Stimulus-responsive MOF–hydrogel composites: Classification, preparation, characterization, and their advancement in medical treatments

2.1.1 Physical stimuli

2.1.1.1 Light-responsive MOFs

Light-mediated therapy is recognized for its non-invasive nature and precise control over time and space, making it a valuable therapeutic and diagnostic tool for specific areas both in vitro and in vivo using light at a particular wavelength. Diring et al. [58] introduced a novel method for releasing carbon monoxide (CO) using an MOF. They achieved this by embedding a manganese carbonyl complex, MnBr(bpydc)(CO)3 (bpydc = 5,5′-dicarboxylate-2,2′-bipyridine), into the walls of a durable Zr(iv)-based MOF via a post-synthetic immobilization process. CO release was triggered by visible light (wavelength of 460 nm and intensity of 300 W) and monitored using infrared and UV–Vis spectroscopy. Upon irradiation, the disappearance of the characteristic CO stretching vibration indicated changes in the Mn(i) coordination sphere, marking the release of CO (Figure 2). This light-responsive MOF is designed for precise targeting of tissues or organs, potentially revolutionizing tissue-specific therapy by providing localized CO release with anti-inflammatory benefits, while avoiding the toxicities associated with systemic CO administration in the bloodstream and lungs [58].

Figure 2

Schematic representation of MnBr(bpy)(CO)₃ loading onto UiO-67-bpy to synthesize CORF-1. Upon light irradiation, the loaded complex undergoes CO release. Reuse under terms of the CC-BY license [58]. Copyright 2016 Royal Society of Chemistry.

2.1.1.2 Temperature-responsive MOFs

Temperature-mediated therapy is used in oncology treatment due to its remote controllability and non-invasiveness. Jiang et al. [59] reported an isoreticular zirconium-cluster-based MOF denoted as ZJU-801 and applied it as a drug delivery system using heat to activate its drug release. The encapsulated diclofenac sodium (DS) drug (DS@ZJU-801) was released when the temperature was raised. The temperature rise caused the breaking of the π–π interactions between MOF and drug molecules (Figure 3). The release rate was reported to be ten times higher at 60°C than at 25°C due to the breakdown of π–π interactions between the MOF and drug molecules [59].

Figure 3

Schematic representation of π–π interactions between DS and MOF. As the temperature increased, the breaking down of π–π interactions released DS.

2.1.2 Chemical stimuli

2.1.2.1 pH-responsive MOFs

Compared to normal tissue, extracellular tumor tissues tend to exhibit a slightly lower pH because of their abnormal cell growth, whereby the cancerous cells consume nutrients and induce carbon dioxide generation, causing a decrease in the local pH of the tumor microenvironment. During the synthesis of MOFs, acidic or basic reaction conditions play a crucial role in the formation of MOFs in influencing the crystallization and coordination of the MOFs. Therefore, the coordination bonds of the MOFs can be modified to detect external pH [60].

Zhu et al. [61] reported on iron-based MIL-100 MOFs formed by a multifunctional core–shell with polypyrrole (PPy) nanoparticles (NPs) as cores and iron(iii) carboxylate MOFs (MIL-100) as the outer shell. MIL-100 was loaded with the anticancer drug, doxorubicin (DOX) by coating the drugs on its outer shell. Over 80% of DOX was released from MIL-100 at pH 5.0. The acidic environment caused gradual degradation of the MIL-100 structure, resulting in the weakening of the electrostatic interactions between DOX and MIL-100 shell (Figure 4) [61].

Figure 4

Schematic representation of multifunctional NPs composed of a PPy core and a mesoporous MIL-100 shell [61] Adapted with permission from Zhu et al. [61]. Copyright 2016 American Chemical Society.

2.1.2.2 Redox-responsive MOFs

Tumor cells flourish in a reducing environment primarily controlled by the oxidation and reduction states of glutathione (GSH) and nicotinamide–adenine dinucleotide phosphate. This is because GSH controls the cellular reducing environment through fragmentation and disulfide bond formation, hence causing an increase in GSH concentration, which is used to regulate the tumor microenvironment. Due to these unique features, drug delivery systems have been designed to detect the reducing environment caused by the tumor cells to trigger drug release via the cleavage of the disulfide bond in GSH-sensitive materials [62]. Lei et al. [63] reported a redox-responsive MOF carrier using metal nodes such as aluminum, iron, or zirconium and 4,4′-dithiobisbenzoic acid (4,4′-DTBA) as the organic ligand where the disulfide bond was cleaved by GSH which is often overexpressed in tumor cells (Figure 5). MOF-Zr(DTBA) with zirconium as the metal ion was identified as a potential carrier for curcumin (CCM). Drug release of CCM-MOF-Zr(DTBA) and CCM@UiO-67 was demonstrated in vitro by exposing them to DL-dithiothreitol that mimics the tumor microenvironment. CCM-MOF-Zr(DTBA) shows a superior release behavior and enhanced cell death compared to CCM@UiO-67, indicating GSH responsiveness in the tumor microenvironment due to the cleavage of the disulfide bond, which accelerated the drug release from MOF-Zr(DTBA) [63]. In summary, the single-responsive MOFs for biomedical applications are driving targeted therapy. These MOFs, which are sensitive to light, temperature, and redox triggers, offer precise drug delivery. Light-responsive MOFs enable localized therapy, bypassing systemic toxicities. Temperature-triggered systems like ZJU-801 facilitate controlled drug release, which is beneficial in oncology. Redox-responsive MOFs like MOF-Zr(DTBA) exploit biomolecules in tumors for triggered drug release. These advancements are demonstrating a shift toward targeted therapy in order to minimize damage to healthy cells.

Figure 5

Redox-responsive degradation of curcumin-loaded MOF (CCM@MOF-M(DTBA)) in tumor cells. Adapted with permission from Lei et al. [63]. Copyright 2018 American Chemical Society.

2.1.3 Dual- and multi-responsive MOFs

To enhance the development of cancer therapies, dual- and multi-responsive stimuli MOFs have emerged to give dual or multiple triggers rather than a single stimulus response due to the complex human body environment. Hence, by harnessing the synergistic effects of different stimuli, such as pH, temperature, light, and redox conditions, dual- or multi-responsive MOFs can provide enhanced control and precision in drug release, leading to more effective and targeted cancer treatments.

Nagata et al. reported dual pH and temperature-responsive MOFs for drug release via surface modification of MOFs using a copolymer of acrylic acid (AA) and N-isopropyl acrylamide (NIPAm) in the post-synthetic process [64]. The polymer exhibited both pH and temperature responsiveness, which allowed the release of procainamide from the MOF in an on–off manner. At pH 6.8 or temperature lower than 25°C, the polymer exhibited a fast release of procainamide. However, at pH 4 or higher temperature (>40°C), procainamide molecule release was suppressed, thereby demonstrating the MOFs were pH and temperature sensitive. This dual stimuli MOF was proposed to be used in controlled drug delivery by responding to the changes in pH and temperature during the therapy.

