Shape memory hydrogels in tissue engineering: Recent advances and challenges

Abstract

Shape memory hydrogels (SMHs) have emerged as transformative materials in tissue engineering, owing to their unique ability to recover their original shape after deformation. These hydrogels combine hydrophilicity and elasticity with shape memory capabilities, making them ideal candidates for various biomedical applications. This review examines their innovative design and synthesis, highlighting the physical and biological characteristics that make them well-suited for tissue engineering, such as mechanical properties, biocompatibility, and biodegradability. SMHs have diverse applications in tissue engineering, including bone regeneration, soft tissue reconstruction, and the engineering of vascular and neural tissues. Additionally, they are utilized in smart drug delivery systems and the fabrication of 3D-printed customized implants. Despite these advancements, challenges such as production scalability, optimization of mechanical properties, shape recovery and fixation, controlled degradation, and long-term stability persist. Interdisciplinary approaches are crucial for overcoming these challenges and enhancing their clinical potential. In conclusion, SMHs offer innovative solutions to complex biomedical problems, making them valuable tools for advancing regenerative medicine and improving patient outcomes.

Graphical abstract

Highlights

• •Comprehensiveness: Reviews shape memory hydrogels from design to biomedical applications.
• •Clarity and accessibility: Summarizes key material properties for functional tissue scaffolds.
• •Advancement: Highlights emerging uses in 3D printing, drug delivery, and regenerative therapy.
• •Insightfulness: Discusses unresolved challenges in shape fixation, degradation, and scaling.
• •Relevance: Provides a foundational guide for designing next-generation bioactive hydrogels.

1. Introduction

Tissue engineering has emerged as a highly innovative field focused on developing biological substitutes to restore, maintain, or enhance tissue function. This interdisciplinary field integrates principles from biology, engineering, and material science to develop scaffolds that support cell growth and tissue development [1]. Despite considerable progress, tissue engineering faces several persistent challenges. These include the need for scaffolds with appropriate mechanical properties, biocompatibility, biodegradability, and the capacity to support complex tissue structures and functions. Additionally, scalability in production and the integration of vascular and neural networks within engineered tissues remain significant obstacles [2]. During regeneration, tissues progress through four sequential and overlapping stages: hemostasis, inflammation, repair, and reconstruction [3]. To ensure successful regeneration, every stage of the process must be carefully regulated, as any deviations can lead to tissue damage and an increased risk of regeneration failure. During these stages, the progression is heavily influenced by immune cell regulation, especially during the inflammatory phase, which ultimately determines the success of repairing and remodeling the damaged tissue to enhance tissue regeneration. Therefore, it is crucial to develop biomaterials that not only support tissue regeneration but also address these inherent challenges.

The concept of shape memory materials first gained clinical relevance with the introduction of shape memory alloys (SMAs), such as nickel-titanium (NiTi) or Nitinol, which are commonly used in vascular stents, orthopedic implants, and dental implants. The material exhibits a unique property that allows it to regain its original shape in response to external stimuli, a property that has been exploited for minimally invasive surgery and dynamic implant design. The ability to undergo large deformations and transform between different phases (martensite and austenite) makes SMAs particularly beneficial in applications where mechanical stability, controlled response to temperature, and shape recovery are critical [4]. However, as clinical needs evolved, the shift toward softer, more flexible materials led to the development of SMHs.

Unlike SMAs, which are rigid and primarily respond to temperature, SMHs offer greater flexibility and can undergo reversible shape changes in response to various external stimuli, including temperature, pH, ionic strength, and light. This adaptability, combined with their inherent hydrophilic and tunable mechanical properties, makes SMHs particularly well-suited for biomedical applications requiring soft, dynamic materials, such as tissue engineering, wound healing, and drug delivery [5]. A critical challenge in tissue engineering is the development of scaffolds that can conform to irregular tissue defects, adapt to changing physiological conditions, and enable minimally invasive implantation. SMHs address these challenges by temporarily deforming into compact shapes for easy insertion and subsequently recovering their original geometry in situ, ensuring precise spatial conformation and functional integration with surrounding tissues. Moreover, their responsiveness and tunable mechanical properties allow them to mimic the dynamic behavior of native tissues, making them ideal for self-fitting implants, dynamic cell culture systems, and responsive drug delivery platforms. By providing controlled shape recovery, biocompatibility, and enhanced tissue adaptability, SMHs represent a promising avenue for the development of next-generation biomedical scaffolds and therapeutic systems [6].

This review provides a comprehensive overview of the current state and prospects of SMHs in tissue engineering, highlighting their innovative design and synthesis strategies, as well as the physical and biological properties that make them suitable for various biomedical applications. SMHs, renowned for their ability to fix quickly or recover permanently following external stimulation, have recently attracted considerable attention. The applications of these technologies span beyond traditional tissue engineering into areas such as bone regeneration, soft tissue reconstruction, vascular and neural tissue engineering, and emerging fields like smart drug delivery systems and 3D/4D-printed customized implants [7]. This review further explores the mechanisms involved in the shape memory effect (SME) in SMHs, examining how SME is activated by a range of stimuli, such as temperature, light, chemicals, sound, electricity, and magnetism. By analyzing current research and developments, this review highlights significant advancements in SMH technology, while also identifying remaining challenges and future research directions to enhance the clinical applicability of these materials (Fig. 1). Overall, SMHs represent a promising frontier in tissue engineering and regenerative medicine, offering innovative solutions to complex biomedical challenges [8].

Fig. 1.

Schematics of SMHs and their role in tissue engineering.

2. Design and synthesis of SMHs

2.1. Basic mechanism and design principles of SMHs

The concept of “shape memory" was first proposed in 1930 to describe a phenomenon found in AuCd alloys [9]. Shape memory materials are a class of stimulus-responsive intelligent materials capable of adopting a temporary shape under specific environmental conditions and subsequently reverting to their original form upon exposure to an appropriate stimulus. These stimuli can include temperature, light, electricity, magnetism, chemical agents, and pH changes, among others. The process of transitioning from a temporary shape back to the original one is known as the SME. To achieve shape memory, three key conditions must be met. (1) A stationary phase: this phase maintains the hydrogel's structure and records the initial shape. (2) A deformation program: this program induces deformation of the hydrogel's elastic network. (3) A reversible phase: also known as the “transition switch," this phase can open and close in response to external stimuli, allowing the temporary shape to be fixed and then returned to the initial shape [10]. The “transition switch" may involve chain segments or crystalline phases with low glass transition temperatures. It can also include dynamic chemical bonds, such as hydrazone, borate, or disulfide bonds, as well as supramolecular interactions, including ionic bonds, hydrogen bonds, hydrophobic bonding, host-guest interactions, coordination with metal ions, and others.

SMHs are polymer materials with shape memory properties, generally with original and reversible cross-linking networks. It changes the reversible cross-linking network in the cross-linked structure of the hydrogel under the action of one or more stimuli, and the temporary shape of the hydrogel is fixed [11]. However, when the external stimulus is withdrawn or changed, the cross-linked structure of the hydrogel responds to this stimulus, and the temporary shape returns to the permanent shape. As shown in Fig. 2, the original cross-linking network can be a physical or chemical cross-linking network. Reversible cross-linking sites are constructed using reversible interactions. The original cross-linking network determines the original shape of the SMH, while its temporary shape is stabilized by generating a new reversible cross-linking network [12]. The hydrogel returns to its original shape, depending on the original cross-linking network, by applying appropriate stimulation to break the reversible cross-linking network. The shape recovery of SMH is mostly carried out in a water environment. The shape memory recovery of SMH can be promoted by introducing water into the system or altering the water environment of the material, such as changes in temperature, pH, ion concentration, or redox potential [12]. Additionally, stimuli such as ultraviolet and ultrasound can also induce shape memory recovery in SMH.

Fig. 2.

Schematic of the changes in crosslinking structure of SMHs. Generated with biorender (Biorender.com) and adopted with permission [19]. Copyright 2017, pubs.rsc.org.

The rational design of SMHs for tissue engineering requires a clear understanding of how molecular structure governs macroscopic properties and ultimately determines biological function. This structure–property–function relationship is central to engineering SMHs with predictable mechanical behavior, shape transformation, biocompatibility, and degradation kinetics suited to specific tissue targets. Tissue-specific applications further require that structure and mechanical behavior align with functional goals. In cartilage engineering, SMHs must endure repeated compressive strain while maintaining extracellular matrix (ECM) compatibility. In neural applications, ultra-soft SMHs fabricated from silk sericin, PNAGAm, or gelatin-based materials promote neural cell viability and alignment while minimizing glial scarring. In smart drug delivery, structure significantly affects loading and release. SMHs with hydrophobic domains or dynamic crosslinks can achieve dual-stage release, for example, 40 % of the drug release within 6 h, with a sustained release over 60 days. In 3D printing, shear-thinning and rapid gelation enable SMHs to be printed in compressed configurations, allowing them to recover their designed architecture upon exposure to physiological conditions [13]. Therefore, the structure and design of SMH play a crucial role in determining its suitability for specific applications.

Researchers have extensively investigated various methods for classifying shape-memory polymers, each offering unique insights into their properties and behaviors [14]. According to the classification of shape memory function, it can be divided into unidirectional, bidirectional, multiple shape memory, and multifunctional shape-memory hydrogel [15]. However, the design of the SMH molecular structure is inseparable from the choice of action mechanism. Therefore, SMH realizes its shape memory function by introducing reversible action mechanisms, such as hydrogen bonding, metal coordination, and host-guest interactions.

To rationally select a stimulus for a given biomedical task, it is essential to link each trigger to its reversible chemistry. Thermal actuation relies on the melting and recrystallization of semicrystalline segments, such as poly(N-isopropylacrylamide) (PNIPAm) or poly(vinyl alcohol) (PVA), enabling rapid shape recovery at body temperature for injectable bone or cartilage fillers. Triggers based on pH and ionic strength operate through reversible ionic coordination and hydrogen bond clusters, for example, catechol Fe³⁺ or COO⁻ Ca²⁺ interactions, which confer wet adhesion and stable fixation in highly hydrated soft tissue niches. Dynamic covalent bonds, such as boronate esters and imines, enable multiple shape-locking cycles under fluctuating micro-pH conditions. In contrast, photo-responsive motifs such as azobenzene, spiropyran, or photothermal fillers, including gold nanorods and polydopamine, provide remote spatiotemporal control, making them suitable for minimally invasive drug release depots or vascular embolic devices. Aligning the working window of each stimulus with local physiological conditions, therefore, dictates application specificity [16].

Under physiological conditions, the permanent (covalent) network and the transient (reversible) network perform complementary functions. The permanent network supplies the elastic restoring force that drives shape recovery. Increasing its crosslink density enhances the recovery ratio (R_r) but reduces the number of reversible locking points, thus lowering the shape-fixity (R_f). In contrast, a denser transient network, composed of hydrogen bonds, ionic interactions, or coordination bonds, provides additional temporary crosslinks that improve R_f but also partly restrict the release of stored elastic energy, and can therefore reduce R_r. By adjusting the ratio of permanent to transient crosslinks, SMHs can be tailored to specific applications, with a higher proportion of transient network supporting long-term shape retention and a higher permanent network fraction enabling rapid and complete shape recovery [17].

For example, thermally responsive SMHs (triggered by body heat or mild heating) often recover very rapidly; for instance, some polyacrylamide/elastin-like dual-network SMHs revert to their permanent shape within 2–3 s. These can achieve nearly 100 % recovery quickly, which is ideal for applications such as self-expanding stents. However, achieving a very fast response usually requires a flexible, low network or reduced crosslink density, which can limit stiffness. In contrast, high-strength SMHs can approach “rubber-like” mechanical moduli, comparable to natural smooth muscle. Such rigid or double-network SMHs may utilize strong covalent crosslinks or reinforcing fillers to achieve high compressive strength; however, these tougher networks typically recover more slowly or require stronger/more specific stimuli (e.g., complete hydration or higher temperature) to fully return to their original shape. In practice, researchers often balance stiffness and speed; many tissue-engineering SMHs utilize dual networks or nanofillers to enhance strength, accepting a slightly longer recovery time [18].

Moreover, a simple SMH responds to a single trigger (e.g., temperature, pH, moisture). These single-stimulus SMHs are easier to design and often give reliable, rapid shape changes. For example, body-temperature–sensitive SMHs can actuate quickly at 37 °C without additional intervention. In contrast, multi-stimulus SMHs are engineered to respond only when two or more conditions are met (e.g., temperature and pH, or light and ionic strength). Multi-responsive designs offer finer control and reduce false triggers, but at the cost of complexity. Very few multi-stimuli SMHs have been demonstrated in practice; in fact, the incorporation of true multi-responsive or stepwise reconfigurable SMHs is still largely unexplored. In general, multi-stimulus SMHs may have slower or staged actuation, since each trigger condition must be satisfied, and they often require more complex fabrication (combining different chemistries or nanocomponents).

2.1.1. SMHs based on hydrogen bond interactions

A hydrogen bond is a short and directed electronic interaction between a donor and an acceptor, a special intermolecular force that can be constantly broken and reconstructed under the influence of the external environment. However, hydrogen bonding forces are weaker than many covalent and ionic bonds, ranging from 5 to 30 kJ mol⁻¹. However, the synergistic effect of multiple hydrogen bonds can increase the directionality and strength of hydrogen bonds to overcome the weak effect of single hydrogen bonds [20]. The strength and reversibility of hydrogen bonds are influenced by temperature, pH, and the type of solvent. Thus, hydrogen bonding is an ideal switch for constructing shape memory polymers. As one of the most widely used physical interactions, hydrogen bonding has been employed in the synthesis of shape-memory polymers over the past few decades [21]. Feng et al. [22] synthesized a double-network SMH with multiple hydrogen bonds by heating and cooling N-acryloyl glycinamide (NAGA), N-benzyl acrylamide (NBAA), agar, and UV-initiator in the aqueous phase with photoinitiated free radical copolymerization, as shown in Fig. 3A. The physical cross-linking network of P(NAGA-co-NBAA) was formed by hydrogen bond interaction between NAGA and NBAA on the side chain of the P(NAGA-co-NBAA) network obtained by free radical initiated copolymerization. However, linear agarose macromolecules form a double helix network structure due to hydrogen bonding. In addition, the two networks are interconnected by hydrogen bonds. Moreover, the reversibility of hydrogen bonds under temperature changes endows the gel with a temperature-responsive SME.

Fig. 3.

A) Synthesis scheme of self-healing and shape memory P(NAGA-co-NBAA)/agar DN hydrogels and the multi-H bond network of the hydrogel. Reproduced with permission [22]. Copyright 2018, Wiley Online Library. B) The temperature-responsive shape memory mechanism of PVA-TA hydrogel and the three shape memory cycles of PVA-TA hydrogel. Reproduced with permission [23]. Copyright 2016, ACS. C) Five-fold shape memory mechanism of PAA‐GO‐Fe³⁺ hydrogels. Reproduced with permission [25]. Copyright 2017, Wiley Online Library. D) Schematic representation of the synthesis of 4-armed PEG-DA self-healing SMHs and Fe³⁺ coordination regulating self-healing ability and pH-induced shape memory function. Reproduced with permission [26].Copyright 2019, pubs.rsc.org. E) Redox response mechanism of host-guest SMH. Reproduced with permission [27]. Copyright 2015, Wiley Online Library. F) Structure and temperature response principle of SMHs based on the crystalline phase. Reproduced with permission [28]. Copyright 2018, Wiley Online Library. G) Preparation and structure diagram of SMHs based on borate ester bonds. Reproduced with permission [29]. Copyright 2017, ACS.

In addition, Chen et al. [23] prepared an SMH using multiple hydrogen bonds for temperature response. As shown in Fig. 3B, they used PVA and tannic acid (TA) to form a strong hydrogen bond when they were physically mixed at high temperatures. When they were cooled to room temperature, they could form a gel. At this time, the strong hydrogen bond between PVA and TA was utilized as a permanent crosslink to stabilize the permanent shape, while the PVA itself formed a weak hydrogen bond between the chains. It serves as a switch that locks the temporary shape into place. The breaking and reformation of the hydrogen bond represent the recovery of the permanent shape and the stabilization of the temporary shape, respectively. Moreover, this type of shape memory behavior, based on strong and weak hydrogen bonds, exhibits strong repeatability, demonstrating excellent shape memory performance.

2.1.2. SMHs based on metal-ligand coordination

Metal coordination refers to the complexation between metal ions and organic or inorganic ligands. Under various supramolecular interactions, the coordination bonds between metal and ligands are extremely strong, and the bond strength can vary significantly from 50 to 400 kJ mol⁻¹, which is determined by the type of metal ion employed and the type of ligand used. Usually, these types of strong bonds are capable of being broken and re-established in response to external stimuli. Consequently, metal-ligand binding interactions have been widely utilized in SMH synthesis [24]. Ligands of metal ions are generally nitrogenous and oxygenic functional groups, such as imidazole, phosphoric acid, and carboxyl. Zhao et al. [25] mixed GO and acrylic acid and then performed free radical copolymerization to obtain hydrogen-bonded cross-linked polyacrylic acid hydrogel. Then, the hydrogels were immersed in a Fe³⁺ solution to coordinate Fe³⁺ and COO⁻ to form an ion-crosslinked structure. By adjusting the type and time of immersion in solvent, the ionic crosslinking degree and the coordination type of Fe³⁺ and COO⁻ were changed, and the five-fold SME was realized, as shown in Fig. 3C.

In addition, Lu et al. [26] were inspired by the marine mussel epidermis, and they prepared a hydrogel with four-arm polyethylene glycol modified with dopamine terminal iron ions, as shown in Fig. 3D. The metal coordination between catechol and Fe³⁺ can be reversibly changed by altering the pH value, which functions as a switch to fix the temporary form and restore the original form. Moreover, the presence of dynamic coordination bonds between the metal and ligand also facilitates self-healing of the hydrogel.

2.1.3. SMHs based on host-guest interactions

Additionally, the interaction between the host and the guest plays an important role in the preparation of SMHs. Host molecules of hydrophobic cavities within an object form a coating structure with hydrophobic molecules. As a result, the formation of the gel can be influenced by different stimuli, depending on the host molecules and the object, thus allowing a variety of SMHs to be prepared through the interaction between the subject and object. Host-guest interactions endow materials with unique functions, such as macroscopic self-assembly of polymeric gels, stimuli-responsive self-repair, and macroscopic movement of biomimetic artificial muscles. This kind of host-object interaction has received much attention from researchers. Kohei Miyamae et al. [27] synthesized SMHs by radical copolymerization of host-guest inclusion complexes using acrylamide-modified ferrocene, adamantane, and cyclodextrin. Fig. 3E shows adamantane-cyclodextrin inclusion as a stationary phase and ferrocene-cyclodextrin inclusion as a stimulus-responsive reversible phase. The transition of hydrophilicity and hydrophobicity of ferrocene in the REDOX state leads to the reversible destruction and formation of ferrocene-cyclodextrin inclusions in the REDOX state, which eventually leads to the fixed temporary shape of the gel in the reduction state and the shape recovery in the oxidation state.

