Alg backbone, transforming it into an open-chain polymer and thereby accelerating its degradation rate.
2.1.1.6. Silk fibroin. SF is a natural high-molecular-weight biological material primarily extracted from silkworm silk. It contains 18 amino acids that are also found in the human body and is mainly composed of many interconnected light and heavy chain polypeptides.45 Remarkably, SF not only boasts exceptional biocompatibility, biodegradability, hydrophilicity, and low immunogenicity but also exhibits extraordinary mechanical properties.46 These unique mechanical attributes are predominantly governed by the heavy chains, which are capable of forming β-sheet microcrystals, whereas the contribution of the smaller light chains is minimal.36 SF has the capacity to form hydrogels through a variety of physical interactions, including hydrogen bonding, hydrophobic interactions, electrostatic interactions, ionic interactions, and chain entanglement. Moreover, it can establish stable crosslinking points via chemical bonds (for instance, utilizing phenolic hydroxyl or amino groups on its molecular chains), thereby fostering the development of a spatial network structure for hydrogels.47 The gelation process of SF hydrogels can be expedited by diverse external stimuli (such as pH, shear, and US),48 offering valuable insights for the advancement of SF-SRHMS. Notwithstanding the formidable mechanical properties of SF hydrogels, their intrinsic bioactivity is somewhat limited. Consequently, SF is frequently amalgamated with other natural polymers that, while possessing lower mechanical strength, offer superior bioactivity, to create hybrid hydrogels.30 Chen et al. adeptly engineered a novel pH- and temperature-sensitive sodium alginate (SA)-modified SF microsphere embolic agent through an emulsion crosslinking technique.49 The research elucidated that SA-modified SF microspheres have the potential to function as biodegradable arterial embolic agents in the treatment of liver cancer, equipped with drug-loading and release functionalities that stimulate the proliferation of fibroblasts and human umbilical vein endothelial cells.
2.1.2.1. Polyethylene glycol. PEG is a synthetic polyether obtained through the polymerization of ethylene oxide and polyethylene. By carefully adjusting the proportions of chemical reactants, we can achieve a more uniform molecular weight distribution, thereby effectively controlling the length of its molecular chains.22 PEG, renowned for its great biocompatibility and hydrophilicity, is a non-toxic, low-immunogenic, and anti-fouling synthetic polymer that is extensively utilized in the tissue engineering strategy. It stands as one of the most prevalent synthetic materials for the preparation of hydrogels. Furthermore, PEG serves as an ideal heat storage carrier with commendable photothermal conversion efficiency, making it applicable as a phase change material (PCM) in tissue engineering.52 However, due to the inherent bio-inertness of PEG, pure PEG-based HMS fall short of providing an optimal environment for cell proliferation, differentiation, adhesion, and migration. To address this limitation, natural polymers endowed with diverse functionalities and high bioactivity, such as ECM components, CS, and Col, can be incorporated into PEG-based HMS.30,50 The long hydrophilic chains of PEG play an important role in preventing unwanted protein adsorption while enhancing the biocompatibility and mechanical strength of other materials, which is why PEG is frequently employed in modifying various materials.50 Cheng et al. developed an innovative composite hydrogel-EVs system, designated as DHPM(4APPC)_EVs.53 This system is based on Sparchigh Treg-derived extracellular vesicles (EVs) that overexpress CXCR2, combined with pH/H2O2/enzyme MMP-9-responsive HMS. The microspheres within this system were ingeniously fabricated by linking distearoylphosphatidylethanolamine (DSPE) and PEG through hydrazone bonds. This design enables rapid targeting of the acute myocardial infarction (AMI) region and facilitates the responsive release and capture of EVs in the presence of pH/H2O2/enzyme MMP-9. Such a system is beneficial for sustaining the long-term repair function of EVs in damaged hearts, offering a promising therapeutic approach for cardiac tissue regeneration.
2.1.2.2. Polyvinyl alcohol. PVA is a long-chain synthetic polymer synthesized through the polymerization of vinyl acetate followed by hydrolysis. Its structure is characterized by an abundance of hydroxyl groups, which not only impart mechanical stability but also govern its physicochemical properties.54 PVA hydrogels have garnered significant attention in the field of tissue engineering due to their remarkable attributes, including high water content, low toxicity, excellent tissue compatibility, and superior elastic modulus and mechanical strength. The cross-linking of PVA chains to form hydrogels can be achieved via physical methods—such as freeze–thaw cycles, hydrogen bonding, and electrostatic interactions—or chemical methods, including irradiation and enzymatic reactions.55 However, given the potential use of harmful substances in chemical methods, physical methods are often the preferred approaches. PVA's mechanical properties closely resemble those of human articular cartilage, making it an ideal candidate for load-bearing applications such as the repair of craniofacial deformities and bone tissue engineering.56 Moreover, PVA hydrogels exhibit structural similarities to human intervertebral disc tissue, positioning them as a promising material for spinal tissue engineering.22 Notably, the semi-crystalline framework of PVA, derived from its monomeric structure, facilitates the efficient transport of oxygen and nutrients to cells, further enhancing its utility in biomedical applications.56 Despite these advantages, PVA hydrogels are not without limitations. Challenges such as poor elongation at break, low fatigue resistance, and high coefficient of friction (COF) can hinder their performance. These issues, however, can be effectively addressed through an innovative approach, chemical double-cross-linking (CDC).54 Beyond these mechanical concerns, the lack of inherent bioactivity remains a critical challenge in the development of PVA HMS. To overcome this, researchers often cross-link PVA with natural polymers—such as CS, Alg, and fibrin—or hybrid polymers incorporating natural materials, such as phenylboronic acid (PBA)-modified HA (HA-PBA) and GelMA. These modifications significantly enhance the bioactivity of the system. In a groundbreaking study, researchers utilized emulsification techniques to develop a glucose-responsive, cell-loaded HM system composed of CS and PVA mixed with 3-aminophenylboronic acid (APBA), termed APBA-GEL-PVA.57 This innovative system leverages glucose-responsive solubilization to release growth factors produced by cells, thereby promoting cell survival and functional improvement in tissues.
2.1.2.3. Poly(lactic-co-glycolic acid). Polylactic acid (PLA) and polyglycolic acid (PGA) are synthetic aromatic polyesters known for their temperature-sensitive properties. The ester groups on their long chains undergo hydrolysis, resulting in the degradation of the polymer chains into lactic acid (LA) and glycolic acid (GA), respectively. Through further polymerization, a more stable copolymer, PLGA, can be synthesized.22 PLGA exhibits excellent biosafety, lacking immunogenicity and antigenicity, while the abundance of carboxyl groups on its side chains allows for various chemical modifications.38 PGA is a crystalline hydrophilic polymer, while PLA is a rigid hydrophobic polymer. Consequently, during the synthesis of PLGA, the PLA
![[thin space (1/6-em)]](https://www.rsc.org/images/entities/char_2009.gif)
:PGA ratio and the molecular weight of the polymer significantly influence its properties, including hydrophobicity, crystallinity, mechanical performance, size, and biodegradation rate.58 As a result, the hydrophilicity and biodegradation rate of PLGA hydrogels increase as the PLA:PGA ratio decreases. However, the hydrophobic segments in the polymer often lead to poor hydrophilicity and low biodegradation rates in PLGA hydrogels, and they inherently lack bioactivity. To address these limitations, PLGA is typically cross-linked with natural polymers to prepare hydrogels. SRHMS based on PLGA have attracted considerable attention due to their ability to respond to specific endogenous and/or exogenous stimuli, enabling precise spatiotemporal control and advancing targeted therapy and precision medicine. Yu et al. developed an injectable thermosensitive PLGA-PEG-PLGA (PPP) triblock copolymer HM system (Gel-MS/TIIA) for treating myocardial damage caused by AMI.59 This system demonstrates excellent mechanical properties at body temperature, and the HMS encapsulating tanshinone IIA (TIIA) exhibit superior ROS-responsive characteristics under high ROS conditions. This innovative system provides localized anti-thermal and anti-inflammatory effects, effectively reducing cardiomyocyte (CM) apoptosis and pyroptosis. The formation of block copolymers between PLGA and synthetic polymers, such as PEG, has become one of the most extensively studied approaches for creating thermosensitive hydrogel systems. Moreover, by specially modifying these block copolymers with photosensitive or conductive polymers, dual-sensitive hydrogel systems responsive to both heat and light or heat and electricity can be achieved, further expanding their potential applications in tissue engineering.
2.1.2.4. Polyacrylamide. PAM is a colorless, transparent, and homogeneous gel-like polymer that was initially employed as a soft tissue filler and has since found extensive applications in tissue engineering. It is renowned for its exceptional elasticity, hydrophilicity, and resistance to heat, acids, and bases; it may be a homopolymer or copolymer derived from acrylamide and its derivatives.22 PAM hydrogels can be fabricated by cross-linking PAM or acrylamide monomers through either physical or chemical methods, endowing them with biocompatibility, hydrophilicity, and tissue-like mechanical properties. The abundance of amide groups in PAM facilitates the formation of hydrogen bonds, which researchers often exploit to achieve the desired mechanical strength in PAM hydrogels.60 Based on the polymer's network structure and chemical composition, PAM hydrogels can be engineered as smart responsive polymer systems. Although PAM itself lacks ionizable functional groups, the incorporation of ionizable monomers during hydrogel synthesis can produce pH-responsive hydrogels. Interestingly, the COF of copolymer gels decreases with increasing pH, a phenomenon attributed to the conversion of side-chain amide groups to carboxyl groups under alkaline conditions and/or swelling induced by cross-linker hydrolysis.61 PAM can also form semi-interpenetrating or interpenetrating polymer networks with suitable polymers to create thermally responsive hydrogel systems.62 By encapsulating inorganic or organic nanoparticles—such as silica, carbon nanotubes, Gel, and cellulose—the mechanical strength, adhesion, and self-healing properties of PAM hydrogels can be finely tuned and enhanced.60 However, like other synthetic polymer hydrogels, PAM hydrogels inherently lack bioactivity. To address this, incorporating bioactive materials during preparation remains the most effective strategy to mitigate their bio-inertness. Juby K. Ajish and his team utilized inverse emulsion polymerization technology to synthesize a glycosylated stimulus-responsive PAM HM system loaded with silver nanoparticles (AgNPs), known for their antibacterial properties.63 This innovative system exhibits remarkable capabilities in capturing, detecting, and eliminating pathogenic bacteria, thereby promoting the regeneration and repair of infected tissues.
2.2. Fabrication techniques
The remarkable properties of HMS render them exceptionally well-suited for addressing the growth needs of tissues, organs, and cells in tissue engineering, thereby aiding in the repair, enhancement, and preservation of their functions. Consequently, the development of HMS capable of replicating the extracellular microenvironment is of paramount importance. Generally, the process begins with the preparation of precursor solutions to induce droplet formation, which is then followed by solidification or polymerization to yield HMS.64 The characteristics of the resulting HMS are intricately linked to the materials used, their structures, and the methods of fabrication. In the following section, we will provide an overview of several techniques employed in the fabrication of HMS, encompassing microfluidics, emulsion technique, electrospray, mechanical fragmentation, lithography, and 3D bio-printing.To successfully prepare HMS, the precursor solution must fulfill two critical requirements: (a) the viscosity of the solution must be sufficiently low to prevent clogging of the microchannels; and (b) to prevent droplet aggregation, cross-linking must occur rapidly during droplet collection.64,65 A variety of cross-linking strategies have been employed to solidify or polymerize droplets, including the application of external light for rapid photo-cross-linking and the use of thermally responsive cross-linkers that cool the droplets, enabling either “on-chip” or “off-chip” cross-linking, with the former being the preferred approach. The primary advantage of microfluidics lies in its ability to generate precisely controlled droplets, achieved through the use of specifically designed microfluidic channels and meticulous control of input liquid flow rates.64 Currently, two main types of microfluidic devices are utilized for preparing HMS: capillary microfluidic devices and polydimethylsiloxane (PDMS) devices. Capillary microfluidic devices ensure complete separation of internal and external fluids throughout the process, facilitating efficient and stable encapsulation while allowing for precise control over microfluidic emulsion generation.68 However, they encounter challenges in achieving large-scale repetitive manufacturing.69 In contrast, PDMS devices can be tailored to specific sizes using soft lithography or 3D printing to create microfluidic junctions.64 PDMS is renowned for its excellent reproducibility, low cost, good gas permeability, and transparency, making it highly suitable for large-scale production. Nonetheless, it has a tendency to absorb small hydrophobic molecules, which can impact solute concentration.70 Intriguingly, a study has introduced an air microfluidic method in which two micro-scale liquids are ejected to collide and form droplets71 (Fig. 2A). While such techniques show considerable promise, further exploration is necessary to determine their viability as alternatives to chip-based microfluidics.
 | ||
| Fig. 2 (A) Air microfluidics employed for the preparation of HMS. Reproduced with permission.71 Copyright 2018, American Association for the Advancement of Science. (B) HMS prepared via the emulsion technique. Reproduced with permission.90 Copyright 2021, Wiley-VCH GmbH. (C) Alg HMS synthesized through electrospray. Reproduced with permission.79 Copyright 2022, Elsevier Ltd. (D) HMS formed by the fragmentation of hydrogels using gas shear stress. Reproduced with permission.80 Copyright 2021, Wiley-VCH GmbH. (E) Schematic diagram of the workflows for imprint lithography, photolithography, and flow lithography. Reproduced with permission.83 Copyright 2011, Elsevier Ltd. (F) Cell-laden Col HMS generated by the 3D bio-printing technique and cross-linked with TA solutions of varying concentrations. Reproduced with permission.86 Copyright 2022, Wiley-VCH GmbH. | ||
While microfluidics excels in precisely controlling the size, shape, and distribution of microspheres, it still faces several challenges, including low production efficiency, high production costs, stringent technical requirements, complex device configurations, and labor-intensive maintenance and cleaning. Moreover, the relatively low throughput of microfluidics makes the production of HMS with smaller diameters particularly challenging.64 A pioneering study demonstrated the development of a step-emulsification device featuring meticulously engineered channel geometries with hundreds of parallelized nozzles, enabling scalable and parallelized droplet generation in microfluidic systems.69 This significant technological advancement provides a solution for the high-throughput production of HMS while minimizing batch-to-batch variation. Moreover, the synergistic combination of microfluidics and the emulsion technique, complemented by innovations in three-phase microfluidics, has facilitated the precise engineering of sophisticated core–shell microspheres with tunable multicompartmental structures.10,12 These cutting-edge developments not only overcome the inherent throughput constraints of traditional microfluidics but also empower researchers with unprecedented control over microsphere architecture (including core–shell configurations and multicompartmental designs) through versatile modular fabrication strategies. Consequently, before HMS can be mass-produced and broadly integrated into clinical practice, substantial efforts are required to refine and enhance microfluidics.
The emulsion technique is unable to produce controlled droplets, which stands as its most significant limitation, leading to the polydispersity of the resulting microspheres.64 In comparison to microfluidics, HMS fabricated through the emulsion technique demonstrate reduced lubricity and exhibit more pronounced variability in drug release profiles. When it comes to cell delivery, polydispersity hinders the ability to ensure a consistently fixed number of cells encapsulated within each individual microsphere. Nevertheless, there have been promising reports of employing filters with specific pore sizes to refine HMS, resulting in a more uniformly dispersed suspension and presenting a potential remedy to the challenge of polydispersity.75 Moreover, residual oil phases and surfactants can provoke adverse immune responses, imparting toxicity to cells and tissues, thereby undermining the biosafety and biocompatibility of the microspheres.10,72 Consequently, the selection of non-toxic or minimally toxic materials during the fabrication of HMS is imperative to ensure their safer and more effective application in tissue engineering.
Electrospray stands out as an efficient method for producing uniform HMS by fine-tuning various parameters, all while circumventing the detrimental effects of high temperatures, excessive mechanical rates, toxic oil phases, and surfactants on the HMS. Consequently, the resulting HMS demonstrate superior encapsulation efficiency and enhanced biocompatibility. Reports indicate that microspheres fabricated via electrospray often exhibit a high dispersion index (>5%) and suboptimal monodispersity; however, the use of filtration techniques can yield more uniformly dispersed microspheres.64,65 Although electrospray is simple and efficient, its drawbacks are evident. It demands a high level of technical expertise, specialized equipment, and a high-voltage power supply. Moreover, the multitude of preparation parameters and the intricate process of microsphere collection add layers of complexity to the production workflow.30,78
Mechanical fragmentation is a straightforward and rapid approach for preparing HMS, characterized by its low cost and high efficiency, which makes it well-suited for large-scale production. More significantly, this method eschews the use of reagents that could compromise cell viability, such as toxic oil phases or surfactants, thereby preserving some of the intrinsic properties of the hydrogel. However, the limitations of this technique are also quite pronounced. Firstly, the microgels produced via mechanical fragmentation may exhibit irregular shapes, and there is no assurance that the resulting products will be uniformly spherical. Secondly, the crude nature of the preparation process often results in polydispersity among the obtained HMS, with significant variations in size. This polydispersity can markedly impair the encapsulation efficiency of cells or drugs within the microspheres, as well as their ability to function effectively in controlled release applications.
