Stimuli-responsive Hydrogels: Smart State of-the-art Platforms for Cardiac Tissue Engineering

doi:10.21203/rs.3.rs-2011475/v1

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Stimuli-responsive Hydrogels: Smart State of-the-art Platforms for Cardiac Tissue Engineering

https://doi.org/10.21203/rs.3.rs-2011475/v1

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Biomedicine and tissue regeneration have made significant advancements recently, positively affecting the whole healthcare spectrum. This opened the way for them to develop their applications for revitalizing damaged tissues. Thus, their functionality will be restored. cardiac tissue engineering (CTE) using curative procedures that combine biomolecules, biomimetic scaffolds, and cells plays a critical part in this path. Stimuli-responsive hydrogels (SRHs) are excellent three-dimensional (3D) biomaterials for tissue engineering (TE) and various biomedical applications. They can mimic the intrinsic tissues' physicochemical, mechanical, and biological characteristics in a variety of ways. They also provide for 3D setup, adequate aqueous conditions, and the mechanical consistency required for cell development. Furthermore, they function as competent delivery platforms for various biomolecules. Many natural and synthetic polymers were used to fabricate these intelligent platforms with innovative enhanced features and specialized capabilities that are appropriate for CTE applications. In the present review, different strategies employed for CTE were outlined. The light was shed on the limitations of the use of conventional hydrogels in CTE. Moreover, diverse types of SRHs, their characteristics, assembly and exploitation for CTE were discussed. To summarize, recent development in the construction of SRHs increases their potential to operate as intelligent, sophisticated systems in the reconstruction of degenerated cardiac tissues.

Stimuli-responsive hydrogels

cardiac tissue engineering

tissue engineering

drug delivery

heart

nanocomposites

nanomaterials

stem cells

myocardial infarction

Cardiovascular diseases (CVDs) are the leading cause of mortality worldwide, with the number of individuals dying from CVDs growing every year (1). Many CVDs may now be detected and even predicted early because of advancements in heart function measurement tools (2–10). Given the restricted ability of human cardiac cells to regenerate completely, several tissue engineering (TE) techniques are used to circumvent this obstacle. To heal cardiac and vascular tissues, a variety of materials have been used (11–13). Furthermore, a variety of hydrogels have been employed to overcome this constraint. The basic goal of TE is to repair damaged tissues and replace them with new biological ones (14–16). Cell biology and biochemistry are required for this interdisciplinary process. For clinical applications, clinical medical and material sciences investigations are also included (15). The physiologically active platforms are porous, 3D structures that allow biologically active components such as biomolecules, proteins, and growth factors to be attached to their surface. These biosystems' ability to contribute particular bioactivity to scaffold construction verifies its unique promise in TE and other biomedical applications (17). They can be used to deliver medicines and bioactive peptides, as well as filling agents and 3D structures. They can also regulate the regeneration process and encourage the creation of the needed tissue (18). Here, the biomaterial must have the right characteristics to fit the needs of tissue regeneration. Cell seeded bioactive materials serve an important part in the development of newly formed tissue by directing self-seeded cell growth or encouraging cell migration. They also serve as cell delivery matrices to certain bodily tissues. Furthermore, they are involved in the integrity of freshly formed tissue structure and function (19). They must have the essential physico-chemical characteristics that cause cell adhesion to their surfaces, cell growth, multiplication, differentiation, and migration for these objectives (20) as presented in Fig. 1, and to avoid the adverse consequences of a lack of these qualities, such as cell necrosis and impaired tissue regeneration (17, 20).

Several biological platforms have been developed over the years from natural sources such as algae (20), and animal tissues (21, 22), as well as synthetic sources such as lactic acid (23), and glycoside monomers (24), and caprolactone (25, 26). Even though many polymeric scaffolds utilized for CTE might provide critical support and assets, they lack several critical qualities such as appropriate cell mimicking and acceptable contact with stromal cells. Among them were collagen-fibrin-based hydrogels implanted with cardiomyocytes (CMs) produced from human-induced pluripotent stem cells (hiPSCs-CM) to restore the damaged cardiac tissue (27). This combination of stem cell-laden hydrogel scaffolds has shown promise and played a key role in CTE. Another micro-channeled 3D printed gelatin (Gt)-based hydrogel system was created to stimulate cardiac cell proliferation while also allowing stem cells to be used to improve heart regeneration (28). However, given the rapid advancement of hydrogel-based systems for heart regeneration, more research is needed to overcome the current constraints. Hydrogel-based scaffolds might provide a suitable, less invasive alternative that could support the qualities that conventional scaffolds lack (17). Hydrogels are water-based hydrophilic 3D polymeric structures. They might be natural, synthetic, or semi-synthetic. Extracellular matrix (ECM), epidermis, mucous, cartilage, meniscus, Gt, collagen, tendons, and vitreous humor are some of the bodily structures that include them (25, 26). They were proposed as innovative materials with promising results in TE. They may transport a large amount of water or biological fluids due to their hydrophilic nature, in which nutrients dissolve and diffuse to cells. Within primarily aquatic settings, their cross-linking structure maintains their integrity and avoids dissolution. Furthermore, they provide support to neighboring cells by adopting a high level of elasticity and flexibility, similar to the original ECM (29). The perfect bioscaffold can bridge tissue defects and improve repair by stimulating new tissue development and neovascularization while also demonstrating high levels of incorporation and biodegradation since they should vanish during or shortly after healing (30). Hydrogels are the preferred scaffold for a variety of biomedical and TE applications because of their unique structure and properties (31). However, increasing their therapeutic use by making them more adaptable to ongoing body functioning and pathological changes remains a challenge. Thus, new hydrogel materials with smart characteristics capable of promoting clinical applications in biomedicine and TE are still needed. Researchers are increasingly recognizing that recapitulating original tissues' intrinsic reactive capacity to biophysical and biochemical cues is a key component for improving functional tissue restoration. This rapidly expanding notion is leading to the development of smart hydrogel platforms that can respond to stimuli on demand (32, 33). When realizing and responding to either internal physiological or external applied stimuli, such smart designs provide unprecedented monitoring over network assembly/disassembly, selective biomolecule presentation, and other customizable qualities. (16). The development of hydrogel polymeric scaffolds in various biomedical applications will be discussed in this review, with a focus on TE. Aside from current advancements in stimuli-responsive hydrogels (SRHs) and their future perspectives in regenerative medicine.

For CTE, diverse techniques have been employed successfully as shown in Fig. 2. In this section, we will discuss them in detail with an emphasis on the materials used and the limitations of each strategy.