Jiang et al. [65] reported a dual ATP/pH-responsive MOF (ZIF-90), synthesized by a fast assembly process. ZIF-90 was loaded with DOX and conjugated to neuropeptide Y Y1 subtype receptor, as a new target site for breast cancer treatment via surface modification. The modified ZIF-90 reduced the premature DOX release and effectively triggered faster DOX release in the tumor cells in response to low pH conditions and high ATP in the tumor cells. By combining targeted delivery of DOX and dual-responsive DOX release, the study reported an 80% survival rate in MDA-MB-231 tumor-bearing mice after 40 days of treatment with minimal side effects on the liver and renal functions. The modified ZIF-90 was proposed to be used as a targeted triple-negative breast cancer treatment [65].

Tan et al. also reported on a multi-Ca²⁺/pH and thermo triple-responsive mechanized Zr-based MOFs for drug delivery in bone disease [66]. The loading of the drug 5-fluorouracil (5-Fu) into the CP5-based mechanized UiO-66-NH2 nanocarrier system was carried out by introducing the drug into the nanopores of the scaffold at room temperature. This was followed by the introduction of CP5 rings via host–guest complexation, forming pseudorotaxanes as the movable elements of the mechanized Zr-MOFs, thus achieving drug encapsulation. The drug release from the mechanized MOFs can be triggered by changes in calcium ion concentration (Ca²⁺), pH, and temperature. These triggers caused changes in the MOF structure or environment that resulted in the release of the encapsulated drug. The addition of Ca²⁺ ions led to an increase in the encapsulation capacity. The release was triggered in an acidic environment where the neutralization of CP5 sodium salts weakened the host–guest interactions, leading to the unblocking of the nanopores. Elevating the temperature to 60°C weakened these interactions and promoted the gradual release of 5-Fu. At normal body temperature or room temperature, the premature release of the drug was not significant. This aided the targeted delivery of the drug to the appropriate site, such as a tumor subjected to hyperthermia treatment. Overall, these reported stimuli-responsive MOFs contributed to a controlled and targeted drug delivery system capable of responding to specific physiological conditions or treatment protocols.

To summarize, the emergence of dual- and multi-responsive MOFs represents a significant step forward in cancer therapy, as compared to single-responsive MOF. The dual- and multi-responsive MOFs offer promising avenues due to their tunable properties in response to various stimuli such as pH, temperature, light, redox, and even combinations of stimuli-responsive MOFs such as pH/temperature-responsive and ATP/pH-responsive conditions. The dual- and multi-responsive MOFs present enhanced precision in targeted drug release for oncology therapy. The multi-responsiveness of MOFs holds potential for advancing combined therapies, such as chemoradiation therapy, complementing localized irradiation for a synergistic treatment effect.

2.2.1 Direct precipitation synthesis

Stimulus-responsive MOFs can also be synthesized via direct precipitation synthesis. The responsiveness can be imparted using a stimulus-responsive ligand [23]. In this method, MOFs are built using an organic ligand containing ionizable chemical groups that have the potential to undergo protonation and exhibit charge reversal when exposed to acidic conditions. Wang et al. [15] reported a synthesized ZIF-8 loaded with 5-FU (anticancer drug) that exhibited a pH-triggered controlled drug release property. ZIF-8 was synthesized by dissolving Zn(NO₃)₂·6H₂O in deionized water and 2-methylimidazole in methanol. The zinc nitrate solution was added to the 2-methylimidazole solution while stirring at room temperature. The mixture was then mechanically stirred at 500 rpm for 30 min. ZIF-8 was collected by centrifugation, washed with methanol, and dried overnight in an oven at 80°C [15].

2.2.2 Surface modification

MOFs that are not inherently responsive to changes in the environment can be modified via surface modification by functionalizing the surface of MOFs with amine groups [72], responsive polymer materials such as chitosan [75], carboxymethylcellulose (CMC) [76], pectin [77], or other photo-responsive material [78] for biomedical applications.

Javanbakht et al. reported on the development of oral vehicles by coating the nano-encapsulating MOFs with a biopolymer. The bio-nanocomposite beads functioned as pH-responsive oral drug delivery systems for anti-inflammatory purposes. This method utilized a combination of different biopolymers, including gelatin, CMC, and κ-carrageenan (κ-Cr), along with various MOFs such as UiO-66 and Cu-MOF. When subjected to the conditions of the gastrointestinal tract, the biopolymer-coated MOFs exhibited responsiveness corresponding to the stomach’s acidic pH. Additionally, the biopolymer-coated MOFs contributed to the enhanced stability of the drug molecules, leading to improved therapeutic efficacy [78–82].

3.1.1 Direct mixing method

Most of the reported stimulus-responsive MOF–hydrogel composites were prepared by the direct mixing method due to convenience and cost-effectiveness. In this approach, the preparation of hydrogels is altered by incorporating stimulus-responsive MOF particles into the liquid precursor or colloidal solution prior to gelation, as shown in Figure 7a. Following with gelation of the hydrogel matrices, the stimulus-responsive MOF particles are trapped within the hydrogel matrices to form stimulus-responsive MOF–hydrogel composites [106,107]. Javanbakht et al. [78] have successfully prepared a bio-nanocomposite carrier for oral delivery of the anticancer drug, 5-Fu, where 5-fluorouracil@MOF-5 NPs were coated with CMC via direct mixing method. The blend of CMC and 5-FU@MOF-5 was introduced into distilled water and stirred at room temperature. To crosslink the resulting mixture, it was gradually added dropwise into 100 mL of ferric chloride hexahydrate solution (0.01 M) [78].

Figure 7

Preparation of MOF–hydrogel composite via (a) direct mixing and (b) in situ synthesis.

3.1.2 In situ MOF synthesis

In situ synthesis is another commonly used preparation method for MOF–hydrogel composites whereby the hydrogel is first synthesized as the supporting material for the MOFs. The MOF crystals are grown inside the pores of the hydrogels by sequential adding in the metal precursor and organic linker to obtain the MOF–hydrogel, as shown in Figure 7b [106]. The in situ preparation method is employed to overcome the challenges associated with processing MOF particles into specifically shaped structures for practical industrial applications [108]. Zheng et al. [92] developed a porous ZIF-8 (HZIF-8) with pH responsiveness and antibacterial activity using a bicomponent hydrogel as a template through an in situ synthesis method. They transferred the transparent hot gel solution into 280 mL of preheated n-hexane, which contained 10 mL of Span 85 and stirred the mixture at 1,000 rpm for 1 h to emulsify it. The emulsion was then stirred at 0°C under mechanical stirring at 300 rpm for 1 h to induce gel formation. To the resulting gel, 20 mL of a 0.4 mol/L Hmim (1-hexyl-3-methylimidazolium) solution was added and stirred for 12 h. Afterward, an equal volume of ethanol was added to the sample to promote demulsification, and the resulting product was collected by centrifugation. The collected ZIF-8/gel was suspended in water at 70°C for 3 h, washed with methanol, and dried under vacuum at 80°C for 24 h [92].

