The discovery of the first hydrogel intended for biological use dates back to 1960 [1]. Since then, hydrogel systems have undergone significant improvements. This development has seen the transition of hydrogels from conventional “dumb” gels into “smart” gels capable of responding to unique external stimuli or environmental changes. In recent decades, smart hydrogels have received increasing attention in the field of drug delivery due to their stimuli-responsive properties; low invasiveness; ease of administration; and controlled, sustained drug release capabilities ^{[2][3][4][5][6][7]}[2,3,4,5,6,7]. While conventional “dumb” gels can swell or shrink in response to osmotic pressure, their responsiveness is often inefficient, leading to imprecise drug release and limited control over dosage timing [8]. Thanks to advancements in polymer science and nanotechnology, researchers have been able to develop “smart” hydrogels, endowed with tunable and “on-demand” drug release patterns ^{[3][4][7][9][10]}[3,4,7,9,10]. These smart gels, also referred to as stimuli-responsive hydrogels, are engineered to respond to various stimuli such as temperature, pH, electromagnetic radiation, magnetic field, or the presence of specific biological factors. The fundamental aspect of smart hydrogels is their ability to modify their mechanical properties, swelling ability, hydrophilicity, or bioactive molecule permeability, influenced by various stimuli. This property enables them to be triggered to release drugs in a controlled and targeted manner, thus enhancing the precision and effectiveness of drug delivery.

Smart hydrogels exhibit versatile applications, spanning tissue engineering, cell cultures, and innovative drug delivery systems (DDSs) ^{[3][11][12][13]}[3,11,12,13]. By responding to physical, chemical, and biological stimuli, these hydrogels can effectively release active substances, making them ideal for disease treatment. Furthermore, their mechanical properties are comparable to those of biological tissues, making them well-suited for mimicking natural living tissues [6].

In DDSs, smart hydrogels deliver substantial advantages ^[3][14][15][3,14,15]. They respond to stimuli such as pH, temperature, light, redox, and biomolecules, leading to improved drug efficacy and reduced side effects. Moreover, they enable precise drug release over time, reducing the need for frequent dosing. This highlights the immense potential of smart hydrogels in revolutionizing targeted drug delivery by improving efficacy, minimizing side effects, and enabling personalized treatments.

2. pH-Sensitive Hydrogels

Biological fluid pH is a significant chemical property that holds immense potential in the drug delivery field. pH-responsive hydrogel and nanogels have been developed as intelligent drug delivery carriers, capable of exhibiting swelling or shrinking behavior in response to changes in pH ^{[16][17][18][19]}[71,72,73,74]. Thus, these hydrogels are able to release encapsulated drugs in a site-specific manner. The pH-responsive mechanisms can be divided into two types: First is employing polymers with ionizable moieties like amines and carboxylic acids, protonated or deprotonated at various pH values ^[20][21][38,39]. For instance, when the pH level is below the pKa of the basic functional groups, such as poly[(2-dimethylamino)ethyl methacrylate] (PDMA), the hydrogel becomes protonated and forms a positively charged polymer chain, leading to swelling due to electrostatic repulsion between the charged groups. Conversely, when the pH level is above the pKa of the functional groups in the hydrogel, the hydrogel is deprotonated and shrinks due to electrostatic attraction between the charged groups. Polyacids, such as poly(methacrylic acid) (PMAAc), behave inversely, accepting protons at low pH and releasing protons at neutral and high pH. Therefore, pH shifts can alter the interaction of hydrogel polymer chains, triggering drug release from pH-sensitive hydrogels. However, it is important to note that the response of ionizable moieties to pH can be influenced by other factors, such as temperature and ionic strength ^[22][40], making it challenging to control drug release in complex conditions. Second is utilizing polymers that contain acid-labile linkages ^[23][24][41,42]. These covalent linkages are stable at physiological pH but cleave as pH decreases, leading to polymer chain degradation or aggregate dissociation. Various chemical moieties, including anhydrides (DMMA, succinic anhydride, cis-aconitic anhydride, and cis-cyclohexene-1,2-dicarboxylic anhydride) ^[25][26][43,44], hydrazone ^[27][28][45,46], and imine ^[29][47], can be adopted to create acid-labile linkages. Compared to the first approach, this approach is more flexible, as it allows for the selection of different acid-labile linkages conjugated with various polymers. As a result, it offers more precise control of drug release in response to the acidic environment in pathological conditions ^[16][23][41,71]. However, note that acid-labile linkages can be unstable, potentially releasing drugs before reaching the acidic target sites. For example, Zou et al. reported that the drug conjugate poly(ethylene oxide)-block-polyphosphoester-graft-PTX (PEO-b-PPE-g-PTX G2) with acid-labile linkages exhibited 20% PTX release in 8 days under neutral conditions, even though it showed accelerated drug release under acidic conditions (approximately 50% PTX release in 8 days under acidic conditions) ^[30][48]. It is noteworthy to mention that the conventional polymerization methods for acid-labile linkages may raise toxicity issues ^[18][73]. Over the past few years, one of the main advancements in this field has been the development of pH-responsive hydrogels and nanogels with improved biocompatibility and drug-loading capacity. Researchers have explored natural polymers like chitosan and alginate, along with biocompatible polymers such as hydroxypropyl methylcellulose (HPMC), to create pH-sensitive hydrogels that are both biodegradable and safe ^[31][32][33][75,76,77].