2.1 Scaffoldings and cells:

Multipotent cells capable to self-renew are called stem cells (SCs). They are crucial for tissue homeostasis and regeneration (34). However, adult hearts have a limited capacity for regeneration, and cardiac SCs (CSCs) could not replace the missing CMs. Recent advancement has been made in inducing pluripotent SCs (iPSCs) out of adult cells via reprogramming with specified transcription factors (35–37). Various cardiovascular cells have been obtained from iPSCs inside the cardiovascular system, and a recent study found that postnatal cardiac or dermal fibroblasts could be directly converted into cardiomyocyte-like cells (37, 38). While personalized iPSCs or cardiomyocyte-like cells opened up a new avenue for autologous stem cell therapy, the risk of tumor development has arisen (39, 40). Several sorts of SCs, including mesenchymal SCs (MSCs), embryonic SCs (ESCs), skeletal progenitor cells, hematopoietic SCs, and cardiac progenitor cells (CPCs), have presented a unique potential to develop functional CMs in the heart (34, 37, 41–43). Even though exogenous stem cell transplantation therapy has received a lot of attention, most allogeneic cells go through necrosis or apoptosis after transplantation due to immunoreaction and inadequate surroundings. Furthermore, the employment of ESCs therapy has been restricted by ethical issues (37, 44). Although several investigations have reported that cytokines can mobilize autologous SCs, there is no functional way to selectively attract SCs to cardiac damage sites. Furthermore, SCs need a sequence of coordinated interactions with their biological surroundings, whether they were injected directly or attracted by mobilization factors. Compared to the original environment of SCs, the microenvironment changed dramatically, making it harsher for these SCs to survive in the region of injury (45). Heart failure was treated with tissue-engineered cardiac patches such as ECM, however, the potency of myocardial healing was restricted due to the restricted ability for cell infiltration (46, 47). To imitate the various interactions between heart cells and the ECM, decellularized platforms from porcine or rat myocardium have recently been used (47, 48). Although decellularized matrices are found to be favorable when it comes to CTE, their application in transplantation has high risks, including a scarcity of human donors, and other immunological concerns, (47). Hence, researchers are now focusing their efforts on developing ECM biomimetic platforms made of synthetic and/or xeno-free biomaterials (47, 49). Biomaterials are being used more and more to target repair (21, 22, 50–52). They could function as a network for cell endurance, propagation, and multiplication, and as a guide for reestablishing 3D tissue (37, 45). Biomaterials are being established as cardiac patches to increase heart function via supporting and repairing damaged regions by replacing damaged myocardial tissues or scar tissue (53, 54).

2.2 Engineering of the cardiac tissue:

The classic TE process involves seeding target cells into a scaffold in vitro, sometimes with modification (e.g., special conditioned culture (55, 56), and then implanting the construct in vivo. Biomedical alternatives such as biomaterials are constantly being explored in the field of TE for the entire (or partial) replacement of injured tissue. The advance of a 3D matrix as a platform is a marked role for biomimetic materials. The biomaterials must also be suitable for the conservation of the cells and the signals necessary for tissue or organ regeneration. Following that, regenerated tissues need to maintain, reinstate and augment function (57–60). Whenever cardiac constructs had been implanted in a damaged myocardial area, neovascularization from the epicardium infiltrated the graft, and distributed fetal CMs survived the implantation process (56, 61). Upon implantation in the patient, tissue-engineered materials may become functional at the implantation time or be capable to integrate and accomplish the predicted function following implantation. As in either instance, the biomaterial needs to integrate well with the recipient or transplanted cells to effectively share in the tissue regeneration via cell-cell signaling and the release of GFs, propagation, multiplication, and development of ECM (59, 62–67). Intrinsic tissue regeneration for the heart isn't a portion of existing therapy for multifaceted cardiovascular injury (59, 68). Relationships between tissue regeneration, engineered biomaterials, and our immune system have to be entirely understood. The objective of cellular and TE is the evolvement of treatments that will stabilize, alter, or improve cardiovascular physiology and anatomy. Polymeric systems used in CTE have been outlined and constructed using different approaches (59, 69). They might be employed in the fabrication of degradable cardiac patches for example. In the long term, these polymeric biodegradable cardiac patches can provide excellent circumstances for cellular growth (70). Studies on elastomeric biodegradable poly(glycerol sebacate) (PGS), like Gt Nano-fibrous scaffolds, and PGS/fibrinogen core/shell fibers, have been conducted. Anisotropy was established in these materials, imitating the left ventricular (LV) myocardium. This can be employed as a construct for myocardial regeneration (71, 72). The cells' cytoskeletal organization was influenced by the scaffolds' structural features. For example, the amalgamation of synthetic Poly(lactic-co-glycolic acid) (PLGA) with natural Gt polymers was produced via electro- spinning by Prabhakaran et al. to create PLGA/Gel Nano-fibers. The potential of these scaffolds as biomimetic cardiac patches was highlighted by culturing the cardiomyocyte cells on them (59, 73). Engineered heart tissue (EHT) is a spontaneously contractile construct created using neonatal rat CMs, collagen I matrix, Matrigel^™, and a mechanical stretching device. It is amongst the most promising CTE approaches (56, 74–77). It was originally designed as a planar structure (77), which was then transformed into a circular structure (78) that had greater contractile qualities and a more distinct CM structure. For in vivo testing, the construct's 3D geometry was further changed into a pouch-like structure (74), which was then used to "encase" a failing heart to function as a biological ventricular assist system. There have been improvements to the in vitro states to lessen the utilization of serum and Matrigel^™ as well. Although the model looks to exhibit significant potential, core tissue viability is minimal, and a more advanced version of the EHT is still needed (56).

2.3 Cell sheet technique (scaffold-less cell):

Due to the challenges of biomaterial systems, such as integration impairment with host tissues, immunogenicity, and undesired degradation products (79, 80). Such dangers cannot be eliminated even though innovative biomaterial features have been devised to fix these issues. Scaffold-free engineered tissues known as cell sheets could be created using intact cell monolayers non-enzymatically isolated from substrates in vitro as an alternative (80). It is a strategy that has therapeutic application potential. Poly(N-isopropylacrylamide (PNIPAAm) is used for covering a thermosensitive cell culture layer. This polymer is cell adherent at 37°C and changes characteristics in reverse below 32°C. It is also possible to extract cells off a culture dish as a cell sheet once they have aggregated and developed gap junctions by lowering the temperature (56, 81). Furthermore, the capability to layer individual CM cell sheets into a 3D contractile cardiac tissue was established in vivo (82). The tissue endured subcutaneous implantation for one year, and cell sheets were functionally integrated with the host's heart when applied to rat hearts (83). Also using an orbital shaker, Stevens and his group have created a scaffold-free CM cell patch. In both scenarios, the result is tissue that is identical to a compact myocardium but without the scaffold. Limitations of this method are the contemporary need for certain vascularization strategies and nutrient diffusion supply to maintain the patch's durability and adaptability for the development of thick myocardial tissue constructs (56).

2.4 Cell Assembly:

It is not necessary to plant cells into 3D porous scaffolds. Instead, cells can be suspended in hydrogel-based scaffolds, that provide a suitable environment for them to travel and arrange into contractile tissues whether in vitro, by gravity-enforced methods to form sphere-like tissues (56, 84, 85), or in vivo, by an arteriovenous loop (AVL) embedded chamber for the vascularization of the arranged CMs (56, 86, 87).