3.2.1 Natural-polymer-based hydrogel

Natural-polymer-based hydrogel is biocompatible, biodegradable, and has good cell adhesion properties. The natural polymers used in the preparation of stimulus-responsive MOF–hydrogel composites are classified into protein-based hydrogel and polysaccharide-based hydrogel. The protein-based hydrogel can be prepared using gelatin, while polysaccharide-based hydrogel can be prepared from alginate, cellulose, pectin, chitosan, and carrageenan [110].

3.2.1.1 Gelatin hydrogel

Gelatin (Figure 8a) can be derived from various natural sources and modified to produce hydrogels with excellent cytocompatibility and desirable biodegradable characteristics [111]. Gelatin-based hydrogels find applications in tissue engineering and drug delivery owing to their capacity to enhance cell adhesion and proliferation. Moreover, these hydrogels are suitable for wound dressing applications due to their impressive fluid absorbance properties [112]. For instance, Du et al. successfully synthesized a porphyrin MOF (PCN) encapsulated with photothermal-sensitive SNP modified with gold NPs to form a near-infrared (NIR) light-triggered stimulus-responsive MOF. The MOF was incorporated into a nontoxic gelatin–chitosan hydrogel via direct synthesis method [91].

Figure 8

General chemical structure of protein-based hydrogel used in stimulus-responsive MOF–hydrogel composites: (a) gelatin, (b) alginate, (c) pectin, (d) carrageenan, and (e) chitosan.

3.2.2 Alginate hydrogel

Alginate (Figure 8b) is a long-chain hydrophilic polysaccharide that is found in seaweeds, where it exists within the cell wall of the seaweed to provide strength and flexibility. It is a well-known natural polymer that can form hydrogel via ionotropic gelation under mild conditions [113]. Liu et al. [90] demonstrated a liquid metal micro/nanocomposite by combining eutectic gallium indium (EGaIn) LM NPs (NPs) and ZIF-8 NPs into the alginate hydrogel using an in situ synthesis approach. The resulting composites were used as light-activating nanofillers with functionalized light stimulus-responsive MOFs in the alginate-based hydrogel. The MOF–hydrogel was reported to exhibit physiological responsive sol–gel transformation, good injectability, controllable release of zinc ions, and stable photothermal performance for antitumor and antibacterial biomedicine [90].

3.2.3 Cellulose hydrogel

Cellulose is a hydrophobic polysaccharide found in natural fibers and plants such as linen and cotton. Specific bacteria such as Acetobacter xylinum can biologically synthesize cellulose [110]. Cellulose possesses unique properties such as biocompatibility, biodegradability, and insolubility in most solvents. Cellulose hydrogel has been a popular candidate for wound treatment and drug delivery applications due to its nontoxic nature [114]. Zhang et al. [86] reported a Cu@ZIF-8 MOFs biomimetic nanoreactor by assembling glucose oxidase (GO _x ) onto the Cu ions doped graphitized imidazolate framework (ZIF-8), imparting glucose responsiveness to the overall Cu@ZIF-8/GO _x . The resulting Cu@ZIF/GO _x complex was encapsulated within a biocompatible hybrid hydrogel composed of bacterial cellulose (BC) and guar gum. The sustained release of GO _x catalyzed the decomposition of glucose in the blood and triggered the in situ MOF-mediated catalytic activity by generating ˙OH radicals for bacteria eradication. Additionally, Cu@ZIF/GO _x encapsulated in the hydrogel displayed a high swelling ratio for water absorption, excellent biocompatibility, and rapid hemostatic properties [86].

3.2.4 Pectin hydrogel

Pectin (Figure 8c) is an anionic, water-soluble polysaccharide that can be found in the cell walls of most plants. It can be extracted from different types of fruits through chemical or enzymatic methods. Due to its unique properties, such as biocompatibility, biodegradability, bioactivity, stabilizing, and gelling abilities, pectin has piqued the interest of researchers to further explore its medical and biomedical applications in wound healing, tissue engineering, drug delivery, and gene delivery [115,116]. Nabipour and Hu [79] synthesized a bio-MOF encapsulated with curcumin. The curcumin@bioMOF-11 was then functionalized by pH-sensitive pectin biopolymer to form a pH-responsive-MOF hydrogel and used as an anticancer drug carrier for controlled delivery [79].

3.2.5 Carrageenan

Carrageenan (Figure 8d) is extracted from a type of red algae called Irish Moss known as Chondrus crispus. It is a natural, linear, sulfated polysaccharide that exhibits unique gelling properties, with a strong negative charge, water absorption, and multiple functional groups, making it an ideal candidate for controlled and sustained drug delivery applications as well as wound healing and tissue engineering [117]. Feng et al. [89] synthesized a light and temperature-responsive hydrogel comprising poly(N-isopropylacrylamide) (PNIPAm) and methacrylated-κ-carrageenan (MA-κ-CA), incorporating PPy-polydopamine NPs (PPy-PDA NPs) and ZIF-8. The ZIF-8-hydrogel generated localized heat and gradually released Zn²⁺ to ensure safe and effective synergetic photothermal–chemical bactericidal capability upon application of NIR light [89].

3.2.6 Chitosan hydrogel

Chitosan (Figure 8e) is derived from chitin after undergoing the deacetylation process. It can be found as a component in the skeletons of invertebrates. It is a semicrystalline, cationic polysaccharide and is considered one of the most abundant natural biopolymers after cellulose [110]. Chitosan has great biocompatibility making it an ideal candidate in biomedical applications such as wound healing material. Chitosan hydrogel has been reported to possess mild antibacterial properties. Therefore, to improve its antibacterial property, the chitosan can be chemically modified to accurately design into a hydrogel structure [118].

Han et al. [18] prepared a photosensitive MOF–hydrogel composite composed of Prussian blue NPs (PBNPs) and chitosan using the direct mixing method. Chitosan was chemically modified with quaternary ammonium and a carbon–carbon double bond (C═C) to impart the hydrogel with an electropositive surface and robust structure. Within the hydrogel, PBNPs served as crosslinking points, absorbing polymers and reinforcing the structure, thereby enhancing the overall mechanical properties of the NIR light-responsive MOF-based hydrogel composite. Under NIR light irradiation, the stimulus-responsive MOF-based hydrogel composite exhibited antibacterial properties through the synergistic effect of heat and the electropositive surface of the chemically modified chitosan hydrogel. The presence of PBNPs facilitated NIR light absorption and light production through charge transfer between Fe³⁺ and Fe²⁺ within the PBNP structure [18].