Lithography facilitates the precise construction of microstructural features in HMS at the microscale.73 The evolution of nano-fabrication techniques has expanded the capabilities of lithography, enabling the preparation of HMS with intricate structures at the nanoscale through the use of molds and masks featuring nanoscale details.64 Duan et al. harnessed asymmetric UV lithography to create HMS with complex 3D porous structures, which were subsequently used to assemble nanoscale photonic crystals.84 It is noteworthy that while flow lithography can be combined with microfluidics to achieve a relatively high throughput compared to the other two lithography routes, lithography's most notable limitation, when juxtaposed with other methods for fabricating HMS, remains its relatively low throughput.64
The 3D bio-printing technique not only facilitates the direct preparation of HMS but also enables combined preparation when synergized with other strategies. For example, Col microspheres embedded with cells, directly fabricated through 3D bio-printing, can be cross-linked with tannic acid (TA), thereby overcoming the limitations inherent in techniques that rely on surfactants and/or oil phases86 (Fig. 2F). The electro-assisted 3D bio-printing technique excels in the precise generation of HMS, which are pivotal for applications in cell therapy, drug delivery, and the construction of microsphere-based organoids.72 Zhang et al. pioneered a microfluidic-based printing nozzle designed for the preparation of monodisperse HMS using high-viscosity bio-inks.87 Furthermore, the 3D bio-printing technique has been instrumental in functionalizing HMS and constructing composite scaffolds. By integrating modified substances into the bio-inks during the 3D bio-printing process, it is feasible to enhance or introduce functionalities such as antibacterial properties and mechanical strength.88,89 Despite the vast potential of the 3D bio-printing technique in the fabrication of HMS and its applications in tissue engineering, there remain areas necessitating refinement. These include the optimization of various equipment parameters, the printing environment, and the maintenance of bioactivity.
2.3. Smart responsiveness
The core of SRHMS lies in their ability to dynamically respond to endogenous stimuli (e.g., pH, ROS, enzymes) and/or exogenous stimuli (e.g., light, magnetic fields, ultrasound), a pivotal feature underpinning their versatile applications. The stimulus-responsive mechanisms of smart materials provide a fundamental theoretical framework for the advancement of SRHMS. Early theoretical research on material response mechanisms laid a critical foundation for SRHMS design. For instance, Tanaka et al. first proposed the volume phase transition theory of gels in the 1980s, elucidating the molecular mechanism behind the reversible swelling–shrinking behavior of gels under external stimuli such as temperature and pH.93 This theory offered key guidance for the subsequent development of pH- and temperature-responsive hydrogels. In the 1990s, Yoshida team's “self-oscillating gel” model further clarified the dynamic process of chemical energy conversion into mechanical energy within gels, providing theoretical support for designing enzyme- and redox-responsive systems.94 Recent studies on response mechanisms based on dynamic covalent chemistry (e.g., Schiff bases, boronate esters) have further expanded the intelligent regulatory dimensions of SRHMS, enabling more precise spatiotemporal controlled release.95,96 These theoretical breakthroughs have not only deepened the understanding of material–environment interactions but also provided a scientific basis for the application and development of SRHMS in the tissue engineering strategy.2.3.1.1. pH-responsive. pH is a critical biochemical parameter that often undergoes significant changes in the microenvironment of damaged tissues, particularly in characteristic disease states such as tumors and inflammation.15,101 As a result, pH-responsive systems are designed to dynamically modify their structural properties, enabling them to activate site-specific delivery functions in the presence of pathologically induced acidic pH. The development of pH-responsive delivery systems primarily revolves around two key strategies: (i) the utilization of polymers equipped with ionizable groups (polyacids or polybases), which experience conformational and/or solubility changes in response to fluctuations in environmental pH; and (ii) the integration of acid-labile bonds into the polymer carrier, the cleavage of which can release molecules anchored to the polymer backbone, modify the polymer's charge, or unveil targeting ligands.97,102
A hallmark pathological characteristic of the tumor microenvironment is the reduction in extracellular pH.98,103 A multitude of anticancer pH-responsive delivery systems have been engineered to capitalize on the subtle pH differences between healthy tissues and the tumor microenvironment. Typically, the extracellular pH in the tumor microenvironment (TME) ranges from 6.5 to 6.8, whereas the lysosomal pH hovers around 5.0. In contrast, the pH of normal tissues is maintained between 7.2 and 7.4.104 This pH discrepancy stems from the ability of cancer cells to reprogram the metabolic environment of the tissue, enhancing glucose uptake and converting glucose into lactate, even in the presence of fully functional mitochondria—a phenomenon famously known as the “Warburg effect”.105 Dong et al. have synthesized pH-responsive HMS by utilizing self-assembled silica colloidal crystals as templates within microfluidic droplets.106 These injectable multifunctional HMS are endowed with responsive photonic structures, specifically designed for in vivo pH detection within the TME. The pH detection mechanism relies on measuring the characteristic reflection peak of the microspheres in the near-infrared (NIR) wavelength range, which is sensitive to environmental pH changes. A decrease in pH induces a blue shift in the characteristic reflection peak (Fig. 3A(i)). Leveraging the deep tissue penetration capability of NIR, this system facilitates continuous in vivo monitoring of pH. Furthermore, pH-responsive HMS can function as drug carriers for controlled drug release and tumor treatment (Fig. 3A(ii)), while simultaneously enabling the assessment of treatment efficacy through the monitoring of in vivo pH changes.
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| Fig. 3 (A) pH-responsive HMS are designed to monitor changes in the TME while simultaneously releasing drugs in response to pH variations. (i) These microspheres enable in vivo detection of TME pH, with a decrease in pH resulting in a blue shift of the characteristic reflection peak. (ii) A schematic representation illustrates how pH-responsive HMS release the chemotherapeutic drug doxorubicin in response to pH changes. Reproduced with permission.106 Copyright 2020, Elsevier B.V. (B) A schematic of the detailed mechanism of the ROS-responsive MTK-TK-drug HM system, which is prepared from PEG-TK-AOA, PEG-TK-PEG, and PEG-UPy, is presented for the treatment of MI. Reproduced with permission.117 Copyright 2024, Elsevier Ltd. (C) Enzyme-responsive KGE HMS release Exos in response to MMP-1, thereby promoting the migration and differentiation of bone marrow BMSCs. Reproduced with permission.129 Copyright 2023, Elsevier Ltd. (D) Three distinct pathways to achieve glucose responsiveness are outlined: (i) the GOx-catalyzed oxidation of glucose generates gluconic acid, which lowers the microenvironment pH and triggers drug release from glucose-responsive HMS. Reproduced with permission.138 Copyright 2024, Elsevier B.V. (ii) ConA binds to glucose to form a complex, causing glucose-responsive HMS to swell and release drugs. Reproduced with permission.139 Copyright 2018, Elsevier B.V. (iii) PBA complexes with glucose to form a compound, leading to the swelling of glucose-responsive HMS and subsequent drug release. Reproduced with permission.140 Copyright 2022, MDPI. | ||
In the context of employing pH-responsive drug delivery systems for the treatment of OA, protonation-induced endosomal/lysosomal escape can result in the leakage of hydrolytic enzymes into the cytoplasm, triggering autophagy and cell death.107 Consequently, further investigation into the pH-responsive mechanisms of chemical group protonation must be approached with prudence, with a dual focus on efficacy and safety during testing. The advancement of pH-responsive materials based on acid-labile bond strategies is also subject to certain constraints. For example, the pH-responsive delivery system that employs the acid-labile hydrazone bond is limited by the presence of ketone or aldehyde functional groups in therapeutic drugs and the undesirable cytotoxicity of cationic polymer residues.13,108 While a study on pH-responsive HMS has substantiated their efficacy and stability, it is equally imperative to evaluate their biosafety, a factor of paramount importance in clinical practice.
2.3.1.2. Redox-responsive. Redox stimuli can be elicited through the application of voltage or the action of oxidizing and reducing agents.109 The distinct redox conditions that exist between intracellular and extracellular compartments, as well as between healthy and diseased cells, serve as a catalyst for redox-responsive systems to release their therapeutic cargo.97 In accordance with the pathological mechanisms of various diseases, redox-responsive materials undergo processes such as degradation, structural alterations, property transformations, and functional regulation, thereby ensuring the targeted transport and release of therapeutic agents at the pathological site.110 These materials are adept at precisely mediating the release of active molecules within various cellular sub-compartments, including the cytoplasm, mitochondria, and nucleus, thereby facilitating therapeutic interventions. Significantly, redox-responsive materials offer effective cellular protection by eliminating corresponding oxidative substances.111,112 The reaction mechanisms of these materials are broadly categorized into two types: the amphiphilic transformation of functional groups and bond cleavage. In the case of amphiphilic transformation, the equilibrium of amphiphilic properties in redox-responsive materials is disrupted by a transition from hydrophobic to hydrophilic states, leading to the dissociation of structures such as polyphenylene sulfide, hydrophobic monosulfides, monosilyl, or monotellurium polymers, and the subsequent release of their cargo.113 Bond cleavage, on the other hand, involves the complete severance of functional bonds in response to redox substances like ROS or glutathione (GSH), which disrupts the carrier's structure.114 This category predominantly includes materials that contain boronic esters, thioesters, disulfide bonds, and diselenide bonds.
ROS responsiveness specifically pertains to the role of ROS in cellular signaling, whereas redox responsiveness is a more comprehensive concept that includes all oxidation and reduction reaction processes within cells. In biological systems, ROS encompass oxygen-derived radicals and non-radicals, such as superoxide anion (O2˙−), hydrogen peroxide (H2O2), hydroxyl radical (˙OH), and singlet oxygen (1O2), which are vital for various physiological processes.115,116 Nonetheless, an overproduction of ROS can contribute to a spectrum of diseases, including inflammatory disorders, cancer, atherosclerosis, and cognitive dysfunction.17 In the design of ROS-responsive hydrogels, functional groups are typically grafted onto polymers based on ROS-sensitive units. Commonly used covalent bonds for this purpose include Schiff bases, disulfide bonds, boronate esters, acylhydrazone bonds, and Diels–Alder reactions, which facilitate controlled drug release in specific environments. Wang et al. have developed an in situ ROS-responsive HM system that releases MTK-TK-drug in response to ROS stimuli117 (Fig. 3B). Within this system, (aminooxy)-acetic acid (AOA) in MTK-TK-drug promotes the transdifferentiation of Th17 cells into anti-inflammatory Treg cells. This HM system markedly reduces cell apoptosis and inflammation during the early stages of myocardial infarction (MI), thereby mitigating fibrosis, fostering angiogenesis, and preserving cardiac function.
For redox-responsive HM systems, several critical considerations must be meticulously addressed. Firstly, the growing emphasis on personalized medicine necessitates tailored approaches, as the sensitivity of redox-responsive materials is contingent upon the fundamental chemical bonds, as well as the structural and hydrophilic properties of the polymers. Secondly, different diseases inherently require unique formulations of redox-responsive materials, and to optimize therapeutic efficacy, the rate of drug release—mediated through solubility changes and bond cleavage—must be finely tuned to align with the pathological context. This ensures that the drug delivery is both effective and harmonious with disease progression. Thirdly, a thorough evaluation of the biocompatibility and biosafety of redox-responsive materials is imperative, especially for their translation into clinical applications. Lastly, given that normal cellular activities generate relatively low levels of ROS, the design of these materials must possess the capability to differentiate between the basal ROS levels characteristic of normal cellular functions and the elevated ROS levels associated with pathological conditions.115
2.3.1.3. Enzyme-responsive. Enzymes are indispensable catalysts renowned for their high specificity and selectivity toward substrates in enzyme-catalyzed reactions, playing a pivotal role in a myriad of biological and chemical processes within living organisms.118,119 As a result, enzyme-responsive hydrogels have emerged as a highly promising approach for directly targeting and responding to specific molecules. The enzyme-responsive strategy leverages smart polymers as carriers, wherein enzyme-sensitive moieties are designed to mimic the enzyme's substrate. Upon enzymatic catalysis, these moieties undergo significant macroscopic physical or chemical transformations, facilitating the controlled release of therapeutic payloads.120 Among the most notable strategies in tissue engineering is enzyme-catalyzed bond cleavage. This approach involves the strategic incorporation of specific peptide sequences or enzymatically cleavable bonds—such as phosphates or esters—into the hydrogel network or its precursor.121 Enzyme-responsive hydrogel systems, constructed by using peptide building blocks, feature peptide fragments with precise amino acid sequences that are selectively degraded or digested by particular enzymes, ultimately leading to the degradation of the hydrogel.122
In the realm of tissue engineering, the first documented application of an enzyme-responsive hydrogel system featured a PEG-based hydrogel designed to be sensitive to MMPs. In its synthesis, MMP-sensitive peptides were employed as cross-linking agents, allowing the hydrogel system to undergo biodegradation when exposed to MMPs such as gelatinases and collagenases.122,123 MMPs constitute a group of 26 endopeptidases that are adept at cleaving peptide bonds, and they play crucial roles in metabolic processes. These enzymes are classified into gelatinases, collagenases, stromelysins, and membrane-type MMPs.97,124 During the process of tendon healing, several MMPs, including MMP-2, MMP-9, and MMP-13, are overexpressed, thereby influencing the repair and remodeling of tendons.125–127 Gel, a natural polymer, is susceptible to degradation by a variety of enzymes, including MMPs, rendering it an ideal candidate for on-demand drug release in response to MMP activity. A research study has demonstrated that exosomes (Exos) derived from bone marrow mesenchymal stem cells (BMSCs) can activate paracrine pathways, thereby recruiting BMSCs to the site of injury to facilitate tissue repair.128 Yang and his team engineered an injectable MMP-1-sensitive HM (KGE) that responds to neovascularization during the angiogenesis phase, enabling the spatiotemporally controlled release of Exos.129 This HM system was constructed using microfluidics, integrating self-assembling peptides (KLDL-MMP-1), GelMA, and BMSC-Exos on a microfluidic chip. Both in vitro and in vivo experiments have shown that KGE exhibits minimal cytotoxicity and can recruit CD90+ stem cells through neovascularization, thereby promoting bone repair during angiogenesis (Fig. 3C).
In comparison to other endogenous stimuli, enzyme-responsive HMS present unparalleled advantages, primarily due to their superior bio-recognition capabilities and highly efficient catalytic performance. Owing to their specificity in both location and function, enzyme-responsive systems can ideally surpass other conventional stimuli (such as pH, temperature, and light), which may be limited in terms of efficiency and specificity. For example, while pH gradients are present within the intracellular compartments of each cell, enzymes offer enhanced selectivity as a result of their evolutionarily optimized substrate recognition.4 Nevertheless, challenges such as unstable controlled release, high costs, and biosafety concerns related to enzyme-responsive hydrogel systems still necessitate careful consideration.130
2.3.1.4. Glucose-responsive. Glucose stands as the principal energy source for human life activities, typically maintained in the blood at physiological levels through a stringent feedback regulation mechanism. In this distinctive scenario, the human pancreas operates as an intrinsic glucose-responsive platform, where β-cells detect variations in blood glucose levels and release endogenous insulin (INS) in response.131 Nevertheless, this delicate equilibrium is profoundly disrupted in certain pathological conditions, such as diabetes, which is marked by a substantial elevation in circulating glucose concentrations (exceeding fourfold) due to either a deficiency in INS production or the emergence of INS resistance.132 This disruption triggers a series of complications and inflicts damage on tissues and organs. The discomfort and inconvenience stemming from frequent INS injections and monitoring significantly diminish the quality of life for these patients.30,133 Drawing inspiration from bio-responsive delivery systems, glucose-responsive hydrogels that undergo structural changes in reaction to surrounding glucose concentrations are particularly well-suited for diabetes treatment.
Currently, the glucose responsiveness of injectable hydrogels, including injectable HMS, is predominantly realized through three distinct pathways: glucose oxidase (GOx) hydrogels, lectin hydrogels (primarily concanavalin A (Con A)), and PBA hydrogels (Fig. 3D): (a) the foundational approach based on GOx entails embedding GOx into the hydrogel matrix. The GOx-catalyzed oxidation of glucose produces gluconic acid and H2O2, which lowers the pH of the microenvironment and induces the swelling of pH-responsive hydrogels, thereby facilitating controlled drug release30,102,121,134 (Fig. 3D(i)). Notably, leveraging the GOx-catalyzed glucose oxidation reaction, Yu et al. devised an in situ hypoxia-responsive system,135 while Xu et al. engineered an H2O2-responsive system.136 (b) The second principal pathway involves the incorporation of lectins (such as Con A) into hydrogels. Lectins are carbohydrate-binding proteins that form complexes with the carbohydrate chains of glycoproteins and glycolipids on cell surfaces. Glucose and glycosylated polymers engage in competitive binding to Con A, as Con A exhibits a higher affinity for glucose than for glycosylated polymers. The binding of free glucose to Con A triggers polymer chain expansion, leading to swelling and subsequent INS release via diffusion133,134,137 (Fig. 3D(ii)). However, a notable drawback is that Con A is susceptible to leakage from the hydrogel during the swelling process.131,137 (c) The third principal pathway is founded on the complex formed between PBA and polyol compounds (sugar alcohols). In an aqueous environment, phenylboronate groups maintain an equilibrium between uncharged (non-ionic) and charged (ionic) forms. Upon glucose recognition, PBA engages in reversible complexation with polyol compounds. The reaction between PBA and glucose proceeds through the cationic charged form, and due to dissociation equilibrium, the reaction shifts the overall gel charge density toward cationic charged species. The augmented charge within the hydrogel fosters polymer chain repulsion and increased hydrophilicity, culminating in swelling131,133,134,137 (Fig. 3D(iii)).