2.1.4. SMHs based on crystalline phase

Crystallization is a physical interaction between molecular chains with special conformations. When the gel system contains components capable of crystallization, the modulus of the hydrogel is affected by crystallization, enabling the realization of shape memory performance. Zhao et al. [28] prepared organic hydrogels by in situ emulsion polymerization using acrylic acid, acrylamide, silica nanoparticles, oleophilic octadecanoate methacrylate, and n-eicosane. As shown in Fig. 3F, there are chemical cross-linked structures, a metal coordination structure formed by the coordination of Fe³⁺ and COO⁻, and a crystalline phase formed by the hydrophobic interaction of the aliphatic chains. The SME triggered by temperature is accomplished by utilizing the crystalline phase as a reversible component, utilizing its ability to undergo phase transitions in response to temperature fluctuations.

2.1.5. SMHs based on other mechanisms

In addition to the above interactions, shape memory functions can be achieved by introducing other reversible mechanisms of action, such as SMHs based on borate ester bonds. For example, phenylboronic acid is negatively charged when the pH of the aqueous solution is greater than the pKa value of phenylboronic acid (pKa≈8.5). The negatively charged phenylboronic acid can form reversible borate ester bonds with molecules that have a diol structure (such as PVA, glucose, and catechol), which can be reversibly formed and broken under pH changes. Le et al. [29] mixed gelatin powder, acrylic acid, acrylamide, double-bond modified phenylboronic acid, and glucose, heated and cooled it to form hydrogen-bonded crosslinked hydrogels. After radical copolymerization, the hydrogel with a semi-interpenetrating network structure was prepared, in which the polypropylene backbone was grafted with carboxylate, glucose, and phenylboronic acid. Under alkaline conditions, phenylboronic acid will form a dynamic, reversible borate ester bond with glucose. In the Fe³⁺ solution, the coordination of Fe³⁺ with COO⁻ will form a dynamic reversible ionic bond, as shown in Fig. 3G. Thus, three reversible phases that respond to temperature, pH, and Fe³⁺, respectively, are introduced, and a quadruple SMH with a triple stimulus response is obtained.

SMHs based on hydrophobic interaction are a kind of effective physical interaction, which refers to the tendency of the non-hydrophilic part of a molecule to aggregate within or between molecules in an aqueous medium. Hu et al. [30] carried out UV-initiated copolymerization of acrylamide, methylcellulose, and cross-linking agent to obtain a PAAM chemical cross-linking network. Based on the fact that under heating conditions, hydrophobic repeating units in methylcellulose aggregate to form hydrophobic interaction regions, and under cooling conditions, these hydrophobic interactions are disrupted. The cooling-induced SME is achieved by using PAAM as the stationary phase and the methyl cellulose hydrophobic network as the reversible phase, which overcomes the disadvantage of the high recovery temperature required for the temperature-induced SMH.

Additionally, SMHs based on Schiff base bonding are also widely utilized. Xiao et al. [31]blended chitosan, oxidized glucan, acrylamide, and cross-linking agent and then carried out free radical copolymerization to obtain a hydrogel with a semi-interpenetrating network. Chemically crosslinked polyacrylamide (PAM) networks can impart a permanent shape to the hydrogel. The Schiff base bonds formed between the amino groups on chitosan and the aldehyde groups on oxidized glucan under alkaline conditions stabilize the temporary shape of the hydrogel, and the shape is restored in response to changes in pH, amino acids, and vitamin B6. In addition, the reversible metal coordination bond formed by metal ions and chitosan can also fix the temporary shape and respond to EDTA, H₂SO₄, and oxidants to achieve a dual SME. Furthermore, the gel can achieve triple shape memory by programmed formation and destruction of metal coordination bonds and Schiff bases.

2.1.6. Hydration and diffusivity in SMHs

The role of hydration and diffusivity in SMHs is very crucial. Hydration directly affects the swelling behavior and mechanical properties of SMHs, which are vital for their application in tissue engineering [32]. For example, hydration can adjust the swelling of the hydrogel to mimic the characteristics of human tissues, thereby enhancing its compatibility with cells and its adhesion to tissues. Hydration not only affects the volume change but also significantly improves the softness and plasticity of the hydrogel, which is important for applications in complex environments such as soft tissue and bone tissue repair. The degree of hydration in hydrogels is determined by various factors, including polymer composition, crosslink density, and hydrophilicity, which determine how closely it mimics the extracellular matrix.

Diffusivity is a key factor that determines the functionality of hydrogels in transporting various molecules. Hydrogel network structure, porosity, and hydration influence the transport of molecules within it (e.g., nutrients, drugs, and signaling factors). Diffusivity is generally enhanced by high water content, but mechanical stability can be compromised by excessive swelling. The diffusivity of hydrogels influences the distribution and release rate of drugs or other biomolecules within them [33]. For example, in smart drug delivery systems, the hydration state of the hydrogel affects the diffusion rate of drug molecules inside the hydrogel, thereby determining the timing and duration of drug release. Therefore, good hydration and diffusivity enable the precise release of drugs and biomolecules, enhancing therapeutic efficacy while minimizing side effects.

2.1.7. Kinetics of shape recovery

The recovery time of SMHs typically depends on several factors, including the material composition, the type of stimulus used (such as temperature, pH, light, etc.), and external environmental conditions (such as humidity and pressure). For example, the shape recovery time of temperature-responsive hydrogels may vary between low and high temperatures, typically ranging from seconds to minutes. In contrast, the recovery time of pH-responsive hydrogels may be longer, especially when significant pH changes are required, with recovery times potentially extending from minutes to several hours. Specifically, the duration of recovery time greatly impacts the choice of applications. For example, in the case of stent applications, it is necessary for the hydrogel to rapidly recover its original shape within the body, which typically requires a recovery time ranging from seconds to minutes. On the other hand, in tissue engineering, slower recovery times may be more beneficial, particularly when the hydrogel is used to support cell growth and proliferation. Slower shape recovery can better mimic the natural regenerative process of biological tissues.

In the medical field, SMHs are commonly used in devices such as vascular stents and airway stents. These devices typically must recover their predetermined shape within the body to effectively fulfill their supportive function. Rapid recovery (ranging from seconds to minutes) is especially critical for these devices. A quick recovery not only ensures the smooth use of the stent but also provides stable support to blood vessels and other tissues in the shortest possible time. For example, the HydroCoil endovascular system has been extensively studied for its application in the treatment of the internal carotid artery (ICA). The HydroCoil endovascular aneurysm embolization device utilizes the property of hydrogel swelling upon contact with liquids, which increases the volumetric filling of the aneurysm, thereby enhancing the therapeutic effect [34]. For SMHs in tissue engineering, slower shape recovery helps to mimic the natural tissue repair process. In this process, hydrogels may be used to construct scaffolds that provide the physical support needed for cell growth. In such cases, slower shape recovery times (ranging from minutes to hours) can provide a more stable environment, facilitating cell attachment and proliferation. Slow recovery also aids in controlling the release of growth factors or drugs, further promoting tissue repair and regeneration.

2.2. Classification of stimulation modes of SMHs

SMHs are a type of hydrogel that can fix and maintain a temporary shape under specific stress conditions and then return to the initial shape under certain temperatures, light, or other conditions. Shape-memory hydrogels typically contain two sets of cross-linking networks: a permanently fixed cross-linking network that determines the original shape of the hydrogel, and a variable network that fixes the temporary shape of the hydrogel in response to stimulation. Various conditions can regulate the temporary morphology of the SMH. For example, in the case of temperature-sensitive SMHs, the stimulus condition is the transition temperature. When the temperature exceeds the transition temperature, the weak interaction is disrupted, molecular chain flexibility increases, and the hydrogel can transition from its initial shape to a temporary shape. When the temperature is lowered again below the transition temperature, weak interactions reform, and the movement of the molecular chain is restricted, allowing a temporary shape to be maintained.

SMHs are classified according to the type of external stimulation they receive. Accordingly, numerous external stimuli can be used to induce shape change, such as temperature, water, light, electricity, magnetic fields, and pH stimulation, among others. Here, SMHs are categorized according to the mechanisms by which they are activated and the external stimuli to which they are subjected. Table 1 provides an overview of the various crosslinking techniques that can be applied to polymeric hydrogel networks (e.g., AAm, NIPAM, PEG) to produce a range of SMHs. Additionally, to facilitate the classification of SMHs from both fundamental chemistry and actuation mechanisms perspectives, the role of additives (such as ions, hybrid polymers, and nanofillers) is discussed as a means of obtaining SMHs with diverse stimuli-responsive characteristics.

Table 1.

Overview of the stimulus-response characteristics of various types of SMH.

Material	Stimulus	Mechanical properties	Recovery Mechanism	Rf. Rr.	Recovery time (s)	Application	Ref.
P(AAm) and ELP	Temperature (20 °C)	∼56 kPa	H-bonds → thermoresponsive	78.1 %, 75.6 %	2–3	Temperature driver	[66]
Collagen	Chemical (water)	∼70 kPa	H-bonds → water-responsive	–	–	Cartilage defect repair	[67]
P(AAm) and PEA	Temperature (37 °C)	5.1 MPa	H-bonds → thermoresponsive	100 %, 95 %	36	Embolization plug	[68]
PVA and TP	Chemical (water)	∼0.25 MPa		90 %(Rr)	30	Wound dressing and human skin	[69]
P(AAm) and MAA	Chemical (pH 2)	–	H-bonds → pH-responsive	80 %(Rf)		Soft robots and bionics	[70]
P(AAm), AAc, and Ad-Am	Chemical (EtOH)	64 MPa	Ionic bonds→ solvent-responsive	–		Artificial muscle	[71]
PEG, MDI, and IU	Temperature (4 °C–50 °C)	∼2.3 MPa	H-bonds → thermoresponsive	95 %	1	Flexible electronic devices	[72]
PEG, MDI, and IU	Temperature (37 °C)	1.1 MPa	H-bonds → thermoresponsive	100 %		Invasive surgery	[73]
P(AAm), PDA, AA, and alginate	Chemical (Ferric Chloride)	0.021 MPa	H-bonds → chemical-responsive	100 %		Drug delivery	[74]
P(AA-co-AN)	Chemical (pH)	∼23 MPa	H-bonds → pH-responsive	100 %, 100 %		Flexible electronic products	[75]
P(NIPAm) and PDMS	Light (NIR = 800 nm)	–	H-bonds → NIR light-responsive	–		Soft robots	[76]
PDA,CMC-Fe3+,and PVA	NIR/Temperature (2W, 60 °C)	163.48 kPa	H-bonds → thermoresponsive/NIR-responsive	–		Artificial intelligence	[77]
P(NIPAm-AA), PPy and alginate	Light (NIR = 800 nm)		H-bonds → NIR-responsive	–		Conjugated polymers	[78]
PCL	Electrical (2 V)	–	H-bonds → electroresponsive	95 %, 99 %		Soft robotics	[79]
PCL and PEG	Temperature (37 °C)	–	H-bonds → thermoresponsive	90 %, 80 %		Minimally invasive implantation	[80]
P(SSA-co-DEAEMA)	Chemical (pH)	42.5 MPa	Ionic bonds→pH-responsive	100 %, 70 %		Actuating soft machines	[81]
P(NaSS-DMAEMA-co-MPTC)	Chemical (pH)	680 kPa	Ionic bonds→pH-responsive	70 %, 85 %		Flexible wearable devices	[82]
(PF127-DA -co-AAm)	Chemical (Ethanol)	∼180 kPa	H-bonds → solvent-responsive	97 %. 100 %		Stress sensors	[83]
PVA and chitosan	Temperature (60 °C)	12.79 MPa	H-bonds → thermoresponsive	80 %, 95 %		Human sensors	[84]
PVA and cornstarch	Chemical (water)	–	H-bonds → water-responsive	97–99 %,100 %		Crohn's fistula treatment	[85]
F127DA and SA	Chemical (Ca2+)	300–800 kPa	Ca2+ coordination bonds → ion-responsive	81.47 %, 98.15 %		Drug delivery and tissue engineering scaffold	[86]
PVA and TA–Fe3+	Light (NIR = 808)	22.7 kPa	H-bonds → NIR-responsive	87 %(Rr)		Biomedicine and intelligent engineering	[87]
PVA and chitosan	Electricity (40 V)	3.68 MPa	H-bonds → electroresponsive	100 %(Rr)		Tissue engineering and smart sensors	[88]
PU	Temperature (10 °C)	0.44 MPa	Van der Waals bonds → thermoresponsive	–		Peripheral nerves treatment	[89]
PU, CHTMA, LAMMA and GELMA	Temperature (37 °C)	–	H-bonds → thermoresponsive	50–90 %, 100 %		Biological and nonbiological	[90]
P(AN-co-AAm)	Temperature (20–40 °C)	–	H-bonds → thermoresponsive	97.5 %, 100 %		Embolic agents	[91]
P(AN-co-ACG)	Temperature/Chemical (37 °C and pH 6)	12.44 MPa	H-bonds & ionic bonds → thermoresponsive/pH-responsive	94.4 %, 90.9 %		Embolic agents	[92]
PCL diol - PDLLA	Temperature (37 °C)	430 MPa	H-bonds → thermoresponsive	95 %, 98 %		Bio-ink	[93]

Abbreviations:R_f, shape fixity ratio; R_r, shape recovery ratio; P(AAm), polyacrylamide; ELP, elastin-like polypeptide; PEA, 2- phenoxyethyl acrylate; SA, sodium alginate; IU, imidazolidinyl urea; PEG, poly (ethylene glycol); MDI, methylenediphenyl 4, 4-diisocyanate; PDA, poly (diacetone acrylamide); PAA, Poly(acrylic acid); PAN, Polyacrylonitrile; P(NIPAM), poly(N-isopropylacrylamide); PDMS, poly (dimethylsiloxane); PPy, polypyrrole; PCL, polycaprolactone; PSSA, poly(styrene sulfonic acid); DEAEMA, 2-(diethylamino)ethyl methacrylate; NaSS, sodium p-styrenesulfonate hydrateMPTC, [3-(methacryloylamino)propyl]trimethylammonium chloride; DMAEMA, 2-(dimethylamino)- ethyl methacrylate; PVA, poly vinyl alcohol; TA, tannic acid; PU, polyurethane; ACG, N-acryloyl 2-glycine; NaAC, sodium acrylate; DMAEMA Q, quaternized dimethylaminoethyl methacrylate; XNBR, carboxylated nitrile rubber; CMC, carboxymethyl cellulose; TP, tea polyphenol; PDLLA, poly(D, L-lactide).

2.2.1. Temperature-responsive SMHs

Temperature, as a stimulus source, exhibits certain controllability and holds great application prospects in the biomedical field. Temperature-responsive SMH has become the most widely studied SMH. Temperature-responsive SMHs are typically constructed in one of two ways: (1) by embedding a temperature-sensitive group within the polymer chain. As the temperature increases, the polymer undergoes a transition in which its hydrophobicity changes, defined by the upper or lower critical solution temperature, resulting in a hydrogel with a critical transition temperature. Additionally, (2) it utilizes temperature responses through reversible hydrogen bonding or crystallization processes in interpenetrating networks.

In 1995, Osada, Y et al. [35] prepared the first temperature-responsive SMH with polyacrylic acid as the main chain and stearyl acrylate as the short side chain. The hydrophobic backbone enables the hydrogel to absorb water and swell. The short-side chains undergo crystallization and aggregation at temperatures below their transition temperature (T_trans), and the formed microcrystals act as temporary cross-linking points to fix the temporary shape of the hydrogel. When heating increases the temperature of the hydrogel above T_trans, the microcrystals that act as temporary crosslinking points are destroyed and become amorphous, allowing the hydrogel to return to its original form. Similarly, other crystallizing side chains can be introduced into the hydrogel system to act as temporary forces, fixing the shape of the hydrogel and facilitating the fabrication of temperature-responsive SMHs. All the above-mentioned side chains respond to temperature stimulation, and the shape memory function of the hydrogel is realized under a single network. If the crystallizing units are not coupled in the crosslinked network structure, a second network that immobilizes the hydrogel morphology can be introduced, thereby enabling the shape memory function.

Polyvinyl alcohol (PVA) possesses a significant number of hydroxyl groups in the molecular chain and is capable of forming numerous hydrogen bonds with water. PVA has the characteristics of good water absorption, low toxicity, and biocompatibility, making it an ideal material for the preparation of hydrogels. After freeze-thaw treatment, PVA molecules can produce crystal structures. However, at high temperatures, the crystal structure will be destroyed. This process is reversible, thereby facilitating the realization of a hydrogel shape memory function. Li et al. [36] developed a dual-network SMH by chemically cross-linking polyethylene glycol (PEG) to fix the initial shape of the hydrogel, as illustrated in Fig. 4A, while incorporating an interpenetrating polyvinyl alcohol (PVA) network. The hydrogel adopted a temporary shape when a deforming force was applied at low temperatures. Upon removing the external force, the hydrogel was immersed in hot water at 90 °C. Within 15 s, the PVA chains melted and recrystallized, enabling the hydrogel to recover its original shape, thus demonstrating effective shape memory behavior. Moreover, the crystallized PVA network also conferred self-healing properties to the hydrogel.

Fig. 4.

SMHs with thermal responsiveness. A) PVA/PEG double-network hydrogel with shape memory and self-healing functionalities. Reproduced with permission [36]. Copyright 2015, ACS. B) Channeled PNIPAm (ch-PNIPAm) hydrogels: (I) Chemical structures of PEGDA, NIPAm, and agarose; (II) Hydrogel synthesis schematic; (III) Photographs of ch-PNIPAm and standard PNIPAm at 25 °C and 60 °C; (IV) Total reflectance at both temperatures; (V) Temperature-dependent transmittance. Scale bar: 5 mm. Reproduced with permission [38]. Copyright 2020, Wiley Online Library. C) Thermally induced dual- and triple-shape memory in Dns15-PAAM demonstrated by sequential recovery steps at elevated temperatures. Reproduced with permission [39]. Copyright 2016, ACS. D) Multi-shape memory behavior triggered by Fe³⁺, pH, and temperature; stabilized via coil-helix transitions and PBA-diol ester bonds. Reproduced with permission [29]. Copyright 2017, ACS. E) Triple-shape memory at macro- and microscale, achieved via ionic interactions and pH-triggered transitions between CaCl₂, Gly–NaOH, and K₂CO₃. Reproduced with permission [40]. Copyright 2016, pubs.rsc.org. F) SMHs showing shape memory and self-healing: (I–II) Editable 2D/3D shapes; (III–IV) Recyclability before/after healing; (V) Rheological behavior during heating/cooling; (VI–VII) Two- and multi-step programmable shape recovery. Reproduced with permission [43]. Copyright 2024, pubs.rsc.org.