The viscoelastic properties of hydrogels and their adaptability to chemical and physical changes have attracted substantial attention as cardiac tissue constructions (88, 89). They are water-insoluble polymers that can absorb a large quantity of water or biofluids, causing swelling and an expansion of their dimensions while keeping their shape. This feature makes them very close to soft tissues in their structure and function (88, 90, 91). It is likely to modify the surface of a hydrogel to have it respond to a certain stimulus, such as temperature, pH, molecules, magnetic or electric signals, and ionic strength (92).

Since typical hydrogels are quite often formed into larger sizes that have low surface-to-volume ratios, they have slow degradation values and limited cell infiltration along with weak vascularization. Hydrogels of this type have only nanoporous meshes within the cross-linked networks and lack micropores, indicating that nutrient transfer and cell vitality are insufficient within hydrogels (93–95). It was found that replacing bulk hydrogel with microporous annealed particle (MAP) that possesses a larger surface/volume percent and shorter diffusion distance can boost the mass movement of nutrients and promote long-term cells survival. Its pores can help guide cell multiplication and tissue development before the hydrogel breaks down (94, 96–98). Blood-derived MSCs are essential infiltrating cells that have a predisposition to relocate to the myocardial infarction (MI) region. It was hypothesized that vascular endothelial growth factor (VEGF)-encapsulated MSCs aimed at MI tissue could enhance the cardiac activity via angiogenesis and the MSCs' tropism to the MI area. Angiogenesis and heart function was improved by employing self-assembled alginate (Alg.) and Gt polyelectrolytes in the first stages of development. SDF-1 was found to be an attractive target for the VEGF-encapsulated MSCs in vitro with a stable release of VEGF. In vivo, angiogenesis was stimulated in the MI region by VEGF-encapsulated MSCs, and cardiac functions were enhanced. For MI treatment, these preclinical data imply that this VEGF-loaded layer-by-layer self-assembled encapsulated MSCs may be an effective and minimally invasive treatment option for MI (99). When self-assembling peptides are situated in a physiological environment, they create stable nanofiber hydrogels (88, 100). As a consequence of the in-situ injection of RAD16-II peptide gels that self-assembled in the myocardial, it induced an appropriate microenvironment (45, 88). Endothelial cells, nonvascular cells, and smooth myocytes were recruited by this microenvironment. RGDSP sequence with a cell-adhesive domain was connected to the self-assembling peptide RAD16-I to produce a biomimetic self-assembling peptide. The produced scaffold enhanced the adherence and viability of marrow-derived CSCs and facilitated their propagation to develop CMs, which as a result, improved heart activity and repair (101).

2.5 Heart matrix decellularization:

When cells are removed from organs or tissues while the ECM is left intact, it is called decellularization (102). Decellularized ECM has been established based on the idea that native ECM might be a better substitute for a tissue's complex environment (88). This can be achieved in diverse ways: chemically, physically, and enzymatically (21, 22, 102). Although this could alter the ECM's biochemical and morphological components, it has the advantage of removing cellular antigens which can trigger a foreign body reaction via inflammation, antibody activation, and probable transplant rejection (88). Biologic scaffolds used in clinical applications are produced using this technique. However, it has been proven that using perfusion decellularization preserves the organs' 3D geometry while removing the cells in a more uniform manner (102–104). To develop an entire-heart ECM scaffold, rats' hearts were decellularized. The scaffold maintained its 3D structure, with the vascular tubing skeleton preserved by the presence of vascular basement membranes (56, 105). A spontaneous contractile complete heart was generated after planting neonatal rat EC and CMs under physiological circumstances. This method has the potential of creating a big human or animal heart to replace the human heart functionally. Hydrogels made from the decellularized extracellular matrix (dECM) have gained a lot of interest in recent years due to significant advances in hydrogel technology and theory, as well as advances in the use of dECM hydrogels as a novel regenerative and TE medicine approach. There are structural and stimulatory features of hydrogel responsiveness that are retained along with ECM functionality. dECM hydrogel preserves cell GFs such as transforming GF, Fibroblast GF, and hepatocyte GF, which can improve the seed cell's proliferation, migration, differentiation, and angiogenesis (106). They offer the following benefits. (1) The capability to inject. At physiological temperatures, viscous fluids can be polymerized to create hydrogels that adjust to the form of the defective location. (2) The bioactivity of the native matrix is found in dECM hydrogels (106, 107). (3) There is no immunogenic cellular content in dECM hydrogels. (4) The mechanical assets can be altered. (5) Crosslinking or modifying the hydrogel concentration can be utilized to modify the mechanical strength of the dECM (108). (6) A gelled dECM has a 3D conformation that is beneficial to cell growth. (7) dECM hydrogels are machinability friendly. It is likely to customize 3D geometric shapes with 3D printing (109). Human iPSCs-derived CMs have a significant potency for disease categorizing and drug monitoring when they are exploited to construct human tissues. Hybrid hydrogels comprised of porcine cardiac dECM and reduced graphene oxide (rGO) were produced to support the normal development of cells and tissues (Fig. 3). EHTs developed using hiPSC-derived CMs and dECM-rGO hydrogels showed a significant increase in the tension powers and the expression of genes that control contractile function. It also improved many electrophysiological functions, including calcium handling, conductance speed, and action potential period (110).

2.6 Neo-vascularization:

Most novel blood vessel development occurs in mature organisms because of angiogenesis. Angiogenesis is described as the sprouting and formation of novel microcirculatory vessels from pre-existing vessels (56, 111). This happens through the breakdown of basement membranes and EC proliferation, which is stimulated by a large variety of growth agents (112).

There are only 100–200 mm layers of 6–12 cm thick that can survive in an in vitro designed construct that is implanted with CMs and relies only on diffusion for oxygenation and nutrition. Infarcted myocardium's epicardial surface will be significantly more difficult to neo-vascularize (56, 113). There have been diverse ways of vascularizing a clinically significant-sized construct. In general, these procedures can be divided into two categories: in vitro and in vivo vascularization approaches (56, 114, 115).

2.6.1 In vitro neo-vascularization:

In bioengineered tissue constructs, this is described as the growth and manipulation of biological components to generate microvasculature outside of a patient's own body. This technique is extensively employed in classical TE (56, 114). A CM-seeded cardiac construct is similar to an avascular transplant. For the transplant to survive, it must be infiltrated by the host's blood vessels. Consequently, the core of the 'transplant' will go through necrosis owing to the time it requires for the host vessels to vascularize it. By creating the intrinsic microvasculature in vitro, it is conceivable to rejoin the steward vasculature by either inosculation or by surgically connecting the graft to host vessels (56, 116).

2.6.2 In vivo neo-vascularization:

It can be further divided into extrinsic/external and intrinsic/internal vascularization: Extrinsic vascularization is the utilization of the patient's ectopic vasculature as subcutaneous fat, omentum, peritoneum, and axial vessels having the high-angiogenic potential to generate functional microvessels in a non-vascularized engineered construct (56, 114). Omentum was successful in CTE, while peritoneum failed in this regard (117). As a clinically significant vascularized graft still requires 3–4 days of 'taking', it relies on diffusion for survival. Therefore, the perfusion timing of these transplants should be carefully considered. If the ECT's size is large, this may be insufficient to support it (56). Regarding intrinsic vascularization, a central macrovascular conduit is put in a protected region to vascularize an endogenously produced or transplanted scaffold. It is employed to develop a microcirculatory network. During reconstructive surgery, the principle of prefabrication of flaps prompted the development of intrinsic vascularization in CTEs (118, 119). When an arterial-vein pedicle was implanted within or under the tissue graft, the tissue would become vascularized and generate an entirely new flap that is based on the pedicle. According to research, the AVL design is the most angiogenic of all pedicle implant configurations (56, 120).