3.3.1 PAM hydrogel

PAM (Figure 9a) is a synthetic polymer of the acrylamide monomer. It is a colorless hydrogel material that is nontoxic, stable, non-immunogenic, and non-resorbable with hydrophilic, cohesive, viscoelastic, and biocompatible properties [120]. Tian et al. introduced a PAM hydrogel wound dressing incorporating a bimetallic MOF (Fe-Cu-MOF) loaded with GO _x , termed MOF(Fe-Cu)/GO _x -PAM gel. This PAM hydrogel offers an effective cascade-catalyzed system for accelerating wound healing, leveraging synergistic antibacterial and anti-inflammatory effects provided by the Fe-Cu-MOF. The authors noted that the catalytic activity of the bimetallic MOF(Fe-Cu) is approximately five times greater than that of the monometallic MOF(Fe), attributed to charge transfer from Cu(i) to Fe(iii). This accelerates the regeneration of Fe(ii) and enhances the subsequent production of hydroxyl radicals (˙OH) from the dissociation of H₂O₂. Within this hydrogel cascade-catalyzed system, an abundant supply of gluconic acid and H₂O₂ is continually generated by the breakdown of glucose through the enzyme GO _x . This gluconic acid significantly enhances the peroxidase performance of MOF(Fe-Cu), enabling efficient decomposition of H₂O₂ to achieve effective antibacterial properties [100].

Figure 9

General chemical structure of synthetic polymers used in stimulus-responsive MOF–hydrogel composites: (a) PAM, (b) poly(polyethylene glycol-co-N-isopropylacrylamide) (PPCN), and (c) poly(d,l-lactide-coglycolide)-poly(ethy-leneglycol)-poly(d,l-lactide-coglycolide) (PLGA-PEG-PLGA).

3.3.2 PPCN hydrogel

PPCN (Figure 9b) is a biodegradable, temperature-responsive gel with intrinsic antioxidant properties. Temperature-responsive polymers are sensitive toward the balance of hydrophilic and hydrophobic groups as well as charge interactions, leading to a phase transition whenever there are temperature changes. PPCN can be synthesized through copolymerization by sequential polycondensation and radical polymerization by reacting citric acid, poly(ethylene glycol) (PEG), and poly-N-isopropylacrylamide (PNIPAAm) [121]. Xiao et al. [44] embedded copper-MOF NPs (HKUST-1 NPs) within an antioxidant temperature-responsive citrate-based hydrogel (PPCN). The incorporation of the HKUST-1 NPs into the hydrogel accelerated the wound healing of the diabetic mice and reduced the copper ion toxicity. The PPCN prevented the HKUST-1 NPs from disintegration, which also enabled the sustained release of copper ions without causing copper ion toxicity [44].

3.3.3 Poly(d,l-lactide-coglycolide)-poly(ethy-leneglycol)-poly(d,l-lactide-coglycolide) (PLGA-PEG-PLGA)

PLGA-PEG-PLGA (Figure 9c) is a triblock copolymer that can act as temperature-responsive copolymers, also known as thermogel. The block copolymers are composed of hydrophobic PLGA segments and hydrophilic PEG segments. The copolymer solution can form a high-viscosity gel at body temperature due to the weakening of the hydrogen bond of the PEG segments. At low temperatures, PLGA-PEG-PLGA is a free-flowing solution due to the interaction of PEG hydrophilic segments with water molecules. PLGA-PEG-PLGA has been heavily studied by researchers in clinical applications due to its biodegradability and biocompatibility [122]. Tan et al. [96] have prepared a hybrid nanocomposite in which MOFs were modified with temperature-sensitive PLGA-PEG-PLGA hydrogel to devise an injectable implant. DOX and celecoxib (Cel) were co-loaded into the MOF–hydrogel system for localized oral cancer therapy (DOX/Cel/MOFs@Gel). This formulation exhibited a high drug loading capacity, stable and pH-responsive release of both drugs, and enhanced toxic effects against oral cancer cells [96].

3.3.4 N-isopropylacrylamide/poly(ethylene glycol) dimethacrylate (NiPAm/PEGMA)

Poly(N-isopropylacrylamide) (PNiPAm) polymer-based hydrogels are recognized for their temperature-responsive behavior, particularly in proximity to a lower critical solution temperature (LCST) [123]. Hence, it demonstrates excellent lubricating performance and temperature-sensitive tribological properties [124]. Furthermore, by integrating hydrophilic polymer chains like PEG into the PNiPAm matrix, the LCST of the PNIPAm microgels can be elevated [123]. PEG is used for surface modification to enhance colloidal stability as well as reduce cytotoxicity in drug delivery systems [125]. Wu et al. [104] reported a MIL-101(Cr)@PEG-g-PNIPAm hybrid by growing the poly(ethylene glycol)-graft-poly(N-isopropylacrylamide) (PEG-g-PNIPAm) microgel layer using NIPAm, PEGMA, and N,N′-Methylenebisacrylamine (MBA) as the hydrogel materials on the MIL-101(Cr) surface with 2,2′-azobis(2-methyl-propionamidine) dihydrochloride (V50) to initiate polymerization via a one-pot soap-free emulsion method. At the osteoporosis site (∼40°C), the temperature exceeds the LCST of the MIL-101(Cr)@PNIPAm, which compromises its lubricating performance and affects drug release. Additionally, the colloidal stability of MIL-101(Cr)@PNIPAm is influenced by the transition of PNIPAm chains from hydrophilic to hydrophobic above the LCST. Therefore, the newly synthesized MIL-101(Cr)@PEG-g-PNIPAm hybrid, with the addition of PEG, elevates the LCST and stability of the microgel by creating a robust hydration layer around the ethylene glycol groups through hydrogen-bonding interactions with adjacent water molecules [104].

5.1.1 Single-responsive MOF–hydrogels

The reported single-responsive MOF–hydrogels were mostly pH-responsive MOF–hydrogels for controlled release of ibuprofen, fluorouracil, exendin-4, itaconate, and natural polyphenolic compounds, such as curcumin. Nabipour and Hu [79] reported an MOF, namely bioMOF-11, synthesized from cobalt acetate tetrahydrate and adenine. The bioMOF-11 was encapsulated with curcumin followed by functionalizing it using a pH-sensitive pectin biopolymer to protect the drug formulation against the harsh conditions of the gastrointestinal tract (GIT). The release of 100% curcumin from curcumin@bio-MOF-11 was observed within 90 min due to the decomposition of bio-MOF-11 at pH 1.2. When curcumin@bio-MOF-11 was prepared in pectin hydrogel composite form, the controlled release rate was found to be 93% over 410 min at different pHs of 1.2, 6.8, and 7.4. The observed improvement in release behavior was attributed to the pH sensitivity of pectin, as shown in Figure 12. Furthermore, pectin functions as a protective layer, aiding in regulating the release of curcumin [79].

Figure 12

Curcumin release profile of cucurmin@bio-MOF-11 and pectin/cucurmin@bio-MOF-11 hydrogel. Reuse under terms of the CC-BY license [79]. Copyright 2022 Springer Link.