3. Thermoresponsive Hydrogels

Due to their inherent physiological condition and convenient administration, temperature or thermoresponsive hydrogels have been widely utilized in the field of drug delivery. These hydrogels can undergo phase transitions from a swollen state to a collapsed or shrinking state, or vice versa, in response to changes in temperature. Using monomers such as N-isopropylacrylamide (NIPAM) and cross-linkers such as methylene bisacrylamide (MBA) or poly(ethylene glycol) diacrylate (PEGDA), temperature-sensitive hydrogels are commonly synthesized through free radical polymerization ^[34][35][49,50]. Sol–gel phase transitions in these hydrogels are driven by changes in the interaction between their hydrophobic and hydrophilic segments with water molecules, leading to changes in the solubility of the cross-linked network and resulting in sol–gel phase transition ^[36][78]. The sol phase is a flowing fluid, while the gel phase is non-flowing and maintains its integrity. Hydrogels can form either above the lower critical solution temperature (LCST) or below the upper critical solution temperature (UCST), depending on the specific composition and ratio of hydrophilic and hydrophobic components. The polymer is soluble below the LCST, but as the temperature rises above the LCST, the hydrogel begins to shrink, becoming hydrophobic and insoluble, resulting in the formation of a gel ^[37][79]. Conversely, cooling the polymer solution below the UCST triggers the formation of a hydrogel. Near the critical temperature, the polymer undergoes a phase change from a soluble state (random coil) to an insoluble state (collapse or micelle form) ^[38][80]. The phase transition temperature (PTT) of a temperature-sensitive hydrogel can be adjusted by changing the chemical composition of the polymers, the concentration of anionic monomer, or the ratio of hydrophilic/hydrophobic groups in the gel materials ^[39][81]. While natural thermoresponsive polymers, such as some polysaccharides (e.g., agarose, amylose, amylopectin, and some cellulose derivatives) and certain proteins (e.g., gelatin, collagen, and elastin-like polypeptides), can form thermo-reversible hydrogels, they generally exhibit weak mechanical strength and slow temperature responses, necessitating chemical modification to improve their properties ^[40][41][51,82]. In contrast, synthesized polymers, such as poly(N-isopropylacrylamide), poloxamers, and PLGA-PEG-PLGA triblock polymers, offer greater adjustability in physical properties. They can be incorporated into natural or synthetic polymers to introduce thermoresponsive qualities, enabling controlled drug release. For example, N-isopropylacrylamide (NIPAM) can be grafted with polymers like alginate ^[42][83], chitosan ^[43][84], hyaluronic acid ^[44][85], or PEG ^[35][50] to impart thermoresponsive properties to these polymers. Regarding biocompatibility, it is widely acknowledged that the NIPAM monomer possesses toxicity. However, the PNIPAAm polymers with high molecular weight or the grafted pPNIPAAm with other polymers have exhibited notable biocompatibility in several studies ^[45][46][47][86,87,88]. Examining the cytocompatibility and hemocompatibility of thermoresponsive PNIPAAm and PNIPAAm-PLLA-PNIPAAm triblock copolymers, Su et al. discovered that the latter demonstrated exceptional biocompatibility, thus offering potential for targeted drug delivery ^[45][86]. In a separate study, Yogev et al. investigated the biocompatibility of thermoresponsive polymers, including PNIPAAm, poly(ethylene glycol)-poly(propylene glycol)-poly(ethylene glycol) triblock copolymer, poly(lactic acid-co-glycolic acid), and poly(ethylene glycol) triblock copolymer, both in vitro and in vivo. They concluded that all tested materials demonstrated satisfactory biocompatibility in vivo ^[45][86]. The toxicity of commercially available PNIPAAm probably results from the release of NIPAM monomer and impurities in the pNIPAAm.