3.1 Classifications of hydrogels:

Hydrogels are usually prepared from many natural, artificial (synthetic) polymers, or a hybrid mixture of both types. Alginate (Alg.), chitosan (CH), hyaluronic acid (HA), collagen (Col), and Gt represent the frequently used natural polymers. While poly(acrylic acid) (PAA), poly(acrylamide) (PAAm), polyethylene glycol (PEG), and poly(2-hydroxyethyl methacrylate) are among the frequently used synthetic polymers comprise (121). Hydrogels could be categorized depending on various principles including their polymeric configuration (homopolymeric (122), copolymeric (123), multipolymeric (124, 125)), their physico-chemical arrangement (amorphous, semicrystalline, and/or crystalline), sort of crosslinking (chemically, physically, or hybrid crosslinked) (124), their physical appearance (film, matrix, and/or microsphere), and the electric charging of their platform (ionic, nonionic, zwitterionic, and/or amphoteric electrolyte) (25).

3.2 Limitations to using conventional hydrogels in cardiac tissue engineering:

For years, hydrogels were involved as appealing platforms in many TE applications (126). However, the utility of conventional hydrogels for this purpose is sometimes still encountered as a prominent challenge. For instance, several pristine hydrogels lack adequate mechanical power (127). While stable, mechanically robust hydrogels are generally favored for long-term, tension-bearing implementations (128). They display spatial inhomogeneity where the distribution of their crosslinking density is not homogenous which lessens the potency of their networks. Also, since they are fragile, their loading ability and manipulation are problematic. They may need support with further dressing materials to improve their weak adherence capacity. Too, their reduced absorption power renders their therapeutic role in exudative wounds not efficient (129). Synthetic hydrogel polymers present limited biocompatibility, and biodegradability that restricted their employment in cardiac tissue engineering (CTE). Difficult sterilization of conventional hydrogels is an additional restriction to use them as they are sensitive to the traditional sterilization techniques accounted to their hydrophilic characters. The strategy of crosslinking effects hydrogels' release aptitude of biomolecules. Additionally, the chemical cross-linking agents present another threat of toxicity (130). The previously mentioned challenges have directed the research toward fine-tuning the hydrogels' characters and supporting their situation as functional systems for applications in CTE. In that context, electrical conductivity is a novel trait that has been recently acquainted with hydrogels to expand their pertinence and additionally the acknowledgment of innovative aptitudes while keeping up their original ones (e.g., tenderness and hydrophilicity) (18, 20, 131–133).

The core objective of TE is to renew the degenerated tissues with the restoration of their functions (18). SRHs could express unique characters that made them commonly utilized in TE applications. They serve as platforms to bridge cell migration to the seat of injury, possess an outstanding capacity to provide conditions that mimic the ECM surrounding environment, and could effectively modulate their mechano-physical properties to fit the required application (Fig. 4). Besides their great potential as tissue defects repair scaffolds (92). SRHs have been consumed for the reconstruction of different body tissues, including cardiac tissue (134–136), neural tissue (137, 138), skin (139–141), the cornea (142–145), bone (146–149), cartilage (150–152), tendon (153, 154), meniscus (155), and intervertebral disc (156, 157). Scientists have struggled to design SRHs by amending their physico-chemical features. These smart hydrogels can respond to various physical (temperature, light, electromagnetic fields, pressure, and ultrasound (US) radiation), chemical (pH, glucose, and ionic strength), or biological (enzymes and antigens/antibodies) stimuli and are described as SRHs (16, 158). Different SRHs exploited for CTE are presented in table 1.

Table (1)

Different stimuli-responsive hydrogels employed for cardiac tissue engineering.