Poly(N-isopropylacrylamide) (PNIPAm) exhibits excellent temperature sensitivity due to the intra- and inter-chain interactions between hydrophilic and hydrophobic groups and water molecules. At lower temperatures, hydrogen bonding between the amide groups of PNIPAm and water molecules is the primary interaction. This results in a highly ordered solvation shell, stabilized by hydrogen bonds and van der Waals forces, surrounding the polymer chain. However, as the temperature increases, the interaction between PNIPAm and water undergoes significant changes. Hydrophobic interactions between the polymer chains intensify, leading to the disruption of hydrogen bonds. Consequently, the hydrophobic regions of the polymer disrupt the solvation layer, expelling water molecules and triggering a phase transition. This process underlies PNIPAm's temperature sensitivity. The temperature at this point is defined as the lower critical solution temperature (LCST) of PNIPAm. Because PNIPAm exhibits good chemical stability, low toxicity, low pH dependency, is electrically neutral, has an LCST (∼32 °C) close to body temperature, and is easily adjustable with monomers, it has been used in many applications. Temperature-responsive SMHs, often based on LCST transitions in polymers like PNIPAm (LCST ∼32–42 °C), enable rapid entropy-driven shape recovery (typically 5–30 s) through reversible hydrophobic collapse, making them advantageous for soft tissue regeneration where they support cell encapsulation (>90 % viability) and dynamic scaffold modulation, as demonstrated in rabbit cartilage models achieving 80–85 % defect closure over 12 weeks [37]. As shown in Fig. 4B, Amanda Eklund et al. [38] demonstrated the use of switchable channel PNIPAm hydrogels (ch-PNIPAm) that displayed vibrant white colors as the gels were heated above their phase transition temperatures. A ch-PNIPAm hydrogel was formed by removing the agarose network through a semi-interpenetrating network, which was created by the physical cross-linking of agarose and the chemical cross-linking of the PNIPAm network. The agarose is used as a template to form nanoscale channels within the hydrogel that are beneficial for transporting water and forming smaller pores during phase transition, resulting in bright white coloration and fast kinetics. These hydrogels were found to respond to heat 18 times more rapidly than conventional PNIPAm hydrogels.

Gong et al. [39] synthesized a versatile pH- and thermal-responsive multi-SMH by introducing dansyl groups into the network of chemically crosslinked polyacrylamide (Dns-PAAM). As shown in Fig. 4C, hydrophobic aggregations of dansyl groups function as molecular switches. Under changes in pH or temperature, they undergo a reversible transition from aggregation to association, and revert to their original shape when the stimulus is removed. Interestingly, since electrolysis can alter the pH of the electrolyte solution, the current in the electrolyte solution can also trigger the Dns-PAAM hydrogel to exhibit its shape-memory properties.

In addition, Le et al. [29] prepared SMHs with thermal, Fe³⁺, and pH responses by mixing agar, acrylic acid, acrylamide, glucosamine, and aminophenylboronic acid (APBA). As shown in Fig. 4D, the prepared gel contains a variety of functional groups, including a boric acid group, a cis dihydroxy group, and a carboxyl group, among others. Under alkaline conditions, boric acid groups can combine with cis-dihydroxy groups to form dynamic borate ester bonds. The carboxyl group can chelate Fe³⁺ and form a stable tridentate bond. Agar molecules can undergo the coiled-to-helix transformation by increasing or decreasing their temperature. As a result, the three reversible cross-links are useful for preserving the temporary shape of the hydrogel, which facilitates its shape memory capabilities. The design and preparation of SMHs with multiple stimulus responses has become a research hotspot.

Li et al. [40] proposed the development of a new mechanically stretchable supramolecular hydrogel exhibiting a triple SME at both macro and micro scales. As shown in Fig. 4E, their unique design, featuring two non-interfering supramolecular interactions, namely, the dynamic interaction between the phenylboronic acid (PBA)-diol ester bond and alginate/Ca²⁺ chelation, enabled the hydrogel to exhibit excellent triple-shape memory function. In addition, Hua et al. [41] used acrylic acid and calcium acetate to prepare a stimulus-response SMH with low-temperature induced shape recovery. Firstly, PAA hydrogel was prepared using acrylic acid (AA) as the raw material, N, N′-methylene diacrylamide as the crosslinking agent, and ammonium persulfate as the initiator. The prepared PAA hydrogel was then immersed in a calcium acetate solution, where the Ca²⁺ ions in the solution would complex with the carboxyl groups in the gel. As shown in Fig. 4F, due to the low solubility of the acetic acid group at high temperatures, which leads to hydrophobic aggregation and hardening of the polyacrylic acid network, the temporary shape of the hydrogel can be fixed at 70 °C and restored to the initial shape at 20 °C. At the same time, shape restoration can be achieved by adjusting the shape fixation time. This unique shape memory mechanism enables the application of such SMHs in environments with high temperature and heat resistance.

Despite these advances, temperature-responsive SMHs face certain limitations, including the risk of hyperthermia-induced cellular apoptosis (15–25 % at >43 °C via heat shock protein disruption) and the need for precise in vivo temperature control, often requiring external devices like inductive heaters, which reduces practicality for deep-tissue applications [42]. Overall, the clinical feasibility is moderate, favoring ex vivo preconditioning; however, recent designs tuned to physiological temperatures mitigate overheating risks.

2.2.2. Chemically-responsive SMHs

SMHs, which are chemically responsive, can change shape in response to external chemical stimuli. These chemical factors include solvent, ions, pH, redox reaction, and special gases such as CO₂ and NH₃ molecules. Chemically-responsive SMHs, triggered by pH shifts (e.g., protonation at pKa ∼4.5–6.5) or ions (e.g., Ca²⁺ coordination with Kd ∼10⁻⁵ M), provide endogenous actuation without external energy, facilitating targeted drug release (zero-order kinetics over 24–72 h, efficiency 70–85 %) and vascular regeneration with 90–95 % endothelialization in rat models [44]. Their main benefits lie in their biomimetic design and cost-effective production, achieving synthesis yields greater than 80 %.

Metal ions can complex with groups bearing anions in polyelectrolytes. The SMH with ion response can be prepared by the complexation of a metal and a ligand within the hydrogel. The complexation between metal ions in the solution and ligands on the polymer chain can be utilized as reversible cross-linking points to maintain the temporary shape of the gel. The metal ions commonly used are Ca^{2 +}, Fe³⁺, Zn^{2 +} and Cu ^{2 +}. Ren et al. [45] prepared a hydrogen-bonded and Ca²⁺-crosslinked hydrogel from 2-vinyl-4, 6-diamino-1,3, 5-trioxide (VdT), acrylic acid (AAc), and polyethylene glycol diacrylate (PEGDA). As shown in Fig. 5A, reversible Ca²⁺ cross-linking endows the hydrogel with shape memory properties, allowing it to firmly retain various temporary shapes and revert to its initial shape upon addition of EDTA 2Na.

Fig. 5.

Examples of chemically responsive SMHs. A) PVDT-PAA-PBS hydrogel exhibiting Ca²⁺-induced reversible shape memory. Protocruciform hydrogel strips were transformed into box, pyramid, and spring shapes, then fixed in 100 mmol/L Ca²⁺ solution. Reproduced with permission [45]. Copyright 2015, pubs.rsc.org. B) Programmable shape memory in a “Butterfly Loves Flowers” hydrogel system: (I) Illustration of self-actuation, memory, and recovery; (II) Shape transformations triggered by pH and Fe³⁺; (III) Gradual shape recovery induced by increasing H⁺ concentration. Reproduced with permission [46]. Copyright 2018. Wiley Online Library. C) CO₂-triggered shape memory mechanism. Reproduced with permission [49]. Copyright 2015, Wiley Online Library. D) Photographs show the shape recovery behavior of Alg-PBA/PVA hydrogel in (Ⅰ) 0.2 M Gly (pH 6) solution, (Ⅱ) 0.2 M glucose solution, and (Ⅲ) alkaline aqueous solution (pH 10.6). Reproduced with permission [50]. Copyright 2015, Wiley Online Library.

Similar to ionic SMHs, pH-responsive SMHs can be endowed with pH-responsive shape memory properties by introducing chemical or physical regulators with pH sensitivity. Generally, pH-responsive SMHs contain acid or base groups that are easily hydrolyzed [46]. Moreover, pH is very important in biological and natural systems. Different organs of the human body have different environmental pH levels. It can be seen that pH is an attractive source of stimulation when designing SMHs of medical value. Le et al. [46] used PAAm as the main body through free radical polymerization and then introduced pH-responsive AAc monomers via UV photopolymerization. The PAAm-PAAc hydrogel with the best anisotropic structure can be obtained by controlling the photopolymerization time. As shown in Fig. 5B, in a 0.2 M NaOH solution, the carboxylic acid groups of polyacrylic acid absorb a significant amount of moisture, leading to swelling. This causes the ‘butterfly’ structure to bend downwards, opening its four ‘petals.’ When H⁺ ions are reintroduced, the structure recovers its original form. If Fe³⁺ ions are introduced, the temporary shape can be stabilized through chelation between the carboxylate groups and the iron ions, creating a ‘shape memory’ effect. Gradually increasing the concentration of HCl from 0.03 M to 0.05 M, and eventually to 0.1 M, disrupts the interaction between AAc and Fe³⁺, allowing the material to return to its original shape.

Zhang et al. [47] synthesized zwitterpolymer hydrogel by one-step copolymerization of cationic monomer 3- (methylacryloamino) propyl trimethylammonium chloride, anionic monomer sodium p-styrene sulfonate (NaSS), and methacrylic acid (MAA) without salt addition and formation of cross-link points. The hydrogel exhibited pH-responsive shape memory behavior. The hydrogel was soaked in a NaOH solution and then transferred to an HCl solution for fixation, forming a temporary shape. This shape could be restored by soaking the hydrogel in a NaOH solution again. Interestingly, this kind of hydrogel also has spontaneous actuation behavior. By alternately soaking in an acid-base solution, the hydrogel can transform spontaneously between its temporary and original shapes without the need for external force to deform, which overcomes the drawback of traditional SMHs that require external force to obtain temporary shapes.

Currently, gases such as CO₂ can also be used as an external stimulation mode for shape memory polymers. CO₂ is a crucial gas for cell metabolism, exhibiting good biocompatibility and cell membrane permeability [48]. Additionally, CO₂ can react with stimulus-responsive groups, such as amino groups or amidines, to form hydrophilic complexes. The initial group can be recovered when the CO₂ is removed by inert gas or heating, so using CO₂ as a trigger allows many conversion cycles to occur without producing by-products. Therefore, using the non-toxic gas CO₂ as a new trigger to realize the shape memory function of hydrogels further increases the diversification of SMHs. Xu et al. [49] prepared hydrogels by copolymerization of 2-vinyl-4, 6-diamino-1, 3, 5-triazine, N, N-dimethylacrylamide, and polyethylene glycol diacrylate through photoinitiator as shown in Fig. 5C. This hydrogel polymerization network structure contains diaminotriazine-diaminotriazine (DAT-DAT) hydrogen bonds and amide-amide hydrogen bonds, so this hydrogel is called a double hydrogen bond (DHB) hydrogel. Because the presence or absence of CO₂ leads to the destruction and reconstruction of DAT-DAT hydrogen bonds, DHB hydrogels show reversible CO₂ responsive volume and modulus changes. Based on this mechanism, CO₂ can induce the SME of DHB hydrogels without introducing external acids to make a CO₂-responsive SMH.

In addition, biomolecules such as glucose can also induce deformation in SMHs. For instance, Meng et al. [50] developed a glucose- and pH-responsive supramolecular SMH. At pH levels above 8.5, phenylboronic acid dissociates, acquiring a negative charge, and forms dynamic borate ester bonds with molecules that possess o-hydroxyl groups. As illustrated in Fig. 5D, phenylboronic acid grafted onto the alginate backbone (Alg-PBA) forms dynamic borate ester bonds with polyvinyl alcohol (PVA), allowing for shape memory and recovery through the reversible formation and breaking of these bonds. Upon exposure to glucose, fructose, or acidic conditions, the borate bonds break, erasing the temporary shape memory.

Despite these advantages, chemically-responsive SMHs face challenges, including slower responses (1–10 min) and potential cytotoxicity from leached ions (10–20 % oxidative stress via reactive oxygen species). Clinical feasibility is high for superficial sites, such as wounds; however, long-term ion accumulation may provoke inflammation. Nevertheless, targeted nanoparticle delivery systems can be incorporated to improve specificity.

2.2.3. Light-responsive SMHs

Light serves as a remote-controlled, non-contact stimulus, offering the ability to be switched in two-dimensional and three-dimensional spaces as well as over time through the simple modulation of the light source. Light energy can be adjusted based on its fundamental properties, such as wavelength, intensity, and other factors, with common categories including visible (Vis), near-infrared (NIR), and ultraviolet (UV) light. Light-responsive SMHs, incorporating chromophores for near-infrared (NIR, 700–900 nm) activation with penetration depths of 1–2 cm, allow spatiotemporal precision via photoisomerization or photothermal effects, achieving recovery in 10–60 s at low doses (1–5 mW/cm²) and supporting neural regeneration with >95 % neuron survival and 75–85 % axon guidance in mouse models. Their non-contact nature enhances feasibility, but phototoxicity (resulting in 5–15 % cell death from singlet oxygen) and device requirements limit their use in deep tissue [51]. Clinical feasibility is high, and recent 4D-printed NIR systems optimize dosage to reduce risks, paving the way for minimally invasive implants. During stimulation, the structure of the hydrogel network is altered, resulting in changes to its physical and chemical properties, which leads to gel transformation (as shown in Fig. 6A) [52]. Based on the underlying mechanism of action, the impact of light stimulation on hydrogels may be divided into three specific categories. The use of the light source can induce or disrupt chemical bonds, which affects the density of the cross-links in the hydrogel. Additionally, photosensitive molecules, such as azobenzene and spiropyran, can undergo isomerization when exposed to specific wavelengths of light, including cis-trans isomerization and ring-opening/ring-closing transitions, respectively. These processes may result in sol-gel transitions or alter the swelling behavior of a hydrogel. Additionally, the introduced hydrogel system incorporates nanomaterials, such as carbon nanotubes, graphene, and gold nanoparticles, which exhibit unique thermal properties. These nanomaterials can absorb light energy and convert it into heat, triggering a thermal response in the hydrogel. Dai et al. [53] prepared a high-toughness nanocomposite hydrogel by adding gold nanorods to the copolymerization network of 1-vinylimidazole and methacrylic acid. As shown in Fig. 6B, localized light irradiation was applied to the hydrogel to achieve precise regulation of its network structure and mechanical properties. The heating of gold nanorods caused a rapid increase in the temperature of the nanocomposite hydrogel, promoting the formation of denser hydrogen bonds, which in turn enhanced the mechanical properties in the targeted region.

Fig. 6.

Light-responsive SMHs. A) Changes in hydrogel structure induced by light stimulation. Reproduced with permission [52]. Copyright 2019, Wiley Online Library. B) Ⅰ) Synthesis of a tough nanocomposite hydrogel from MAA, VI, MBA, and gold nanorods in DMSO, followed by water exchange, forming dense hydrogen bonds; (II) Site-specific stiffening and shape retention under optical stimulation due to enhanced hydrogen bonding. Reproduced with permission [53]. Copyright 2020, ACS. C) Light-triggered shape recovery in SA/PAAm hydrogel via Fe³⁺ reduction, enabling optically programmable deformation. Reproduced with permission [54].Copyright 2019, ACS. D) Photographs of PCL2.8K-DA-15 % Fe3O4 network in near-infrared light showing one-way SME. Reproduced with permission [55]. Copyright 2018, ACS. E) PVA–PDAP composite hydrogel with dual noncovalent crosslinks exhibiting NIR-induced shape memory behavior. Reproduced with permission [56]. Copyright 2017, Wiley Online Library.

Among the many photostimuli-responsive hydrogels, near-infrared photostimuli-responsive SMHs have attracted more attention and favor due to their remote adjustability. This is mainly because the shape, size, and mechanical properties of the near-infrared light-responsive SMH can be adjusted in time and space without physically interacting with the hydrogel, thereby eliminating the limitation of contact stimulus-response. For example, Li et al. [54] prepared an SA/PAAm hydrogel with light-stimulated shape recovery by a radical polymerization reaction in alginate (SA) solution using acrylamide as raw material, ammonium persulfate (APS) as initiator, and N, N′-methylene diacrylamide (MBA) as crosslinker. Fig. 6C illustrates the process of forming a temporary shape of the gel using an external force, followed by immersion in a Fe3+ solution to facilitate reversible metal coordination and fix the shape. Subsequently, hydrogels with temporary shapes were exposed to UV light for a specified period. Since UV light reduces Fe3+ to Fe2+, the metal coordination that maintains the temporary shape will be disrupted, and the gel will revert to its initial shape.

Du et al. [55] constructed a metal-based supramolecular polycaprolactone (ε-caprolactone) network centered on pyroquinone. Fig. 6D shows photoresponsive SMEs due to Fe3O4 NPS acting as a nanoscale heat source that effectively absorbs near-infrared (NIR) light and converts it into thermal energy. In addition, Yang et al. [56] prepared polyvinyl alcohol (PVA) composite hydrogels doped with polydopamine (PDA) particles by a cyclic freeze-thaw method. During the freeze-thaw cycle, the hydrogen bond between PVA and PDA forms a physical cross-linking network, which can be used to fix the temporary shape. When using near-infrared light irradiation, the excellent photothermal conversion efficiency of polydopamine particles increases the local temperature of the gel, destroying the hydrogen bonds, and the gel gradually returns to its original shape. The near-infrared light response enables local focusing and quantitative stimulation, extending the range of use of SMHs.

However, in vivo implementation faces hurdles: visible and UV light lack sufficient penetration, while NIR exposure can cause localized overheating or phototoxicity. Implantation of optical fibers or probes increases surgical complexity and infection risk. Additionally, clinical light sources often entail bulky equipment with high power demands, limiting portability and widespread adoption. Future efforts should focus on optimizing light delivery and heat dissipation systems, engineering efficient photosensitive nanocarriers, and integrating multimodal (e.g., photo-thermal, magnetic, chemical) strategies to enhance the translational potential of light-responsive SMHs for deep-tissue and rapid-response applications.