4. Photo-Responsive Hydrogels

Light exposure serves as a stimulus for drug release from hydrogels, encompassing UV, visible, and infrared light ^{[48][49][50][51]}[52,53,54,55]. As a non-invasive and efficient external trigger, light optimally controls drug release from hydrogels, enhancing therapeutic effectiveness and minimizing side effects by regulating drug distribution within the body. Photo-responsive hydrogels undergo phase transition, stiffness alteration, or biochemical activation upon light exposure, prompting drug release exclusively in illuminated areas. This method offers accurate and non-contact drug delivery, applicable in a range of medical scenarios such as chemotherapy, immunotherapy, photodynamic therapy, gene therapy, and wound healing ^{[52][53][54][55]}[89,90,91,92]. Various photosensitive chemical moieties can be used for photo-responsive functionality, including o-nitrobenzyl ester linkers [56], arylazopyrazole [57], azobenzene [58], and photocleavable proteins such as PhoCl [59]. These chemical moieties allow hydrogels to respond to either UV light or visible light, leading to reversible changes in their properties. Photo-responsive hydrogels are fabricated by integrating these photosensitive components into their polymeric structures through diverse approaches. These approaches can be categorized into three groups based on their mechanisms: photoisomerization, photochemical reactions, and photothermal reactions ^[51][55]. Photoisomerization is a process in which a molecule undergoes a structural transformation upon light absorption. Upon exposure to light, the photosensitive components within the hydrogel experience photoisomerization, which induces a change in the hydrogel’s structure and initiates drug release ^[49][53]. Photochemical reactions, on the other hand, refer to chemical reactions initiated by light absorption. In light-responsive hydrogels, photosensitive moieties undergo photochemical reactions upon exposure to light. This results in the cleavage of chemical bonds, leading to drug release ^[60][93]. Photothermal reactions involve the absorption of light by a material, causing an increase in temperature. In light-responsive hydrogels, this resulting temperature rise brings a change in the hydrogel’s structure and then triggers drug release [3]. It should be noted that photo-responsive processes may be reversible or irreversible. Reversible photo-responsive hydrogels can undergo a gel-to-sol transition or sol-to-gel transition upon light exposure. This can be used to accelerate drug release and achieve a step-by-step release pattern. Irreversible photo-responsive hydrogels undergo a permanent change upon light exposure, which can be used for one-time drug release or gradual release through hydrogel degradation. Photo-responsive hydrogels can respond to varying light wavelengths, including UV, visible, and near-infrared (NIR) light, depending on the adopted photosensitizers. UV light provides greater energy for photo curing compared to the other two longer-wavelength lights, resulting in a higher curing speed and efficiency ^[61][94]. However, UV light is associated with DNA, tissue damage, and limited tissue penetration due to light absorption and scattering by water and other substances ^[62][95]. UV light photocuring can be used when enhanced photochemical reactions are required, such as in vivo 3D printing of bone substitutes through photo-fabrication technologies ^[61][94]. On the other hand, near-infrared (NIR) light within the 700 to 1000 nm range is more appealing for DDSs compared to other wavelength spectra. It does not cause harm to living cells or tissues, and it also possesses superior tissue penetration capability ^[63][96]. As a result, an increasing number of NIR-sensitive photo-responsive hydrogels are being developed. Recent advances in material science have led to diverse new photosensitizers for NIR-sensitive photo-responsive hydrogel preparation, such as rare metal nanostructures and black phosphorus nanoparticles ^[48][50][52,54]. Qiu et al. have developed an innovative photothermal hydrogel by integrating black phosphorus into hydrogel nanostructures ^[50][54]. The hydrogel can be activated by NIR light with a wavelength of 808 nm, which causes the drug-loaded hydrogel nanostructures to soften and melt, ultimately resulting in drug release. Qiu used a power density of 1 W/cm², while Auge et al. designed a more efficient photothermal hydrogel, notably lowering the power density to 0.16 W/cm². They formulated a nickel-bis(dithiolene) complex that can undergo a volume phase transition and release loaded hydrophobic dye molecules upon NIR light exposure. They also extended the working NIR spectral region to 1000 nm ^[48][52].