Type of stimuli-responsive hydrogels		Stimuli-responsive hydrogel polymer		Stimuli	Study Model	Purpose of use		Refs.
Physical stimuli-responsive hydrogels	Temperature-responsive hydrogels	-Chitosan (CH)	- Chitosan-gold nanoparticles GNP loaded with mesenchymal stem cells (MSCs) (CH-GNP/MSCs)	-Temperature	-In vitro assessment	-Cardiac tissue engineering	2016	(159)
		-poly-lactic-co-glycolic acid (PLGA) -poly(N-isopropylacrylamide (PNIPAAm )	-PNIPAM containing PLGA-encapsulated PVP/H2O2 core /shell microspheres	-Temperature	-In vitro (cardiac fibroblast, cardiomyocyte, and endothelial cell)	-Treatment of myocardial infarction	2018	(160)
		-Gellan gum	-Gellan gum/reduced graphene oxide hydrogel	-Temperature	-In vitro (rat myoblasts (H9C2))	-Development of myocardial tissue engineering scaffold	2019	(161)
		-Chitosan (CH) -Dextran (DEX	-Chitosan (CH)/dextran (DEX)/β–glycerophosphate (β-GP) loaded with umbilical cord mesenchymal stem cells (UCMSCs)	-Temperature	-In vitro (3T3 cells and human umbilical vein endothelial cells (HUVECs))	-Cell delivery carrier for therapy of myocardial infarction	2020	(162)
		-Chitosan (CH) -Hydrolyzed collagen (HC)	- β-glycerophosphate (β -GP) and different kinds of hydrolyzed collagen HC-chitosan hydrogel	-Temperature	-In vitro (Fetal human ventricular cardiomyocytes cell line RL-14)	-Regeneration of the cardiac tissue	2021	(163)
	Light/Photo-responsive hydrogels	-Collagen (Col) -polydopamine (PDA)	-Collagen-polydopamine hydrogel (Col-PDA)	-Light	-In vitro assessment	-Control of the cardiomyocyte and neuron activity	2021	(164)
		-poly(2-alkyl-2-oxazoline) (POx)	-Cell-degradable poly(2-alkyl-2-oxazoline) hydrogel	-Light	-In vivo (rat myocardial infarction model)	-Epicardial placement of mesenchymal stem cells for myocardial repair	2021	(165)
		-Gelatin methacrylate (GelMA)	-Carbon nanotube (CNT)-incorporated photo-cross-linkable gelatin methacrylate (GelMA) hydrogels	-Light	-In vitro assessment	-Cardiac engineering and bio actuators	2013	(166)
		-Gelatin methacrylate (GelMA) -Reduced graphene oxide (rGO)	-GelMA-rGO nanocomposite hydrogels	-Light	-In vitro (cardiomyocytes cell culture)	-Cardiac Tissue Engineering	2016	(167)
	Electro-responsive hydrogels	-Gelatin (Gt) -poly-3-amino-4-methoxy benzoic acid (PAMB)	-PAMB crosslinked Gel hydrogels	-Electric	-In vivo (rat MI model)	-Propagation of the electrical impulse at the MI site to prevent cardiac arrhythmia and preserve ventricular function	2020	(168)
	Electro-responsive hydrogels	-Oxidized alginate (OAlg.) -Gelatin (Gt) -Polyacrylic acid (PAA)	-PAA mixed with OAlg. /Gel hydrogel	-Electric	-In vivo (rat MI model)	-MI repair	2021	(169)
	Magnetic-responsive hydrogels	-Chitosan (CH) -Carbon nanotubes (CNTs)	-CH/CNTs nano scaffold hydrogel	-Magnetic	-In vivo (neonatal rat heart cells)	-Cardiac tissue engineering	2014	(170)
		-Collagen (Col) -Polyethylene glycol (PEG)	-Col/magnetic Fe₃O₄ nanoparticles coated with PEG	-Magnetic	-In vitro assessment	-Cardiac tissue engineering	2017	(171)
		-Alginate (Alg.)	-Magnetic Alg. hydrogel scaffolds	-Magnetic	-In vitro assessment	-Cardiac tissue engineering	2012	(172)
		-polyethylene glycol (PEG)	-PEG diacrylate magnetic nanoparticles hydrogels	-Magnetic	-In vitro assessment	-Cardiac muscle cells engineering	2018	(173)
		-Gelatin methacrylate (GelMA)	-Cryogels based on GelMA and elastin adapted with carbon nanotubes (CNTs) and magnetic nanoparticles (MNPs)	-Magnetic -Light	-In vitro assessment	-Cardiac tissue engineering	2021	(174)
	Pressure-responsive hydrogels	-polyaniline (PAni)	Polymer polyaniline (PAni) hydrogel	-Pressure	-In vitro (cardiomyocytes culture)	-Supports cardiomyocyte organization into a spontaneously contracting system	2019	(175)
	Pressure-responsive hydrogels	-Hyaluronic acid (HA) -Cyclodextrin (CD) -Adamantane (Ad)	-CD-HA and Ad-HA hydrogels	-Pressure	-Ex vivo (porcine cardiac tissue)	-Cardiac tissue engineering	2017	(176)
	Ultrasound/acoustic-responsive hydrogels	-Silk sericin (MSS)	-MSS-Fe₂O₃ nanocomposite hydrogels loaded with secretome (Sec) biomolecules (Sec@MSS)	-Ultrasound	-In vitro (cardiomyocytes culture)	-Reduction of the Doxorubicin (DOX) induced cardiotoxicity in human stem cell-derived cardiac muscle cells	2021	(177)
	Ultrasound/acoustic-responsive hydrogels	-Polyethylene glycol (PEG)	-Heparin-binding based, Gd(III)-tagged PEG hydrogels	-Ultrasound	-In vivo (mouse myocardium)	-To deliver and monitor cardiac progenitor/stem cell engraftment for implantation	2017	(178)
Chemical stimuli-responsive hydrogels	PH-responsive hydrogels	-Poly-N-isopropyl-acrylamide (PNIPAAm) -Butyl acrylate (BA) -Propyl-acrylic acid (PAA)	-PNIPAAm-BA-PAA composite hydrogels	-pH -Temperature	-In vivo (rat MI model)	-Improvement of the angiogenesis in infarcted myocardium	2011	(179)
		-Poly-N-isopropyl-acrylamide (PNIPAAm)	-PNIPAAm with mono carbon nanotubes hydrogel entrapping stem cells	-pH	-In vivo (rat MI model)	-MI treatment	2018	(180)
		-Polyethylene glycol (PEG)	-Hydrogen bond crosslinked ureido-pyrimidinone group to PEG	-pH	-In vivo (pig MI model)	-MI treatment	2014	(181)
		-N-isopropyl-acrylamide (NIPAAm) -Di(ethylene glycol) divinyl ether (DEGDVE)	-NIPAAm hydrogel cross linked with DEGDVE (p(NIPAAm-co-DEGDVE))	-pH -Temperature	-In vitro assessment	-Cardiac Tissue Engineering	2019	(182)
	Ionic strength-responsive hydrogels	-Dopamine-gelatin (GelDA) -Dopamine-polypyrrole (DA-PPy)	-Fe-GelDA and DA-PPy composite hydrogels	-Ionic	-In vivo (rat MI model)	-MI treatment	2020	(183)
		-Polypyrrole (PPy) -Chitosan (CH)	-PPY-CH hydrogel	-Ionic	-In vivo (rat MI model)	-Prevention of heart failure	2020	(184)
		Polyacrylic acid (PAA)	-Self-healing ionic hydrogel (POG1) with biocompatible polyacrylic acid (PAA)	-Ionic	-In vivo (rat MI model)	-MI repair	2021	(185)
Biological stimuli-responsive hydrogels Physical stimuli-responsive hydrogels	Enzyme-responsive hydrogels	-Proline-Leucine-Glycine-Leucine-Alanine-Glycine (PLG\|LAG) polypeptides	-Matrix metalloproteinases (MMP-2) and elastase combined with PLG\|LAG polypeptides o form biopolymer hydrogels	-Enzyme	-In vivo (rat MI model)	-MI treatment	2019	(186)
		-Hyaluronic acid (HA)	-MMP-injectable hydrogels utilizing HA	-Enzyme	-In vitro assessment	-MI treatment	2022	(187)
		-Collagen (Col) -Proline-Leucine-Glycine-Leucine-Alanine-Glycine (PLG\|LAG) polypeptides	-Recombinant protein glutathione-S-transferase (GST)-TIMP-bFGF by combining bFGF, MMP-2/9-degradable peptide PLGLAG (TIMP), and GST entrapped in a GSH-modified collagen hydrogel (GST-TIMP-bFGF/collagen-GSH) hydrogels	-Enzyme	-In vivo (rat MI model)	-Growth factor delivery	2019	(188)
	Antigen/antibody-responsive hydrogels	-Collagen (Col)	-Sulfated glycosaminoglycan-like ECM-mimetic injectable hydrogel loaded with Artificial apoptotic cells (AACs) and vascular endothelial growth factor (VEGF)	-Antigen	-In vivo (rat MI model)	-MI treatment	2021	(189)
	Antigen/antibody-responsive hydrogels	-Polyethylene glycol (PEG)	-Magnetic basic structure nanoparticles (Fe₃O₄-SiO₂) hydrogel augmented with hydrazine hydrate and aldehyde-PEG to improve antibody conjugation (Fe₃O₄@SiO₂-PEG)	-Antibody -Magnetic	-In vivo (rabbit and rat models of MI)	-MI treatment	2020	(190)