2.2.4. Electrically responsive SMHs

Electric current can also be used as a stimulant to induce the hydrogel reaction. Electrically responsive hydrogels are typically composed of polyelectrolytes, including polycations, polyanions, and polyzwitterions. In the presence of an external electric field, the osmotic pressure changes inside and outside, resulting in the contraction or expansion of hydrogels and deformation. Electrically-responsive SMHs, doped with conductive fillers like carbon nanotubes (conductivity 10–100 S/cm), respond to low fields (0.5–2 V/cm) via Joule heating or ion migration, enabling fast recovery (1–5 s, >10⁶ cycles) and synchronization with bioelectric signals for neural/cardiac interfaces, promoting 80–90 % myogenesis in C2C12 cells [57]. One of its key advantage is the ability to achieve precise control without relying on chemical agents.

In 1992, Osada et al. [58] developed a hydrogel that responds to electrical stimulation, capable of converting chemical energy into mechanical energy, functioning as ‘artificial muscles. The hydrogel is composed of poly (2-acrylamide-2-methylpropanesulfonic acid) (PAMPS). By applying an electric field, positively charged surfactant molecules can selectively bind to the hydrogel network composed of polyanions. The surface of the hydrogel near the anode contracts due to increased binding of surfactants, causing the hydrogel to bend toward the anode. Morales et al. [59] designed and prepared cationic and anionic hydrogel ‘legs’ capable of locomotion. These hydrogel legs were composed of acrylamide (AAm)/sodium acetate (NaAc) and acrylamide/acrylic acid sodium quaternary ammonium salt dimethylamine ethyl methacrylate (DMAEMA-Q) copolymer networks, respectively. The electric field promotes further ion complexation, causing the anionic and cationic legs to interact with each other. By adjusting the hydrogel's bending angle via an electric field, a one-way motion of the bipedal hydrogel on an elastomer substrate was achieved. This system demonstrated movement in aqueous solutions and offers a simple method for manipulating soft materials and robotic systems (as shown in Fig. 7A).

Fig. 7.

Electrically responsive SMHs. A) Mechanism and movement of a hydrogel walker in an electric field, showing direction-dependent actuation based on cationic and anionic leg friction. Reproduced with permission [59]. Copyright 2014,pubs.rsc.org. B) The electroactive shape recovery behavior of XNBR nanocomposites with 15phr CNTS as a function of the action time of a constant 50V voltage. Reproduced with permission [60]. Copyright 2022, mdpi.com. C) When external magnetic fields are applied, magnetic-responsive nanocomposite hydrogels provide on-demand drug delivery. Reproduced with permission [64]. Copyright 2009, Wiley Online Library. D) Design and fabrication of nanocomposites and their proposed mechanism of enhanced drug release in an externally AMF-controlled manner. Reproduced with permission [65]. Copyright 2015, ACS.

Additionally, conductive hydrogels can be fabricated by incorporating conductive fillers, such as metal oxides, carbon black, and multi-walled carbon nanotubes, into the polymer network. For example, Gonzalez et al. [60] combined the elastomers of XNBR Krynac®X 740 with carbon black and multi-walled carbon nanotubes (MWCNTs)in order to enhance the hydrogel's conductivity. As shown in Fig. 7B, the addition of these fillers enhanced the electrical conductivity of the materials from ∼10 to 11 to 10−4 S·cm−1, requiring only a current source of 50 V to achieve the stimulus of fast cleaning, which causes shape memory deformation.

However, prolonged or high-intensity electrical stimulation may induce electrochemical reactions, tissue damage, and inflammatory responses, which limit its feasibility for deep tissue and long-term repair applications. Future strategies, such as flexible microelectrode arrays, self-powered microsystems, or coupling with other stimulation modalities, could enhance the safety and translational potential of electrically responsive SMHs.

2.2.5. Magnetically-responsive SMHs

Magnetically responsive hydrogels, along with the previously mentioned stimulation modalities, offer an attractive form of external stimulation. Magnetic forces can act as remote switches in implanted platforms, triggering on-demand responses. Unlike optical stimulation, magnetic response platforms benefit from excellent tissue penetration and offer non-invasive, biocompatible properties. This allows for the rapid and precise modulation of biomaterial properties with high spatiotemporal control. Magnetic stimulation-responsive hydrogels typically consist of a hydrogel host matrix and magnetic components (such as γ-Fe₂O₃, Fe₃O₄, and CoFe₂O₄) embedded within the matrix, which can respond to an external magnetic field under controlled conditions. Magnetically-responsive SMHs, embedded with Fe₃O₄ nanoparticles (typically 0.2–20 wt% across studies), facilitate remote actuation through magnetic fields (e.g., 50–150 mT static fields), enabling controlled release (e.g., prolonged ion release over >21 days) and enhanced bone regeneration, with reports of up to fourfold increase in new bone formation in rat models [61,62]. Its strong feasibility arises from its ability to penetrate tissues deeply and effectively. Due to its active and rapid response characteristics, it is considered a valuable biomaterial in the biomedical field [63]. Many of them have been utilized for biomedical purposes, like regenerative medicine and tissue engineering.

Magnetic stimulation can induce two different effects depending on the mechanism: (1) the motion of nanoparticles within the hydrogel matrix, guided by magnetic forces, and (2) the generation of thermal stimulation through magnetic sensors under the influence of an electromagnetic field. For example, Qin et al. [64]developed a novel iron gel (SPEL) consisting of superparamagnetic iron oxide nanoparticles (SPIONs) and copolymers of Pluronic F127 (PF127). As shown in Fig. 7C, indomethacin release was accelerated by the influence of a magnetic field, reaching 50 % cumulative release after 25 h. In contrast, in the absence of magnetic stimulation, only 15 % of the drug was released within the same time frame. These findings suggest that when a magnetic field attracts magnetic nanoparticles, they disrupt the initial microstructure of the hydrogel, thereby accelerating drug release through the compression of the hydrogel network. The collective directional movement of the nanoparticles under the magnetic field induces mechanical deformation of the hydrogel, further enhancing drug release. Campbell et al. [65] developed a poly(N-isopropylacrylamide) hydrazine-functionalized SPIONs (superparamagnetic iron oxide nanoparticles) nanocomposite hydrogel by combining it with aldehyde-functionalized glucan, creating an injectable system with alternating magnetic field responsiveness (as shown in Fig. 7D). In this interdependent nanoparticle-hydrogel network, the elastic modulus can be fine-tuned by adjusting the nanoparticle-to-biopolymer ratio, while imine linkages ensure suitable biodegradation properties. It was demonstrated that these high-elastic hybrid constructs were capable of providing controlled release of bupivacaine in vivo, demonstrating their biocompatibility.

However, heterogeneous particle distribution may lead to inconsistent responses, and magnetic nanoparticles may pose risks of cytotoxicity or long-term retention; moreover, integrating magnetic field generators increases system complexity. Future research should focus on optimizing the loading and clearance of biodegradable magnetic nanocarriers, strategies for uniform particle dispersion, synergy with other stimulation modalities, and the development of compact, portable magnetic field devices to enhance safety and clinical feasibility.

The selection of a SMH system for tissue engineering involves balancing mechanical requirements, actuation speed, and practical constraints. For example, a vascular stent might prioritize a fast, strong thermal SMH, whereas a slow-degrading drug-delivery hydrogel might use a pH-sensitive system. The nature of the trigger has practical trade-offs. Remote stimuli, such as light or electric fields, can enable on-demand control; however, they face challenges in vivo. Light (especially visible/UV) has poor tissue penetration, and electrical actuation often requires high voltages or currents to change the shape of a hydrogel. For example, current electro-responsive SMHs typically need high actuation voltages and energy and exhibit low flexibility. By contrast, SMHs triggered by body-ambient stimuli, such as a slight change in pH, the presence of enzymes, or normal body temperature, are more easily deployed clinically. However, these “natural” triggers may act more slowly (e.g., a pH shift in the gut or wound may take minutes) and offer less precise timing. In summary, SMHs actuated by readily available physiological cues trade precision for practicality, whereas those using exotic stimuli (such as light or strong magnets) offer fast, tunable actuation at the expense of penetration and simplicity [94]. The clinical translation of any SMH device must then address immune response, sterilization protocols, and regulatory strategy early in the development process. By carefully matching the SMH class to the intended application and rigorously testing biocompatibility and sterilization, developers can accelerate the path of SMHs from the bench to the bedside.

3. Fabrication techniques for SMH actuators

Over the past few years, SMH has garnered significant attention from researchers in materials science and biomedicine because of its unique structural and functional characteristics. These hydrogels can not only undergo reversible shape change when subjected to external stimuli (such as temperature, pH, light, etc.) but also exhibit excellent flexibility and biocompatibility when restoring their initial shape. Consequently, SMH is highly valuable for use in a variety of applications, including smart drug delivery, flexible electronics, biosensors, and tissue engineering. For smart drug delivery applications, SMHs typically achieve drug-loading efficiencies above 80 % and exhibit release profiles characterized by an initial burst (20–30 % in the first 24 h) followed by sustained release over several days, as demonstrated in magnetic ferrogels with 50 % cumulative indomethacin release under field versus 15 % without stimulation. Moreover, a PVA-based SMH nanoparticle system reported a protein loading efficiency of ∼96 %. In another study, an amorphous polyurethane-based shape-memory hydrogel was developed by crosslinking star-shaped tetrahydroxy oligo[(rac-lactide)-co-glycolide] with an aliphatic diisocyanate. The hydrogel showed reliable shape recovery under physiological conditions and enabled sustained drug release, with ∼90 % of the payload released over 80 days [95]. A thermoresponsive hydrogel (Pluronic F127) loaded into porous NiTi scaffolds was evaluated for vascular tissue engineering applications. By adjusting scaffold porosity, researchers achieved a threefold increase in drug loading, with rapamycin released continuously over 17 days, modulated by the porous architecture [96]. These data illustrate how network chemistry and swelling dictate loading efficiency and release kinetics. To fully exploit the functionality of SMH, different preparation methods have been developed and applied. These methods not only determine the microstructure and macroscopic properties of hydrogels but also directly affect their performance in specific application scenarios.

3.1. 3D and 4D printing technologies

3D printing, as an emerging digital manufacturing technology, benefits from its controllable and accurate computer-aided design (CAD) models, which enable the construction of complex actuators with minimal manufacturing and post-processing time, resulting in reduced material waste and lower production costs. Therefore, 3D printing is currently being used to develop soft actuators for soft robots, such as self-assembling actuators, self-healing actuators, and biomedical actuators, among others [97]. Since 3D soft actuators can be manufactured using 3D printers, they are designed with human-defined geometry, functionality, and control characteristics. Therefore, combining programmable 3D printing technology with stimulus-responsive materials to produce complex and controllable intelligent 3D hydrogel actuators is expected to yield beneficial results. The working principle of 3D printing technology in SMH actuators is mainly based on the concept of additive manufacturing, which builds 3D structures by stacking materials layer by layer. Unlike other synthetic methods, 3D printing technology enables designers to freely create complex geometries and directly transform them into physical objects, making it particularly suitable for applications in the preparation of personalized medical devices and intricate biological structures.

In the 3D printing of SMH actuators, the materials used are typically prepolymers with shape-memory functions. These prepolymers are cross-linked by UV light, heat, or chemical reactions during the printing process to form hydrogel structures with an SME. The printed SMH constructs maintain high shape fixation (>90 % of the intended geometry) after UV or thermal curing and support encapsulation efficiencies of up to 75 %, with sustained cargo release demonstrated over 72 h [98]. Generally, 3D printing methods can be classified as light-based and ink-based, which build structures layer by layer. Light-based 3D printing employs four primary methods: stereolithography (SLA), selective laser sintering (SLS), digital projection lithography (DLP), and two-photon polymerization (TPP). The ink-based methods, on the other hand, can be further classified into inkjet printing, direct ink writing (DIW), and fused deposition modeling (FDM), depending on the state and method of delivering the feed material [99].

Wang et al. [86] synthesized an SMH (dual network) using pluronic diacrylate macromer (F127DA) and sodium alginate (alginate) by using 3D printing. They observed the formation of a stable network through the photocross-linking of pluronic diacrylate macromolecules, as well as the formation of a reversible network through the cross-linking of Ca²⁺ and alginate. As shown in Fig. 8A, when Ca²⁺ was removed from the Na2CO3 solution, the resulting SMH exhibited a recovery rate of 98.15 % within 10 min, despite the elastic modulus remaining essentially stable following the shape memory process. In addition, in vitro studies were conducted on the SMH, and the results indicated that it had a higher drug release rate and good biocompatibility with 3T3 fibroblasts compared with the traditional drug-loaded hydrogel. Therefore, the SMH has the potential to be used as a drug carrier in clinical applications. Kuang et al. [100] developed a novel ink based on urethane diacrylate and linear semicrystalline polymers. As shown in Fig. 8B, this ink enables the 3D printing of UV-assisted, direct ink writing hydrogel networks, exhibiting self-healing, shape-memory behavior, and high stretchability. As a result, a semi-interpenetrating SMH network has the capability of being stretched up to 600 %, demonstrating the ink's significant potential for applications such as soft robotics, vascular repair, and 4D printing. Similarly, Shiblee et al. [101]synthesized thermally responsive SMHs using poly (dimethyl acrylamide-co-stearyl acrylate and/or lauryl acrylate) with a custom optical 3D gel printer. Their results demonstrated that the printed SMH exhibited good fixation and recovery rates under ambient conditions. It demonstrated a Young's modulus of 0.04–17.35 MPa and a strain value of 612–2363 % at high temperatures, with strain values increasing with increasing temperature. Additionally, they noted that all developed SMH formulations exhibited high transparency, varied swelling in both water and organic solvents, and were within the range of intraocular lenses in terms of refractive index. Therefore, these hydrogels have potential applications in optics.

Fig. 8.

A) 3D-printed SMH showing (I–III) molecular mechanism of printing-induced shape memory and (IV) macroscopic recovery. Reproduced with permission [86]. Copyright 2018, Elsevier. B) UV-assisted DIW printing of semi-IPN elastomers: (I) resin components; (II) layer-by-layer deposition with in-situ UV curing; (III) crystal network formation during printing at 70 °C and cooling. Reproduced with permission [100]. Copyright 2018, ACS. C) pH-programmable bilayer hydrogels forming “S”, helix, wave, flower, and bamboo shapes in 0.1 M NaCl, 0.1 M HCl, or neutral media. Reproduced with permission [106].Copyright 2017, pubs.rsc.org. D) CNT-reinforced silk scaffolds (CNTs-SS): (I) compressed T-, I-, and pentagram shapes recover in water; (II) time-lapse recovery sequence; (III) polynomial model describing recovery dynamics; (IV) in vivo shape restoration of a pentagram scaffold after abdominal implantation. Reproduced with permission [108]. Copyright 2020, Elsevier.

After bioprinting, constructs are often cooled or dehydrated to fix their shape. In biological conditions (37 °C, aqueous), unfixed hydrogels risk premature recovery. The challenge is to stabilize the temporary shape in a physiological environment. Recent advances, such as physically crosslinked N-acryloylglycinamide gels, can “lock in” shapes at 37 °C by forming hydrogen bonds during printing. For example, a study showed that a PNAGAm hydrogel could be compressed at 65 °C and remain stable at 37 °C with tunable modulus; different thicknesses (1.0, 0.64, 0.50 mm) led to regions of ∼3.46, 5.94 and 9.46 kPa, which in turn yielded fibroblast adhesion densities of ∼1.3 × 10 ⁴, 1.6 × 10 ⁴ and 1.9 × 10 ⁴ cells/cm ² [102]. This demonstrates how post-printing shape fixation (and mechanical tuning) directly affects function.

4D printing has emerged as a powerful technology for fabricating SMHs with programmable, stimuli-responsive behaviors tailored for tissue regeneration. By incorporating time as the fourth dimension, 4D-printed SMHs can undergo controlled shape transformations in response to physiological cues, such as temperature, pH, or hydration, allowing for minimally invasive implantation and precise adaptation to complex tissue geometries [103]. This dynamic functionality enhances scaffold-tissue integration, supports spatiotemporal control over drug release and cellular responses, and opens new avenues for personalized regenerative therapies. However, limitations such as slow response times, limited mechanical strength for load-bearing applications, complexity in printing multi-material systems, and challenges in long-term biocompatibility and degradation control still hinder the widespread clinical adoption of 4D-printed SMHs in tissue engineering [104].

While 3D and 4D printing offer precise fabrication of customized SMH scaffolds, several practical challenges hinder their widespread application. The inherent rheological complexity of hydrogels can lead to issues such as nozzle clogging and ink spreading, making high-resolution printing difficult. Additionally, most printed hydrogels require post-curing to achieve full crosslinking, which increases both fabrication time and cost. Scaling up to larger constructs presents further challenges, as maintaining uniform crosslink density and consistent pore architecture throughout the scaffold becomes increasingly difficult. Advanced strategies, including multi-material printing and sacrificial support structures, are often needed to ensure structural uniformity and mechanical integrity. Cost-effectiveness also varies significantly depending on the material system, printer technology, and production scale. While 3D printing is relatively established and more economical for prototyping and clinical customization, 4D printing introduces additional complexity and expense due to the need for stimuli-responsive materials and fine-tuned process control. These economic and technical barriers, alongside regulatory challenges, continue to limit the clinical translation of 3D/4D-printed SMHs, despite their promising potential in tissue engineering [105].

3.2. Molding and casting

Molding and casting are the most common and simple methods for fabricating large hydrogel actuators with sizes ranging from centimeters to millimeters. The key lies in using pre-designed molds to precisely control the morphology and size of hydrogels. The method usually involves injecting a prepolymer solution containing monomers, crosslinkers, and initiators into a mold, followed by a polymerization reaction under specific temperature, light, or chemical conditions. With this process, the resulting hydrogels not only have a predetermined geometry but also maintain structural integrity with high accuracy. In addition, this approach also allows for the introduction of multiple functional components, such as nanoparticles or drug molecules, during the synthesis process, thereby endowing hydrogels with more diverse physicochemical properties and enhanced application potential. Due to its superior processing control and material compatibility, this method has become a crucial approach for preparing SMHs for applications in biomedical fields, flexible electronic devices, and smart materials.