Javanbakht et al. reported on κ-Cr and CMC used for surface modification on three different MOFs, namely, UiO-66, Cu-MOF, and MOF-5 [82]. It was reported that the drug release of carrageenan/tramadol@UiO-66 at pH 1.2 was the lowest because of the attraction between UiO-66 and tramadol at acidic pH, which hindered the drug release. However, as the pH increased to 7.4 (simulating the intestinal fluid), the drug release rate dramatically increased. Ion-exchange between the buffered solution ( PO 4 3 − ) and tramadol is related to the sufficient diffusion of the anions from the buffered solution into the κ-Cr network, which triggered the release of tramadol@UiO-66 from the κ-Cr matrix, resulting in the diffusion of tramadol from the κ-Cr hydrogel bead into the aqueous medium. Tramadol@UiO-66, with its well-organized interaction with tramadol, led to a more controlled release of the drug under gastrointestinal conditions. The authors also reported the pH-sensitive nature of CMC biopolymer served as a protective layer against digestion within the stomach [77,78,81].

In another study, Wu et al. reported thermosensitive microgel nanocarriers of PEG-g-PNIPAm copolymer functionalized on nanoMOFs (MIL-101 (Cr), namely MIL-101(Cr)@PEG-g-PNIPAm) [104]. This thermosensitive MIL-101(Cr)@PEG-g-PNIPAm-hydrogel composite was loaded with DS, aiming for the treatment of osteoarthritis. It was found that an initial rapid drug release was observed in both DS-MIL-101(Cr) and DS-MIL-101(Cr)@PEG-g-PNIPAm within the first 4 h, followed by a gradual release over time (Figure 13). After 72 h, DS-MIL-101(Cr) released approximately 67.4% of the drug at 37°C, with a similar release profile at 32°C. In contrast, DS-MIL-101(Cr)@PEG-g-PNIPAm exhibited reduced drug release, with approximately 43.1% released at 32°C and 36.6% at 37°C. This reduction was attributed to drug loss during the polymerization and washing processes. As the temperature increased to 42°C, the cumulative drug release of the DS-MIL-101(Cr)@PEG-g-PNIPAm further decreased to 31.7%, likely due to the partial collapse of PEG-g-PNIPAm microgel layers, which restricted drug release from the MOFs. Therefore, improving the anti-inflammatory drug-delivery efficiency in a controlled manner. Overall, the investigation revealed the temperature-dependent drug release profile of these materials, demonstrating the potential of DS-MIL-101(Cr)@PEG-g-PNIPAm as a controlled drug delivery system influenced by temperature variations.

Figure 13

(a) Release profile of DS-MIL-101(Cr) in PBS at 32, 37, and 42°C. (b) DS-MIL-101(Cr)@PEG-g-PNIPAm in PBS at 32, 37, and 42°C. Reuse under terms of the CC-BY license [104]. Copyright 2023 Elsevier.

Apart from the physical stimulus responsiveness, Chen et al. [101] developed an ATP-responsive nucleic acid-based PAM hydrogel functionalized on DOX-loaded UiO-68 nano-MOF (Figure 14). In the presence of ATP, which is known to be overexpressed in cancer cells, the nucleic acid-based PAM hydrogel coating underwent degradation through the formation of the ATP–aptamer complex. This degradation process led to the release of the DOX. The resulting hydrogel-coated UiO-68 nanoMOFs also exhibited a high degree of drug loading with a loading capacity of DOX at 79.1 nmol mg⁻¹ and low background leakage (∼28%) of the drug loads as compared to the non-hydrogel-coated nanoMOFs. The cytotoxicity of the DOX-loaded hydrogel-coated UiO-68 nanoMOFs toward cancer cells was substantially higher (∼40%) as compared to the cytotoxicity of the non-hydrogel-coated nanoMOFs toward cancer cells (∼25%), indicating the efficiency of the nucleic acid-based PAM hydrogel.

Figure 14

(a) Illustration of the synthesis of ATP-responsive DNA@PAM hydrogel NMOFs loaded dye/drugs. (b) SEM image of the nucleic acid with promoter-functionalized UiO-68 NMOFs before the deposition of the hydrogel. (c) SEM image of the hydrogel-coated NMOFs. Reuse under terms of the CC-BY license [101]. Copyright 2017 John Wiley and Sons.

5.1.2 Dual- and multi-responsive MOF–hydrogels

Dual-responsive MOF–hydrogels consist of the combination of two stimuli to enhance the effectiveness of its applications. The reported dual stimulus-responsive MOF–hydrogels are thermo- and photo-responsive, thermo- and shape-memory-responsive, thermo- and enzyme-responsive, or shape-memory and water-responsive hydrogels (Table 2).

Zeng et al. reported on an injectable pH and ATP-responsive MOF-based hydrogel composite, which possessed structurally dynamic properties such as shear-thinning and self-healing ability [97]. This stimulus-responsive MOF-based hydrogel composite was proposed to be a local depot for the sustained release of DOX drugs at tumor sites. The DOX-loaded MOF-based hydrogel composite was prepared by using a bisphosphonate (BP)-functionalized hyaluronic acid (HA-BP) macromolecule, which was mixed with a suspension of DOX-embedded Zn-MOF (MOF@DOX). The hydrogel network formed through dynamic coordination bonds between the BP group of HA-BP and Zn²⁺ ions on the MOF surface. In the presence of protons, these coordination bonds between Zn²⁺ ions and the imidazole groups in the MOF were cleaved, leading to drug release as the MOF framework collapsed. Additionally, ATP exhibited stronger coordination with Zn²⁺ ions than the imidazole groups, allowing it to competitively bind with Zn²⁺ ions, further promoting the collapse of the MOF and accelerating DOX release (Figure 15). In comparison to the HA-BP·MOF@DOX delivery system, MOF@DOX demonstrated faster drug release kinetics. The HA-BP·MOF@DOX hydrogel exhibited multiple stimuli-responsive drug release performances. For instance, only 2.6% of DOX was released from the HA-BP·MOF@DOX gel in 97 h in PBS of pH 7.4. In contrast, within the same period, 5.7, 13.4, 10.2, and 26.9% of DOX was released from HA-BP·MOF@DOX in mediums with pH 6.0, pH 5.0, pH 7.4 + 2 mmol/L ATP, and pH 5.0 + 2 mmol/L ATP, respectively. These data suggested that the presence of protons in the medium can cleave the coordination bond between Zn²⁺ ions and the imidazole group in the MOF, leading to drug release after framework collapse. Additionally, ATP exhibited stronger coordination with Zn²⁺ ions than the imidazole group, thereby competitively binding with Zn²⁺ ions and expediting MOF collapse and DOX release.

Figure 15

Mechanism of drug release after the cleavage of coordination bond between Zn²⁺ ions and the imidazole group in the MOF [97].