4.1 Physical stimuli-responsive hydrogels:

4.1.1 Temperature-responsive hydrogels:

Thermo-induced hydrogel use change in temperature as the only trigger for their gelling tendency, with no other external influences. These hydrogels are appealing for biomedical applications because they can change to semisolid form in situ under physiological circumstances and are easy to administer (191). Poly (N-isopropylacrylamide) (PNIPAAm) hydrogels have been widely used as delivery payloads for a variety of treatments. However, due to their poor power to promote encapsulated cell growth, they showed inferior bioactivity for encapsulated cells (192). Xia and colleagues effectively inserted single-wall carbon nanotubes (SWCNTs) into the base PNIPAAm hydrogel, resulting in a thermo-responsive SWCNTs-modified PNIPAAm (PNIPAAm/SWCNTs) hydrogel with improved cytocompatibility (134). They could test the PNIPAAm/SWCNTs hydrogel system's bioactivity in brown adipose-derived stem cells (BASCs) and its effectiveness in delivering BASCs to the MI site. The PNIPAAm/SWCNTs hydrogels not only showed significant-high bioactivity to encapsulated BASCs in vitro, with enhanced cell proliferation and adhesion, but they also demonstrated a satisfactory capacity to deliver the encapsulated BASCs to the infarct myocardium in vivo, with enhanced engraftment of seeded cells at the MI site. In another work, an electroconductive gold nanoparticle (GNP)-loaded CH thermal-induced hydrogel for cardiac repair was explored. The electrical connection between stem cells and neighboring cardiac cells could be increased by raising the concentration of GNPs (159). A PNIPAM thermo-induced hydrogel containing PLGA-encapsulated PVP/H₂O₂ microspheres was used to heal cardiac cells following MI in another investigation. After 4 weeks of injection in the MI location, cardiac cells showed lower TGF expression, and cardiac fibrosis was decreased, indicating that myocardial cells were healed due to oxygen uptake (160).

Another thermo-induced hydrogel made of gellan gum and rGO has been investigated. The findings of the MTT experiment showed that gellan gum thermos-induced hydrogels comprising 1% and 2% rGO weren't cytotoxic. Gellan gum/2 percent might be a good option for mending and recovering infarcted cardiac tissue, according to these findings (161). For MI therapy, a chitosan (CH)/dextran (DEX)/ β-glycerophosphate (β-GP) parenteral thermo-induced hydrogel occupied with umbilical cord mesenchymal stem cells (UCMSCs) was created. UCMSCs develop towards myocardial and have significant potential for clinical applications of cardiac repair, according to the expression of cardiac markers cTnI and Cx43 and signaling pathways p-Akt and p-ERK1/2 (162). In another study different quantities of beta-glycerophosphate (-GP) and different kinds of hydrolyzed collagen HC, as well as CH, were used to synthesize thermo-induced hydrogels. The hydrogel showed cell viability of more than 75%. These properties make it appropriate for injectable and easy to perform cardiac therapeutic uses, in which cells might be trapped and directed to the wounded spot of the heart muscle via the porous structure. Based on the gelation time and biocompatibility results, it can be inferred that these hydrogels have good physical-chemical and biological properties, making them ideal prospects for TE treatment modalities against infarcted or ischemic cardiac tissue damage (163). An injectable, thermo-induced hydrogel of poly(lactide-co-glycolide)-poly(ethylene glycol)–poly(lactide-co-glycolide) (PLGA–PEG–PLGA) was created and mixed with colchicine (Col) to form Col@Gel network (Fig. 5A). The injectable solution formed in situ gel at 35°C and remained in a gel form at body temperature (37°C) (Fig. 5B). In vitro evaluation of the smart hydrogel system revealed that they reduced the viability and the migration of the Raw264.7 macrophages which are the most vigorous inflammatory cells present in the MI (Fig. 6A). The injection of this hydrogel ameliorated cardiac inflammatory response effectively, hindered myocardial necrosis (Fig. 6B), improved cardiac performance, and enhanced mouse viability in an animal model of MI without inducing severe cytotoxic effects, as seen after administration of conventional Col solution (193).

4.1.2 Light/photo-responsive hydrogels:

Light is an excellent stimulator for polymer shape alteration. Light allows for temporal and spatial modulation and distant manipulation of polymer. The quantity of energy bending deformations can be well-adapted by adjusting various illumination parameters including intensity of light, length of exposure, and wavelength. Furthermore, controlled light irradiation or changing polarization direction may be used to create well-defined and complicated 3D hydrogel structures (194). UV-induced hydrogels and visible photo-induced hydrogels are the two types of photo-induced hydrogels. Visible light, apart from UV light, is readily available, inexpensive, safe, clean, yet simple to use (195). The physiological functioning of the brain and cardiac tissues was controlled using near-infrared (NIR)-sensitive polydopamine NPs. By subjecting the cells to polydopamine NPs encapsulated in collagen foam and thereafter illuminating the NIR light source on the NP-foam mixture, the electrical activity of both types of cells was regulated. photo-induced NPs were placed in highly porous hydrogel frameworks that are physiochemically and physiologically adjustable. The highly porous design of the photo-induced hydrogel can improve cell movement (196). A poly(2-alkyl-2-oxazoline) (POx) derivative was used for the preparation of photo-induced hydrogel. it was shown that the cell-degradable characteristics of the resultant POx hydrogels allow for the control of cell outgrowth in 3D matrices. In partly cell-degradable POx hydrogels, the transcription of pro-angiogenic genes was elevated. It was demonstrated that the epicardial insertion of MSC-loaded POx hydrogels improved the restoration of cardiac function in arteries and heart walls due to the outstanding tissue adhesion characteristics of thiol-ene polymerized hydrogels (165). Another research found that incorporating carbon nanotubes (CNTs) into a photo-cross-linkable gelatin Methacrylate (GelMA) hydrogel can result in high-performance cardiac scaffold materials. In comparison to pure GelMA, the cardiac cells in CNT-GelMA were found to be elongated, while F-actin fibers were shown to be viable and more homogenous. Synergistic beating activity can be generated by both CNT-GelMA and pure GelMA. The CNT-GelMA, on the other hand, demonstrated a beating frequency average that was both more steady and more consistent when the impulsive beating rhythms were collected over 6 days (166). Reduced graphene oxide included in GelMA photo-induced hydrogels is used to create cardiac tissue constructions. The addition of rGO to the GelMA matrix improves the material's conductance and mechanical behavior dramatically. Furthermore, relative to GelMA hydrogels, cells cultivated on rGO-GelMA scaffolds demonstrate improved biological functions such as cell survival, proliferation, and development. During the rGO-GelMA hydrogel period, CMs exhibited improved contractility and a quicker natural beating rate (167).

4.1.3 Electro-responsive hydrogels:

The development of non - conducting fibrotic tissue post-CM mortality due to ischemia is a typical feature of MI. The growth of scar tissue obstructs the transmission of electric signals between CM resulting in abnormal heart rhythms and functional decompensation. Restoring electric signals within the heart allows the contraction to be resynchronized, preventing additional remodeling and ventricular dysfunction (168, 197). Various hydrogels having conductive qualities, dubbed electro-induced hydrogels, have been developed to repair the heart's electric pulses (169, 198, 199). For instance, conductive hydrogels have been made by attaching conductive polymer poly-3-amino-4-methoxy benzoic acid (PAMB) on Gt and cross-linking it using carbodiimide. The self-doped (PAMB-G) hydrogel was found to be compatible with CM in vitro and effective in synchronization. An ex vivo experiment was conducted in which two rat hearts, one beating and the other not, were segregated by either PAMB-G hydrogel or a non-conductive hydrogel to investigate the hydrogel's capacity to carry an electrical impulse (Gt). In this experiment, PAMB-G had a substantially greater measured improved amplitude in the non-beating cardiac muscle than the non-conductive hydrogel (168).