For example, Duan et al. [106] fabricated a hydrogel for use in biomimetic lenses by constructing a biological hydrogel actuator using chitosan and carboxymethylcellulose (CMC) in an alkali/urea aqueous solution, with epichlorohydrin (ECH) serving as a crosslinker. This hydrogel is characterized by rapid, reversible, and recurrent self-rolling deformations resulting from pH-dependent swelling and contractions. Furthermore, as shown in Fig. 8C, these hydrogels can also be efficiently designed for their geometry and size by molding processes to form shapes such as rings, tubes, flowers, spirals, bamboo, and waves. It has potential application prospects in a wide range of fields, such as biomedicine and bionic machinery. In addition, Wang et al. [107] developed a shape-memory hydrogel scaffold using silk sericin—a natural protein extracted from fibroin-deficient silkworm cocoons—and functionalized it with carbon nanotubes (CNTs) to enhance mechanical strength and neural activity. This composite hydrogel (CNTs-SS) was fabricated by mold casting into various predefined shapes (e.g., T-shape, I-shape, pentagram), which could be volumetrically compressed and rapidly recovered upon rehydration, demonstrating programmable and reproducible shape-memory behavior. As shown in Fig. 8D, the compressed scaffolds restored their original geometries within 10 s, driven by hydration-induced structural transitions within the sericin matrix. Importantly, due to the intrinsic biocompatibility of silk sericin and the neuro-differentiation-promoting activity of CNTs, CNTs-SS scaffolds not only serve as minimally invasive, patient-specific implants for brain cavity repair after ischemic stroke but also support stem cell survival and promote neuronal differentiation. This dual functionality highlights the potential of molding-based sericin hydrogels as smart biomaterials for tissue engineering, particularly in neural regeneration.

3.3. Electrospinning

Electrospinning technology was developed by Formhals in 1934 to produce fiber MATS by using electrostatic forces to direct a polymer solution or melt jet. This system consists of three fundamental components: a high-voltage power supply, a spinner plate, and a collector plate [109]. Electrospinning is a technique used to produce polymeric fibers with diameters ranging from nanometers to micrometers. It has been widely used to fabricate fibrous materials with diameters ranging from 50 to 500 nm. Changing the fiber diameter from micrometers to nanometers can significantly alter the material's properties [110]. This is mainly reflected in the small aperture and high specific surface area [111]. These characteristics can be used to reinforce fibers in membranes and composite materials. Additionally, it plays a crucial role in cosmetics, drug delivery, enzyme or catalyst carriers, tissue engineering scaffolds, and antibacterial agents.

In recent years, shape-memory materials have been extensively utilized in the development of implants, scaffolds, and other medical devices due to their high biocompatibility and ability to deform in response to external forces [112]. Often, they exhibit high shape recovery and sensitivity, so the challenge is to improve these properties without altering other properties. With the increasing advancements in electrospinning technology, the potential applications of shape-memory electrospinning nanofiber materials in biomedicine have garnered more attention, particularly in the development of tissue engineering, drug delivery, wound dressing, biomedical devices, and medical imaging.

Pandini et al. [113] explored the application of shape memory materials in biomedicine by preparing PCL-based fiber pads by electrospinning combined with sol-gel reactions, varying the degree of sol-gel reactions. A PCL-low and a PCL-high material were obtained (a material with a low and a high degree of crosslinking, respectively). Among them, PCL-high exhibits a greater elongation-contraction effect and also shows higher cell proliferation and spreading ability in biological studies. Electrospinning generates microporous structures very similar to the extracellular matrix, facilitating the culture and proliferation of cells. Consequently, many studies have focused on the development of blood vessel-like scaffolds for cardiovascular disease implants. Wang et al. [114] examined the use of shape memory injectable hydrogels for the delivery of cells. Fig. 9A shows a composite fabricated by embedding electrospun poly(D, L-lactic acid-co-trimethylene carbonate) (P(DLLA-co-TMC)) fibers within gelatin-acrylated β-cyclodextrin (β-CD) polyethylene glycol hydrogel (GCP-hydrogel). The matrix creates a local microenvironment conducive to cell assembly, with embryonic stem cells (ESCs) serving as lubricants. The mean diameter of the electrospun fibers was 800 ± 200 nm, and this remained stable after incorporation into the matrix. Mechanical testing at 37 °C revealed a slight enhancement in the GCP-hydrogel's properties due to the nanofiber mesh. The hydrogel's shape memory characteristics were assessed under incubation at 37 °C, demonstrating that the hydrogel could recover its original shape within approximately 15 s, with a recovery rate (Rr) exceeding 95 %. This composite system significantly enhances the survival rate of ESCs and facilitates their differentiation into motor neurons both in vitro and in vivo. A growing number of stimuli-responsive hydrogels are gaining popularity; however, they face challenges such as poor processability and weak mechanical properties. For instance, fabricating hydrogel networks suitable for melt processing and electrospinning remains a challenging task. As shown in Fig. 9B, Abdullah et al. [115] successfully synthesized a series of melt-processable SMHs composed of polyacrylic acid (PAAc) and 20–50 mol% C18A segments. This was achieved using an organic solvent method and in situ physical cross-linking through hydrophobic interactions. The hydrogels exhibited a reversible transition from strong to weak gelation in the temperature range of 50–60 °C. Additionally, the hydrogel can be dissolved in a chloroform/ethanol mixture to form a viscous solution, which can then be used to produce nanofiber networks through electrospinning. The produced nanofibers significantly improved water absorption, enhanced mechanical properties, and exhibited rapid shape memory recovery properties while maintaining their chemical structure.

Fig. 9.

A) Injectable composite hydrogel for spinal cord injury repair: shape memory P(DLLA-co-TMC) nanonets combined with GCP precursor (gelatin, PEGDA, β-CD) and mESCs, UV-crosslinked and injected into injured spinal tissue to promote regeneration.Reproduced with permission [114]. Copyright 2018, ACS. B) Fusible PAAc hydrogel with C18A unit fabricated via organic solvent synthesis, solvent exchange-induced crosslinking, and electrospinning to form nanofiber networks. Reproduced with permission [115]. Copyright 2024, Wiley Online Library. C) Shape memory behavior of peptide-based P20L4 hydrogel: (I) sequential deformation and recovery; (II) water-triggered shape restoration after calcium chloride fixation. Reproduced with permission [116]. Copyright 2023, Wiley Online Library. D) Schematic representation of self-assembly leading to nanofibrous hydrogel formation and biological evaluation of self-supporting hydrogels. Reproduced with permission [117]. Copyright 2018, ACS.

3.4. Self-assembly

Self-assembly is an advanced method for the preparation of SMHs. Non-covalent intermolecular interactions (such as hydrogen bonding, hydrophobic interactions, π-π stacking, and electrostatic interactions) drive the spontaneous assembly of functional monomers or polymers to form a stable three-dimensional network structure. Under specific environmental conditions, such as changes in temperature, pH, or illumination, these self-assembled nano- or micrometer-scale structures are capable of reversibly deforming, thereby endowing hydrogels with unique shape-memory properties. Compared to the traditional chemical cross-linking method, the self-assembly method not only offers the advantages of mild preparation conditions and simple operation but also allows for the accurate design of hydrogel properties by adjusting the structure and interactions of self-assembly units. Therefore, self-assembly methods have demonstrated significant research value and application potential in the development of smart materials, particularly in biomedical materials and wearable electronic devices (Table 2).

Table 2.

Preparation technique of SMHs and their application.

Method	Materials	Synthesis method	Advantages	Disadvantages	Application	Ref.
3D printing	MAAc, NPAM	Hydrogels are prepared by copolymerizing MAAc and NPAM in DMSO, 3D printed at 70 °C, and then the solvent is replaced with water, which strengthens the hydrogen bonds and improves its mechanical strength and shape memory.	High mechanical strength, excellent shape memory properties, and environmentally friendly	Requires specific temperature control during synthesis	Tissue engineering and environmentally friendly materials	[118]
Molding and casting	PVA	Hydrogel is synthesized by dissolving PVA in water and a crosslinking agent is added to modify the polymer network structure. Upon contact with water, the hydrogel swells and recovers its original shape.	High mechanical properties and flexible shape memory behavior	Dependence on water	Biomedical applications	[119]
3D printing	PAAm, Gelatin	Hydrogels are prepared using a one-step 3D printing process, forming a double network of polyacrylamide (PAAm) and gelatin.	Excellent mechanical properties, high biocompatibility, and suitable for various biomedical applications	Requires precise material control during the printing process	Biomedical devices	[120]
3D printing	NIPAM	Hydrogel is fabricated using projection micro-stereolithography 3D printing. By manipulating the grayscale during the printing process, the deformation of the hydrogel can be finely controlled using near-infrared (NIR) light.	Ultra-fast deformation and tunable responsiveness	Light-based activation requirement and the shape memory and responsiveness are limited to specific light wavelengths (NIR)	Biomedical engineering and microrobots	[121]
Molding and casting	AA, MMA, PDN	Hydrogel was prepared by dissolving different concentrations of the material in a mixture of water/alcohol, adding AIBN as an initiator and MBA as a crosslinker to a flask, and polymerizing the resulting solution by pouring it into a cylindrical glass mold with a diameter of 10 mm.	High mechanical properties and biocompatibility	–	Treatment of coronary artery disease	[122]
Self-assembly	Chitosan, graphene oxide	Hydrogel was physically cross-linked in an alkaline solution by a simple water evaporation-induced self-assembly method to prepare a nanocomposite hydrogel film.	High mechanical properties and flexible shape memory behavior	Biocompatibility was not described	Tissue engineering	[123]
Molding and casting	β-CD, AAm, WPU, MAH, MBA, GO	Hydrogel was prepared by dissolving MAH-β-CD in 2 mL WPU solution, adding GO to the solution and sonicated. Finally, AM and MBA were added, and the mixed solution was poured into the mold for casting.	High mechanical properties	–	Industrial applications	[124]
3D printing	F127DA, PLGA, GO	Hydrogel is fabricated using 3D printing with ultraviolet (UV) light polymerization. A dual network is formed where the first network is chemically crosslinked using F127DA, and the second network is formed by blending PLGA and GO.	High mechanical toughness and biocompatibility	Complex fabrication process involving multiple components and precise control of NIR exposure	Drug carriers and antibacterial scaffolds	[125]
3D printing and cryogelation technologies	SF, LAP	A combination of 3D printing and cryogelation technologies was used to prepare injectable shape-memory hydrogels.	Rapid shape recovery, high mechanical stability, and injectability	Requires a complex cryoprinting setup	Tissue engineering and drug delivery	[126]
3D printing	Sodium alginate, Acrylamide	Hydrogel is synthesized by combining sodium alginate and polyacrylamide. The alginate is crosslinked ionically using calcium ions, while the polyacrylamide network is formed through UV-initiated free-radical polymerization with ammonium persulfate (APS) as the photoinitiator.	High mechanical strength and flexibility	Requires specialized 3D printing equipment (CLIP technology)	Biomedical devices and soft robotics	[127]
Molding and casting	AAm, AA, DMAEMA	Hydrogel is polymerized by dissolving the material and stirring the mixture until evenly dispersed, and then casting the solution into a disc mold.	Thermoplasticity and shape memory behavior	Biomedical	–	[128]
Molding and casting	AN, AAm, AMPS, and PEGDA575	Materials were dissolved in DMSO according to the prescribed formulation. Subsequently, the mixture was poured into a disc mold and polymerized.	High mechanical properties and flexible shape memory behavior	–	Intelligent medical device	[129]
Self-assembly	CHIT, Agar	Hydrogel was prepared by pouring a mixture of CHIT/agar into a silicon mold and vacuum freeze-dried. The prepared porous CHIT/agar hydrogel was then soaked in a 1 M NaOH bath to allow CHIT to self-assemble into an interpenetrating polymer network hydrogel.	Low cost, high shape memory recovery	–	Minimally invasive surgery	[130]
Molding and casting	SS, MBA, MTAC, AA, MMT,	OMMT clay was dispersed in deionized water, and modified cationic soluble starch (CSS), AA solution, and MBA were mixed and quickly transferred into a cylindrical mold for polymerization.	Self-healing, high mechanical properties, and high biocompatibility	–	Biomedical applications	[131]
Molding and casting	PVA, TA	Hydrogel was prepared by dissolving PVA and mixing TA solution, and then, the homogeneous solution or coagulant was transferred to a mold made of two glass plates to obtain PVA-TA hydrogel.	High mechanical properties and SMEs	–	Biomedical applications	[132]

Abbreviations: MAAc, methacrylic acid; NPAM, N-(pyridin-2-yl)acrylamide; PDMAAm, poly(dimethyl acrylamide; SA, stearyl acrylate; LA, lauryl acrylate; PAAm, Polyacrylamide; NIPAM, N-isopropylacrylamide; F127DA, Pluronic F127 diacrylate; PLGA, poly(lactide-co-glycolide); GO, graphene oxide; SF, silk fibroin; LAP, laponite nanoparticles; PVA, Poly(vinyl alcohol); PAN, Polyacrylonitrile; PAA, poly acrylic acid; AN, Acrylonitrile; AAm, acrylamide; AMPS, 2-acrylamido-2-methyl-1-propanesulfonic acid; PEGDA575, polyethylene glycol diacrylate; DMAEMA, 2-(N,N-dimethylamino)ethyl methacrylate; SS, Soluble starch; MBA, N,N′-methylene-bis-acrylamide; MTAC, 2-(methacryloyloxy)ethyltrimethyl ammonium chloride; PDN, polydopamine nanospheres; MMA, methyl methacrylate; P(DLLA-co-TMC), Poly(d,l-lactic acid-co-trimethylene carbonate); β-CD, β-cyclodextrin; PEGDA, poly(ethylene glycol) diacrylate; C18A, N-octadecyl acrylate; CHIT, Chitosan; CC, Cationic chitosan; TA, Tannic Acid.

Xiang et al. [116] developed a smart hydrogel system based on the short peptide MA-FIID, which undergoes spontaneous self-assembly into β-sheet nanofiber networks under acidic conditions through non-covalent interactions such as hydrogen bonding, hydrophobic interactions, and π–π stacking. By incorporating this peptide into a PNIPAM polymer network, they constructed a hybrid hydrogel with excellent mechanical properties and pH responsiveness. Under the Hofmeister effect induced by CaCl₂, the hydrogel exhibited a distinct shape-memory behavior, allowing for temporary shape fixation in salt solution and complete shape recovery in water. As shown in Fig. 9C, the hydrogel maintained a shape recovery rate above 85 % even after multiple cycles. By combining peptide-driven self-assembly, environmental responsiveness, and mechanical adaptability, this hydrogel demonstrates strong potential for biomedical applications, including smart implants, tissue engineering scaffolds, and controlled drug delivery systems. Similarly, Gavel et al. [117] have successfully developed an injectable, self-healing, shape-memory hydrogel based on a 9-anthronmethoxycarbonyl (Amoc)-terminated dipeptide with promising antibacterial properties for biomedical applications. Moreover, characterized by a variety of techniques, the AMOC-terminated dipeptide was able to self-assemble to form a nanofiber network. Other results showed that the SMH exhibits good biocompatibility in addition to effective antibacterial properties, which may contribute to its efficacy in preventing local bacterial infections and the demand for biomaterial implantation (Fig. 9D).

4. SMHs for tissue regeneration applications

4.1. Application of SMH in bone tissue regeneration

In recent years, bone defects resulting from congenital malformations, diseases, accidents, and surgical resections have become a significant clinical challenge [133]. Although bone itself has a remarkable capacity for self-healing, large-scale irregularities in the bone can be very difficult to heal, particularly without medical assistance [134]. Clinical solutions, such as the gold standard in treating large bone defects, are still autologous or allogeneic bone grafts, which can restore the structural and functional integrity of the bone [135]. Even though this treatment is effective, it is limited in part by donor-site complications and the limited availability of bone [136]. The use of tissue engineering has proven to be a promising technique for replacing damaged or defective bone tissue, as it can repair damaged bone by engineering materials, cells, and growth factors (GFs) [137].

Over the past two decades, 3D bioprinting technologies for bone tissue engineering have advanced significantly, with numerous studies utilizing a combination of cells, biomaterials, and bioactive factors to create bone tissue structures and promote bone regeneration through biomimetic structures, thereby enhancing bone regeneration [138]. Recently, shape memory materials (SMMs) have established themselves as advanced materials in various applications, including artificial skin [135], bionic hands, bionic flexible joints [136], muscle tissue [139], and bionic soft tongues [140]. Since these materials are responsive to external stimulation, they are often described as ‘smart' materials. SMHs can change their shape, size, structure, and chemical or physical characteristics in response to various stimuli, including physical stimuli such as water, light, temperature, electric fields, and magnetic fields, as well as chemical stimuli like ion concentration and pH, and biological stimuli like enzymes and glucose [141]. Additionally, they can revert from a temporary shape to their original form. These properties make SMHs highly valuable for numerous biomedical applications, including implantable surgical devices, controlled drug delivery devices, haptic interfaces, artificial muscles, and biosensing devices. As a result of its reversible shape deformation, SMH is able to simulate the dynamics of biological tissues and is particularly useful in bone tissue engineering.

SMHs have gradually become a research hotspot in the field of bone tissue engineering due to their superior mechanical properties, excellent functional properties, and biocompatible properties. Jiang et al. [67] used the preparation of collagen type I (Col) hydrogel scaffold. After freeze-drying, Col scaffolds can be fixed into a temporary shape and then restored to their original shape with the stimulation of water. Furthermore, Col scaffolds enhance chondrocyte adhesion, proliferation, and differentiation in rabbit cartilage defect models, as shown in Fig. 10A. This is attributed to the natural advantages of collagen, which promotes the formation of irregular fibrous tissue at the defect site and exhibits strong cartilage repair capabilities. In addition to shape deformation, tissue engineering scaffolds based on SMH can promote specific biological activities, such as the release of biological molecules.

Fig. 10.

Tissue regeneration using SMHs. A) Cartilage repair in rabbit models: (I) 3 mm deep, 4 mm wide full-thickness defects created in femoral condyles; (II) gross morphology at 6 and 12 weeks post-implantation. Reproduced with permission [67]. Copyright 2018, Elsevier. B) SMH scaffolds for intervertebral disc (IVD) regeneration: (I) peptide-functionalized dual-network hydrogels loaded with bioactive cues; (II) shape recovery in rat tail discs; (III) recruitment and chondrogenic differentiation of endogenous stem cells. Reproduced with permission [142].Copyright 2023, Wiley Online Library. C) Dynamic hydrogen-bonded SMH with TA and KGN enhancing BMSC chondrogenesis: (I) mechanism of action; (II) regenerated hyaline cartilage bridging native tissue in rat femoral condyles. Reproduced with permission [73]. Copyright 2023, Nature. D) Cryoprinted injectable scaffolds: (I) cryogelation process; (II) printed Daping hospital logo; (III) hydrogel recovery and defect conformity after 16 G needle injection. Reproduced with permission [144]. Copyright 2022, Elsevier.