5. Redox-Responsive Hydrogel

Innovative DDSs have emerged with the development of redox-responsive hydrogels ^[64][65][66][60,97,98]. These hydrogels, when subjected to specific biological redox stimuli, can rapidly release encapsulated drugs at the target site ^[67][62]. The redox-responsive behavior of the hydrogel/nanogels is achieved through the incorporation of specific chemical moieties. One such chemical moiety is the disulfide linker, which can be cleaved in the presence of reducing agents such as glutathione (GSH) ^[64][68][60,61]. Another widely used chemical moiety is the selenide group, which can be responsive to reactive oxygen species (ROS) such as hydrogen peroxide (H₂O₂) ^[69][63]. The redox-responsive hydrogel carriers can be specifically designed for specific purposes. For gene drug delivery, Zhao et al. devised gelatin/silica-aptamer nanogels that can selectively release siRNA into the cytosol in nucleolin-positive cells (A549) triggered by GSH ^[64][60]. To encapsulate therapeutic proteins, Schotz et al. crafted polyglycerol-based redox-responsive nanogels using inverse nanoprecipitation and inverse electron-demand Diels–Alder cyclizations between methyl tetrazines and norbornene. The encapsulated cytochrome C was released at the action site under physiological reductive conditions ^[70][64]. For intracellular delivery of cationic drugs, Maciel et al. synthesized redox-sensitive nanogels (AG/Cys) through in situ cross-linking of alginate using cystamine as a cross-linker via a mini-emulsion method. The cationic doxorubicin was encapsulated via electrostatic interactions, and the encapsulation efficiency was up to 95.2 ± 4.7% ^[51][56][55,56]. Natural polymer-based materials, such as hydroxypropyl cellulose (HPC)-based grafted copolymers, were also adopted for biodegradable and biocompatible nanogels ^[65][71][97,99]. These advancements in redox-responsive hydrogels/nanogels hold more promise for effective and targeted DDSs.

6. Biomolecule-Responsive Hydrogels

6.1. Enzyme-Responsive Hydrogels

Enzyme-responsive hydrogels have been developed to respond to enzymatic activity in specific environments, enabling the release of encapsulated therapeutics ^[72][73][65,66]. These hydrogels offer a distinct advantage due to their reliance on endogenous enzyme expression. One mechanism involves designing the hydrogel matrix to degrade under specific enzymes, allowing for controlled therapeutic release. Yang et al. devised enzyme-responsive nanogels (EPNGs) cross-linked with cinnamyloxy groups in PEGylated hyaluronic acid, which are sensitive to hyaluronidase ^[73][66]. These EPNGs exhibit high loading efficiency and excellent stability in various biological media. However, they degrade rapidly within tumor cells that overexpress hyaluronidase, allowing for rapid release of encapsulated cytochrome C. Another mechanism involves covalently linking therapeutics to the hydrogel scaffold with enzymatically sensitive cross linkages. Amer et al. created a PEG hydrogel delivering the anticancer drug doxorubicin, which was covalently attached to the hydrogel via the MMP-sensitive peptide linker, C-VPLS↓LYSG-C ^[74][67]. The hydrogel was able to release doxorubicin in the presence of MMP-2 and MMP-9, significantly reducing tumor growth in a breast cancer mouse model [57].