Cai et al. harnessed microfluidics to engineer a glucose-responsive three-compartment HM system loaded with INS, specifically tailored for diabetes treatment.138 In this system, glucose reacts with GOx to produce gluconic acid, which induces the degradation of the microspheres under acidic conditions, thereby facilitating the controlled release of INS. A significant advantage of the three-compartment microspheres over their single-compartment counterparts is the ability to fine-tune the CS concentration in each compartment. This adjustment achieves a gradient pH-responsive range and ensures that the structural integrity and bioactivity of the drug are preserved over an extended duration. Bai et al. employed emulsification technology to design glucose-responsive core–shell HMS based on glycidyl methacrylate-modified dextran (Dex-GMA)-Con A, enabling the controlled release of INS for diabetes management.139 Zhu et al. developed a multifunctional spherical microgel system through a one-pot free radical polymerization process, utilizing 4-vinylphenylboronic acid (VPBA) and acrylamide (AAm) as monomers, and hydroxyl- and carboxyl-functionalized carbon dots (CDs) along with ethylene glycol dimethacrylate (EGDMA) as cross-linkers.140 The boronic acid-based spherical microgels exhibited exceptional glucose recognition capabilities, thereby facilitating the precise and controlled release of INS.
2.3.2.1. Thermo-responsive. Thermo-responsive delivery systems stand out as one of the most extensively researched smart responsive strategies. Within the tissue microenvironment, the majority of tissue injuries manifest pathological features of temperature elevation due to inflammation. For example, the articular cartilage of patients with OA often exhibits a slight temperature increase compared to healthy tissue.141 The underlying mechanism of thermo-responsive HM delivery systems involves the application of external thermal stimuli, such as radiofrequency ablation (RFA), microwave thermotherapy, and high-intensity focused ultrasound (HIFU), which create a temperature differential between normal and diseased tissues.142 Concurrently, the amphiphilic polymer matrix that forms the system, comprising both hydrophilic and hydrophobic components, allows the thermo-responsive HM system at the target site to undergo a sol–gel phase transition.4 This transition leads to significant alterations in physical and chemical properties, thereby enabling the system to execute its delivery function effectively. The core elements of thermo-responsive delivery systems are polymers that can respond to local temperature changes, referred to as thermo-responsive polymers.143 These include poly(N-isopropylacrylamide) (PNIPAM),119,121,144,145 poly(N-vinyl amides) (PVAM),146,147 agarose (AG),148 CS,97,149,150 and poloxamers,151–153 among others. It is essential that such materials maintain stability in healthy tissues to ensure their efficacy and safety in therapeutic applications.
Thermosensitive matrices are broadly classified into systems that exhibit a lower critical solution temperature (LCST) and those that display an upper critical solution temperature (UCST). In LCST hydrogel systems, the solubility of the polymer demonstrates a non-linear dependence on temperature, decreasing as the temperature rises. When the temperature remains below the LCST, the hydrophilic segments of the polymer chains engage in hydrogen bonding with hydrophilic molecules in the microenvironment, thereby conferring high solubility to the hydrogel. As the temperature rises, the interactions among the hydrophobic elements of the polymer chains intensify, while the hydrogen bonds weaken, culminating in a diminished solubility of the hydrogel.100,118,143,154 In contrast, UCST systems exhibit an increase in matrix solubility with increasing temperature, accompanied by a gel–sol transition. However, UCST systems are operational only at elevated temperatures, which poses a risk of denaturing certain loaded bio-agents, especially proteins and peptides.143,154 As a result, their application as thermogels in the biomedical domain has been sparingly explored. The gel mechanism of amphiphilic polymers can be elucidated within the framework of micelle formation: in LCST systems, as the temperature increases, hydrophobic interactions gain strength, propelling the aggregation of micelles and instigating a sol–gel transition; conversely, when the system temperature attains the UCST, the micellar structure disassembles, and the matrix commences to flow.155,156
Mild heat can stimulate the upregulation of cellular heat shock proteins (HSP-70),157 which play a crucial role in inhibiting the secretion of caspase-3 in mitochondria,158 thereby suppressing both mitochondrial and focal chondrocyte apoptosis. Lin et al. utilized dynamic spiral mosaic technology to develop a spiral mosaic micro/nano HM heat-transfer microneedle (ST-needle) system159 (Fig. 4A). This innovative system harnesses the molecular chain motion responsive to thermal stimuli to modulate the adhesion of non-directional triblock polymers, functioning as a dual delivery system for both thermal energy and biological factors. The ST-needle system capitalizes on its physical properties to effectively penetrate physical barriers and precisely target deep-seated lesions, facilitating the transfer of heat to the affected cells. The applied heat activates the dynamic spiral mosaic mechanism of the ST-needle system, eliminating the mosaic state of the HMS and enabling them to specifically target mitochondria for the release of the PARKIN protein, a pivotal regulator of mitochondrial autophagy. This ensures the simultaneous delivery of both heat and HMS to the lesion site. The synergistic interplay between the applied heat and the release of biological factors from the HMS induces long-term chondrocyte autophagy, effectively removing senescent and damaged mitochondria and inhibiting mitochondrial apoptosis. This process significantly suppresses chondrocyte apoptosis, enhances cell function, and mitigates the symptoms of OA. Presently, the majority of thermogels are confined to responding solely to temperature variations. To address more complex biological applications, there is a pressing need to develop composite thermogel systems that integrate additional sensing capabilities and other bioactive components. A further challenge is the development of thermogels that retain their original injectable properties while achieving high levels of biocompatibility and mechanical strength. Nonetheless, with the advent of advanced technologies such as 3D printing, thermogels are set to make a substantial impact on the burgeoning field of tissue engineering, offering new avenues for therapeutic innovation and application.160
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| Fig. 4 (A) The thermo-responsive ST-needle system, upon reaching deep lesions, removes the mosaic state of the HMS and releases growth factors and thermal energy, which synergistically regulate mitochondria to inhibit chondrocyte apoptosis. Reproduced with permission.159 Copyright 2023, Wiley-VCH GmbH. (B) The structure and response of light-responsive hydrogels. Possible locations of light-responsive groups include (1) at cross-linking points, (2) on the polymer backbone, (3) on side chains, or (4) dissolved in the aqueous medium of the hydrogel. Reproduced with permission.161 Copyright 2019, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. (C) The incorporation of magnetic nanoparticles (MNPs) into hydrogels enables the construction of magneto-responsive hydrogel systems. Reproduced with permission.186 Copyright 2021, Elsevier B.V. (D) US-induced mechanical and thermal effects are harnessed to control drug release and activation. Reproduced with permission.193 Copyright 2021, Henan University and John Wiley & Sons Australia, Ltd. (E) A simplified schematic of the conjugated structure in conductive polymers, featuring alternating double and single bonds with conjugated electrons. Reproduced with permission.236 Copyright 2014, Elsevier Ltd. (F) Mechano-responsive hydrogels applied in biomedical engineering demonstrate diverse response characteristics under external mechanical stimuli. Reproduced with permission.215 Copyright 2020, American Chemical Society. | ||
2.3.2.2. Light-responsive. Light represents an easily accessible and remotely controllable external stimulus, providing a safe, non-invasive, manageable, and highly efficient method for spatiotemporal control.30,151,161–163 This positions it as an exceptionally promising strategy in the design of SRHMS. The properties of light-responsive hydrogels are altered upon exposure to various forms of light, including visible light, UV radiation, and NIR radiation.161 These alterations can induce phenomena such as sol–gel phase transitions and the formation or degradation of hydrogels (Fig. 4B).
One of the most prominent methods for creating light-responsive systems involves the integration of photothermal agents into the hydrogel matrix, a strategy that has garnered extensive research and widespread adoption. Photothermal agents, such as indocyanine green (ICG),164 black phosphorus (BP) nanosheets,165 organic nanoparticles166 and rare metal nanostructures,167 are capable of generating heat through the absorption and emission of light radiation.30 As the temperature increases to the gel phase transition threshold, the hydrogen bonds between the polymer chains and the surrounding water molecules are weakened, precipitating alterations in the hydrogel's volume and water absorption capacity, and triggering a phase transition.118,168 An alternative approach employs ionizable functional groups that respond to incident light. When exposed to light, these gels generate a substantial number of ions. The resultant ionization elevates the osmotic pressure within the matrix, ultimately causing the hydrogel to swell. Furthermore, chromophores can be embedded into the hydrogel matrix. Under the influence of light, the physical and chemical properties of these groups, such as dipole moment and geometry, undergo changes, and the ensuing differences modify the structure of the hydrogel.118 Drawing inspiration from the natural process of “sowing and harvesting”, Ji and his team have innovatively designed artificial hair follicle (HF) seeding light-responsive HMS using microfluidics and photo-crosslinking technology for the purpose of skin tissue repair.169 These microspheres are laden with tideglusib (Ti) and tamibarotene (T), which have the capacity to regulate the fate of fibroblasts, steering their differentiation into dermal papilla cells (DPCs) via the PI3K/AKT pathway, thereby fostering wound healing and HF regeneration.
Before light-responsive hydrogels can be transitioned into clinical applications, several critical challenges must be meticulously addressed. Firstly, the development of photochemical reactions that can be activated by low-energy light remains a formidable obstacle. Secondly, although NIR-responsive hydrogels have achieved remarkable precision in control and depth of penetration, their energy efficiency still falls short of optimal standards. Moreover, the light responsiveness of hydrogels is influenced by a myriad of factors, including their chemical composition, physical properties, 3D architecture, and porosity. The impact of these hydrogels on cell fate is yet to be fully elucidated. Additionally, the construction of precise tissue engineering constructs and the on-demand delivery of live cells present ongoing challenges. Of paramount importance is the need for a comprehensive understanding of the reaction kinetics of these gels to ensure their prompt and dependable response to light stimuli. Furthermore, the multi-wavelength responsiveness of photosensitive hydrogels, the rigorous assessment of their biocompatibility, and the mitigation of photothermal overheating toxicity are pressing issues that demand immediate attention.170
2.3.2.3. Magneto-responsive. Among the diverse substances integrated into hydrogel matrices, systems incorporating magneto-responsive materials deserve particular attention. In contrast to optical stimuli, magnetism emerges as a highly promising exogenous stimulus, characterized by its exceptional tissue penetration and negligible harmful ionizing effects.171 As a result, magneto-responsive materials have found applications in magnetic resonance imaging (MRI),172 magnetic drug delivery,29 enzyme quantification,173 and magnetic hyperthermia therapy.174 Magneto-responsive hydrogel systems, as a form of smart responsive hydrogel, have been introduced into biomedical applications to augment the bioactivity of cells, tissues, or organs. Notable examples include magnetic hyperthermia therapy (MHT) for tumors,175 the repair of neural tissues,176 and targeted drug delivery in cancer treatment.177 This is largely due to their responsiveness to external magnetic fields, which facilitates the acquisition of functional structures capable of remotely modulating the physical, biochemical, and mechanical properties of the environment surrounding cells, tissues, or organs.178–180
By integrating magneto-responsive materials into hydrogels, magneto-responsive hydrogels are constructed (Fig. 4C). These materials are activated by external magnetic fields to induce desired responses in the environment.181,182 Based on the mechanism of magnetic response, external magnetic fields can be classified into two types: a constant magnetic field provided externally and an alternating magnetic field (AMF) applied externally.13,102,183 Magneto-responsive materials can be oxides (such as iron oxide), metals (such as nickel, iron, and cobalt), or nanoparticles of electroplated oxides and metals.102,174,184 Among these, superparamagnetic iron oxide nanoparticles (SPIONs) are the most extensively utilized due to their superior superparamagnetic or ferromagnetic properties, high specific surface area, and biocompatibility.184 When a specific external magnetic field is applied to a biological target, the incorporated magneto-responsive materials typically vibrate or aggregate under the influence of the field, resulting in a sharp increase in local temperature.30,102,154,168 This facilitates the sol–gel transition of the hydrogel and the controlled release of the encapsulated drugs. Consequently, magneto-responsive hydrogel systems often integrate magnetic responsiveness with thermal responsiveness. Zhong and his team innovatively designed a magneto-thermo-responsive HM system by integrating magneto-responsive Fe3O4 nanoparticles with a thermo-sensitive PLGA polymer.185 This intelligent delivery platform achieves synergistic therapy for bacteria-infected skin wounds through the combined action of nitric oxide (NO) smart-release and magnetothermal therapy.
To date, no smart responsive system has achieved perfection, and magneto-responsive HM systems are no exception. While numerous studies have demonstrated the usability of magneto-responsive HMS in both in vitro and in vivo settings, the therapeutic efficacy reported in current research remains confined to animal trials and short-term outcomes. Consequently, the potential application of magneto-responsive HMS in clinical treatment is highly anticipated. A comprehensive evaluation of their long-term pharmacokinetics, toxicology, biodegradation, and metabolic processes in vivo is essential for enhancing their biosafety and minimizing immunogenicity. Moreover, further optimization of their performance is crucial to meet the elevated demands of practical applications. In the realm of drug delivery, magneto-responsive HMS must exhibit accelerated magnetic responsiveness and precise drug-controlled release capabilities to enhance drug delivery efficiency. When employed in cancer hyperthermia, these microspheres require more accurate heating control to precisely target and eliminate tumors while minimizing collateral damage to normal tissues. Simultaneously, the development of multi-responsive magneto-responsive HMS has emerged as a significant trend. By integrating magnetic fields with other external stimuli, these microspheres can achieve more sophisticated and intelligent therapeutic modalities, thereby amplifying their potential in clinical applications.174,186 In summary, the customization and performance enhancement of magneto-responsive HMS will be pivotal to their successful future clinical deployment.
2.3.2.4. Ultrasound-responsive. US refers to mechanical vibration waves that surpass the upper threshold of human hearing, typically with frequencies above 20 kHz.30,187 In biomedical applications, US frequencies range from 0.1 to 50 MHz, with peak negative pressures of 0.01 to 10 MPa.188 Owing to its exceptional tissue penetration, biosafety, non-invasiveness, cost-effectiveness, controllability, and high spatiotemporal resolution, US has been extensively employed in clinical and biomedical fields, including biomedical imaging, cancer treatment, real-time disease monitoring, and targeted drug delivery.30,142,189 A research study has demonstrated that US can also modulate signaling pathways to regulate cell proliferation and differentiation.190 In the realm of injectable smart responsive hydrogels, US, akin to light, heat, and magnetism, is a frequently utilized and highly promising exogenous stimulus. US-responsive hydrogels, as an innovative class of smart hydrogels, amalgamate the properties of US and hydrogels, showcasing multifunctionality, high biocompatibility, and operability. They possess significant potential for achieving targeted, precise, and controlled drug release, thereby demonstrating unparalleled prospects in tissue engineering.189
The mechanisms by which US-responsive hydrogel systems deliver and release therapeutic substances can be primarily categorized into two distinct aspects (Fig. 4D): (a) mechanical effects: the transformation of acoustic energy into mechanical energy, predominantly through cavitation effects, generates mechanical forces that disrupt the structure of the hydrogel carrier, thereby achieving localized drug release.30,142,189,191–193 Cavitation involves the formation of microbubbles within the carrier under US stimuli. These microbubbles oscillate continuously, resulting in high-speed motion or rapid collapse (the former is referred to as stable (or non-inertial) cavitation, while the latter is termed inertial cavitation), thus generating mechanical forces.189,192–194 Additionally, mechanical effects can encompass the utilization of US's mechanical forces, such as acoustic radiation force, to cleave mechanochemically unstable bonds within the carrier, thereby facilitating drug release.191,193,194 (b) Thermal effects: as US traverses cells and tissues, acoustic energy is converted into thermal energy, which can regulate the release of drugs from thermo-responsive hydrogels.30,142,189,191–193 Furthermore, the movement of microbubbles during US cavitation also induces thermal effects.189,193,195 The heat generated by US can further promote drug uptake by perturbing cell membranes, thereby increasing their permeability.142,196,197
Chen and his team, capitalizing on the spatiotemporal oxygen imbalance characteristic of bone defects and the benefits of US technology, utilized microfluidics to incorporate oxygen-loaded nanobubbles, prepared through double emulsification, into the macromolecular network of GelMA/HepMA microspheres.198 Concurrently, they non-covalently attached BMP-2 to create a spatiotemporal HM. This innovative spatiotemporal HM not only boasts superior oxygen-carrying capacity and US responsiveness but also showcases exceptional vascularization and osteogenic potential. Nevertheless, research on US-responsive HMS remains in its nascent stages, and the associated theoretical framework necessitates further exploration and refinement. Firstly, there is a lack of definitive research data on the tolerance range of US intensity for various human organs and tissues, which will constrain the practical application of US-responsive HMS in clinical settings. Secondly, the design of the structure and cross-linking methods of HMS to simultaneously achieve robust mechanical strength and biosafety presents a significant challenge. Moreover, mitigating energy dissipation during US propagation is another critical issue that requires careful consideration.189
2.3.2.5. Electro-responsive. Electrical activity plays a pivotal role in sustaining normal cellular functions, facilitating intercellular communication, transmitting neural signals, and driving numerous complex life processes, such as the activities of cardiac and skeletal muscles.17,30 Despite hydrogels demonstrating exceptional biocompatibility and biosafety, their inherent non-conductive nature restricts their application in the repair and reconstruction of several electrically dependent tissues, including bone, cartilage, heart, nerves, and skeletal muscles.199–201 As a result, electric fields are also harnessed as exogenous stimuli for the design and regulation of smart responsive hydrogel systems. Electro-responsive hydrogels, as multifunctional biomaterials201 with unique conductive properties,202 have garnered considerable attention for their capacity to emulate the electrophysiological environment of biological tissues in electrically dependent tissue repair applications.17 By employing techniques such as grafting, cross-linking, and copolymerization, conductive polymers are polymerized within the porous network structure of hydrogels to create electro-responsive hydrogel systems.202 This amalgamation of hydrogel properties and electroactive characteristics amplifies their suitability for tissue engineering. The incorporation of inorganic conductive nanomaterials into the hydrogel matrix bestows the hydrogels with superior conductivity, thereby facilitating their utilization in a myriad of biomedical applications, including the repair of MI and damaged nervous systems.203,204
The most prevalent conductive polymers include polypyrrole (PPy), polyaniline (PANi), and polythiophene (PTh).30,151,202 In their unmodified state, these polymers demonstrate relatively low conductivity. To regulate and optimize the electroactivity of hydrogels to the desired level, appropriate dopants (oxidizing or reducing agents205) can be incorporated into the conductive polymer network, thereby enhancing the polymer's conductivity by several orders of magnitude.206 Typically, to further augment the electroactivity of electro-responsive systems, inorganic conductive nanomaterials (such as metal nanoparticles207 and carbon nanotubes208) are introduced as auxiliary agents into the conductive polymer-based hydrogel network, a strategy that is commonly employed.17,202 The conductivity of conductive polymer-based hydrogels originates from their continuous and ordered conjugated backbone (Fig. 4E), characterized by alternating double and single bonds and delocalized conjugated electrons, which promote electron flow along the polymer chain.17,30,182 When electro-responsive hydrogels are subjected to an external electric field, electrochemical oxidation/reduction reactions occur within the polymer, inducing directional migration of free ions and generating ion concentration differences, which lead to hydrogel deformation.202 Inorganic conductive nanomaterials within the hydrogel matrix, such as carbon nanotubes, form extensive delocalized π-bonds from the p-electrons of carbon atoms under electrical stimuli.208 This facilitates the delocalization and mobility of electrons between atoms, causing the carbon nanotubes to straighten and align, resulting in the abrupt release of drugs encapsulated within the hydrogel network. The deformation of electro-responsive hydrogels in electrolyte solutions is attributed to the directional migration of free ions in the solution under the electric field, leading to an uneven distribution of ion concentration inside and outside the gel, thereby creating osmotic pressure differences.118 This innovative approach underscores the potential of electro-responsive hydrogels in advancing therapeutic strategies and improving patient outcomes.