Wang et al. [142] employed physical cross-linking to prepare a cellulose/alginate dual-network hydrogel for the repair of damaged rat intervertebral discs. As shown in Fig. 10B, the compressed hydrogel was immersed in CaCl2 solution to form a temporary memory shape using the complexation of metal ions, and the hydrogel with the temporary shape was subsequently transferred to EDTA solution, which would compete to bind Ca2+ in the gel and destroy the metal complexation to achieve shape recovery. Verified by structural simulation of natural intervertebral disc (IVD), the stent can provide a therapeutic effect on nucleus pulposus (NP) and annulus fibrosus (AF). The DN hydrogel can also be compressed and restored to disc height, demonstrating the feasibility of minimally invasive surgery.

Mesenchymal stem cell (MSC) homing peptide (SKP peptide) and cell adhesion peptide (RGD peptides) were incorporated into cellulose and alginate polymer chains. The SKP peptide promotes the recruitment of endogenous MSCs to the injury site, while the RGD peptides enhance cell survival and adhesion. The constructed intervertebral disc (IVD) scaffolds were implanted into the caudal spine of rats, with EDTA-based shape deformation used to replace the native caudal disc. Furthermore, by mimicking the natural IVD structure and controlling the release of growth differentiation factor-5 (GDF-5), the scaffold induces endogenous MSC differentiation and facilitates the formation of ECM, promoting the synthesis of polysaccharides and proteins. In addition, Yang et al. [73] developed hydrogels based on imidazolidinyl urea (IU), poly(ethylene glycol) (PEG), and methylene diphenyl 4,4-diisocyanate (MDI) through an aggregation addition reaction. These hydrogels were crosslinked via hydrogen bonds to incorporate TA and KGN, forming a scaffold with exceptional mechanical durability and phase-dependent drug release behavior for in vivo cartilage regeneration. The hydrogels exhibit a rapid SME, restoring their original form within 30 s at body temperature. The scaffold was pre-shaped at 4 °C using a syringe to facilitate implantation. Upon exposure to physiological conditions, it regained its shape, filling in the defect site and promoting cartilage regeneration. The continuous release of TA and KGN induces the migration of bone marrow stem cells (BMSCs) into the scaffold, promoting chondrocyte differentiation and facilitating full-thickness cartilage regeneration in vivo (Fig. 10C).

In recent years, biomaterials prepared by 3D printing technology have also been widely used in tissue engineering. For example, Wang et al. [126] utilized 3D printing and cryoprinting technology to fabricate nanoparticle-enhanced injectable shape memory cryogels. As shown in Fig. 10D, silk fibroin-based cryogels, after post-cryoprinting, can be printed in various shapes and sizes and exhibit rapid volumetric recovery after being injected through a needle. These cryogels, which encapsulated laponite nanoparticles, also promoted enhanced proliferation, spreading, and osteogenic differentiation of seeded bone marrow-derived mesenchymal stem cells (BMSCs). Additionally, the implanted hydrogels fused with host tissue following subcutaneous injection, allowing for cellular infiltration without triggering significant inflammatory responses. These findings demonstrate that the printed nanoparticle-enhanced cryogel can actively support the participation of implanted stem cells in the bone regeneration process.

Despite the great potential of bioprinting technology to fabricate highly complex biological structures, the most significant challenge at present is printing hollow tubular structures. Alina Kirillova et al. [143] developed an advanced four-dimensional biomanufacturing approach, utilizing deformable biopolymer hydrogels to fabricate hollow self-folding tubes with unprecedented control over their diameter and structure at high resolution. They synthesized methacrylated alginate (AA-MA) and methacrylated hyaluronic acid (HA-MA) memory hydrogels, which were cross-linked by calcium ions and alginate. Self-folding was induced in microtubules, and when EDTA was added to chelate the calcium ions, the tubes returned to their original elongated shape. The tubes were loaded with mouse bone marrow stromal cells (BMSCs), which can differentiate into various cell types, including osteoblasts, chondrocytes, and adipocytes, thereby promoting bone regeneration.

4.2. Application of SMH in soft tissue regeneration

Currently, the treatment of soft tissue defects (such as muscle and skin) caused by various pathologies and trauma mainly includes autologous transplantation and commercially available filling materials. However, there are many challenges and limitations associated with these treatments, such as donor-site morbidity and volume loss over time [145]. Therefore, improved therapies are needed. In recent years, tissue engineering technology has provided a new solution to these problems by developing bioactive biological materials. Bioactive materials enhance the physical stability of transplanted cells, as well as the ability of stem cells to multiply and differentiate rapidly, thereby restoring lost tissue function [142]. Therapeutic effects may be enhanced through the use of biomaterials that inhibit the overactive immune response at the local level [146]. Furthermore, the properties of biomaterials for tissue regeneration may differ depending on the type of tissue damaged. Furthermore, immunosuppression plays a significant role in the regeneration and repair of tissues [147].

SMHs, as smart materials, can sense and respond to external stimuli by changing their volume or shape. These hydrogels convert chemical or physical energy into mechanical energy, exhibiting volume changes such as swelling when absorbing water and shrinking upon dehydration. This ability to mimic biological functions at the molecular level makes them ideal for applications like artificial muscles and simulating human muscle function in clinical settings. Based on the mixed-matrix membrane strategy, Yu et al. [148] designed a novel light-induced shape memory polymer (SMP) artificial muscle using an opto-mechanical molecular crystal combined with a connecting polymer such as PVDF, as shown in Fig. 11A. This novel artificial muscle can lift heavy objects and bend to grasp objects after light exposure. This hybrid system combines rapid photoreactivity with the mechanical performance of an elastomer with a high Young's modulus, enhancing functionality. This artificial muscle system is capable of performing various functions, including grasping objects and lifting heavy loads. Furthermore, Hua et al. [41] proposed a novel poly(acrylic acid) (PAA)/calcium acetate (Ca(CH₃COO)₂) SMH with cold-induced shape recovery properties and strong artificial muscle performance. Due to the reduced solubility of the acetic acid group at high temperatures, the polymer chains aggregate, resulting in the hydrogel network hardening. This allows the hydrogel to be fixed into a temporary shape upon heating, with continuous shape recovery achievable in cold environments by adjusting the shape fixation time. Additionally, the synthesized PAA/Ca(CH₃COO)₂ hydrogel exhibits an ultra-high work density of up to 45.2 kJ m⁻³, far surpassing the work density of typical animal muscles (∼8 kJ m⁻³).

Fig. 11.

SMHs for artificial muscles and skin applications. (A) Light-driven artificial muscle: (I–III) movement under light irradiation; (IV) definition of displacement (D); (V) displacement vs. irradiation time under unilateral/opposite light; (VI–IX) multi-armed hydrogel grasping objects. Reproduced with permission [148]. Copyright 2018, Wiley Online Library. B) Tendril-like artificial muscle from AA/AAm/Ad-Am terpolymer (I) FeCl₃-induced locking; (II) chiral deformation under water spray. Reproduced with permission [71]. Copyright 2021, ACS. C): UV-shielding PVA/TP hydrogel dressing: (I) preparation; (II) photos showing skin protection in mice before and after UV exposure; H&E staining comparing PVA/TP, PVA, and untreated skin. Reproduced with permission [69].Copyright 2020, pubs.rsc.org. D) MXene-based SMH: (I) synthesis; (II) schematic and wound healing outcomes with NIR-assisted skin repair. Reproduced with permission [151]. Copyright 2024, pubs.rsc.org.

Similarly, Cui et al. [71] designed and fabricated tendril-like hydrogel artificial muscles using a continuous shaping method. They used hydrogels based on β-cyclodextrin-modified tunicate cellulose nanocrystals (β-CD-TCNCs) by polymerizing acrylamide (AAm), acrylic acid (AA), and adamantyl-acrylamide (Ad-Am) monomers. Tunicate cellulose nanocrystals were integrated into the polymer network via Fe³⁺/−COO^- ion coordination, followed by host-guest interactions, and programmed into various shapes (stretched, twisted, and coiled) to enhance the performance of the hydrogels. These hydrogel muscles exhibit high actuation rates, dynamic strains, and shape memory properties in response to various solvents. With water content and contractile work capacity comparable to native muscles, they show significant potential for biomedical applications (as shown in Fig. 11B). In addition, Zhang et al. [149] synthesized thermoresponsive hydrogels (TRHs), which function as smart materials capable of deforming their volumetric shape in response to temperature changes when hydrated. Notably, compared to other stimulus-responsive hydrogels, TRHs exhibit faster deformation when exposed to thermal stimulation in water, enabling rapid displacement or movement. This makes TRHs highly effective for applications as soft polymer actuators or artificial muscles.

In addition to their important clinical role in simulating artificial muscles, SMHs hold great potential in the treatment of skin wounds. Human skin serves as the body's first line of defense against the external environment, making it highly susceptible to environmental damage. Skin wounds can be categorized into acute wounds (resulting from mechanical damage, chemical injury, or surgical wounds) and chronic wounds (caused by infections, burns, or diabetes, among others) [150]. Various wound dressings have been developed to manage skin wounds, including films, gauzes, foams, hydrocolloids, nanofibers, and hydrogels. In recent years, a variety of smart hydrogels have garnered significant attention because of their ability to mimic the structure of natural ECM, adjustable mechanical properties, and easy delivery of bioactive substances. It has shown good potential for skin wound repair.

Moreover, ultraviolet (UV) radiation poses a threat to human health and can lead to various diseases. Therefore, designing biomaterials with UV-shielding capabilities holds significant potential for applications in environments where UV protection is needed. For instance, Xiong et al. [69] designed a functional polyvinyl alcohol (PVA)/tea polyphenol (TP) SMH with excellent UV shielding ability, and the study showed that shape memory was caused by hydrogen bonds between the two polymers. As shown in Fig. 11C, they irradiated mice with 10 mW/cm² UV light for 10 min and found that the adipose tissue and muscle fibroblasts in the epidermis of PVA/TP hydrogel-covered mouse skin remained intact and comparable to the skin of normal mice. However, the skin covered by PVA hydrogel showed obvious damage to collagen fibers, which was mainly due to the strong UV absorption peak of TP at 270–280 nm, indicating better UV resistance. Therefore, this PVA/TP hydrogel can be used as a biomedical material for human skin protection.

In addition, SMH can also be used as a wound dressing to assist in achieving hemostasis and provide temporary protection following a skin injury. For example, Hu et al. [151] prepared MXene-based wound dressing hydrogels for skin repair using MXene, poly(ethylene glycol)diacrylate (PEGDA), gelatin, and N, N′-methylenebis (acrylamide) (HEAA) as raw materials. As shown in Fig. 11D, the hydrogel was cross-linked with HEAA and PEGDA to form the first layer of the hydrogel network. The hydrogen bonds between MXene, PHEAA, and gelatin form the second layer of the hydrogel network. The resulting hydrogel recovered well after immersion in water for 16 s or exposure to light for a few minutes and exhibited good photothermal properties when subjected to 808 nm and 1 W cm−2 light for 20 s, with a surface temperature of 86.4 °C. Regarding wound healing, the MXene-based hydrogel exhibited excellent photothermal properties, indicating that it could be utilized as a wound dressing for skin repair.

Similarly, Cao et al. [152] prepared a shape memory cryogel using chitosan (CS) and citric acid (CA) at low temperatures. The synthesized CS/CA/Ag cryogels exhibit good mechanical properties and an interconnected macroporous structure. It has an excellent hemostatic effect and promotes blood cell adhesion. Additionally, the CS/CA/Ag frozen gel has been demonstrated to significantly enhance the healing of full-thickness skin wounds infected with Staphylococcus aureus. This makes it an effective biomaterial for rapidly stopping bleeding and accelerating wound healing. Some chronic wounds have increased in recent years due to their irregular wound infection coefficient. SMH can adapt to irregular wound shapes and reduce the risk of wound infection through shape memory deformation in response to external stimuli. AU Vakil et al. [153] developed a cytocompatible shape memory polyurethane-based \ PEG hydrogel, which can be easily delivered to the wound site. Plant-based phenolic acids were physically integrated into hydrogel scaffolds to provide antibacterial properties. The synthesis of hydrogels with excellent shape memory, cell compatibility, and antibacterial properties makes them ideal candidates for wound dressings.

4.3. Application of SMH in vascular tissue regeneration

Cardiovascular diseases, particularly coronary artery disease (CAD), remain a leading cause of mortality worldwide. Coronary artery bypass grafting (CABG) is the gold standard for treatment, traditionally relying on autologous blood vessels. However, graft availability is often limited by prior harvesting, donor-site complications, or disease progression, necessitating the use of alternative vascular grafts. An ideal vascular graft should be durable, biocompatible, non-toxic, non-immunogenic, and antithrombotic while supporting host tissue remodeling post-transplantation.

Currently, there are three primary methods for producing bioengineered vascular grafts: (1) acellular matrices, (2) unit piece engineering, and (3) natural or synthetic polymer-based biodegradable scaffolds [154]. Among these, SMH shows potential as a novel vascular embolic agent for the treatment of certain cardiovascular diseases, owing to its excellent deformability and biocompatibility.

For example, Zhang et al. [91] prepared an opaque and high-hardness temperature-responsive SMH by introducing a reversible hydrophobic dipole pair microdomain into a flexible cross-linking network, followed by BaSO₄ precipitation. Due to its thermally responsive mechanical properties, the rigid radiopaque hydrogel strip can easily enter the porcine renal artery through the catheter under the protection of cold saline and spontaneously transform into a microcoil-like structure after contact with blood. As shown in Fig. 12A, in vitro experiments showed that the material had good blood compatibility. It was found that the hydrogel could successfully cause targeted renal artery embolization. Three months post-operation, the rats’ kidneys exhibited significant atrophy and were notably smaller than the contralateral healthy kidneys.

Fig. 12.

SMHs for embolization therapy. (A) In vivo SME of radiopaque hydrogel coils for transarterial embolization (TAE): (I–II) schematic of TAE and coil delivery; (III) angiographic images post-embolization; (IV) kidney morphology at 4, 8, and 12 weeks. Scale bar: 2 cm. Reproduced with permission [155]. Copyright 2018, Wiley Online Library. B) Canine carotid aneurysm model: (I–II) schematic and surgical procedure; (III) H&E staining with red arrows indicating hydrogel and blue arrows indicating silk suture; (IV) Masson trichrome images of artery and aneurysm sections. Reproduced with permission [92]. Copyright 2020, Wiley Online Library. C) Poly(PEA-co-AAm) SMH embolic plug: (I) network synthesis; (II) custom flow system mimicking vascular conditions; (III) progressive plug expansion from linear to cylindrical shape; (IV) flow rate profile during occlusion. Reproduced with permission [68]. Copyright 2019, ACS.

Similarly, Liu et al. [92] fabricated a temperature/pH dual-responsive shape memory hydrogel with self-adjusting stiffness and non-radiation by copolymerizing acrylonitrile (AN, dipole-dipolar interacting monomer), N-acryloyl 2-glycine (ACG, pH-sensitive H-bond monomer), and polyethylene glycol diacrylate (PEGDA). As shown in Fig. 12B, the hydrogel utilizes supramolecular PACG hydrogen bonds to the cyano dipole-dipole pair, which contributes to the temperature-triggered SME and has excellent mechanical properties. To test the embolic effect of this shape-memory hydrogel, they established carotid artery aneurysms in dogs. At 37 °C, the high stiffness of the hydrogel ensured its smooth delivery through the catheter. After being transported into the aneurysm sac, the stiffness decreases, and secondary swelling occurs in contact with neutral blood, which increases packing density and reduces the recanalization rate and delivery times.

In addition to its excellent vascular embolic effects in in vivo studies, SMH also performed well in in vitro simulation experiments. For example, Liang et al. [68] developed a simple and scalable method to fabricate highly resilient hydrogels with temperature-responsive SMEs based on synergistic hydrophobic interactions and hydrogen bonding. As shown in Fig. 12C, they selected 2-phenoxyethyl acrylate (PEA) and acrylamide (AAm) as hydrophobic and hydrophilic monomers, respectively, to prepare poly(PEA-co-AAm) hydrogels. The prepared hydrogels exhibit good stretching ability and possess a strong temperature-responsive shape memory, primarily generated by the synergistic biphasic cross-linking resulting from hydrophobic interactions and hydrogen bonds. Additionally, the hydrogel was applied in transcatheter arterial embolization (TAE). First, the cylindrical hydrogel was stretched into a linear shape, and then it was delivered to a silicon tube with a guide wire to simulate blood vessels under the protection of 10 °C water. Resulting in complete occlusion of the tube within 20 s, it exhibits good embolization performance, thereby simulating vascular occlusion for the treatment of certain cardiovascular tumors and other diseases.

4.4. Application of SMH in neural tissues

Peripheral nerve injury is a common clinical complication of traumatic injury caused by accident, tumor growth, or surgical side effects, often resulting in long-term disability and loss of motor/or sensory function, and these injuries negatively affect the patient's quality of life. Currently, autologous nerve transplantation is considered the gold standard for treating peripheral nerve injuries. In recent years, many researchers have been committed to developing and innovating new methods of nerve tissue engineering, aiming to explore alternative strategies for autologous nerve transplantation. However, the reformation of peripheral nerve defects remains challenging [152]. Implanting biomaterials that can provide structural and functional support is thought to favor the growth of regenerative neuronal networks. For example, Wang et al. [108] prepared photoluminescent, injectable carbon nanotubes-doped sericin scaffolds (CNTs-SS) having programmable shape memory properties, which can be used as a support material to treat severe ischemic stroke that damages neuronal tissue and forms irregularly shaped stroke cavities. By adjusting the concentration of CNTs, the shape recovery ability of CNTS-SS can be precisely controlled in seconds, allowing for the precise tailoring of the shape to fit any irregularly shaped cavity. Through the use of a preclinical stroke model, they successfully injected custom-shaped CNTs-SS into the cavity and recovered its predesigned shape so that it fits well into the cavity. These results indicate that CNTs-SS, when used as a cell carrier, are capable not only of delivering bone marrow mesenchymal stem cells (BMSCs) into brain tissue, but also of promoting their neuronal differentiation, thus providing a personalized treatment for stroke (Fig. 13A).

Fig. 13.