PAA was mixed with oxidized alginate (OAlg.)/Gt to create a self-healing highly porous electrically conductive hydrogel. This hydrogel was shown to exhibit stable rheological characteristics, resilience, and flexibility, as well as the ability to maintain its network after or during deformation. Furthermore, CMs sown inside the hydrogel synchronized their beating and exhibited a tendency for directed development due to the hydrogel's conductive qualities. Furthermore, the hydrogel was loaded with donor CM before being implanted rather than injected. While the hydrogel kept the cells in the heart and improved functional recovery compared to a control lacking CMs, no experiments were done to see if it might reduce arrhythmias. PAMB-G is more conductive than the non - conducting hydrogel (169).

Conductive hydrogels can potentially be effective in synergistic conjunction therapy. One of the studies reported a conductive hydrogel filled with adipose-derived stem cells (ADSCs) and plasmid DNA expressing endothelial nitric oxide synthase (eNOS). The hydrogel was designed to improve electrical signal conductivity inside the infarcted heart, reduce inflammation, and stimulate NO generation and angiogenesis. To produce the conducting hydrogel, tetraaniline-polyethylene glycol diacrylate (TA-PEG) NPs were self-assembled and interacted with thiolated HA through an acrylate-thiol Michael Addition reaction (HA-SH). ADSCs elevated NOx levels in the myocardium and, as a result, stimulated angiogenesis, which improved recovery (198). Gold nanoparticles (GNPs) were mixed in the chitosan–glycerol phosphate (CH-GP) gels to create a composite that mimicked the electromechanical capabilities of the heart without sacrificing thermosensitivity. Comparative to cells loaded on pure CH hydrogels, MSCs loaded on the CH–GP–GNP hydrogels demonstrated increased cardiomyogenic development (159). Poly(pyrrole) (PPy), an electro-responsive substance, was integrated into acid-modulated silk fibroin to create an electron-induced hydrogel polymer substrate. The sarcomere size and Z-bandwidth of CM cultured on this substrate increased, as well as the expression of cardiac-specific genes. Carbon nanotubes have been studied extensively as conductive substances for cardiac engineering applications (200–202).

4.1.4 Magnetic-responsive hydrogels:

Magnetic-responsive hydrogels are constructed by incorporating magnetic nanoparticles (MNPs) in the hydrogel that are triggered by an exterior magnetic force to obtain the desired reaction in the environment (203, 204). The number of nanoparticles incorporated in the hydrogel affects the magnetic characteristics of the hydrogel, the release behavior of the loaded medications, and the sensitivity of the cultured cells (205). MNPs can be in the form of oxides (for example, iron oxides), metallic (for example, nickel, iron, and cobalt), or plated oxides and metallic nanoparticles (203). Scaffolds for cardiac engineering must resemble the mechanical qualities of the original tissue, give signals for cell alignment and elongation to duplicate certain contractile characteristics, and enable electrical conductivity. Conductive polymers such as polypyrrole (PPy) or polyaniline, as well as noble metals or carbon-based particles, have been investigated (174). Chitosan-CNTs, with their rod-shaped architecture, are particularly well-suited for cardiac engineering applications. These NPs were employed to make conducting polymeric hydrogel scaffolds with an anisotropic design that supports the development of elongated cells like those found in cardiac original tissues (170). MNPs for cell stimulation in cardiac-engineered constructions at a distance were utilized by co-precipitated magnetite MNPs combined with freeze-dried Alg. nanoporous hydrogels. The scaffolds were inoculated with bovine endothelial cells of the aorta, and a weak AFM (1.5 mT, 40 Hz) was applied using Helmholtz coils. The stimulation of magnetic scaffolds to AFM stimulated the formation of early capillary-resemble structures in endothelial cells, replicating the biological structure of real heart tissue (172). Magnetic Fe₂O₃ nanoparticles produced by co-precipitation were inserted into collagen hydrogel film for cardiac TE scaffolds. It was feasible to obtain micropatterns of MNPs within the collagen matrices by using external magnetostatic fields (created by two parallel magnets or current wire configurations). This method enables the conductivity of the specified scaffolds to be increased without the use of highly conductive materials. The scientists highlighted the relevance of anisotropy in accurately mimicking the electroconductive characteristics of cardiac tissues, resulting in the development of a viable tool for triggering stem cell cardiac differentiation (171). Another relevant study, PEG diacrylate hydrogels having two stacked sheets modified with widely viable MNPs as prospective cardiac and muscular TE scaffolds. The sheets were bent into 3D tubes and CMs were implanted within, resulting in high adhesion as well as survivability, and preservation of contractility over 7 days (173). Another CTE, magnetically induced cryogels based on GelMA and methacrylate adapted with CNTs and MNPs were explored. Under the externally applied magnetic fields, the resulting scaffolds coupled exceptional elasticity, plasticity, and strain conductivity, making them ideal options for use as controlled conductive actuators for the heart (174).

4.1.5 Pressure/mechano-responsive hydrogels:

Pressure-responsive hydrogels have made great advancements in the biological field (206). They are a special sort of hydrogel that has a high sensitivity, a lot of flexibility, and a lot of repeatabilities, enabling them to be employed in cycling and single pressure measurements. Character modifications in these hydrogels are common in reaction to external stimuli. The strength of the stimulus and the impact of external stressors may be determined by examining system parameters (207). The capacity of these hydrogels to respond to recognized pressure stimuli is particularly noticeable in highly elastic hydrogels. A conducting pressure-responsive hydrogel was created using a fortified dipeptide as a supramolecular gelator, owing to its inherent biocompatibility and outstanding gelation capabilities, along with the conductive polymer polyaniline (PAni), which has been polymerized in the site. The hybrid hydrogel is physically stiff, and the rigidity may be adjusted by varying the peptide content. The hydrogel has ohmic conduction, pressure responsiveness, and, perhaps most crucially, self-healing properties. the hydrogel's innate conductivity may be restored owing to its self-healing ability after the macroscopic separation of its block. Cardiomyocytes cultured on the hybrid hydrogel have a high cell survival, indicating that it is noncytotoxic. The hydrogel's combination of properties allows it to be used for a dynamic range of pressure sensing as well as a conductive interface for electrogenic heart cells. Cardiomyocyte organization into a naturally contracting system is aided by the hybrid hydrogel (175). Another study created a pressure-responsive hydrogel by combining HA either with -cyclodextrin (CD) or adamantane (Ad) to form CD-HA and Ad-HA, and assembling the hydrogels thru supramolecular hydrophilic groups interaction to create shear-thinning injectable hydrogels in an animal model with MI (176).