Sun et al. developed a novel conductive HM loaded with dental pulp stem cells (DPSCs) and integrated with a wireless generator, incorporating BP nanosheets into the HM matrix.209 The enhanced angiogenic behavior and upregulated anti-inflammatory macrophage polarization demonstrated that electrical signals fine-tune the regenerative niche by improving DPSC-mediated paracrine patterns, thereby further promoting advanced autologous mandibular bone defect regeneration. Electrical activity offers substantial flexibility in the design and preparation of SRHMS. In recent years, extensive research has been conducted on electro-responsive HMS. These systems, characterized by their stimulus responsiveness, biocompatibility, biodegradability, and conductive properties, have exhibited distinct advantages in the repair and reconstruction of electrically dependent tissues. Despite their unique benefits, several challenges remain to be addressed in the future. Firstly, the limited tissue penetration depth of electrical stimuli and the potential for unnecessary tissue damage may constrain the application of electro-responsive HM systems in therapy.210 Additionally, how to further enhance and balance the system's biocompatibility, biodegradability, and mechanical performance warrants careful consideration.208 The underlying principles of how electro-responsive HMS promote cell proliferation and differentiation still require exploration, and extensive research is needed to extend the lifespan and improve the portability of power sources.202 Notably, electro-responsive hydrogels can achieve mutual conversion of electrical and mechanical energy under an electric field, exhibiting excellent piezoelectric effects. Furthermore, non-electrical exogenous stimuli can be combined with electrical stimuli to achieve thermoelectric or photoelectric conversion, or multiple stimulus-responsive systems. This provides innovative ideas for the design and preparation of electro-responsive HM systems.
2.3.2.6. Mechano-responsive. Mechanical forces, including compressive, tensile, and shear forces, are pervasive in biological systems.13,211 At the microscopic level, mechano-responsive proteins such as Piezo1/Piezo2 function as transducers, swiftly converting mechanical stimuli into electrochemical signals and engaging in a spectrum of biological activities, including bone formation, neuronal stimulation, angiogenesis, tumor tissue stiffening, and airway stretch sensing during respiration.212–214 At the macroscopic level, mechanical responsiveness is evident in phenomena such as heartbeat, lung expansion, facial expressions, chewing during eating, and joint/muscle compression during daily activities.215 Drawing inspiration from these natural mechanisms, hydrogel systems endowed with exogenous mechanical responsiveness have increasingly captured attention in the tissue engineering strategy. These systems emulate the adaptive behaviors of tissues and can modulate their physicochemical properties in response to exogenous mechanical cues, including strain hardening,216 self-healing,217,218 shear thinning,219,220 and mechanochromism221–223 (Fig. 4F).
Although these responses all originate from the deformation of the hydrogel network, each one adheres to specific mechanisms involving polymer conformation, intermolecular interactions, and micro/nanostructural changes.215 Strain-hardening hydrogels typically emulate the filamentous networks found in biological materials, with the most straightforward method being the construction of a 3D network based on biomimetic filaments. The formation of semi-flexible fibers can be achieved through self-assembly, where amphiphilic molecules stack into core–shell fibers via hydrogen bonds and hydrophobic interactions.224,225 Strain hardening can also be achieved through the limited extensibility of flexible linkers226 and the repositioning of nanofillers.227 By utilizing non-covalent cross-linking (such as hydrogen bonds,228 metal–ligand coordination229) or dynamic covalent cross-linking (such as Schiff bases,230 boronate esters231), mechano-responsive hydrogels can exhibit self-healing and/or shear-thinning properties. When shear force is applied, the relatively weak cross-links are the first to break, leading to high mobility of the dissociated polymer chains and a reduction in the viscosity of the hydrogel, a phenomenon known as shear-thinning.219 Upon removal of the mechanical stimulus, due to the excellent reversibility of the dissociated bonds, they can spontaneously recombine when the broken ends come into contact, thereby re-establishing the hydrogel network and achieving self-healing.232 Typically, shear-thinning and self-healing properties are combined. The ability of mechanochromic hydrogels to change color under mechanical force usually stems from two mechanisms: the color change of incorporated mechanophores (molecules containing mechanically unstable bonds with color-regulating properties)233 and the switchable structural coloration234 in photonic gels.
Strain-hardening and self-healing hydrogels, which inherit the physiological properties of biological tissues, stand out as one of the most promising options in the tissue engineering strategy. The shear-thinning property endows hydrogels with injectability, making targeted regional delivery highly significant for minimally invasive treatments. Lei et al. developed self-regenerating HA HMS loaded with rapamycin-liposomes using microfluidics and photopolymerization processes.24 Notably, this system demonstrates self-renewing lubricity. Within the frictional forces of the OA microenvironment, the system achieves self-renewal, and the mixed HMS enhance joint lubrication through a rolling mechanism. The visualization of mechanical stress in mechanochromic hydrogels offers substantial utility in biological detection, monitoring, and diagnosis.215 Drawing inspiration from jellyfish tentacles, Guo et al. utilized microfluidic spinning technology to design hydrogel microfibers embedded with discrete structural colored inverse opal HM units.235 When local forces (such as tension or pressure) are applied to the microspheres, the hydrogel microfiber system exhibits synchronous changes in the photonic bandgap (PBG) and structural colors, enabling spatial tactile sensing. This innovation provides valuable inspiration for the application of mechano-responsive HMS in dynamic monitoring and spatial sensing within tissue engineering. However, the application of mechano-responsive HM systems still encounters challenges, such as the limited range of mechanical responsiveness and the need to prevent burst release or leakage release in the absence of stimuli.
3. Characteristics of smart responsive hydrogel microspheres
It is widely recognized that SRHMS exhibit a multitude of exceptional biochemical and physical properties in tissue engineering, including biocompatibility, biodegradability, bioactivity, hydrophilicity, mechanical strength, smart responsiveness, loading and delivery capabilities, and lubricity, among others. While the aforementioned characteristics have been discussed, the aspects of lubricity and bio-agent delivery function have not yet been addressed. In the following section, we will provide a brief overview of the lubricity and the bio-agent delivery function of SRHMS.3.1. Lubricity
The lubricity of biomaterials plays a significant role in influencing the repair efficacy of damaged tissues.10 Efficient biolubrication at biological contact surfaces is essential for the restoration of tissue functions.65,237 Tissue engineering is in dire need of promising biolubricants to mitigate friction at biological contact surfaces, such as joints, the oral cavity, eyes, and the gastrointestinal tract, thereby alleviating wear and discomfort in epithelial and cartilage tissues, among others.238,239 HMS can provide exceptional lubricity, offering effective hydration lubrication.240 Hydration lubrication forms a hydrated layer on biological contact surfaces that can endure high pressures in the microenvironment, attributed to the electrostatic interaction between microspheres and water molecules when they contact biological surfaces.240 The ball-bearing effect provides novel insights into the lubrication mechanism of microspheres. HMS, being micron-sized spherical hydrogels with relatively uniform dimensions, can diminish friction at biological contact surfaces through their ball-bearing mechanism, rendering them highly attractive as biolubricants65,240,241 (Fig. 5A). | ||
| Fig. 5 (A) Superlubricating HMS engineered based on the ball-bearing effect. Reproduced with permission.241 Copyright 2020, Wiley-VCH GmbH. (B) RAPA@Lipo@HMs integrate hydration lubrication and ball-bearing lubrication while preserving chondrocyte homeostasis to treat OA. Reproduced with permission.24 Copyright 2022, American Association for the Advancement of Science. (C) Schematic representation of the hydration lubrication mechanism of CAP/FGF18-hyEXO@HMs. Reproduced with permission.67 Copyright 2024, Wiley-VCH GmbH. | ||
In comparison to traditional hydrogels, HMS possess a larger biological contact area and superior dispersibility, offering greater natural advantages for biolubrication.65 In terms of lubrication performance, the key distinctions between HMS and conventional hydrogels lie in interfacial interactions, friction mechanisms, and rheological properties. HMS exploit a ball-bearing effect, wherein microscale rolling motion minimizes interfacial friction, facilitating elastohydrodynamic lubrication coupled with hydration lubrication.67,241 By contrast, conventional hydrogels depend predominantly on hydration lubrication at biological interfaces, encompassing boundary lubrication and fluid-film lubrication mechanisms.242–244 The spherical morphology of HMS promotes uniform stress distribution and enhanced load-bearing capacity, yielding a lower COF (0.02–0.03) under dynamic shear,24,67,240 while conventional hydrogels exhibit a comparatively higher COF (0.02–0.06).239 Furthermore, the mechano-responsive properties of HMS allow for dynamic self-replenishment of hydration layers, enabling real-time lubrication tuning and friction reduction,24 whereas conventional hydrogels rely mainly on static swelling for lubrication.245 Regarding interfacial interactions, the relative fluidity and flexibility of uniformly structured HMS enhance the stability of their electrostatically bound hydration layers in high-pressure microenvironments.11,246 Conversely, conventional hydrogels, owing to their heterogeneous network architecture, display inferior mechanical robustness, with hydration layers susceptible to shear-induced degradation and consequent lubrication failure.239 Furthermore, the natural gaps between HMS can augment their lubricity in tissue engineering applications.10 The degree of biolubrication can be regulated not only by the size, dispersibility, volume fraction, and surface roughness of the microspheres but also by their mechanical responsiveness.67,241 Consequently, SRHMS emerge as formidable contenders for effective biolubricants in tissue engineering.
Effective biolubrication in joints is essential for daily activities, and is primarily sustained by synovial fluid and articular cartilage.65 However, in the pathological environment of OA, the lubricating properties of cartilage degrade, resulting in persistent joint wear, degenerative changes, and ultimately irreversible damage.240 Lei et al. developed self-regenerating liposome-HMS loaded with rapamycin24 (Fig. 5B). The smooth rolling mechanism of the microspheres, combined with the continuous self-regenerating lipid hydration layer formed through friction, enhances joint lubrication, thereby mitigating OA. In a wear test, the Lipo@HM group exhibited narrower and shallower wear tracks compared to both the PBS group and the HM group. The Lipo@HM group, characterized by its optimally exposed lipid vesicles, demonstrated superior lubricity, resulting in a lower COF (0.03) relative to the PBS group (0.06) and the HM group (0.04). Chen et al. employed microfluidics to design a self-regenerating HM system (CAP/FGF18-hyEXO@HMs) that promotes the expression of fibroblast growth factor 18 (FGF18) for the treatment of OA67 (Fig. 5C). This system not only provides sustained lubrication to address the frictional wear of articular cartilage but also prevents ECM degradation and reduces inflammation.
In addition to mechano-responsive HM systems being employed for the treatment of OA, HM systems with other response mechanisms have also shown significant promise in providing lubrication and promoting cartilage repair. Drawing inspiration from the structure of chocolate-covered peanuts, Miao and his team designed a MMP-responsive double-layer chondroitin sulfate methacrylate (ChsMA) HM system.246 This system is designed to respond to the OA microenvironment, facilitating the localized release of celecoxib-loaded liposomes and Chs. It not only demonstrates anti-inflammatory and cartilage repair-promoting properties but also offers lubrication. Inspired by the wind-dispersal mechanism of dandelion seeds, Bi et al. employed microfluidics to develop a pH and ROS dual-responsive HM system loaded with cyclic peptide cortistatin-14 (CST-14).247 In the OA microenvironment characterized by inflammatory progression (pH < 5), the microsphere system responsively releases the drug and scavenges ROS, inhibiting TNF-α signaling to suppress OA while providing lubrication to further protect cartilage tissue.
3.2. Delivery of cells/drugs
Biocompatible SRHMS, upon activation by endogenous or exogenous stimuli, can precisely target and release their cargo, facilitating the intelligent delivery of therapeutic cells or drugs (non-cellular agents) for a wide range of tissue engineering strategies (Fig. 6). | ||
| Fig. 6 Schematic representation of the role of SRHMS in the delivery of bio-agents (cells or drugs) for tissue engineering. Created in https://BioRender.com. | ||
As previously discussed, a range of SRHMS fabrication techniques are compatible with cell encapsulation. Microfluidics and lithography are commonly utilized to produce cell-laden HMS, as they enable precise control over the size and shape of the microspheres, accurate regulation of the number of cells within each microsphere, even down to the single-cell level, and ensure uniform cell distribution.64 The controllable porosity and structure of the microspheres are essential for optimal cell loading and regulated release.248 The pore spaces within the microspheres facilitate the rapid supply and diffusion of nutrients, supporting the vital activities of the encapsulated cells.65 Moreover, the microspheres demonstrate excellent biocompatibility and can function as a 3D culture system, mimicking the natural ECM and promoting robust cell–cell and cell–matrix interactions, thereby serving as a scaffold for cell proliferation and differentiation.253 As effective cell carriers, the microspheres provide mechanical and physical support to the encapsulated cells, shielding them from shear forces during injection and offering protection after injection into the recipient tissue.64,65 In damaged tissues, under endogenous or exogenous stimuli, the microspheres gradually degrade in vivo, enabling the controlled and sustained release of cells. Consequently, SRHMS can efficiently deliver cells and are widely applicable in various tissue engineering strategies.