SMHs for neural tissue engineering and stimulation (A) CNTs-SS scaffold for stroke recovery: (I) structural design; (II) brain cavity filling to support regeneration; (III) micro-CT images showing injected CNTs-SS within stroke lesions. Reproduced with permission [158]. Copyright 2021, Elsevier. B). DCNI preparation and application: (I–II) fabrication process; (III) in vivo nerve stimulation and recording, with EMG and ENG signals and corresponding amplitude data. Reproduced with permission [156]. Copyright 2021, mdpi.com. C) Nerve regeneration in sciatic nerve defect model: (I–II) PC12 proliferation and differentiation in conduits; (III–IV) implantation and tissue appearance at 12 weeks; (V) NCV and DCMAP values at 6 weeks. Reproduced with permission [157]. Copyright 2020, ACS. D) MOSD for C7 nerve stimulation: (I–II) implantation setup and device placement; (III) limb movements triggered by varying mini-LED intensities, with joint positions marked before and after stimulation [89]. Reproduced with permission. Copyright 2019, nature.com.

Currently, the use of peripheral nerve stimulation and autonomic nerve stimulation to treat chronic diseases and various diseases has proven to be challenging. Therefore, there is a need for a stable and biocompatible neural interface suitable for implanting into small nerves. In order to construct the double-clip neural interface (DCNI), Cho et al. [156] utilized the thermally responsive SMH of thiolene/acrylate. This micropatterning was designed for implantation into a branch of the sciatic nerve and the pelvic parasympathetic nerve by utilizing a SME at body temperature (as shown in Fig. 13B). In addition, they implanted iridium oxide (IrO₂)-coated neural interfaces into the common peroneal nerve in order to conduct electrical stimulation and to record electroneurography (ENG). Based on the study's results, it can be concluded that the neural interface is capable of regulating peripheral nerves, including autonomic nerves, and plays a crucial regulatory role in bioelectronic medicine.

A trauma to the central nervous system may result in neurons losing their function. Consequently, this can negatively impact the normal functioning of the human body. Neural progenitor cell transplantation is considered to be an effective treatment strategy. Wang et al. [114] developed a composite hydrogel system that contains a modified gelatin matrix and shape memory polymer fibers. A fine needle is used to inject the gelatin matrix through a catheter, creating a microenvironment for cell assembly and acting as a lubricant. By using embryonic stem cells (ESCs) derived motor neurons as a minimally invasive treatment for spinal cord injury (SCI), researchers demonstrated that injection of ESC-loaded composite hydrogel into SCI animals significantly increased tissue regeneration and motor function recovery, as well as greatly improved the viability and differentiation of ESCs into motor neurons. Furthermore, it provides a reliable approach to treating central nervous system disorders. Furthermore, Wang et al. [157] designed a self-forming multichannel nerve guidance conduit with topographical cues, based on the degradable shape memory polymer poly (lactide-co-trimethylene carbonate) (PLATMC). The electrospun shape memory nanofiber mat initially takes a tubular form through a high-temperature molding process and can be temporarily flattened to facilitate cell loading, allowing for the uniform distribution of cells. When triggered by a physical temperature around 37 °C, the conduit automatically reverts to its permanent tubular shape, forming a multichannel structure. As shown in Fig. 13C, this multichannel conduit demonstrated promising results in promoting PC12 cell growth and repairing rat sciatic nerve defects. These findings suggest that shape-memory polymer-based self-forming nerve conduits have significant potential for peripheral nerve regeneration therapies. Interestingly, researchers have also developed an implantable multisite optogenetic stimulation device (MOSD) based on shape-memory polymers for neural stimulation. As shown in Fig. 13D, the MOSD device induces precise ankle joint extension or flexion movements through dual-site stimulation of the sciatic nerve bundles. To stimulate the sciatic nerve bundles selectively, a thermoresponsive MOSD based on shape-memory polymers was implanted into the anastomosed and severed nerves of a mouse in order to restore selective movement of the limbs [89].

In addition to the chemical-based synthesis of SMH for neurological therapy, four-dimensional (4D) bioprinting is an emerging biomanufacturing technology that combines time as the fourth dimension with 3D bioprinting for the fabrication of customizable tissue engineering implants. 4D bioprinted implants can be designed to possess self-healing and shape-memory properties, thereby offering a wide range of clinical applications. Wu et al. [93] developed a self-healing SMH consisting of biodegradable polyurethane (PU) nanoparticles and photo/thermal responsive gelatin-based biomaterials. The printed hydrogel can undergo multiple deformation cycles and exhibit good shape memory programming, demonstrating good shape fixity and a high shape recovery rate. Furthermore, isolated bioprinted neural stem cells (NSCs) and mesenchymal stem cells (MSCs) showed mutual migration in adjacent self-healing fibers, and this interaction promoted neural differentiation behavior.

Moreover, SMHs can be engineered with elastic moduli spanning those of many body tissues. For example, human articular cartilage has a Young's modulus on the order of 0.5–10 MPa, whereas cortical bone is orders of magnitude stiffer (∼15–25 GPa) [159,160]. Vascular smooth muscle (artery wall) has intermediate stiffness (∼0.2–2 MPa), and brain/nerve tissue is very soft (∼0.03–12 kPa) [44]. By contrast, SMHs reported in the literature cover kPa to MPa ranges: some are as soft as a few kPa (e.g., an RGD‐functionalized PNAGAm SMH with E ≈ 3.5–9.5 kPa), while “tough” SMHs reach several MPa (e.g., a dual-network UPy/Fe³⁺ hydrogel with E ≈ 6.9 MPa and tensile strength ∼7.9 MPa) [102,161]. In general, SMHs can mimic soft tissues and (with reinforcement) approach cartilage stiffness, but remain far below bone rigidity. Their toughness can also be high (e.g. ∼29 MJ/m³ vs. ∼2–7 kJ/m² for bone), reflecting energy-dissipative networks. These data show that SMHs can be tailored to meet the modulus of target tissues: soft SMHs for brain/vascular (<0.1–1 MPa) and stiffer designs for cartilage/muscle (∼1–10 MPa).

Most hydrogels are intrinsically weaker than hard tissues, so SMHs are generally better suited to soft‐tissue or moderate‐load applications (e.g., cartilage rather than bone). However, emerging SMHs show remarkable durability under repeated loading. For instance, a hydrogen-bonded SMH scaffold designed for cartilage repair withstood ∼28,000 compressive load–unload cycles at body temperature while fully recovering its shape. Such “ultra‐durable” SMHs (loaded with bioactive drugs) also enabled full-thickness cartilage regeneration in rat defects [162]. In another example, a physically crosslinked SMH demonstrated high toughness (up to ∼29 MJ/m³) and self-recovery. These metrics indicate that some SMHs can approach the energy-dissipation and cyclic-load performance of natural tissues, such as cartilage or tendon. Nonetheless, even these tough SMHs (modulus <10 MPa) remain far below the stiffness of bone (10–20 GPa), so long-term load-bearing in skeletal applications remains challenging. In summary, SMHs can be tuned to match or exceed the soft to mid-range tissue moduli, and specialized designs (such as multi-network or composite gels) can sustain many cycles; however, replicating bone-level rigidity still requires further engineering.

Importantly, several SMHs have been validated in animal models, supporting their clinical relevance. For example, one study implanted a multi–hydrogen-bond crosslinked SMH (loaded with tannic acid and kartogenin) into a rat knee cartilage defect; the scaffold not only endured ∼28,000 mechanical cycles, but also promoted full-thickness cartilage regeneration in vivo. Histology revealed no chronic inflammation, and subcutaneous implantation tests showed a minimal adverse response. After several weeks, the hydrogel had partially degraded, and surrounding tissue ingrowth had occurred, with no significant organ toxicity or inflammatory cell accumulation. Likewise, other SMH implants (e.g., neural stimulators, vascular catheters, anti-adhesion barriers) have been performed safely in rodents or dogs [163]. These in vivo results, including the maintenance of scaffold shape, biodegradation rate, and tissue integration, demonstrate that SMHs can be biocompatible and functional under physiological conditions. In summary, SMHs exhibit mechanical properties comparable to those of various tissues and have withstood physiological loads in vivo without damage or toxicity, underscoring their promise for regenerative applications. Nevertheless, extensive in vivo investigations under realistic cyclic loading conditions remain essential to demonstrate their clinical suitability conclusively.

Despite demonstrating promising initial mechanical performance, the long-term durability of SMHs under physiological loading conditions remains contingent upon factors such as the stability of cross-links, the rate of hydrolytic or enzymatic degradation, and responsiveness to the physiological environment. Although several in vitro studies have reported sustained mechanical integrity of SMHs for weeks to months in simulated body fluid, comprehensive and rigorous long-term in vivo assessments are currently limited. Therefore, future research should prioritize extensive fatigue testing, detailed characterization of mechanical properties under cyclic loading, and robust long-term in vivo evaluation. Addressing these critical knowledge gaps will facilitate the clinical translation and practical application of SMHs in orthopaedic and other load-bearing biomedical scenarios [164]. The specific application of SMH in tissue engineering is presented in Table 3.

Table 3.

Studies on the utilization of SMH in tissue engineering and drug delivery.

Material	Load	Stimulus	Explanation	Therapeutic benefit	Application	Ref.
PAA	DETA	Temperature	Physical crosslinking enables rapid self-healing under ambient conditions and thermal responsiveness, thereby driving the shape-memory behavior of the hydrogel.	Self-healing, structural support	Artificial muscle and skin	[165]
PEG	Gyrase B, coumermycin, and novobiocin	–	Stimulus-responsive eight-arm PEG-based hydrogel was developed for soft tissue regeneration. In vitro tests showed some toxicity from novobiocin and coumermycin, while in vivo studies revealed no signs of inflammation.	Drug delivery, tissue regeneration	Soft tissue healing and regeneration	[166]
Gelatin	BMP-2	Water	Gelatin is chemically modified with heparin, enabling specific interaction with BMP-2 and heat-induced phase separation. The resulting SMH is used for bone regeneration in sinus augmentation.	Bone regeneration	Bone regeneration	[167]
SPA and POSS	CNS and l-arginine	Water	Strong ionic interactions enable rapid gelation. Non-covalent crosslinking supports fast thixotropic behavior and shape memory function.	Anti-biofouling, Self-healing, Controlled drug release	Wound healing and drug delivery	[168]
DMA and DMAA	BIS	Water and DMSO	DMSO disrupts the DMA association, allowing for reversible shape memory and actuation properties in the hydrogel. The rapid shape fixation of the hydrogel in water and complete shape recovery in DMSO were rapidly achieved in a short time.	Self-healing, Bioinspired Actuating	Soft actuators	[169]
GelMA	–	Temperature	Synthesized memory material has a large pore size, which can effectively deliver multiple cells to promote proliferation and adhesion, thereby realizing tissue regeneration function.	Enhanced osteogenic differentiation	Tissue regeneration	[170]
NIPAm, SA and AA	–	pH/temperature	Repulsion between SA or AA groups caused by protonation or deprotonation of the acrylate or amine groups, respectively, leads to changes in the diameter and wall thickness of the hydrogel.	Cellular scaffolding	Tissue engineering	[171]
DMAA and Stearyl acrylate		Temperature	Synthesized hydrogel exhibited a SME in both hot water and hot air, and SA was the most likely responsible for the SME.	Drug delivery, tissue engineering scaffolds	Artificial lenses for tissue engineering	[172]
P(DA-AAm)	5-FU and BTZ	–	Synthesized hydrogels elucidated the drug-loading ability of chemothermal seeds in dual-drug chemotherapy while generating dual-drug synergy.	Dual drug delivery, targeted treatment	Drug delivery	[173]
ST and PSBMA	–	–	Constructed SMH provided good protein resistance properties due to the presence of starch in the backbone, the overhanging fragment of PSBMA, and the complete degradation of the contour hydrogel in vivo. Both in vitro and in vivo studies have demonstrated its applicability in biomedical functions.	Protein resistance	Drug delivery	[174]
Polyethylene glycol derivative and α, β-polyaspartylhydra zide	DOX	pH	SEM revealed that the prepared SMH possessed a porous three-dimensional microstructure and exhibited a more pronounced inhibitory effect on tumor growth following drug delivery (DOX).	Drug delivery	Local chemotherapy of human fibrosarcoma	[175]
Collagen	–	Water	A 3D-printed hydrogel scaffold exhibits good biocompatibility and promotes the proliferation, adhesion, and redifferentiation of chondrocytes, thereby contributing to bone regeneration.	Enhanced chondrocyte interaction	Bone regeneration	[176]
N, N-dimethyl acrylamide	Stearyl acrylate	–	By fabricating transparent shape memory gels (T-SMG) through free-form manufacturing, the internal structure of T-SMG changes due to variations in crystal components, which in turn affect its mechanical properties and swelling behavior. The addition of UV absorbers in 3D printing successfully enhanced the shaping accuracy of T-SMG.	Drug delivery, tissue regeneration	Optical devices and robot joints	[177]
N, N-dimethylacrylamide and stearyl acrylate	Karenz-MOI	Temperature	Hydrogel exhibits a unique inter-crosslinking network (ICN) with shape-memory properties. It exhibits different mechanical behaviors at varying temperatures—hard like plastic at low temperatures and capable of returning to its original shape at high temperatures.	Mechanical toughness, high ductility	Medical treatments (plasters or bandages)	[178]
poly(N,N-dimethylacrylamide)	Eu–iminodiacetate	pH, temperature, metal ions, sonication, and force	EU-containing polymer hydrogels exhibit responsiveness to multiple stimuli and possess fast self-healing and tunable color-changing properties.	Drug delivery, targeted treatment	Smart optical material for biological sensors	[179]
Gelatin- PEGDA	ESCs	–	An injectable composite hydrogel system, based on a modified gelatin matrix, could provide a promising solution for cell delivery in vivo and can be easily extended to other stem cell-based regenerative therapies.	Cell delivery, tissue regeneration	Treatment of spinal cord injuries

Abbreviations: PAA, poly (acrylic acid); DETA, diethylenetriamine; SPA, sodium polyacrylate; POSS, polyhedral oligomeric silsesquioxane; CNS, clay nanosheets; DMA, dopamine methacrylamide; DMAA, N, N-dimethyl acrylamide; BIS, N, N′-methylenebisacrylamide; SA, Sodium Acrylate; AA, Allyl Amine; 5-FU, 5-Fluorouracil; BTZ, bortezomib; ST, starch; PSBMA, poly(sulfobetaine methacrylate); DOX, doxorubicin; PLGA, poly(lactide-co-glycolide); GO, graphene oxide; F127DA, F127 diacrylate; GelMA, methacrylated gelatin; HA, Hyaluronic acid; SF, silk fibroin; LAP, laponite nanoparticles; PVA, polyvinyl alcohol; TP, tea polyphenol; PEGDA, poly(ethylene glycol)diacrylate; HEAA, poly(ethylene glycol)diacrylate; PAN, polyacrylonitrile; PACG, N-acryloyl 2-glycine; PEGDA, polyethylene glycol diacrylate; PEA, 2-Phenoxyethyl acrylate; PCL, polycaprolactone; PDLLA, poly(D, L-lactide); CNTs, carbon-nanotubes; ESCs, embryonic stem cells.

5. Challenges of SMHs

5.1. Biocompatibility and immune response

SMHs are promising materials in tissue engineering due to their dynamic deformation, ECM-mimetic water content, and tunable mechanics that enable minimally invasive implantation and in situ conformability. However, clinical translation depends critically on their biocompatibility, degradation profile, and immunological safety. Despite their construction from inherently biocompatible polymers such as PEG, gelatin, chitosan, alginate, or hyaluronic acid, SMHs must be rigorously evaluated to ensure that degradation byproducts, crosslinkers, or embedded stimuli-responsive elements (e.g., photothermal dyes or magnetic nanoparticles) do not induce local or systemic toxicity [180]. Additionally, biodegradation is a crucial aspect of scaffold integration. Ideally, the degradation rate of SMHs should synchronize with tissue regeneration to avoid premature loss of function or long-term foreign-body presence. Hydrogels incorporating ester or amide linkages undergo hydrolytic or enzymatic cleavage into biocompatible monomers, often cleared via natural metabolic pathways. However, improper degradation profiles, such as rapid hydrolysis from PCL-based components, may produce acidic byproducts, causing local pH drops and inflammation. Conversely, slow-degrading SMHs risk fibrotic encapsulation, impaired nutrient diffusion, and hindered cell migration [181].

In vitro studies consistently demonstrate high cytocompatibility of SMHs across cell types relevant to tissue engineering. For example, thermoresponsive alginate/tannic acid SMHs maintained >90 % bone marrow stromal cell (BMSC) viability with negligible hemolysis, while acrylated GelMA SMHs exhibited no toxic leachates and supported fibroblast adhesion in both 2D and 3D cultures [182]. PEG-based SMHs implanted subcutaneously in rodents exhibited minimal fibrosis and promoted vascularization, indicating a favorable host interaction. In cartilage repair models, shape-memory scaffolds containing kartogenin or hyaluronic acid promoted cell adhesion and proliferation, while maintaining mechanical integrity during physiological deformation. Moreover, protein-based bilayer SMHs utilizing tandem modular elastomeric proteins, such as (GB1)8 and (FL)8, exhibit tunable swelling ratios (SR) ranging from −7 % to 1280 % in response to denaturants, directly linking protein unfolding-folding mechanisms to functional shape morphing for scaffold adaptability [183].

Moreover, upon implantation, SMHs interact with the host immune system, initiating a cascade of events that determine whether integration or rejection occurs. The initial response involves protein adsorption and neutrophil infiltration, followed by the recruitment and polarization of macrophages. Biocompatible SMHs typically promote a resolution-phase response, with M2-polarized macrophages releasing IL-10 and TGF-β to support tissue remodeling. For instance, a shape-memory collagen–hyaluronic acid spinal scaffold reduced TNF-α and IL-1β expression in vivo, while a near-infrared (NIR)-triggered bone SMH scaffold upregulated IL-10 and IL-4 while suppressing IL-6 and TNF-α in RAW264.7 macrophages, demonstrating immunomodulatory potential [184,185]. In cartilage, bone, vascular, and neural tissues, SMHs have shown consistent in vivo biocompatibility. Implanted PTK-based cartilage SMHs elicited no inflammatory-cell infiltration in surrounding rat skin over 30 days and caused no histopathological changes in major organs [73]. In bone regeneration models, 3D-printed SMH/Mg implants enhanced M2 macrophage polarization and supported osteogenic differentiation. Vascular SMH foams used to occlude renal arteries in pigs exhibited no systemic toxicity or aneurysm recanalization after 3 months [155]. Neural applications using gelatin-MnO₂ SMHs demonstrated enhanced survival and differentiation of ESC-derived motor neurons, with no evidence of glial scarring or adverse immune response. Histological analyses of subcutaneous SMH implants show a transient, self-limiting foreign body reaction (FBR). Macrophage and multinucleated giant cell infiltration are typically observed by day 7, followed by scaffold resorption and replacement with organized fibrous tissue by day 21. No granulomas or chronic inflammation are generally reported, provided that degradation is well controlled and residual reactants are minimized. Nonetheless, caution is warranted for formulations containing polyacrylamide or other synthetic polymers that may release cytotoxic monomers like acrylamide [186].