Moreover, compared to the HA-BP·MOF@DOX delivery system, MOF@DOX showed faster drug release kinetics. In 97 h, 4.7, 66.3, 33.1, and 100% of DOX was released from MOF@DOX in mediums with pH 7.4, pH 5.0, pH 7.4 + 2 mmol/L ATP, and pH 5.0 + 2 mmol/L ATP, respectively. The presence of particle-hydrogel superstructures in HA-BP·MOF@DOX inherently slowed down the degradation of MOF@DOX due to changes in proton and ATP molecules of osmotic conditions. This characteristic would prolong DOX release in the tumor microenvironment, even under low pH conditions and high ATP levels.

Reddy et al. [95] reported on a dual-responsive MOF–hydrogel prepared by the in situ growth of a nano-MOF (nZiF-8) embedded into a biocompatible PAM/starch hydrogel (PSH) composite. Photosensitizer, Rose Bengal (RB) (4,5,6,7-tetrachloro-2′,4′,5′,7′-tetraiodofluorescein), was loaded into the nZIF-PSH, and the interaction involving surface adsorption was observed. RBZIF-8 and RB-loaded PSH exhibited a burst release profile, rapidly releasing a significant portion of RB at a pH of 5.5 (∼90% drug release). In contrast, RBZPSH displayed a sustained release pattern over 18 h. This sustained release from RBZPSH was attributed to the π–π interactions between the imidazole ring of the nZIF-8 structure and the phenyl rings of RB embedded within the IPN of PSH. Additionally, ionic interactions between the halogen atoms (chlorine and iodine) of RB and zinc atoms within the nZIF-8 also contributed to the sustained release of RB from ZPSH. RBZPSH demonstrated significantly greater photodynamic antifungal activity compared to its monometallic counterparts [95].

Liu et al. reported on MOF-based injectable in situ gel used for multi-responsive insulin delivery [93]. The Ins@ZIF-8/GO _x -gel was prepared via a one-pot synthesis method. The Ins@ZIF-8/GO _x -gel exhibited several responsive characteristics. It displayed thermosensitivity as an in situ gel. When introduced into the body, the Ins@ZIF-8/GO _x -gel solution transitioned into a semi-solid state due to the poloxamers that form the hydrogel. This is because when the temperature rises above the LCST, the poloxamers become insoluble and undergo phase separation from the solution, thus forming a hydrogel. In this process, the hydrogen bonding between the poloxamers and solvent is weakened, thereby leading to partial dehydration and aggregation of polymeric chains. As the temperature increases, the gradual removal of water molecules will reveal the hydrophobic domains of the polymer macromolecules, thereby promoting the formation of hydrophobic interactions between polymer macromolecules, known as the hydrophobic effect.

As the temperature surpassed the gelation threshold over time, the gel degraded due to interaction with body fluids, gradually releasing both GO _x and Ins@ZIF-8. Additionally, the ZIF-8 component of the composite possessed pH sensitivity. In hyperglycemic conditions, GO _x and glucose interaction led to the production of gluconic acid, resulting in a localized decrease in pH. While in this acidic environment (pH 5.0), the structure of ZIF-8 collapsed, facilitating the release of insulin. Ins@ZIF-8/GO _x -gel was proposed to be an injectable hydrogel that is responsive to temperature and pH changes, which enabled the controlled and targeted release of insulin, making it a versatile candidate for diabetes treatment and blood glucose regulation.

5.2.1 Single-responsive MOF–hydrogels toward physical stimuli

The reported single-responsive MOF–hydrogels for wound healing were mainly aiming for photothermal therapies, where the MOF–hydrogels were proven to have their light or thermal responsiveness. The light responsiveness of the reported MOF–hydrogels was contributed by light-responsive chemicals, such as SNP, Prussian blue, or EGaIn being loaded or integrated into MOF. The hydrogel played the role of carrier matrix, some with injectable properties to aid in the delivery of the active ingredients and coverage of the wound. Some studies reported that the MOFs possessed antibacterial properties contributing to the wound healing properties of the MOF–hydrogel.

Du et al. developed a photothermal-sensitive SNP loaded into a Pt-modified porphyrin MOF (PCN) with gold particles modified in situ to form an NIR light-triggered phototherapeutic nano platform whereby it converted light into heat to kill bacteria [91]. The system was then incorporated into a hydrogel composed of gelatin and chitosan to form a multifunctional injectable hydrogel, known as PSPG hydrogel. The PSPG hydrogel exhibited remarkable catalase-like behavior, promoting the continuous decomposition of endogenous H₂O₂ into O₂ provided by the platinum metal, which is a catalyst for in situ oxygen production due to its long-lasting catalytic and antioxidant properties. This enhanced the photodynamic therapy effect, specifically under hypoxic conditions. When subjected to NIR irradiation, the PSPG hydrogel generated hyperthermia and simultaneously produced ROS from the Pt-modified porphyrin MOF (PCN) while triggering the release of nitric oxide. These combined effects contributed to the removal of biofilms and disruption of the cell membranes of methicillin-resistant Staphylococcus aureus (MRSA) and Escherichia coli (E. coli) bacteria. The study evaluated the photothermal effect of PSPG hydrogel under 808 nm NIR laser irradiation, showcasing a rise in temperature with increasing concentrations of PCN. Notably, the 250 μg/mL hydrogel achieved a temperature of 65.3°C within 10 min, sufficient for bacterial eradication. The photothermal conversion ability of PSPG hydrogels was observed to increase with higher laser powers, demonstrating excellent photostability even after ten cycles of irradiation. The efficiency (η) of the PSPG hydrogel, calculated at 89.21%, surpassed that of PCN (85.73%), possibly due to enhanced material stability conferred by hydrogel formation. Furthermore, quantification of NO release from PSPG hydrogels revealed a positive correlation with hydrogel concentration and irradiation time, with significantly higher NO release observed under NIR irradiation compared to control conditions.

In another work, Han et al. [18] employed quaternary ammonium and double-bond modified chitosan (QCH) and an MOF NP, i.e., Prussian blue MOF NPs (PBNPs) produced by solvothermal synthesis using K₄[Fe(CN)₆] and FeCl₃. The PBNPs were embedded with the modified chitosan hydrogel to form a photo-responsive hydrogel via free radical polymerization and named QCH/PB. The PBNPs were capable of absorbing 808 nm NIR light and generated light through the charge transfer between Fe³⁺ and Fe²⁺ ions within the structure, leading to an increase in the temperature of the QCH/PB hydrogel. Additionally, the modified chitosan hydrogel formed via quaternary ammonium and a carbon–carbon double bond (C═C) in the QCH/PB hydrogel exhibited strong electrostatic absorption, allowing them to capture bacteria and disrupted the surface potential of the bacterial membrane, inhibiting the normal metabolism of the bacteria and consequently suppressing bacterial respiration leading to an antibacterial ratio of up to 99.97% against Staphylococcus aureus and 99.93% against E. coli. As a result, the combination of photothermal effects from the Prussian blue MOF NPs and the inhibition of bacterial respiration provided by the modified chitosan hydrogel led to highly effective and rapid bacteria eradication [18]. These two studies highlighted the incorporation of photo-responsive MOFs into hydrogels to enhance bacteria eradication by generating hyperthermia upon NIR-light irradiation.