In tissue engineering, the selection of optimal cell delivery strategies is critical for achieving desired therapeutic efficacy. Microsphere-based delivery platforms can be broadly categorized into two distinct paradigms: single-cell-on-microsphere and spheroid-on-microsphere systems, each demonstrating unique structural configurations and biological performance characteristics (Table 1). The single-cell delivery approach involves the transplantation of isolated individual cells to target tissues using biocompatible scaffolds, typically HMS.254 In contrast, spheroids are complex spherical cellular aggregates that, within an in vitro 3D matrix microenvironment, can support cell–cell and cell–matrix interactions, and self-assemble into modular cellular units that resemble native tissue in structure and function.72 These pre-formed spheroids can be effectively integrated with biomaterial scaffolds through either surface attachment or encapsulation for subsequent therapeutic delivery. The spheroid-on-microsphere platform offers several distinct advantages: (1) it recapitulates the native tissue architecture by establishing an intricate ECM-mimetic microenvironment that promotes intensive intercellular communication; (2) the resulting multicellular-matrix network enhances cellular functionality and upregulates tissue-specific gene expression profiles;72 and (3) the mechanically robust structure demonstrates superior resistance to physiological stresses, thereby improving cell viability and therapeutic potential in dynamic in vivo settings.253 However, this system faces inherent limitations, including the development of hypoxic cores due to restricted nutrient diffusion in densely packed cellular aggregates, which may precipitate central necrosis.65 While conventional fabrication methods (e.g., hanging drop, rotational culture) often yield spheroids with suboptimal uniformity, emerging technologies such as microfluidic systems, electrohydrodynamic processing, and precision 3D bio-printing have shown promise in generating more homogeneous constructs.248,253 Conversely, single-cell-on-microsphere systems, typically manufactured using high-precision microfluidic techniques, offer superior monodispersity and dimensional consistency. The reduced particle size enhances tissue penetrability, rendering this approach particularly suitable for personalized medicine applications.65,66 Nevertheless, these systems exhibit notable drawbacks, including structural vulnerability under physiological shear forces and the absence of critical intercellular signaling pathways that are essential for proper tissue development and function.72
| Single-cell-on-microsphere system | Spheroid-on-microsphere system | Ref. | |
|---|---|---|---|
| Fabrication | Microfluidics, 3D bio-printing, electrospray | Hanging drop, rotational culture, nonadherent surface, spinning flask, microfluidics, 3D bio-printing | 65, 66, 72, 248, 253 and 254 |
| Advantages | 1. Superior monodispersity and dimensional consistency | 1. Recapitulates native tissue architecture and intercellular communication | |
| 2. Enhanced tissue penetrability | 2. Enhances cellular functionality and gene expression | ||
| 3. Suitable for personalized medicine | 3. Mechanically robust, resistant to physiological stresses | ||
| Limitations | 1. Structurally vulnerable to shear forces | 1. Hypoxic cores and central necrosis due to limited nutrient diffusion | |
| 2. Lack critical intercellular signaling pathways | 2. Suboptimal uniformity in conventional methods | ||
| Applications | Personalized therapies requiring high tissue penetration | Complex therapies requiring native tissue-like functionality |
In various cell-based strategies, stem cells have emerged as the most popular tool for treating a wide range of diseases.237,255 Stem cells cultured in 3D-structured scaffolds can significantly enhance therapeutic efficacy, making the targeted delivery of stem cell spheroids highly promising.55 Mu et al. developed a ROS-responsive conductive HM carrier loaded with adipose-derived stem cells (ADSCs) for the repair of myocardial tissue in MI.256 The carrier also encapsulated the anti-inflammatory agent salvianolic acid B (SalB), demonstrating excellent biocompatibility, promoting cell proliferation and differentiation, and exhibiting antioxidant and anti-inflammatory properties. When exposed to a 0.1% v/v H2O2 solution, the microsphere carrier network begins to disintegrate, resulting in the release of over 90% of SalB within 3 days. This precisely timed release coincides with the therapeutic window for acute inflammatory intervention. By enhancing the expression of connexin 43 (Cx43) and CD31, they promoted ventricular wall regeneration, reduced myocardial fibrosis, and remodeled infarcted myocardium. Sun et al. designed a conductive HM loaded with DPSCs and integrated with a radiofrequency generator, which promotes autologous maxillofacial tissue regeneration, representing a promising electroactive stem cell scaffold therapeutic strategy.209 Yang et al. reported a ROS- and pH-responsive system with self-healing properties, combining HMS loaded with mesenchymal stem cells (MSCs) and a HA hydrogel modified with PBA, for abdominal wall repair.254 This system not only protects MSCs from ROS but also promotes angiogenesis and Col deposition, and exhibits hemostatic and anti-inflammatory properties. A study reported a uniform cell aggregate formed by a glucose-responsive HM scaffold incorporating MSCs.57 This is a highly promising tissue engineering scaffold system, where the cell aggregates can release growth factors on demand under glucose stimuli.
Hydrogel delivery systems have been extensively utilized for drug delivery in tissue engineering, including growth factors, protein drugs, gene drugs, ionic drugs, and other pharmaceuticals. HMS are frequently employed as drug delivery systems (e.g., for small molecule drugs, growth factors, etc.) owing to their capacity to controllably protect drug activity and deliver drugs to damaged tissues, achieving multiple desired local drug release profiles and degradation behaviors.64,237,248 The size of the HMS, the mesh size within the 3D network, and the molecular interactions (covalent bonding, electrostatic, and hydrophobic interactions) between the drug and the microspheres all influence the drug release rate from the microspheres.64 Additionally, combining microspheres with other drug delivery carriers (such as liposomes) can enhance the flexibility of local drug delivery systems.10 Drug delivery is also one of the widely explored areas of SRHMS. Various endogenous or exogenous stimuli can trigger the hydrolysis or conformational changes of drug-loaded carriers, enabling spatiotemporal drug release.13 SRHMS, as a member of the HMS family, are extensively utilized in the field of drug delivery. Compared to conventional hydrogels, their unique structure endows them with a high specific surface area and high porosity, providing abundant drug binding sites and significantly enhancing drug loading capacity. Furthermore, monodisperse microspheres fabricated using advanced techniques like microfluidics can further mitigate issues of uneven drug encapsulation, avoiding drug accumulation or leakage often caused by network heterogeneity in traditional hydrogels.239,248 The drug release from conventional hydrogels primarily relies on matrix swelling/degradation and passive diffusion.265 When introduced into dynamic physiological microenvironments, their release profiles often exhibit a significant initial burst release, while subsequent release rates prove difficult to precisely regulate.64 The smart responsive network structure of SRHMS is central to achieving precise drug release. SRHMS, whether surface-modified or integrated with smart responsive materials, undergo hydrolysis or conformational changes triggered by various endogenous or exogenous stimuli, achieving drug-specific release at the site of damaged tissue.13 Through meticulous design, SRHMS systems can be engineered for sensitivity or specific response thresholds to different stimuli, enabling multistage release regulation. This eliminates the initial burst release and extends the drug's therapeutic duration.53,254 External stimuli facilitate preliminary localized drug release, while internal stimuli enable deep, precise release – a kinetic profile better aligned with therapeutic requirements.24,266 Conventional hydrogels typically lack active targeting capabilities, resulting in limited efficiency and specificity. SRHMS, however, can be functionalized with smart responsive materials to enhance the flexibility of drug delivery and achieve targeted specificity. SRHMS can not only establish “intelligent targeting” based on the pathological microenvironment (e.g., in response to endogenous stimuli like ROS or MMPs),263,267 but can also leverage external physical conditions (such as US or magnetic fields) to spatially enrich drugs at the target location.185,268
Lei et al. incorporated RAPA into a self-regenerating liposome-HM system.24 Cationic liposomes can target negatively charged articular cartilage via electrostatic interactions, releasing the autophagy agonist RAPA to enhance autophagy and maintain cellular homeostasis. During a 3600-second friction test, RAPA@Lipo@HMs initiated a response to mechanical friction in the initial phase (0–300 s), characterized by a gradual decline in the COF. The system achieved a stable COF of 0.03 throughout the subsequent phase (300–1200 s). Furthermore, RAPA@Lipo@HMs exhibited a biphasic drug release profile and extended the release duration to 28 days, significantly longer than the 14-day release observed with liposomes alone. Furthermore, Miao et al. designed a ChsMA HM system that responds to MMPs in the microenvironment, locally releasing CLX-loaded liposomes, offering both anti-inflammatory and lubricating effects.246 In the presence of MMP-13, the microsphere system demonstrated a sustained release profile, with 76.1 ± 3.4% of the encapsulated CLX being released during the initial 7-day period, followed by gradual release of the remaining payload over the subsequent 28 days. The cumulative release reached 93.1 ± 1.6% by day 35. In parallel, the microsphere core exhibited progressive degradation while continuously releasing ChsMA, achieving complete degradation within approximately 28 days. Guo et al. developed an US-responsive multifunctional HM bomb (EMgel).269 Under US stimulation, the microspheres release the encapsulated natural polyphenol EGCG and bioactive MoS2 to repair chronic osteomyelitis caused by methicillin-resistant Staphylococcus aureus (MRSA) infection. Wu et al. encapsulated CUR and ciprofloxacin (CIP) in a pH-responsive and thermo-responsive HM system, which releases the drugs in response to the acidic environment of bacteria and heat generated by NIR for antibacterial treatment, improving the therapeutic effect on infected wounds.258 Under acidic conditions (pH 5.0), CIP and CUR demonstrated significantly improved cumulative release over 48 hours compared to physiological pH (7.4). This pH-dependent release behavior was further enhanced by mild photothermal stimulation from NIR irradiation, which synergistically promoted drug liberation from the delivery system. Moreover, Li et al. encapsulated a mixed compound of Ti and T in light-responsive HMS, providing an effective strategy for skin wound healing and in situ HF regeneration.169 The drug delivery system demonstrated sustained release kinetics, with cumulative release reaching 80–95% of the total payload within a 5-day period. Wang et al. encapsulated SOR and DOX in a photothermo-responsive HM system, significantly reducing tumor cell viability and enhancing the therapeutic efficacy for gastric cancer.260
Zheng et al. engineered MMP-responsive GelMA HMS encapsulating lipid nanoparticles (mRNA@LNP) loaded with ubiquitin-specific protease 26 (USP26) mRNA.263 The released mRNA@LNP upregulates USP26 protein expression to regulate β-catenin and Iκb-α, balancing osteoblast–osteoclast (OB–OC) crosstalk, thereby promoting intervertebral fusion in a rat model of OP. Li et al. formed a G5-AHP/miR-140 nanoparticle complex by combining a multifunctional gene carrier, arginine, histidine, and phenylalanine-modified G5-AHP with miR-140, and subsequently encapsulated the nanoparticles within MMP-responsive HMS, offering a novel non-cellular OA treatment strategy.39 Upon exposure to a 10 ng mL−1 MMP solution, the system displayed rapid release kinetics, achieving near-complete nanoparticle release within five days via enzymatic degradation. Yang and his team designed injectable MMP-1-responsive HMS (KGE) composed of a self-assembling peptide (KLDL-MMP-1), GelMA, and BMSC-Exos, aimed at promoting neovascularized bone healing.129 MMP-1-triggered release profiles demonstrated that KGE underwent rapid initial release, attaining maximal concentrations by day 3, followed by sustained release kinetics that plateaued at 92.29–96.29% cumulative release by day 9. Chen and her team constructed chondrocyte-affinity peptide (CAP) incorporated hybrid Exos loaded with an FGF18-targeted gene-editing tool and encapsulated them within self-regenerating HMS.67 In the OA microenvironment, the system provides continuous lubrication in response to frictional wear while releasing its cargo, promoting cartilage regeneration and preventing ECM degradation. Chen et al. developed an oxygen-carrying US-responsive spatiotemporal HM system loaded with BMP-2.198 Notably, the concentration of oxygen released by the system increases with the intensity of US. Under US irradiation at intensities of 1, 2, 3, and 4 W, the dissolved oxygen concentration increased progressively in both aqueous solution and HMS. Specifically, oxygen levels in water rose by factors of 1.51, 1.90, 2.17, and 2.40, whereas in HMS, the increases were 1.63, 1.95, 2.11, and 2.29-fold, respectively. In hypoxic bone defect models, the spatiotemporal HMS can regulate oxygen homeostasis disorders and maintain high levels of VEGF expression, demonstrating excellent angiogenic and osteogenic capabilities. Sun et al. loaded PTX and VEGF into a photothermo-responsive HM composite system, developing an emerging transplant material for the treatment of spinal cord injury (SCI).262
4. Applications of smart responsive hydrogel microspheres for tissue regeneration
Micron-sized SRHMS can showcase their multifunctionality across a range of applications in the tissue engineering strategy (Fig. 7). As injectable materials, uniformly spherical SRHMS demonstrate exceptional biocompatibility, lubricity, and targeting capabilities, significantly reducing tissue damage. The adjustable internal and external structures, combined with their excellent mobility, enable these microspheres to be evenly distributed in targeted tissues and organs, effectively filling irregular defects. Furthermore, the spherical structure, characterized by a high surface area and porosity, facilitates efficient loading and delivery of bio-agents. Under endogenous or exogenous stimuli, SRHMS regulate the tissue microenvironment and promote tissue repair and regeneration through spatiotemporally controlled cargo release. In this section, we will explore the applications of SRHMS in bone, cartilage, skin, nerve, cardiac, and other tissue engineering fields (Table 2). | ||
| Fig. 7 Multifunctional SRHMS utilized in bone, cartilage, skin, nerve, cardiac, and other tissue engineering fields. Created in https://BioRender.com. Hydrophilicity and Lubricity (Hydration lubrication). Reproduced with permission.24 Copyright 2022, American Association for the Advancement of Science. | ||
| Tissue engineering | Disease | Smart responsive mechanism | Materials | Fabrication techniques | Bio-agents | Ref. |
|---|---|---|---|---|---|---|
| Bone tissue engineering | OP | pH-responsive | Alg and CS | Microfluidics | A BMSC-affine peptide conjugated onto liposomes encapsulating Fisetin | 43 |
| Postmenopausal osteoporosis (PMOP) | pH-responsive | Alg and CS | Microfluidics | Polyhedral oligomeric silsesquioxane (POSS) | 270 | |
| Bone healing | Enzyme-responsive (MMP-1) | GelMA and self-assembling peptides (SAP) | Microfluidics | Exos derived from BMSCs | 129 | |
| OP | Enzyme-responsive (MMP) | GelMA | Microfluidics | USP26 mRNA@LNP | 263 | |
| Bone defects | Magneto-responsive | PLGA | Emulsion technique | SPIONs | 271 | |
| Large bone defects | US-responsive | GelMA and HepMA | Microfluidics | BMP-2 | 198 | |
| MRSA-infected chronic osteomyelitis | US-responsive | HAMA | Microfluidics | Natural polyphenolic EGCG and bioactive MoS2 | 269 | |
| Bone defects | US-responsive and microenvironmental responsive | PLGA | Emulsion technique | BMP-2 and MnO2 | 266 | |
| Mandibular bone defect | Electro-responsive | GelMA | Microfluidics | DPSCs | 209 | |
| Cartilage tissue engineering | OA | pH-responsive | GelMA | Microfluidics | Itaconate (IA) | 272 |
| OA | pH-responsive and ROS-responsive | Dibenzaldehyde PEG (DFPEG), CS and PEG diacrylate (PEGDA) | Microfluidics | CST-14 | 247 | |
| OA | ROS-responsive | HAMA | Microfluidics | Antioxidants | 273 | |
| OA | ROS-responsive and enzyme-responsive (MMP) | GelMA, HAMA and benzenediboronic acid (PBA) | Microfluidics | Dihydromyricetin (DMY) | 274 | |
| OA | Enzyme-responsive (MMP) | GelMA | Microfluidics | G5-AHP/miR-140 nanoparticles | 39 | |
| OA | Enzyme-responsive (MMP) | ChsMA and GelMA | Microfluidics | CLX and Chs | 246 | |
| OA | Electro-magneto-responsive | HAMA | Microfluidics | Sulfhydryl POSS (POSS-SH) linked with PEG, kartogenin (KGN), hydrogenated soya phosphatidylcholine (HSPC), and fluorescein (PPKHF) | 275 | |
| OA | Mechano-responsive | HAMA | Microfluidics | RAPA@Lipo | 24 | |
| OA | Mechano-responsive | HAMA | Microfluidics | CAP/FGF18-hyEXO | 67 | |
| Skin tissue engineering | Infected wounds | pH-responsive and thermo-responsive | PLGA | Emulsion technique | CUR and CIP | 258 |
| Chronic diabetic wounds infected by bacteria | pH-responsive | Alg | Emulsion technique | CUR and tetracycline hydrochloride (TH) | 259 | |
| Diabetic wounds | Enzyme-responsive (MMP-9) | Gel | Emulsion technique | H8 macrophage membrane-derived nanovesicles (H8NVs) | 276 | |
| Chronic diabetic wounds | Enzyme-responsive (MMP-9) and glucose-responsive | Gel | Emulsion technique | INS and CLX | 277 | |
| Bacterially infected wounds | Thermo-responsive | Alg | Microfluidics | Lysozyme and MXene | 278 | |
| Skin wounds | Light-responsive | GelMA | Microfluidics | Ti and T | 169 | |
| Infected wounds | Magneto-responsive | CS | Microfluidics | Zn2+ and VEGF | 29 | |
| Bacteria-infected skin wounds | Magneto-thermo-responsive | PLGA | Emulsion technique | NO | 185 | |
| Neural tissue engineering | Neuropathy and neurological complications | pH-responsive | Carboxylated cellulose | The sol–gel method with freeze-drying technology | Vitamin B12 | 279 |
| SCI | Enzyme-responsive (MMP-2/9) | Gel | Emulsion technique | MG53 protein | 280 | |
| SCI | Photothermo-responsive | Gellan gum (GG) | Microfluidics | PTX and VEGF | 262 | |
| Neurodegenerative diseases and traumatic brain injuries | Magneto-responsive | Poly(ε-caprolactone) (PCL), PLGA, Gel and GelMA | Electrospray | Human neural stem/progenitor cells (hNSCs) | 281 | |
| Cardiac tissue engineering | AMI | pH-responsive, H2O2-responsive and enzyme-responsive (MMP-9) | PEG and DSPE | Emulsion technique | Sparchigh Treg-derived EVs | 53 |
| MI | ROS-responsive | PEG diacrylate (Ac-PEG5000-Ac) and a hydrogen-bonded cross-linker (PEG-UPy) containing ureacymidone (UPy) | Microfluidics | AOA | 117 | |