6.2. Glucose-Responsive Hydrogels

In response to real-time blood-glucose levels, glucose-responsive carriers have been developed to facilitate insulin release ^[75][76][77][68,100,101]. These carriers can be designed using motifs such as glucose oxidase (GOx), phenylboronic acid (PBA), or concanavalin A (Con A) ^[78][102], which can detect glucose levels in their surroundings. By integrating these motifs into the hydrogel matrix, they can be utilized to trigger insulin release in a glucose-responsive manner, leading to improved control of blood-glucose levels and reduced risk of hypoglycemia ^[79][103]. The GOx enzyme catalyzes the conversion of glucose to gluconic acid in the presence of oxygen, causing local physiochemical changes such as pH, H₂O₂ levels, or alterations in oxygen concentrations. These changes are then harnessed to induce insulin release in a glucose-responsive fashion. It should be noted that increasing levels of H₂O₂ and gluconic acid can hinder GOx activity, thereby diminishing the hydrogel’s property changes and reducing sensitivity to glucose. To address this issue, Gordijo introduced the enzyme catalase (CAT) to convert H₂O₂ to water and O₂, enhancing GOx activity ^[80][104]. Based on this, Gu et al. engineered glucose-responsive closed-loop insulin microgels (256 ± 18 μm) containing a pH-responsive chitosan hydrogel and GOx enzyme nanocapsules through a one-step electrospray process ^[81][69]. They observed a decrease in blood-glucose levels in a type 1 diabetes mouse model. Moreover, improving the stability of the GOx enzyme, which is inherently unstable, can be achieved by using high hydrostatic pressure and hydrophobic modification ^[82][105]. Kim et al. developed glucose-responsive hydrogels cross-linked by citric acid, embedding GOx within the hydrophobic β-CD cavity to enhance its stability and achieve long-term glucose monitoring ^[76][100]. To overcome the instability of GOx enzyme, a synthetic glucose-responsive motif, phenylboronic acid (PBA), was adopted. As a derivative of boronic acid, PBA can form reversible covalent bonds with diols like glucose. When glucose levels are high, competition for saccharide binding sites on glucose-binding molecule-polymer complexes lead to separation of the complex, triggering insulin release. Wang et al. introduced the hydrophilic monomer N-vinyl-2-pyrrolidone and the amino-containing monomer N,N-dimethylaminopropyl acrylamide to PBA-based polymers to create glucose-sensitive microgels that respond to physiological temperature and pH ^[75][68]. These microgels, formed by a reversed-phase microemulsion method, create a dense network at low glucose levels that can encapsulate insulin and prevent leaks but disrupt the bond at high glucose levels, thus rapidly releasing insulin. It should be noted that PBA lacks specificity for glucose, despite being more stable than the GOx enzyme. Similarly, concanavalin A (Con A), a natural carbohydrate-binding protein, can reversibly attach to glucose and other saccharides. Con A forms a tetrameric structure and can bind to four glucose molecules, acting as a macromolecular crosslinker. When Con A is integrated into the hydrogel matrix, it forms a glucose-binding element that can release insulin when glucose levels change. Variations in the ratio of uncharged and charged borates, influenced by glucose, impact polymer solubility and facilitate glucose-responsive insulin release. Based on Con A’s glucose-responsive property, Lin et al. developed a pullulan-based glucose-responsive hydrogel by covalently modifying Con A with a pullulan derivative containing COOH groups, allowing for intelligent, controlled insulin release upon glucose concentration shifts ^[83][70]. The hydrogel swells and releases insulin when glucose levels are high, but when glucose levels are low, insulin release is reduced. These discoveries contribute to the progress of hydrogel systems that can adapt to glucose changes and autonomously regulate the release of anti-diabetic medications according to blood glucose levels ^[84][106]. These glucose-responsive hydrogels hold potential for diabetes management, potentially enhancing patient outcomes.

7. Multi-Responsive Hydrogels

The multi-responsive hydrogel/nanogel has been proposed to enable a response to different triggers such as temperature, light, redox conditions, etc. ^[85][86][107,108]. This is crucial for achieving precise control over drug release. In a study by Gao et al., a random copolypeptide was designed through ring-opening copolymerization, incorporating poly(methoxy-diethylene glycol–L-glutamate)-co-poly(S-(o-nitrobenzyl)-L-cysteine). This copolypeptide displayed quadruple thermo–photo–redox-responsive self-assemble behavior, forming nanogels in water ^[87][109]. These nanogels demonstrated excellent biocompatibility and degradability, making them promising for DDSs and tissue engineering scaffolds. Another example by Jo et al. involved the development of a smart hydrogel responsive to multiple stimuli, including pH, reducing agents, oxidizing agents, and NIR irradiation. The hydrogels showed a rapid release of doxorubicin (DOX) in acidic conditions (pH 5), with reducing agents (10 mmol DTT), in oxidizing medium (0.5% H₂O₂), as well as upon NIR irradiation ^[88][110]. These multiple controlled-release mechanisms enhance targeted drug delivery and help mitigate potential side effects.