Liu et al. [90] developed a liquid metal micro/nanocomposite through the integration of EGaIn liquid metal NPs (LMNPs) and ZIF-8 MOFs within an alginate hydrogel. The resulting composites LMNPs@MOFs SupraParticles (LMSPs) were used as light-activating nanofillers with a functionalized light stimulus-responsive MOFs within the alginate-based hydrogel. The LMSPs/Alg hydrogel was prepared by injecting an LMSPs/Alg solution into phosphate-buffered saline containing CaCl₂ to mimic gelation in the physiological environment, resulting in the formation of LMSPs/Alg-Ca²⁺ hydrogel upon injection. The LMNPs within the LMSPs enabled precise photothermal therapy (PTT) in the NIR window, providing localized heating for deep tissue penetration. The in vivo formation of the Alg-Ca²⁺ hydrogel at the injection site entrapped the therapeutic LMSPs, which generated localized heat around the nidus, thereby enhancing the efficacy of the PTT. Additionally, ZIF-8 NPs exhibited strong antibacterial effects owing to the toxicity of Zn²⁺ toward bacterial cells. Hence, the MOF shell within LMSPs underwent degradation, releasing bioactive ions (Zn²⁺) to facilitate a combination of chemo/PTT. Moreover, the hydrogels effectively maintained moisture in proximity to the heated wounds, which played a crucial role in expediting the healing of skin wounds through PTT [90].

A temperature-responsive MOF hydrogel has been gaining attention in the field of wound management due to its capability of undergoing temperature-driven phase transitions. Xiao et al. [44] synthesized a copper-MOF NP (HKUST-1 NPs) and embedded it within a thermoresponsive PPCN hydrogel and is named H-HKUST-1. The HKUST-1 and PPCN composite possessed a synergistic effect, where HKUST-1 did not interfere with the PPCN gelation by shielding the positive charge of the stored copper ions, and PPCN protected HKUST-1 from degradation. Hence, copper was released from H-HKUST-1 in a controlled manner, reducing copper ion toxicity. It was hypothesized that the copper ions of the HKUST-1 undergoing coordination with the carboxyl groups present in H₃BTC molecules during the synthesis process. This coordination process effectively decreased the interactions between these carboxyl groups and those within PPCN, thus having limited influence on the gelation process. As a result of these interactions and coordination, H-HKUST-1 formed a stable hydrogel with only minor effects on the original LCST of PPCN. A comparison was made by treating diabetic mice with wound incisions between PPCN and H-HKUST-1 hydrogel. Figure 16 illustrates the quantitative analysis of digital images of wounds over time, demonstrating that wounds treated with H-HKUST-1 healed significantly faster than those treated with PBS, PPCN, and HKUST-1 NPs, particularly on days 7, 14, 21, and 29. Specifically, diabetic mice treated with H-HKUST-1 hydrogel exhibited accelerated healing compared to those treated with PBS, blank PPCN, and HKUST-1 NPs. The authors reported that the H-HKUST-1 hydrogels showed a higher wound closure rate due to the sustained release of noncytotoxic amounts of copper ions. This release promoted angiogenesis, collagen deposition, and re-epithelialization during wound healing, enhancing the overall healing process [44].

Figure 16

The quantitative analysis of wound closure rates of wounds treated with PBS, PPCN, and HKUST-1 NPs and H-HKUST-1. Reuse under terms of the CC-BY license [106] Copyright 2016 John Wiley and Sons.

Wang et al. [103] reported Ni₃(HITP)₂ MOF nanorods synthesized via ultrasonic-assisted direct mixing (65°C) method using nickel acetic acid tetrahydrate, sodium acetate, and 2,3,6,7,10,11-hexaaminotriphenylene hexahydrochloride. The Ni₃(HITP)₂ MOF nanorods with antioxidant and anti-inflammatory effects, being encapsulated within a temperature-sensitive, nontoxic, and phase-changing synthetic hydrogel, Pluronic F127, was applied as an injectable and sprayable medical dressing. Pluronic F127 displayed distinct transitions between solution and gel states in an aqueous solution. In this work, the F127 hydrogel was formulated with a phase transition temperature of 21°C, considering the skin’s surface temperature is above 21°C. When the temperature exceeded this threshold, the aqueous solution transformed into a gel-like state to cover the wound evenly for a better wound-healing effect. Conversely, the solution reverted to its original state when the temperature dropped below 21°C. The Ni₃(HITP)₂/F127 hydrogel exhibited superoxide dismutase-like activity and anti-inflammatory properties, as well as the ability to promote angiogenesis and fibroblast migration due to the incorporation of Ni₃(HITP)₂ MOFs. This is because Ni₃(HITP)₂ MOF nanorods possessed high conductivity that can promote good free radical scavenging effect and reductive amino moiety in the molecular structure as well as abundant Ni–N₄ active sites for electrical conductivity. When skin defects occur, there is a phenomenon in which negative charges present around the skin interact with positive charges within the wound, resulting in the creation of an endogenous electric field. This electric field is believed to play a significant role in directing and guiding the migration of cells toward the wound during the process of wound healing. As a result, the increased conductivity provided by Ni₃(HITP)₂ served to facilitate the transmission of electrical signals within biological tissues, thereby contributing to the promotion of wound healing [103].

5.2.2 Single-responsive MOF–hydrogels toward chemical stimuli

MOF–hydrogels have been reported to respond to chemical stimulation, such as pH changes, ions, or molecular signals, whereby they will undergo reversible structural alterations. These responsive behaviors enable controlled drug release, selective adsorption, or triggered chemical reactions [21]. The adaptability of MOF–hydrogels to changing chemical environments makes them a compelling avenue for wound treatment. The role of hydrogel in the chemical-responsive MOF–hydrogel composites is mainly serving as a carrier matrix for wound coverage and delivery of the active ingredients.