| MI | ROS-responsive | GelMA | Microfluidics | PEG-TK-AOA | 282 | |
| Myocardial damage resulting from AMI | Thermo-responsive and ROS-responsive | PLGA-PEG-PLGA system and PVA | Emulsion technique | Tanshinone IIA | 59 | |
| MI | ROS-responsive and electro-responsive | CS and DEX | Emulsion technique | SalB and ADSCs | 256 | |
| MI | Thermo-responsive | Poly(l-lactic acid)-b-poly(ethylene glycol)-b-poly(N-isopropyl acrylamide) (PLLA-PEG-PNIPAm) tri-block copolymers | Emulsion technique | hESC-derived CMs | 283 | |
| MI | US-responsive | PLGA-heparin-PEG- cyclic arginine-glycine-aspartate-platelet (PLGA-HP-PEG-cRGD-platelet) and perfluorohexane (PFH) | Microfluidics | Basic fibroblast growth factor (bFGF) | 268 | |
| Other tissue engineering (e.g., tumors, gastrointestinal diseases, etc.) | Hepatocellular carcinoma (HCC) | pH-responsive | Gel | Emulsion technique | Thrombin loaded into calcium carbonate (CaCO3) | 284 |
| HCC | pH-responsive | PLGA | Emulsion technique | Tanespimycin (17-AAG) | 285 | |
| HCC | pH-responsive and thermo-responsive | SA-modified SF | Emulsion technique | Adriamycin hydrochloride | 49 | |
| Tumor malignant ascites | Thermo-responsive | AG | Microfluidics | Methotrexate-packaging tumor-cell-derived microparticles (MTX-TMPs) and BP quantum dots (BPQDs) | 148 | |
| Gastric cancer | Photothermo-responsive | HAMA and GelMA | Microfluidics | SOR and DOX | 260 | |
| Tumor | Photothermo-responsive | Alg | A microfluidic electrospray strategy | NO donors (S-nitrosoglutathione, GSNO), BP and DOX | 286 | |
| Colon cancer | Light-responsive | Glycidyl methacrylate-conjugated xanthan gum | Electrospray | CB1 agonist arachidonoyl 2′-chloroethylamide (ACEA) | 287 | |
| IBD | pH-responsive | Alg | Mechanical fragmentation | Halloysite clay nanotubes (HNTs), EGCG and probiotics | 82 | |
| Diarrhea-predominant irritable bowel syndrome (IBS-D) | ROS-responsive | Alg and CS | Emulsion technique | Puerarin | 267 | |
| IBD | NO-responsive | Poly-γ-glutamic acid | Microfluidics | Probiotics | 288 | |
| IBD | Inflammatory microenvironment-responsive | Gel and HA | Emulsion technique | MXene and L-arginine | 289 | |
| IBD | Inflammatory microenvironment-responsive | Pectin | Emulsion technique | Quercetin | 290 | |
| Abdominal wall defects | pH-responsive and ROS-responsive | GelMA and SF methacrylate (SFMA) | Microfluidics combined with freeze-drying technology | Mesenchymal stem cells (MSCs) | 254 | |
| Type 1 diabetes mellitus | Glucose-responsive | Col and TA | 3D bio-printing | β-Cell | 86 | |
| Type 2 diabetes mellitus | Glucose-responsive | CS | Microfluidics | INS | 138 | |
| Diabetes | Glucose-responsive | Dex-GMA, PEG (600) diamethacrylate (PEGDMA), Con A, cinnamic acid modified dextran (Dex-CCH) and PVA | Emulsion technique | INS | 139 | |
| Tendon adhesion | Enzyme-responsive (MMP-2) | GelMA | Microfluidics | Smad3-siRNA nanoparticles | 291 | |
| Periodontitis | Thermo-responsive | CS and vanillin (VC) | Emulsion technique | Ornidazole and doxycycline hyclate | 292 |
4.1. Smart responsive hydrogel microspheres for bone tissue regeneration
The skeletal system is a crucial organ responsible for weight-bearing, movement, mineral storage, and maintaining the body's integrity. Bone defects resulting from trauma, infection, tumors, and congenital deformities present significant challenges in tissue engineering. Although bone tissue has a certain capacity for self-repair, its healing ability may diminish or be suppressed as the severity of the injury increases, particularly in cases of critical bone defects that lead to severe skeletal damage.293 Clinical treatments for bone defects encompass non-invasive therapies and surgical interventions. Non-invasive therapies stimulate bone tissue regeneration by promoting the proliferation and differentiation of osteoblasts and the expression of related genes through exogenous stimuli, including electromagnetic field therapy, low-intensity US therapy, and thermotherapy.294 Surgical intervention is the most frequently employed method for treating bone defects, with autologous bone grafting being the most effective approach, as the grafted bone shares similar mechanical and biological properties with the host bone.72,295 However, autologous bone grafts are associated with the risk of fracture and may not fully repair the damaged area.17 In some instances, the scarcity of donor tissue and potential secondary trauma also limit the clinical application of autologous bone grafting.72,294Bone repair is a complex and dynamic process that involves intricate interactions between cells, scaffolds, biological cues, and damaged tissues.237 With the advancements in bone tissue engineering, materials encapsulating bio-agents have introduced novel therapeutic strategies for bone repair. In recent years, SRHMS have emerged as highly promising materials in bone tissue engineering, attracting increasing research attention. Compared to traditional bone repair methods, SRHMS offer several significant advantages: (a) as scaffolds, the structure of SRHMS closely mimics the ECM, facilitating cell adhesion, proliferation, differentiation, and stability, while also promoting angiogenesis and mineral deposition, thereby aiding in the repair of defective bone tissue; (b) their excellent hydrophilicity and biodegradability make them ideal carriers for the local delivery and release of bio-agents, eliminating the need for additional surgery to remove them; (c) compared to some bone graft materials, their superior biocompatibility can alleviate or even prevent immune rejection reactions in the body; and (d) they can respond to certain endogenous stimuli in the pathological environment of bone defects, such as ROS, pH, and temperature changes, as well as exogenous stimuli, enabling spatiotemporally controlled release of their bio-agents.
Li et al. employed a POSS nanoplatform to synthesize multifunctional organic–inorganic hybrid nanoparticles (PDAP NPs) via Michael addition reactions.270 They successfully fabricated pH-responsive core–shell structured micro/nano HMS (PDAP@Alg/Cs) loaded with PDAP NPs using gas microfluidics and ionic crosslinking techniques, aimed at preventing and treating PMOP (Fig. 8A). Due to estrogen deficiency, PMOP is characterized by pathological changes such as reduced bone tissue blood circulation and increased osteoclast activation. This HM system activates the HIF-1α/VEGF signaling pathway to promote H-type blood vessel formation and upregulates heme oxygenase-1 (HO-1) expression through the p38 MAPK signaling pathway to inhibit bone resorption caused by increased osteoclast activation, effectively alleviating bone loss. Additionally, it exhibits satisfactory tolerance in the gastric environment and achieves controlled slow release in the intestines, while demonstrating significant targeting ability to bone tissue, thereby effectively treating PMOP. Zheng and his team believe that effectively regulating OB–OC crosstalk is crucial for restoring bone tissue structure and function.263 They loaded USP26 mRNA@LNP and subsequently encapsulated them in MMP-responsive GelMA HMS, forming a bone homeostasis repair microcarrier (BHRC). Notably, USP26 not only stabilizes β-catenin through deubiquitination to promote osteogenic differentiation but also inhibits IκBα ubiquitination and degradation to suppress osteoclast differentiation, enabling the BHRC to coordinate OB–OC crosstalk and effectively promote vertebral growth in rats (Fig. 8B).
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| Fig. 8 (A) Micro-CT of the distal femur (i) reveals that the PDAP@Alg/Cs group demonstrated the most significant increase in bone mineral density (BMD, (ii)) and bone volume fraction (BV/TV, (iii)), along with the largest reduction in trabecular separation (Tb. Sp, (iv)). Reproduced with permission.270 Copyright 2023, Wiley-VCH GmbH. (B) BHRC promotes osteogenic differentiation of BMSCs: immunofluorescence staining indicates that the expression of OPN and Col I was upregulated in BMSCs transfected with Usp26 mRNA@LNP on the 4th and 8th days of osteogenic induction. Reproduced with permission.263 Copyright 2024, Wiley-VCH GmbH. (C) US-responsive EMgel facilitates bone regeneration and eradicates MRSA in SD rats with chronic osteomyelitis through in situ injection. Reproduced with permission.269 Copyright 2024, Elsevier B.V. (D) Schematic representation of the dual-responsive (US and microenvironment) multifunctional MPBP microsphere system for treating bone defects. Reproduced with permission.266 Copyright 2023, Elsevier B.V. (E) Magneto-responsive HMS encapsulating SPIONs repair bone defects under external magnetic field stimuli. Reproduced with permission.271 Copyright 2021, Elsevier B.V. | ||
A study reported an US-responsive HAMA HM bomb (EMgel)269 (Fig. 8C). The microspheres release EGCG and MoS2 nanoparticles, which exhibit excellent anti-inflammatory, antibacterial, and antioxidant properties, promoting osteogenic potential. The synergistic use of EMgel and US accelerates bone healing in patients with chronic osteomyelitis. Song et al. meticulously designed a dual-responsive (US and microenvironment) multifunctional PLGA HM (MPBP microsphere) system, promoting osteogenic differentiation and cell growth at the cellular level266 (Fig. 8D). The PLGA microsphere system releases BMP-2 on demand, enhancing alkaline phosphatase activity and promoting osteocyte regeneration. Notably, the MnO2 buffer in the microsphere system not only neutralizes the acidic environment generated by PLGA degradation to protect BMP-2 but also depletes ROS and acid in the bone injury microenvironment, promoting the polarization of M1 to M2 macrophages and alleviating inflammation. Additionally, the study mentioned that the Mn2+ generated in the reaction can enhance osteogenic effects and accelerate the repair of rat cranial defects. Zhao et al. prepared magnetic microspheres by encapsulating SPIONs in PLGA microspheres.271 This method avoids the toxicity caused by direct contact between SPIONs and host tissues. The results show that the magnetic microspheres regulate the migration and differentiation of BMSCs under a magnetic field and exhibit excellent in vivo bone repair effects (Fig. 8E).
4.2. Smart responsive hydrogel microspheres for cartilage tissue regeneration
In addition to weight-bearing bones, cartilage tissue is also a fundamental component of the skeletal system, functioning as a robust sliding tissue primarily located at joints. Articular cartilage is predominantly composed of the ECM and chondrocytes. The ECM constitutes approximately 95% of the cartilage structure, with its primary components including type II Col and other types of Col (such as types VI, IX, X, and XI), proteoglycans, water, and non-collagenous proteins. Chondrocytes play a pivotal role in cartilage metabolism, synthesizing the aforementioned substances to construct and maintain the structure and function of the ECM.296 Articular cartilage defects are among the prevalent clinical healthcare issues, leading to widespread joint pain, functional deterioration, and long-term OA in patients.55 OA is a progressive joint disease characterized primarily by the degeneration of articular cartilage, including cartilage wear or even loss, joint space narrowing, joint inflammation, and subchondral bone changes.10,30 Current treatments primarily focus on symptom management, including pharmacological interventions and surgical procedures, but often fail to achieve long-term relief or halt disease progression.10,13,97,296 Pharmacological treatments, such as non-steroidal anti-inflammatory drugs (NSAIDs) and corticosteroids, can alleviate symptoms but do not promote cartilage regeneration, and long-term use can lead to adverse effects. Surgical treatments, including bone marrow stimulation, microfracture, autologous chondrocyte transplantation, and allografts, may involve invasive cartilage damage, bleeding, infection, immune rejection, and other adverse events.Hydrogels share similar viscoelastic and biological characteristics with natural cartilage, offering a novel direction for the applications of SRHMS in cartilage tissue engineering. Compared to traditional repair methods, these microspheres demonstrate analogous advantages in bone tissue engineering. During joint movement, cartilage must endure mechanical forces and provide lubrication. SRHMS not only supply lubrication for damaged cartilage but also withstand mechanical forces and intelligently release bio-agents to promote repair. Chen et al. employed microfluidics and photopolymerization processes to prepare HA-based HMS loaded with liposomes and RAPA.24 The HMS displayed a self-regenerating hydration layer on their surface in response to shear stress, enhancing joint lubrication. Additionally, cationic liposomes can target negatively charged cartilage tissue through electrostatic interactions, while the sustained release of RAPA significantly boosts chondrocyte autophagy to maintain cellular homeostasis. The synergistic effects of these three components reduce joint wear and delay the progression of OA. Furthermore, Yao et al. developed an electromagnetic force-responsive lubricating HM system (MHS@PPKHF) encapsulating fluorescent visualization nanomaterials, realizing the concept of “treatment-monitoring integration”275 (Fig. 9A). MHS@PPKHF was injected in situ into the joint cavity, forming a buffering and lubricating layer in the joint space, and releasing positively charged PPKHF under electromagnetic force to promote the differentiation of BMSCs into chondrocytes, accelerating cartilage regeneration. Simultaneously, researchers monitored cartilage repair progress through fluorescence signals.
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| Fig. 9 (A) The electromagneto-responsive MHS@PPKHF releases fluorescent visualization material PPKHF for the treatment and monitoring of OA, realizing the concept of “treatment-monitoring integration”. Reproduced with permission.275 Copyright 2023, Wiley-VCH GmbH. (B) Characterization of PDA@Lipo@HAMA microspheres: (i) microscopic images of PDA@Lipo@HAMA microspheres: dispersed microspheres (a), single microsphere (b), freeze-dried Lipo@HAMA (c), and freeze-dried PDA@Lipo@HAMA (d). (ii) Particle size of PDA@Lipo@HAMA microspheres. (iii) SEM images of microspheres: HAMA microspheres (a), Lipo@HAMA microspheres (b), and PDA@Lipo@HAMA microspheres (c). (iv) In vitro release of gallic acid (GA) from PDA@HAMA-GA and PDA@Lipo-GA@HAMA (n = 3). (v) In vitro release of GA from PDA@Lipo-GA@HAMA in PBS and PBS with 1 × 10−3 M H2O2 (n = 3). Reproduced with permission.273 Copyright 2021, Wiley-VCH GmbH. (C) HAMA/MMP13sp/Lipo@celecoxib releases CLX through MMP-13-promoted degradation, effectively reducing inflammation and reversing the OA process. Reproduced with permission.297 Copyright 2024, Elsevier Ltd. (D) In vivo therapeutic effects of IA-ZIF-8@HM for OA: (i) overview of the in vivo experiment. (ii) X-ray images of different groups after treatment. (iii) Relative joint space width of different groups (**p < 0.01, ***p < 0.001). Reproduced with permission.272 Copyright 2023, MDPI. (E) MMP and ROS dual-responsive DMY@HGP microspheres restore the apoptosis–autophagy balance by activating SIRT3. SIRT3 activation inhibits mitochondrial apoptosis through BCL2/BAX-cytochrome c release-apoptosome arrest and promotes mitochondrial autophagy via the PINK1/Parkin-LC3B-mitophagosomal complex, thereby inhibiting ECM degradation. Reproduced with permission.274 Copyright 2023, Wiley-VCH GmbH. | ||
The overproduction of ROS is one of the critical factors driving the onset and progression of OA, not only promoting the expression of inflammatory mediators, leading to a decrease in pH in the microenvironment, but also upregulating cartilage catabolism and downregulating cartilage anabolism.17,30 Lin et al. designed an injectable ROS-responsive HM system loaded with positively charged nanomaterials for efficient anti-inflammatory drug delivery (PDA@Lipo@HAMA microspheres).273 In a rat OA model, the microspheres effectively penetrated the cartilage ECM and achieved drug release and accumulation near the cartilage ECM under oxidative stress conditions, thereby inhibiting chondrocyte apoptosis (Fig. 9B). Xiang et al. developed MMP-13-responsive HMS (HAMA/MMP13sp/Lipo@celecoxib) encapsulating cationic liposomes loaded with CLX for targeted treatment of OA297 (Fig. 9C). Based on enzyme stimuli, the microspheres enable controlled drug release at specific target sites, demonstrating excellent anti-inflammatory properties and effectively improving joint cartilage degeneration. Han et al. created a pH-responsive HM system (IA-ZIF-8@HMs) anchored with IA-loaded metal–organic nanoparticles (IA-ZIF-8), exhibiting significant anti-inflammatory and antioxidant stress effects within chondrocytes, alleviating the progression of OA272 (Fig. 9D). Beyond single-stimulus responsive systems, multi-responsive HM systems have expanded the development path for cartilage tissue engineering. Inspired by the matrix proteases and ROS in the OA environment, Xia and his team developed a dual-responsive HM (DMY@HGP) loaded with DMY, unveiling a novel therapeutic strategy targeting mitochondria274 (Fig. 9E). DMY, a natural agonist of the mitochondrial sirtuin deacetylase 3 (SIRT3), downregulates mitochondrial apoptosis and upregulates mitochondrial autophagy by activating SIRT3, restoring the balance between the two, and improving cartilage wear and subchondral bone sclerosis.
4.3. Smart responsive hydrogel microspheres for skin tissue regeneration
The skin, the largest organ of the human body, is composed of the epidermis and dermis. The epidermis acts as a barrier between the internal and external environments, while the dermis provides structural support and nourishment to the epidermis. Skin wounds resulting from trauma and pathophysiological conditions represent a significant public health challenge.298 The skin possesses the ability to self-regenerate, a highly intricate and progressive physiological process that encompasses a sequence of events including hemostasis, inflammation, proliferation, and ECM remodeling.17,73,237,298 However, in certain scenarios such as diabetes, infection, immunodeficiency, malnutrition, advanced age, or the presence of foreign bodies, the healing of skin wounds can be impeded, causing many acute wounds to evolve into chronic ulcers and, in severe cases, leading to terminal amputations.237 Research indicates that full-thickness skin injuries exceeding 4 centimeters lose their capacity for self-healing and necessitate surgical transplantation.55,299 Nonetheless, autograft surgeries frequently encounter donor shortages, while allograft or xenograft surgeries may result in infections and immune rejection reactions.One of the most widely utilized strategies in skin repair involves the application of wound dressings as temporary matrices to promote the healing of skin wounds.73 A diverse array of skin wound dressings has been developed, including gauze, films, foams, nanofibers, and hydrogels.300 Skin tissue engineering imposes several critical requirements on wound dressings, encompassing biocompatibility, biodegradability, bioactivity, hydrophilicity, mechanical properties, and the capacity to deliver bio-agents. Beyond fulfilling these multifunctional characteristics, as micron-scale materials, SRHMS also provide high filling efficiency for irregularly shaped and deep wound lesions. Moreover, their smart responsiveness offers enhanced controllability for skin tissue engineering. In summary, SRHMS are regarded as one of the most promising options for skin wound dressings.