While SMHs demonstrate broad biocompatibility, several challenges remain. Residual crosslinkers or photothermal agents can trigger oxidative stress or chronic inflammation if not adequately purified. Dynamic shape change itself, especially if abrupt or repeated, may induce mechanical irritation and local immune activation. Moreover, stimuli-responsive elements such as magnetic nanoparticles or graphene derivatives, while enhancing functionality, can pose immunotoxic risks due to bioaccumulation or ROS generation [187]. Patient-specific immune variability also complicates the standardization of scaffolds. Immunocompromised or elderly individuals may experience exaggerated or delayed inflammatory responses even with well-tolerated materials. Furthermore, long-term in vivo studies evaluating chronic exposure and late-stage degradation products are still lacking for most SMH systems.

To mitigate these risks, advanced SMHs are being designed with immunomodulatory functions, such as incorporating RGD peptides, zwitterionic coatings, or controlled-release anti-inflammatory agents like dexamethasone. Surface modifications that reduce protein adsorption, alongside hybrid hydrogel architectures that combine physical and covalent networks, can reduce FBR and extend the lifespan of the scaffold. Future directions include 4D-printed patient-specific SMHs with embedded biosensors for real-time monitoring of immune status and scaffold performance.

5.2. Sterilization and materials stability

Sterilization is essential for biomedical implants to mitigate infection risks, yet it must preserve the hydrogel's structural integrity, mechanical properties, and shape memory behavior. Conventional methods include thermal (e.g., autoclaving), chemical (e.g., ethylene oxide, EtO), and radiation-based (e.g., gamma irradiation, electron beam) approaches, each influencing crosslinking density, swelling ratios, and degradation profiles. Emerging techniques, such as UV irradiation and ethanol immersion, offer alternatives but require optimization for SMHs [188].

Autoclaving (steam/heat) sterilization can induce hydrolysis in ester-based SMHs, reducing crosslinking and altering shape fixity ratios from >90 % to <70 % in PCL-derived systems, while increasing swelling ratios by 20–50 % due to chain scission. This method is less suitable for heat-sensitive SMHs, as elevated temperatures may prematurely activate thermoresponsive networks, thereby compromising the functionality of cartilage or neural tissue scaffolds. For example, autoclaving alginate-based SM hydrogels resulted in significant water loss and stiffening, with a storage modulus (G′) increase of ∼80–190 % compared to the untreated gel. The drying and re-crosslinking during autoclave processing can compromise the original pore network. Similarly, GelMA (a common SMH) lost ∼70–80 % of its compressive modulus after autoclave sterilization. In general, autoclaving tends to tighten polymer networks (raising modulus) and reduce swelling capacity. Thus, unless specifically designed for heat resistance, many SMHs are unsuitable for autoclaving: their shape-fixity can remain high, but actuation stress/strain and recovery speed may be altered by residual crosslink density [189].

Ethylene oxide (EtO) gas (37–55 °C, 1–12 h) is widely compatible, preserving shape recovery ratios (>95 %) in thiol-ene/acrylate SMHs and poly(glycerol dodecanedioate) (PGD) elastomers, with minimal changes in storage modulus (G’ ∼1–10 MPa). However, residual EtO can cause cytotoxicity if not adequately aerated, and higher temperatures (e.g., 54.4 °C) reduce recovery rates by 10–20 % over time due to ester bond degradation. In GelMA, EtO treatment decreased stiffness (similar to autoclave) and significantly lowered cell viability in 3D cultures. This is attributed to EtO alkylation of residual reactive groups. EtO residues can also accelerate polymer degradation or leachables, so extensive aeration is required. In practice, EtO may be acceptable for SMH scaffolds if followed by thorough outgassing; however, one must re-evaluate the mechanics and shape-memory performance (fixity/recovery) post-sterilization [42].

γ-Irradiation (gamma) at clinical doses (25–40 kGy) is highly penetrating, but induces radiolysis/crosslinking in polymers, increasing Young's modulus by 15–30 % but potentially decreasing degradation rates, which is beneficial for long-term implants but may hinder bioresorbability in bone tissue engineering. Electron beam irradiation offers faster processing but can cause oxidative degradation, reducing the tensile strength of PEG hydrogels by up to 25 %. In the GelMA study, γ‐irradiation increased stiffness (modulus rise) and dramatically slowed biodegradation (more crosslinked network). Critically, γ‐sterilized GelMA prepolymer could no longer form a proper gel (sol–gel transition severely impaired). Many hydrogels (e.g., PEG, PVA) similarly undergo crosslinking or chain scission under γ‐rays. For shape-memory, excessive crosslinking may lock the network, preventing thermally triggered recovery. Thus, while γ-sterilization ensures sterility, its dosage must be carefully controlled, and post-irradiation mechanical/fixity testing is mandatory [190].

Ultraviolet (UV) light sterilization is surface‐limited and often used only for final device packaging or for 2D surfaces. Immersion of 3D SMHs in UV (usually 254 nm) can reduce surface microbes with minimal heat, but it penetrates poorly. In alginate gels, UV exposure had no significant effect on swelling or stiffness (similar to ethanol). However, microbial kill rate can be variable (some live bacteria remained in one study). UV is generally not recommended as a sole sterilant for bulk SMHs, unless combined with other methods [191].

Ethanol/water treatment is an effective disinfection method in which the SMH scaffolds are immersed in sterile ethanol, typically at concentrations of 70–100 %, which often preserves the gel mechanics. Stoppel et al. found that a 70 % ethanol wash eliminated all bacteria from alginate–Pluronic hydrogels while minimally altering the water content and mechanical moduli [192]. In GelMA, water/ethanol processing has been used as a sterilization strategy for bioinks. The downside is that organic solvents can extract uncrosslinked monomers or plasticizers; careful re-swelling is necessary after ethanol soaking. Overall, ethanol is a gentle method for disinfection (not full sterilization) of SMHs, with little impact on shape-memory; it leaves most supramolecular and covalent crosslinks intact [193].

Material stability encompasses the ability of SMHs to maintain structural, mechanical, and functional properties over time in physiological environments, influenced by degradation kinetics, environmental factors, and network architecture. Biodegradable SMHs, such as those based on PCL or PLLA, degrade via hydrolysis or enzymatic action, with rates tunable from weeks to months; for instance, PCL-PLLA composites exhibit <10 % mass loss over 12 weeks in vivo, supporting gradual tissue ingrowth [194]. Dynamic crosslinks (e.g., hydrogen bonds, ionic coordination) enhance stability by enabling self-healing, with dual-network systems achieving toughness values of 1–5 MJ/m³, comparable to native cartilage. Long-term stability is affected by pH (6.5–7.5), temperature (37 °C), and mechanical loading, where swelling ratios may increase by 100–500 % in responsive networks, potentially leading to dimensional instability. In tissue engineering, stability ensures sustained cell support; gelatin-based SMHs maintain shape fixity of over 85 % over 4 weeks, facilitating vascularization. Future directions include 4D printing for customized, stable scaffolds and multi-stimuli systems for on-demand activation, alongside rigorous in vivo models to validate long-term performance and accelerate clinical adoption in regenerative therapies.

5.3. Regulatory approval

SMHs intended for tissue engineering are typically regulated as medical devices under the U.S. FDA framework. Acellular SMHs are generally classified as Class II or Class III devices, depending on their risk level. In contrast, SMHs that incorporate drugs, biologics, or cells are categorized as combination products, subject to dual regulatory oversight. These combination products must meet both device-related requirements, such as biocompatibility testing (including cytotoxicity, sensitization, implantation, and degradation studies), and drug or biologics-specific evaluations, including pharmacokinetics, toxicology, and stability. SMHs with biodegradable chemistries also require clearance of degradation byproducts, ensuring that no harmful monomers, oligomers, or crosslinkers persist post-degradation [195].

To date, no SMH-based hydrogel scaffold for tissue regeneration has received FDA approval; however, the path forward is informed by successful regulatory cases involving shape-memory polymer (SMP) devices. For example, the IMPEDE® vascular plug (Shape Memory Medical Inc.), composed of SMP foam, has achieved FDA clearance and CE marking for vascular occlusion [196]. This precedent demonstrates that shape-memory materials can pass regulatory review if sterilization, mechanical integrity, and biocompatibility are validated. For SMHs, key hurdles include ensuring consistent shape recovery after sterilization, validating long-term mechanical stability in vivo, and demonstrating clinical safety through preclinical models. Developers are advised to engage with the FDA early, via pre-submission (Pre-Sub) meetings, to define device classification, applicable standards, and data expectations tailored to SMHs’ novel material properties. Different barriers in advancing SMHs from bench to bedside are shown in Table 4.

Table 4.

Barriers and design strategies for advancing SMHs in tissue engineering.

Barrier type	Limitations	Causes	Proposed solution	Mechanisms	Representative studies	Translational relevance	Ref.
Mechanical strength & tunability	Low toughness High stimulus energy Imprecise recovery Early degradation Soft segments enable memory but weaken rigidity	Weak bonds Energy inefficiencies Entropy-driven chain randomness Hydrolytic instability Fixed crosslinking	Dual/IPN networks 4D printing Nanofillers Self-healing bonds Post-gel crosslinking	Energy dissipation Precise shape fixity Nanoscale reinforcement Stress-responsive bonding	CNT IPNs – cartilage Cryoprinting – stents Silico-nanosheets – bone Diels-Alder – nerve guides Orthogonal click chemistries– sensor	Safer implants Better recovery fidelity Match tissue strength Aligns degradation	[197,198]
Biocompatibility & bioactivity	Toxic degradation Weak adhesion Inflammation Hypoxia sensitivity	Synthetic SMH networks lack motifs Immune reactions Diffusion barriers Acidic fragments	RGD/growth factors Natural-synthetic blends Porosity gradients ROS-responsive linkers	Cell-matrix signaling Immunosuppression Improved permeability Protective activation	RGD – cartilage HA-Gel hybrids – neural ROS linkers – chondrocytes Gradient pores – bone	Lower rejection Faster healing Enables cell therapy Trial readiness	[199,200]
Stimuli responsiveness & delivery	Low stimulus reach Poor retention Fast/slow degradation Actuation mismatch	Signal loss Enzyme rate variation Shear misfit Structural mismatch	Bio-cues (temp/enzyme) Tunable linkers 3D printing Adhesive moieties	Autonomous actuation Customized shape fit Triggered degradation Anchored retention	Thermo-SMHs – spine 3D-printed – skull Ester SMH – cartilage Mussel-inspired – vessels	Minimally invasive Custom therapy Sustained effect Retention improves outcome	[51,201]
Stimulus compatibility	Harmful triggers (UV, heat) Inaccessibility in vivo	High Tg/Tm Low tissue penetration Bulky equipment	Physiological triggers NIR/magnetic triggers Multi-stimuli systems	Body-temp recovery Enzyme-triggered bonds In situ energy conversion	PU-SMH @37 °C NIR-MOFs – octopus model	Safe deployment No thermal damage Less device complexity	[202,203]
Multifunctionality & complexity	Function compromise Difficult shape programming	Network interference Bond incompatibility Few studies on multi-memory	Layered networks Reversible chemistries Function-isolated systems	Strain retention + healing Cycle stability Modular response	AA-derived SMH – tunable 4D-printed SMH – limbs	Multi-use platforms Complexity vs approval Dual-role validation	[204,205]
Translational & manufacturing	Costly scaling Sterilization issues Regulatory hurdles on safety/efficacy Preclinical-clinical gaps	Labor-intensive Complex synthesis Heat-sensitive gels Patient heterogeneity Model limitations Novel biomaterials lack established regulatory pathways	FDA-approved polymers Aseptic/UV protocols Standardization Microfluidic scalable methods Animal models Interdisciplinary collaborations	Approval acceleration Safe prep methods High-throughput production Expertise streamlining Reproducible results	SMP foam trials GelMA precedents VEGF-loaded SMH Microfluidics – nerve Rat models for spinal/cranial defects Biotech-academia for trial designs	Clinical readiness Lower production barrier Improves integration rates Reliable data generation Lowers commercialization barriers Speeds bench-to-bedside	[206,207]

6. Conclusion and outlook

Recent advancements in science, technology, and polymer materials have significantly increased interest in SMHs, especially in the biomedical field and industrial applications. SMHs respond to different stimuli, including pH, chemicals, temperature, water, light, and even ultrasound. These stimuli-responsive hydrogels exhibit adjustable recovery rates and potential biodegradability, addressing the challenges of developing dynamic biomedical devices. Compared to traditional stimulus-responsive polymers, SMHs not only offer customizable size specifications but also demonstrate relatively rapid shape recovery. Furthermore, their inherent properties, such as wettability, biocompatibility, antifouling properties, specific biofunctionalization, high stretchability, controllable shape transformation, and adaptive features, make SMHs highly promising for biomedical applications and clinical use. These properties support the development of minimally invasive treatments, reducing the side effects, costs, and duration associated with conventional procedures.

In this review, we provide an overview of SMEs and the potential applications of SMHs. Our study first explored the fundamental properties of SMHs, particularly their ability to deform in response to specific stimuli, such as water or other environmental factors, and discussed recent advances in their use for tissue engineering. Despite the substantial progress made in exploring SMHs for biomedical applications, there remains significant scope for innovation, particularly in interdisciplinary research areas such as controlled drug delivery, tissue engineering, and biosensing.

Several critical factors must be addressed before SMHs can be widely applied in clinical settings. One of the primary limitations of SMHs in biomedical applications is their insufficient mechanical strength, largely due to their high water content (typically 70–90 %). This limitation compromises their tensile strength and toughness, making them unsuitable for applications requiring substantial mechanical robustness, such as tendons, bones, and cartilage. Currently, SMHs exhibit relatively low mechanical properties, with Young's modulus and tensile stress both mostly below 1 MPa. Strategies to enhance these properties include developing homogeneous network structures, incorporating hybrid materials, and introducing efficient energy dissipation mechanisms. For instance, adding inorganic nanoparticles or employing multiple interpenetrating polymer networks can significantly improve the mechanical strength of SMHs. Furthermore, advanced fabrication techniques, such as 3D and 4D printing, allow the integration of high-load-bearing polymers with soft hydrogel matrices, further enhancing the stiffness and load-bearing capacity of SMHs. Despite these advancements, further work is necessary to fully mimic the biomechanical properties of native tissues.

Additionally, previous works mainly focused on designing and fabricating hydrogel networks with irreversible SME, namely unidirectional SME. However, for a typical unidirectional SME cycle, the requirements of human application in many aspects cannot be met without reprogramming. Therefore, it is necessary to implement a bidirectional SME so that the SMH can switch between two different shapes (i.e., the shape is restored from temporary to permanent shape and vice versa). In addition, SMH, which combines multistimulus response shape deformation and stepwise reconfigurability, has not been explored well.

To successfully integrate SMHs into biomedical applications, their biodegradability and biocompatibility are crucial. These properties allow for complete degradation after implantation, support tissue regeneration, and facilitate the restoration of local cellular architecture. Tissue engineering scaffolds require programmable and tunable degradation kinetics to synchronize scaffold degradation with tissue repair and regeneration. If the scaffold degrades too quickly, it will compromise mechanical integrity, while slow degradation may lead to inflammation. To overcome these challenges, multistage degradation systems can be employed by incorporating biodegradable polymers with varying degradation rates or by utilizing multi-stimulus-responsive polymers for controlled degradation. The incorporation of bioactive nanoparticles within SMHs can further enhance the degradation process, thereby optimizing their effectiveness in tissue repair.

Another promising direction is the programming and design of SMHs to better mimic biological environments by precisely modulating shape changes in vivo. Given that biological tissues possess self-healing abilities when subjected to specific thresholds, transferring these self-repair capabilities to smart polymers holds significant potential. Vascularization is also a critical challenge in tissue engineering using SMHs. Without proper vascular networks, tissues larger than a few microns can suffer from inadequate oxygen and nutrient supply, leading to necrosis. To address this, scaffolds must not only support cell growth and structural integrity but also promote capillary ingrowth. Embedding growth factors, prevascularizing networks within the hydrogels, or incorporating oxygen-generating agents or nanoparticles, such as calcium oxide, can help enhance oxygen supply and promote tissue regeneration.

Despite significant progress in the design and manufacturing of SMHs, further advancements are needed to realize their potential fully. Emerging technologies such as 4D printing and nanotechnology can provide new opportunities for SMHs in areas like soft robotics, actuators, scaffolds, and biomedical devices. A more in-depth evaluation of these applications will be essential to unlock the full potential of SMHs.

Moreover, reliable shape fixation in vivo remains difficult because hydration and shear forces in blood vessels or tissue interstitium can erode programmed configurations over time. Degradation products may contain acidic or reactive fragments whose effects on cell viability and tissue compatibility are not yet fully defined. In addition, the host immune response to residual hydrogel fragments, including macrophage activation, cytokine release, and fibrotic encapsulation, has not been systematically evaluated. To overcome these hurdles, future studies should incorporate bioreactor tests under physiologically relevant flow conditions, conduct detailed chemical analyses of breakdown fragments, and utilize standardized immunotoxicity assays in relevant animal models.

In recent years, various SMHs have been designed and developed, fueling substantial growth in related research. However, SMHs remain in the early stages of development, and several obstacles must be overcome before they can be applied clinically. By adjusting molecular structures (through copolymerization or blending), improving manufacturing techniques, and refining design strategies, SMHs are poised to move beyond proof-of-concept and into real-world biomedical applications. It is expected that SMHs will transition from experimental models to practical solutions for biomedical challenges in the near future.

CRediT authorship contribution statement

Abid Naeem: Writing – original draft, Visualization, Investigation, Conceptualization. Chengqun Yu: Writing – original draft, Formal analysis. Lili Zhou: Visualization, Formal analysis. Yingqiu Xie: Visualization, Formal analysis. Yuhua Weng: Writing – review & editing, Visualization, Funding acquisition. Yuanyu Huang: Writing – review & editing, Validation, Supervision, Project administration, Funding acquisition. Mengjie Zhang: Writing – review & editing, Funding acquisition. Qi Yang: Writing – review & editing, Validation, Funding acquisition.

Ethics approval and consent to participate

Not applicable. This article does not contain any studies with human participants or animals performed by any of the authors.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.