Zhang et al. [86] reported on a Cu-doped MOF nanocatalyst (GG-Cu@ZIF) loaded with GO _x and encapsulated within a BC reinforced hydrogel for antibacterial and hemostatic therapies. The BC reinforced hydrogel serves as a matrix for encapsulating and immobilizing the Cu-doped MOF nanocatalyst (Cu@ZIF) loaded with GO _x . This encapsulation is crucial for maintaining the stability and controlled release of the nanocatalyst within the wound site. The BC/GG-Cu@ZIF/GO _x hydrogel displayed glucose decomposition capabilities and exhibited peroxidase-mimicking (POD). The Cu-doped MOF nanocatalyst with multienzyme mimetic activity loaded with GO _x catalyzed in a physiological glucose environment into H₂O₂ and glucose acid through a self-activated cascaded reaction. The glucose acid further triggered the POD-mimicking activity, which could overcome the physiological pH limit and demand of exogenous H₂O₂ for POD-mimicking activity, producing exogenous •OH for bacteria inactivation (Figure 17). This BC/GG-Cu@ZIF/GO _x hydrogel demonstrated high water absorption capacity, biocompatibility, and rapid hemostatic capability, both in vitro and in vivo. Experimental data revealed that the hydrogel exhibited a water absorption capacity of ∼3710% in a mass ratio of 2:1 of BC:GG. Additionally, in vitro evaluations demonstrated that the materials were cytocompatible and hemocompatible due to high cell viability in NIH-3T3 cells and low hemolysis ratios under 0.4%, confirming the material’s suitability for biomedical applications. Furthermore, hemostatic assays conducted in vivo showed rapid clotting times of within 1 min using a mouse liver hemorrhage model and a mouse-tail amputation model, underscoring the hydrogel’s efficacy in controlling bleeding. The absorption behavior of BC/GG-Cu@ZIF/GO _x hydrogel contributed to its swift hemostasis by absorbing blood, which led to platelets aggregation and subsequently activated the blood coagulation cascades for hemostasis. Therefore, encapsulating Cu@ZIF/GO _x within the hydrogel provided a synergistic effect of antibacterial properties and hemostasis for wound treatment [86].

Figure 17

Catalytic cascaded reaction for bacteria inactivation and hemostasis of the BC/ Cu-doped MOF nanocatalyst/glucose oxidase (BC/GG-Cu@ZIF/GO _x ) hydrogel [86].

ROS acts as secondary messengers to immunocytes and non-lymphoid cells in the wound repair process. ROS regulates angiogenesis (formation of blood vessels) and contributes to the blood perfusion at the wound site. Furthermore, ROS serves as a crucial defender within the host’s immune system, particularly through the action of phagocytes. These specialized cells initiate an intense ROS burst directed at pathogens within wounds, effectively dismantling and destroying them. However, amidst this targeted assault, an overflow of ROS leaks into the surrounding environment, amplifying its bacteriostatic effects. This excess ROS acts as a formidable deterrent, further inhibiting bacterial growth and proliferation in the vicinity [135].

Zhu et al. reported on substance P (SP)-loaded ZIF-8 (SP@ZIF-8) NPs, which were then coated with PEG-TK to make them responsive to ROS [87]. SP is a neuropeptide composed of ten amino acids produced by both immune and neuronal cells. It is primarily located at the junction between the adventitia and media in the muscular layer, in the connective tissues surrounding blood vessels, and in the epidermis and dermis [136]. The SP@ZIF-8 NPs were incorporated into a hydrogel made of sodium alginate and pectin, which was then crosslinked with calcium chloride to create SP@ZIF-8-PEG-TK@CA dressings. The healing capabilities of SP@CA and SP@ZIF-8-PEG-TK@CA were assessed in vivo using an infected cutaneous wound mouse model. Treatment with SP@ZIF-8-PEG-TK@CA significantly reduced the wound area, showing a 12.6% reduction after 3 days compared to a 5.9% reduction in the SP@CA group. By day 7, SP@ZIF-8-PEG-TK@CA had reduced the wound area by 62.5%, compared to 42.5% in the SP@CA group. After 15 days, nearly complete wound closure (almost 100%) was achieved with SP@ZIF-8-PEG-TK@CA, significantly higher than in the SP@CA group.

Immunofluorescence staining revealed an abundance of CD32-positive cells at the injured sites treated with SP@ZIF-8-PEG-TK@CA, indicating an inflammatory response and accelerated wound healing. Furthermore, higher numbers of CD206-positive cells, indicative of M2 macrophage polarization associated with tissue repair, were observed at the late stage in mice treated with SP@ZIF-8-PEG-TK@CA compared to SP@CA dressings. These results suggested that SP@ZIF-8-PEG-TK@CA in hydrogel form promoted wound healing by stimulating M2 macrophage polarization [87].

5.2.3 Dual-responsive MOF–hydrogels

Dual-responsive MOF–hydrogel composites offer several advantages compared to single-responsive MOF–hydrogel composites due to their ability to respond to two different stimuli in the complex bioenvironment. For instance, Zhao et al. designed a dual light and temperature-responsive MCA-NI-AA/NP/ZIF8 hydrogel composed of PNIPAm and methacrylated κ-carrageenan (MA-κ-CA) incorporated with PPy-polydopamine NPs (PPy-PDA NPs) and ZIF-8 MOF, which offered both shape adaptability and synergistic bacterial elimination for wound healing application. The incorporation of PPy-PDA NPs into the MCA-NI-AA/ZIF8 hydrogel makes it a photothermal agent that converts photon energy into heat energy upon NIR-light irradiation, while the in situ formation of ZIF-8 in the hydrogel allowed the controlled release of Zn²⁺ ions over time. Through prolonged Zn²⁺ release, and when exposed to NIR irradiation, it generated localized heating and accelerated the release of Zn²⁺, thereby enhancing its therapeutic potential for treating highly infected wounds [137].

Nie et al. reported on light and temperature-sensitive polyMOF hydrogel incorporated with curcumin into a copper-tetrakis(4-carboxyphenyl)porphyrin MOF (Cu-TCPP MOF) and a poly(ethylene glycol)-poly(ε-caprolactone-co-lactide) hydrogel (known as CuPP-PELA polyMOF hydrogel) via in situ ring-opening polymerization [99]. Copper porphyrin (CuPP) served as a photothermal agent in the Cu-TPP MOF, where it was formed by the coordination of Cu²⁺ ions with porphyrin, a well-known photosensitive precursor that exhibits NIR absorption characteristics due to the d–d energy band transition of Cu²⁺ ions. When exposed to visible light, Cu-TCPP experienced a substantial local temperature increase, enabling the escape of water molecules from the voids/pores without causing damage to the crystal structure or disrupting proton conduction. Conversely, in the absence of light, water molecules re-entered the voids/pores, leading to reversible changes [138]. When CuPP-PELA polyMOF hydrogel was irradiated by NIR irradiation, the temperature of the CuPP-PELA polyMOF hydrogel increased due to CuPP. Curcumin was then released at a higher dose or released gradually post-irradiation at a lower dose, which functioned as both a sensitizer for PTT and a regulator for inflammation. As for the hydrogel, owing to its temperature-sensitive sol–gel transition properties, the hydrogel responded to body temperature and underwent a transition from a solution to a hydrogel state within the body, allowing it to adhere to the wound and form a complete seal. The study demonstrated the temperature of Cur/CuPP-PELA at the wound surface reached approximately 48.6°C within 5 min of NIR laser radiation. This significant increase in temperature suggested that the high efficacy of Cur/CuPP-PELA in combating pathogenic bacteria and infected surrounding tissue during PTT.