Many chronic skin wounds are strongly associated with diabetes, as the hyperglycemic environment caused by the condition impairs the normal healing process of skin tissue.30,298 The high glucose environment triggers a cascade of adverse events, including wound ischemia and hypoxia, inflammatory cell infiltration, impaired neovascularization, reduced Col deposition, and poor granulation tissue development, ultimately leading to delayed or even halted healing of diabetic wounds.301 Li et al. developed an MMP-9-responsive HM (GMH8NV) loaded with H8NVs.276 H8, a CUR analog, possesses anti-inflammatory properties and is effective in treating diabetes. The microspheres accelerate the healing of diabetic wounds by suppressing inflammation and MMP-9 expression, stimulating angiogenesis, and enhancing Col deposition (Fig. 10A). In another study, CUR was encapsulated in Alg microspheres and incorporated into a bilayer dressing hydrogel-microsphere system (PCL-TH/PVA-Alg@CUR).259 Researchers achieved intelligent controlled release of CUR by leveraging the pH sensitivity of Alg and demonstrated that PCL-TH/PVA-Alg@CUR exhibits superior healing effects on chronic diabetic wounds with bacterial infections (Fig. 10B). Zhou et al. innovatively designed a glucose and MMP-9 dual-responsive HM composite system (CBP/GMs@Cel&INS) encapsulating Gel microspheres containing CLX and INS.277 The system intelligently releases CLX and INS in response to elevated glucose and MMP-9 levels in the microenvironment, thereby exerting anti-inflammatory, hypoglycemic, and angiogenic effects (Fig. 10C). Notably, during its application to chronic diabetic wounds, the system demonstrates temperature-sensitive shape-adaptive behavior, ensuring rapid adaptation to deep wounds while providing mechanical protection.
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| Fig. 10 (A) MMP-9-responsive GMH8NV releases H8NVs to promote the healing of diabetic skin wounds. Reproduced with permission.276 Copyright 2024, American Chemical Society. (B) pH-responsive PCL-TH/PVA-Alg@CUR evaluates wound healing through an in vivo wound infection model: (i) overview of wound formation and subsequent healing steps. (ii and iii) Graphical representation of wound healing in different groups after 15 days of treatment. (iv) Antibacterial capability of different groups against Staphylococcus aureus-infected wounds (exudate diluted 1 × 104 times). (v) Corresponding statistics of colony counts. (vi) Statistics of wound size for infected wound healing. Reproduced with permission.259 Copyright 2023, Elsevier Ltd. (C) CBP/GMs@Cel&INS promotes chronic diabetic wound repair through dual-responsive release of CLX and INS to glucose and MMP-9. Reproduced with permission.277 Copyright 2022, Elsevier Ltd. (D) In vitro antibacterial activity of thermo-responsive i-Lyso@Alg microspheres: (i) bacterial colony images of Staphylococcus aureus after different treatments. (ii) Activity of killing Staphylococcus aureus under different treatments. (iii) SEM images of Staphylococcus aureus under different treatments (scale bar = 300 nm). Reproduced with permission.278 Copyright 2024, Elsevier B.V. (E) NIR and pH-responsive HMS release CIP and CUR to promote the repair of infected wounds. Reproduced with permission.258 Copyright 2024, Elsevier B.V. (F) Schematic illustration of magneto-responsive HMS releasing NO to treat bacteria-infected skin wounds. Reproduced with permission.185 Copyright 2022, American Chemical Society. (G) Characterization of AHFS microspheres: (i) size distribution of liposomes carrying drugs Ti and T. (ii) Size distribution of AHFS microspheres. (iii) Bright-field image of AHFS microspheres. (iv) Elemental mapping image of AHFS microspheres. (v) Morphological changes of AHFS microspheres during degradation tests. Reproduced with permission.169 Copyright 2024, Wiley-VCH GmbH. | ||
The healing process of infected wounds is intricate and necessitates wound dressings that fulfill both antibacterial and repair requirements. Chen et al. successfully developed a thermo-responsive HM (i-Lyso@Alg) containing lysozyme and MXene.278 Owing to the presence of MXene, i-Lyso@Alg stabilizes and enhances the antibacterial activity of lysozyme through photothermal effects, effectively inhibiting bacterial infection and the expression of inflammatory factors in the wound microenvironment (Fig. 10D). Wu et al. designed a dual-drug-loaded, dual-responsive HM system loaded with CIP and CUR.258 In the acidic environment of infected wounds, the system responds to acidic stimuli to release CIP, achieving antibacterial treatment. Subsequently, heating the microspheres with infrared light triggers the thermo-responsive release of CUR, promoting cell migration and growth (Fig. 10E). Zhong et al. developed a magneto-responsive NO-releasing HM.185 The magneto-responsive burst release of NO demonstrated exceptional antibacterial activity. More importantly, in the absence of a magnetic field, the stable release of NO fosters subsequent Col formation and wound healing (Fig. 10F).
Skin appendages, including HFs, sweat glands, and sebaceous glands, play a vital role in skin function and homeostasis.302 Skin scar repair can result in the impairment of skin function and the loss of appendages such as HFs. Drawing inspiration from the natural process of “sowing and harvesting”, Ji et al. employed microfluidics and photocrosslinking techniques to fabricate artificial HF seeding (AHFS) light-responsive HMS169 (Fig. 10G). AHFS drives fibroblasts towards DPCs by activating the PI3K/AKT pathway, thereby promoting in situ HF regeneration and skin wound healing. Notably, they conferred a positive charge to AHFS through polyamide, enabling AHFS to adhere to skin wounds and achieve sustained non-invasive therapeutic effects.
4.4. Smart responsive hydrogel microspheres for neural tissue regeneration
The nervous system consists of the central nervous system (including the brain, cerebellum, brainstem, and spinal cord) and the peripheral nervous system (including cranial nerves, spinal nerves, and their associated ganglia). Nerve injuries are primarily caused by primary injuries (trauma, hemorrhage, compression, surgical complications, congenital defects) and secondary injuries (ischemia, hypoxia, oxidative stress, and inflammation), leading to cellular degeneration, apoptosis, and necrosis at the injury site, ultimately resulting in nerve necrosis, defects, or rupture.78,303–305 In the process of nerve repair, it is crucial to minimize primary injury and mitigate secondary injury, with the core focus on reducing inflammation and oxidative stress, promoting myelin and axon regeneration, and regulating angiogenesis.78,306 Due to the highly differentiated nature of nerves, the self-repair of neurons and axons is limited, and the formation of glial scars at the injury site makes nerve injury repair particularly challenging.305 Current methods for nerve repair include pharmacological treatments,78,307 physical therapies,308 and surgical interventions304,305,309 such as nerve suturing, nerve grafts, and nerve conduits. However, each of these methods has its limitations. Pharmacological treatments are often inefficient due to the lack of targeted delivery, the body's metabolic functions, and physiological barriers.78 Physical therapies, as a conservative treatment strategy, provide only temporary symptom relief, and their long-term efficacy is questionable.308 Surgical nerve suturing is clinically applicable only for short nerve gaps, while the “gold standard” of autografts is limited by donor shortages, multiple surgeries, and secondary surgical trauma (infection, donor-derived neuromas, and functional loss).304,305 “Foreign” allografts may carry pathogens and trigger immune rejection, and nerve conduits still pose uncertain safety concerns.305The objective of neural tissue engineering is to establish a microenvironment for damaged neural tissue that closely mimics the natural nerve growth environment by utilizing available biomaterial scaffolds, and to integrate various bio-agents to facilitate neural tissue repair and regeneration. The nervous system is a highly organized and strongly directional tissue, and the proper alignment of cells and tissues following nerve injury is essential for the functional recovery and reconstruction of neural tissue.30 Stem cell strategies have been extensively employed in neural tissue engineering, with neural stem cells (NSCs), MSCs, and embryonic stem cells (ESCs) all possessing the capability to differentiate into neurons.237 Neurotrophic factors such as nerve growth factor (NGF), VEGF, and INS-like growth factor (IGF) exhibit neurotrophic functions, promoting cell proliferation, differentiation, and intercellular communication, which are advantageous for myelination and axon regeneration, rendering them highly popular in neural tissue engineering.78,305 With the advancements and research of SRHMS, researchers are increasingly concentrating on the delivery of bio-agents by these microspheres and applying stimuli to achieve intelligent controlled release of the cargo within the microspheres to repair the damaged nervous system.
SCI is a condition involving damage to the central nervous system, often leading to significant neurological decline or even permanent paralysis if the endogenous neural tissue repair/regeneration process fails to initiate.17,78 SRHMS systems have been employed in tissue engineering to treat SCI. Sun et al. utilized microfluidics to prepare injectable photothermo-responsive drug-loaded multi-walled carbon nanotube-HMS ([PTX/VEGF]@DMWCNTs/GG).262 The thermoreversible nature of this system enables the controlled release of PTX and VEGF, promoting the differentiation of endogenous NSCs into neurons while inhibiting their differentiation into glial cells and glial scar formation, thereby facilitating spinal cord regeneration (Fig. 11A). The adverse inflammatory microenvironment can hinder the recovery of injured spinal cords. Li et al. combined Gel microspheres with HA and Dex to construct an MMP-responsive neural scaffold loaded with the MG53 protein (MG53/GMs/HA-Dex neural scaffold).280 This scaffold responds to MMP-2/9 proteins in the inflammatory microenvironment and stably releases the MG53 protein, inhibiting the M1 polarization of microglia and reducing neuroinflammation by suppressing the JAK2/STAT3 pathway, thereby promoting spinal cord recovery (Fig. 11B). Another study employed electrospray to prepare injectable biomimetic porous nanomicrospheres (NMs), composed of various short nanofibers made from PCL, PLGA, Gel, methacrylate, bioglass, and magneto-responsive polymer composites.281 Compared to non-porous NMs, porous NMs can load bio-agents, respond to external stimuli, and promote the growth of NSCs within their 3D structure, resulting in a higher number of neurites (Fig. 11C). Gong et al. developed a pH-responsive cellulose HM (CCMs) loaded with drugs through Fe2+/H2O2 oxidative carboxylation279 (Fig. 11D). The released vitamin B12 from the microspheres can prevent neuropathy and neurological complications.
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| Fig. 11 (A) [PTX/VEGF]@DMWCNTs/GG treat SCI by releasing PTX and VEGF in response to NIR. Reproduced with permission.262 Copyright 2023, American Chemical Society. (B) MG53/GMs/HA-Dex neural scaffold responds to MMP-2/9 proteins to treat SCI. Reproduced with permission.280 Copyright 2024, Elsevier B.V. (C) Confocal microscopy images show the expression of neurites (Tuj1, green) in hNSCs cultured in neuronal differentiation medium with non-porous NMs and porous NMs for 10 and 14 days. Neurite growth was first observed at 10 days and continued until day 14. Non-porous NMs supported fewer cells, with sparse and longer neurite growth, while cells in porous NMs exhibited denser neurites throughout the microspheres. Cells were stained with the Tuj1 antibody (green) and counterstained with DAPI (blue). Reproduced with permission.281 Copyright 2020, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. (D) SEM images and EDS patterns of cellulose microspheres (a–d), carboxylated cellulose microspheres (e–h), and vitamin B12-CCM (i–l). Reproduced with permission.279 Copyright 2024, Elsevier B.V. (E) Electrophysiological evidence of functional regeneration promoted by the conductive microsphere system combined with EES: spinal cord motor evoked potentials (SCMEPs) were recorded when rats were stimulated above (T8) and below (S1) the injury at 2 and 6 weeks post-spinal cord transection and after re-transection at 7 weeks post-treatment. In rats receiving combined treatment, stimulation above the injury (T8) at 2 weeks induced small supralesional evoked polysynaptic responses (SEPRs), which became more pronounced at 6 weeks and disappeared after re-transection at 7 weeks. Stimulation below the injury (S1) elicited intermediate responses (MRs) (4.5–5.7 ms), with increased amplitude after re-transection, indicating the influence of regenerating fibers on the sublesional network. In another rat, stimulation at T8 detected SEPRs only at 6 weeks, which also disappeared after re-transection. The average MR peak-to-peak amplitude before re-transection at 6 weeks was 4.28 mV, increasing to 10.67 mV after re-transection. Reproduced with permission.310 Copyright 2021, Springer Nature. | ||
Due to the electrical activity of neuronal cells, the transmission of bioelectrical signals is crucial for the functional recovery of nerves.304 A study reported a drug-loaded and cell-loaded HM system modified with positively charged oligonucleotides.310 The microspheres were loaded with the anti-fibrotic drug rapamycin and Schwann cells (SCs) capable of producing neurotrophic proteins. The conductive microsphere system, combined with epidural electrical stimulation (EES) rehabilitation electrodes, demonstrated excellent neural reorganization and functional recovery (Fig. 11E). In bioelectronics and neural tissue engineering, this also provides a research direction for combining electro-responsive HM delivery systems with neuroelectrical modulation therapy.
4.5. Smart responsive hydrogel microspheres for cardiac tissue regeneration
Cardiovascular diseases, such as AMI, continue to be a leading cause of death worldwide.163,237,311,312 The incidence and mortality rates of cardiovascular diseases in our country remain alarmingly high.237 When MI induces cardiac ischemia, the heart tissue undergoes severe damage due to hypoxia, resulting in the apoptosis and necrosis of CMs, ECM degradation, excessive matrix deposition, and a cascade of myocardial fibrosis and inflammation.311,312 This further damages CMs and impedes the recovery of cardiac function and structural remodeling, ultimately leading to heart failure. Cardiac tissue fibrosis severely disrupts the signal transmission between CMs, cutting off the electrochemical cues for cellular activity, which leads to reentrant arrhythmias and abnormal contractile function.163 This significantly impairs cardiac function and prevents normal blood circulation. CMs are highly differentiated cells with limited regenerative capacity.50,163 Current stem cell strategies and heart transplantation surgeries offer limited long-term effects on the functional recovery of the heart.312 To achieve long-term curative effects for the heart and cardiovascular system, cardiac tissue engineering may be one of the most promising emerging methods for regenerating myocardial tissue.The therapeutic procedure of combining bio-agents with biomimetic scaffolds is crucial for the development of the cardiac tissue engineering strategy. An ideal biomimetic scaffold for cardiac tissue engineering can provide biological cues similar to those in the natural myocardial microenvironment, such as mechanical, electrical, and chemical signals.50 Hydrogels have emerged as a viable material for cardiac tissue engineering. As a member of the hydrogel family, SRHMS are ideal scaffold materials. They not only provide mechanical support for infarcted CMs, promoting their regeneration and repair, but also serve as intelligent carriers for various bio-agents, enabling spatiotemporal controlled release. This helps improve the microenvironment of cells in the damaged area, promotes cell activity, inhibits apoptosis, and enhances myocardial contractility, thereby restoring cardiac function. More importantly, given the unique continuous electrical network organization of the heart, the conductivity of electrically responsive HM systems is highly suitable for the conductive microstructure of myocardial tissue, bridging electrical signals between healthy and damaged tissues.17,163 The restoration of electrical signals within the heart is beneficial for resynchronizing contractions, preventing cardiac remodeling and dysfunction.163 Mu et al. developed a ROS-responsive conductive HM incorporating graphene oxide (GO) for the treatment of AMI.256 The conductive microspheres were loaded with the anti-inflammatory agent SalB and adipose-derived stem cells. In vivo, the conductive microspheres reduced myocardial fibrosis, promoted the regeneration of CMs, and enhanced electrical signal conduction in myocardial tissue, demonstrating excellent myocardial repair capabilities (Fig. 12A).
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| Fig. 12 (A) Fabrication of ROS-responsive conductive HMS delivering ADSCs and the mechanism for treating AMI. Reproduced with permission.256 Copyright 2024, Elsevier Ltd. (B) Effects of microgel treatment on cardiac function in MI rats after 28 days: (i) left ventricle and representative echocardiographic images. (ii) Overview images of Masson's trichrome-stained sections of the heart (top) and magnified views of the areas marked by red boxes in the top images (bottom). (iii) Quantified left ventricular ejection fraction (LVEF), (iv) left ventricular fractional shortening (LVFS), (v) left ventricular end-diastolic volume (LVEDV), and (vi) left ventricular end-systolic volume (LVESV). (vii) Quantitative analysis of the whole heart infarct area, and (viii) left ventricular wall thickness in the infarct area. Reproduced with permission.117 Copyright 2024, Elsevier Ltd. (C) Fabrication process of GTK–TK–drug and its in vivo injection, providing mechanical support, enhancing Treg regulation, and promoting cell ingrowth and proliferation regulation after MI. Reproduced with permission.282 Copyright 2024, Elsevier B.V. (D) pH/H2O2/MMP-9-responsive DHPM(4APPC)_EVs achieve rapid release, capture, and slow release of EVs, ultimately promoting cardiac functional recovery in the MI area. Reproduced with permission.53 Copyright 2022, Wiley-VCH GmbH. | ||
In the pathological processes of heart-related diseases, inflammation and the production of ROS are significant factors leading to myocardial injury. Therefore, anti-inflammatory and antioxidant strategies are crucial for restoring myocardial function. Wang et al. reported a ROS-responsive spherical microgel system for the in situ release of AOA.117 The released AOA modulates the conversion of T cells from pro-inflammatory Th17 to anti-inflammatory Treg, exerting anti-inflammatory and antioxidant effects, reducing CM apoptosis and myocardial fibrosis, promoting angiogenesis, and restoring cardiac function (Fig. 12B). Similarly, Wang and her team developed a microgel scaffold (GTK–TK–drug) that promotes Treg differentiation for the treatment of MI282 (Fig. 12C). This microgel scaffold is composed of microgels containing cyclodextrin (CD) or adamantane (Ad) with diameters exceeding 200 μm; these microgels assemble into the scaffold through host–guest interactions. The scaffold provides a versatile platform for upregulating Treg and promoting cell migration and growth, holding promise for broad applications beyond cardiac tissue engineering.
In cardiovascular diseases such as atherosclerosis, MI, and thrombosis, the upregulation of certain enzymes that promote cardiac matrix deposition can serve as specific stimuli to trigger the release of cargo from SRHMS systems.163 Cheng et al. designed a composite hydrogel-EVs system based on Sparchigh Treg-derived EVs and pH/H2O2/MMP-9-responsive HMS (DHPM(4APPC)_EVs).53 Overexpression of SPARC in Tregs can aid in the treatment of AMI. Through genetic engineering, CXCR2 (inflammatory chemokine) was overexpressed on the EV membrane, promoting the rapid targeting of EVs to infarcted tissue. The microspheres respond to the pathological acidic microenvironment to release EVs. Subsequently, the MMP-9 enzyme substrate peptide-modified microsphere material, 4-arm PEG, based on H2O2-triggered Co2+ oxidation, can capture the previously released EVs in situ in the infarcted area and form a hydrogel. Finally, the in situ formed hydrogel slowly releases EVs under MMP-9 stimulation, inhibiting inflammation in the infarcted area, reducing CM damage, and promoting cardiac repair (Fig. 12D).
Zhao et al. synthesized a biodegradable PLLA-PEG-PNIPAm triblock copolymer, which self-assembled into thermo-responsive gelling microspheres (NF-GMS) composed of nanofibers.283 Notably, these microspheres undergo a thermal stimulus-induced transformation to form a 3D hydrogel, which then carries exogenous CMs for transplantation into the myocardial infarction site. CM transplantation can significantly reduce the infarct area, enhance CM integration, stimulate angiogenesis, and restore cardiac function. Another study reported a thermo-responsive HM combined with embryoid bodies (EBs) to mediate the cardiovascular differentiation of embryonic stem cells.313 Additionally, a study described an US-responsive microsphere system prepared via microfluidics, loaded with bFGF for the treatment of MI.268
4.6. Smart responsive hydrogel microspheres for other tissue regeneration
In addition to the primary directions of tissue engineering, SRHMS are also highly favored in the treatment of other diseases (e.g., tumors, gastrointestinal diseases, etc.), owing to their excellent biocompatibility, hydrophilicity, lubricity, biodegradability, smart responsiveness, and superior loading and delivery capabilities.SRHMS demonstrate significant potential in tumor tissue engineering by enabling multimodal regulation of the TME through integration with synergistic therapeutic approaches. Liang et al. designed a photothermo-responsive HM system encapsulating BP, a nitric oxide donor (S-nitrosoglutathione, GSNO), and the chemotherapeutic drug DOX using a microfluidic electrospray strategy.286 Notably, this system achieved multimodal tumor therapy, including NO therapy, photothermal therapy, and chemotherapy, demonstrating excellent tumor suppression effects (Fig. 13A). Similarly, Zhu et al. reported a thermo-responsive AG HM (MTX-TMPs@MSs) loaded with tumor cell-derived microparticles (TMPs) for the treatment of malignant ascites in tumors.148 The TMPs encapsulating methotrexate (MTX) were released under NIR thermal stimulation, showcasing a multimodal approach combining chemotherapy (CHT), immunotherapy (IMT), and photothermal therapy (PTT) (Fig. 13B). Deng et al. developed light-responsive HMS (ACEA@CL(XGMA)-Fe(III) MSs) fabricated via electrospray for potential colon cancer therapy.287 Upon photo-stimulation, this system simultaneously generates ROS and releases ACEA, establishing a dual-network antitumor mechanism that synergistically remodels the tumor microenvironment (TME) while suppressing tumor invasion. Li et al. also designed a photothermo-responsive HM encapsulating two anti-tumor drugs, SOR and DOX, for the treatment of gastric cancer260 (Fig. 13C). pH-responsive HMS encapsulating drugs can also respond to the slightly acidic TME, enhancing the therapeutic efficacy of transarterial chemoembolization (TAE) for HCC284,285 (Fig. 13D).
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| Fig. 13 (A) Fabrication of photothermo-responsive HMS loaded with BP, GSNO, and DOX, and schematic illustration of the multimodal therapeutic mechanism. Reproduced with permission.286 Copyright 2023, Wiley-VCH GmbH. (B) MTX-TMPs@MSs synergistically combine CHT, IMT, and PTT to form a multimodal treatment strategy, effectively eliminating cancer cells and inhibiting tumor growth. Reproduced with permission.148 Copyright 2024, Elsevier Ltd. (C) Photothermo-responsive HMS release SOR and DOX in mice for tumor treatment. (i) Infrared thermal images of mice after intratumoral injection of photothermal microspheres under NIR irradiation. (ii) Photographs of mice in different groups and their corresponding tumor sizes as shown. (iii) Photographs of tumors in each group on day 21. (iv) Estimated relative tumor sizes of mice receiving different treatments on day 21. Groups: PBS buffer (G I), drug-free composite microspheres (G II), photothermal microspheres with NIR irradiation (G III), DOX-loaded microspheres (G IV), SOR-loaded microspheres (G V), free DOX and SOR (G VI), DOX and SOR-loaded microspheres (G VII), and DOX and SOR-loaded microspheres with NIR irradiation (G VIII). Scale bar: 1 cm (ii). Reproduced with permission.260 Copyright 2024, SCUT, AIEI, and John Wiley & Sons Australia, Ltd. (D) pH-responsive HMS for enhancing the therapeutic efficacy of TAE in HCC treatment. Reproduced with permission.284 Copyright 2024, Elsevier Ltd. (E) Schematic diagram of the synthesis and therapeutic mechanisms of MHBSA HMS. Reproduced with permission.82 Copyright 2025, Elsevier Ltd. (F) Design of P@GH@ML microsphere vehicles. (i) Fabrication of P@GH@ML microsphere vehicles. (ii) Operational process and mechanism of P@GH@ML microsphere vehicles at intestinal inflammatory sites. (iii) Temperature-dependent FTIR spectra (30–200 °C) of P@GH@ML microsphere vehicles. (iv) Synchronous and asynchronous 2D correlation spectra of temperature-dependent FTIR for P@GH@ML microsphere vehicles. (v) SEM images of Gel/HA, GH@ML and P@GH@ML microspheres, and elemental mapping of P@GH@ML microspheres. (vi) TEM images of P@GH@ML microspheres. (vii) Surface morphology of GH@ML and P@GH@ML microspheres after 1 h immersion in deionized (DI) water, artificial gastric fluid (AGF, pH 1.0), artificial small intestinal fluid (ASF, pH 6.8) and artificial colonic fluid (ACF, pH 7.8). Strain scan curves of GH@ML and P@GH@ML microspheres at a constant frequency (1 Hz) in DI water, AGF, ASF and ACF. Reproduced with permission.289 Copyright 2025, American Chemical Society. (G) Characterization of Gel and SF microspheres (GSMs) in pH- and ROS-responsive MSC-GSM@HGR: (i) bright-field images of GSMs before and after freeze-drying. (ii) Diameter distribution of GSMs corresponding to (i). (iii) SEM images of GSMs after freeze-drying. (iv) SEM images of the internal porous structure of GSMs. (v) Representative images of GSMs from freeze-drying to swelling. Reproduced with permission.254 Copyright 2024, Elsevier B.V. (H) Implantation of siRNA@MS@HA at the injured tendon site to reduce peritendinous tissue adhesion. Reproduced with permission.291 Copyright 2021, Wiley-VCH GmbH. | ||
SRHMS have emerged as a promising platform to address critical therapeutic challenges including enzymatic degradation and limited bioavailability, while enabling targeted drug delivery with enhanced pharmacokinetic profiles and smart environmental responsiveness for hepatic and gastrointestinal diseases. Within the field of tumor tissue engineering, our previous discussions have highlighted SRHMS applications in HCC, colon cancer and gastric cancer management. Recent studies have extended SRHMS advancements in other gastrointestinal disorders, particularly IBS and IBD. Zhang et al. developed innovative ROS-responsive HMS incorporating puerarin, demonstrating remarkable resistance to gastrointestinal degradation while achieving targeted therapeutic delivery for IBS-D.267 Chen's research team engineered a sophisticated pH-responsive HM system (MHBSA) combining HNTs, EGCG, and probiotics.82 This system established a novel targeting paradigm for IBD intervention through synergistic antioxidant, anti-inflammatory, and gut microbiota-regulating mechanisms (Fig. 13E). Complementing these advances, Wang et al. designed an NO-responsive HM platform (NRPM) capable of stimuli-triggered probiotic release specifically tailored for IBD microenvironments.288 Furthermore, Zhao et al. utilized an advanced emulsion technique to fabricate an inflammatory microenvironment-responsive microsphere system (P@GH@ML microsphere vehicles) that specifically targets IBD-affected regions, enabling spatiotemporally controlled release of MXene and L-arginine289 (Fig. 13F). This system also modulates gut microbiota composition and exhibits potent anti-inflammatory effects via promotion of epithelial regeneration and intestinal stem cell niche reorganization. Additionally, researchers have reported an inflammatory microenvironment-responsive HM system for colon-specific delivery of quercetin, offering new therapeutic possibilities for IBD management.290
Additionally, Yang et al. combined mesenchymal stem cell-loaded HMS with PBA modified HA hydrogel to form a ROS- and pH-responsive system (MSC-GSM@HG) with self-healing properties, promoting tissue repair of the abdominal wall MSCs254 (Fig. 13G). This represents a highly promising tissue engineering strategy for abdominal wall defects. Cai et al. designed a self-healing HA hydrogel system embedded with MMP-2-responsive HMS (siRNA@MS@HA) to prevent the formation of adhesions after tendon surgery.291 The microspheres respond to MMP-2 to achieve on-demand release of Smad3-siRNA nanoparticles, thereby downregulating Smad3 expression and fibroblast proliferation in the peritendinous region, reducing the degree of adhesion around the tendon (Fig. 13H). Furthermore, compared to non-healing hydrogels, the self-healing system exhibits a lower inflammatory response. Immunoengineering employs biomaterial-based strategies to modulate immune responses for therapeutic applications, particularly in autoimmune disease management. As a representative example, Laura Clua-Ferré et al. pioneered an innovative glucose-responsive therapeutic platform by employing the 3D bio-printing technique to fabricate β-cell-laden HMS. This intelligent system demonstrates dynamic responsiveness to physiological glucose fluctuations, enabling regulated insulin secretion from encapsulated β-cells for effective management of type 1 diabetes mellitus, an autoimmune disorder characterized by pancreatic β-cell destruction.86 Another study reported the use of CS-based thermo-responsive HMS loaded with ornidazole and doxycycline hydrochloride for the treatment of periodontitis in dental tissue engineering.292
5. Summary and prospect
SRHMS, as a novel smart biomaterial, have demonstrated significant potential in the field of tissue engineering due to their unique structures and multifunctional properties. This review systematically reports the latest advancements in the research of SRHMS for tissue engineering, highlighting their role as an effective therapeutic platform for a wide range of tissues. Based on natural and synthetic polymers, researchers have designed and functionalized SRHMS through advanced fabrication techniques such as microfluidics, emulsion technique, electrospray, mechanical fragmentation, lithography, and 3D bio-printing. These methods enhance the controllability and smart responsiveness of SRHMS, driving the innovation of tissue engineering therapeutic solutions. Compared to hydrogels, SRHMS stand out as efficient delivery systems that combine lubrication with smart responsiveness, offering exceptional functional advantages. SRHMS intelligently respond to endogenous and/or exogenous stimuli, releasing bio-agents in a spatiotemporal manner to promote tissue regeneration, making them a crucial component of smart tissue engineering strategies.Despite remarkable progress in laboratory research, the successful translation of SRHMS into clinical applications still faces numerous challenges, particularly in terms of scale-up production and manufacturing processes.69,314 To ensure that SRHMS meet clinical requirements, it is essential to establish production processes that adhere to Good manufacturing practice (GMP) standards, ensuring the quality, safety, and stability of the products.315 Furthermore, scaling up the production of SRHMS requires addressing issues such as cost, production efficiency, and sustainability, necessitating the continuous optimization of manufacturing processes to ensure scalability and consistency during production.69 Current manufacturing technologies, such as microfluidics, face limitations in yield and cost-effectiveness. Techniques like emulsion technique and mechanical fragmentation offer greater scalability but need optimization to ensure monodispersity and batch consistency. As of now, no SRHMS platform has received approval from the U.S. Food and Drug Administration (FDA) or reached clinical trial stages. However, several materials used in the preparation of SRHMS, such as Gel, HA, PLGA, and PVA, have made significant progress toward FDA approval.10,316–318 Moreover, a number of hydrogel products have been clinically translated and commercialized after FDA approval, including Jelmyto, Panretin® Gel, Hyftor, and Vantas®. Additionally, several hydrogel-based delivery systems, such as BIL06v/Alhydrogel® (Phase 1) and SR-T100® Gel (Phase 2), are still undergoing clinical trials and yet to receive FDA approval.319 These advancements indicate that the clinical translation and commercialization of SRHMS hold great promise.
The applications of SRHMS should not be confined to their current fields, as their potential could far exceed expectations. We propose five innovative developmental concepts that not only present novel ideas for advancing SRHMS systems but also offer fresh perspectives on the application of related technologies:
1. Development of multidimensional structures (4D and 5D): as molecular mechanisms are explored more deeply and fabrication techniques continue to advance, there is an opportunity to integrate cutting-edge biomanufacturing techniques such as 4D and even 5D bio-printing. This will enable precise regulation of the internal structures, spatial organization, and dynamic adaptability of SRHMS, expanding their capabilities beyond what is currently achievable.
2. Integration and optimization of multiple response systems: current research has largely focused on single or dual stimulus-responsive SRHMS delivery systems. However, these systems may not be sufficient for the dynamic and complex pathological microenvironments encountered in clinical settings. Future research should prioritize the integration of multiple response systems, enabling SRHMS to respond to a broader range of stimuli (such as temperature, pH, magnetic fields, light, etc.). This would result in more robust, adaptive solutions that are better suited for addressing complex pathological conditions with greater spatiotemporal precision.
3. Enhancement of self-healing and regenerative capabilities: the incorporation of self-healing technologies could significantly improve the stability and long-term functionality of SRHMS in vivo. By embedding functional molecules with self-healing properties within SRHMS, they can autonomously restore their structures after damage, thereby extending their effective lifespan within the body. This is especially important for tissue repair applications in complex environments.
4. Innovation in personalized medical applications: tailoring the response mechanisms of SRHMS to individual patients could enable more personalized treatment approaches, enhancing precision and therapeutic efficacy. Notably, the interdisciplinary integration of self-healing SRHMS with AI-based biosensors could pave the way for minimally invasive therapies and personalized medicine.70 The convergence of materials science, regenerative medicine, and artificial intelligence will be instrumental in driving SRHMS toward successful clinical translation.
5. Visualization of therapeutic outcomes: while SRHMS are primarily designed to promote tissue regeneration, accurately assessing their therapeutic effects remains a significant challenge. Thus, developing detection methods to evaluate therapeutic outcomes is critical. A promising approach could involve using microspheres loaded with fluorescein for cost-effective, non-radiative detection of damaged tissues.275 This would facilitate the establishment of a standardized regulatory framework for SRHMS applications.
In the future, key research directions will focus on refining the regulation of internal structures, enhancing surface topologies, optimizing fabrication techniques, and further exploring molecular mechanisms. Additionally, the establishment of SRHMS system fingerprint databases and overcoming the challenges of developing multidimensional structures (4D and 5D) will be crucial in advancing SRHMS toward clinical success. These efforts will play a pivotal role in the successful translation of SRHMS strategies into clinical practice.
Conflicts of interest
The authors declare no conflict of interest.Data availability
No primary research results, software or code have been included and no new data were generated or analysed as part of this review.Acknowledgements
Part of the elements in Scheme 1, Fig. 1, Fig. 6 and Fig. 7 are created with https://BioRender.com. Funding: the authors express appreciation to the National Natural Science Foundation of China (No. 82372405), the Key Research and Development Program of Hubei Province (No. 2022BCA052), and the Key Research and Development Program of Wuhan City (No. 2024020702030105) for providing support and assistance for the research.References
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Footnote |
| † These authors contributed equally to this work. |
| This journal is © The Royal Society of Chemistry 2025 |