4.1.6 Ultrasound/acoustic-responsive hydrogels:

Attributed to its efficiency and non-invasive feature, the US has widely employed in medicine to image inner tissues, allowing for illness diagnosis and treatment. US waves, like other external stimuli, can be used to remotely initiate on-demand medication administration and enable the development of novel therapeutic platforms with deep tissue penetration (16). For example, US-responsive systems have the potential to deliver bioactive therapeutics in a temporally controlled way in a variety of ways, such as pulsatile or uni-directional burst release, and over extended periods following in vivo application (208). Ultrasound-induced acoustic energy could upset the self-assembly equilibrium of nanocarriers such as lipid nanoparticles, polymeric micelles, and nanocapsules, causing them to disassemble. Silica-based nanoparticles have been widely investigated in US-based applications, benefiting from improved acoustic cavitation and permitting the loading of US disturbed mesopores. Hydrogel networks formed with non-covalent connections can also be spontaneously fragmented with the application of ultrasonic waves, in addition to US-responsive nanomaterials (209). Doxorubicin (DOX)-induced apoptosis CMs were reduced using the innovative US disrupted injected hydrogel architectures mixed with nanoparticles and secretome molecules. A biocompatible silk sericin (MSS) matrix form including Fe₂O₃ nanoparticles was synthesized and employed as an injectable vehicle for secretome for cardiomyocyte metabolism in vivo. The secretome-encapsulated Fe₂O₃ -Silk sericin (Sec@MSS) hydrogel was recommended as a therapy for heart problems (177). Multifunctional US-guided hydrogel consists of a PEG network linked with Gd (III) peptide allowing MRI to monitor the hydrogel's localization and retention in vivo. heparin-binding peptide (HBP) sequencing in the cross-linker design to customize the polymer for cardiac purposes, the formed gel has mechanical characteristics similar to those of heart tissue. The metabolic activity of luciferase-expressing cardiac stem cells (CSC-Luc2) contained inside these gels was sustained lasting for 14 days in vitro. CSC-Luc2 persistence in the mouse heart and hind limbs was enhanced by 6.5 and 12 times, respectively, after encapsulation in HBP hydrogels (178). In a recent elegant study, Fu H. et al., (210) designed ultrasound-sensitive nanomaterials to generate oxygen at the infarct site (Figs. 7A and B). Moreover, they reduce the hypoxemic myocardial milieu to save the cardiac cells following AMI. The construction of the thermosensitive substance heneicosane and polyethyleneglycol was done after the synthesis of calcium peroxide in the mesopores of biocompatible mesoporous silica nanoplatforms. It was possible to achieve US-responsive diffusion of water and release of oxygen thanks to the phase change of heneicosane that was caused by the slight hyperthermia brought on by US irradiation (Fig. 7C). The cardiac cell survival under hypoxic circumstances was markedly enhanced (Fig. 7D). After an AMI, the US-activated oxygen release greatly reduced hypoxia and aided in the reduction of oxidative stress. As a result, the damage to the infarcted myocardial tissue was reduced (Fig. 7E). This nanosystem for US-activated oxygen production might offer an effective AMI therapy option and keeps the door open for future integration of this strategy to produce more sophisticated US-sensitive nanocomposite hydrogels for CTE applications.

4.2 Chemical stimuli-responsive hydrogels:

4.2.1 pH-responsive hydrogels:

pH is a commonly changed biological trait in pathologic conditions including inflammatory disease and tumors as well as anatomical regions including different parts of GIT. PH-responsive hydrogels can readily vary their properties to promote site-specific drug delivery. The prevalent methods for creating pH-responsive hydrogels are the inclusion of a covalent bond that is susceptible to breaking down by changing pH or the employment of polyelectrolytes, having several ionic terminal groups that absorb or give protons in reaction to pH variations in the environment (211). Changes in pH can be exploited as triggers for medication delivery to areas of local acidoses, such as those observed in myocardial ischemia (pH 6,7) (212). Several attempts have been done to create pH-responsive hydrogels utilizing different copolymers that respond to acidic pH, including carboxylic acid-derivative as propyl acrylic acid, and ethyl acrylic acid (213), methacrylic acid (214), or acrylic acid (215).

An injectable pH-responsive hydrogel was developed comprising poly-N-isopropyl-acrylamide (PNIPAAm), butyl acrylate (BA), and propyl-acrylic acid (PAA). At pH 7.4, this polymer is a liquid, but at 37 ^oC and pH 6.8, it transforms into a gel. It was assumed that this polymer's tendency to develop a reversible gel in medium acidic environments permits it to behave as a depot structure for the discharge of angiogenic growth factors to the ischemic myocardial tissue, and then to promote polymer disintegration and removal when the tissue recovered to physiological pH (213). In an attempt to improve the previous hydrogel, reversible fragmentation addition chain polymerization was used for hydrogel synthesis this technique improved the physical strength of the gel and increased the retention time of the growth factor into infarcted rat myocardium by 10-fold promoting 30e40% raised capillary and arteriolar densities (179). PNIPAAm hydrogel also was used to permit the injection of mono CNTs to contain entrapped stem cells derived from adipose cells into the heart of a rat MI model using. It was discovered that the hydrogel improved cell incorporation and provided a therapeutic effect (180). In another study catheter-injectable hydrogel was created by linking hydrogen bonding ureido-pyrimidinone group to PEG through alkyl-urea spacers (181). In a recent study, multiple drugs including phenytoin, indomethacin, and clotrimazole were slipped on glass using polymerization techniques to construct pH-induced hydrogel, and then in vitro release study was conducted, where the blockage of release was measured versus different pH (182).

4.2.2 Ionic strength-responsive hydrogels:

Ionic strength is a frequent and readily regulated stimulus, hence ionic strength-induced hydrogels might be useful in a variety of situations. Ionic-strength-induced hydrogels are frequently made from ionizable polymers that swell differently in water and an electrolytic solution. The degree of ionization and density of charge of the hydrogels have a significant impact on the surrounding solution's expansion and shrinkage behaviors (216). Using Fe³⁺-induced cationic coordination, a uniform hydrogel matrix was formed combining dopamine-gelatin (GelDA) and dopamine-polypyrrole (DA-PPy). As cardiac patching, this conductive hydrogel increases electrical signal conductivity in infarcted myocardial. In another study PPy-chitosan hydrogel (PPy-CH) hydrogel was employed to investigate the role of ionic strength-induced hydrogels in avoiding heart failure. The findings showed that ionic strength-induced hydrogels might improve electrical conductivity to synchronize cardiac contractions by lowering electrical resistance in the infarcted region, laying the groundwork for ionic strength-induced hydrogels to be used in myocardium infraction MI therapy (184). A PPY-CHI hydrogel was also created and employed for MI treatment. The findings showed that, as compared to conventional hydrogels, the ionic strength-induced hydrogels efficiently increased Ca²⁺ conduction in rat CMs in vitro, depressed the QRS interval, and improved electrical pulsed signal transmission and heart function upon MI (217). For the management of MI, ionic strength-induced hydrogels have been progressively utilized. Ionic strength-induced hydrogel choline- bio ionic liquid (Bio-IL) embedded in a GelMA hydrogel, increases primary CMs adhesion, proliferation, and electro-modification in vitro (218). Furthermore, PAA was incorporated into oxidized alginate (OAlg.)/Gt combination to create a new hydrogel (POG) with outstanding conductance and self-healing capabilities due to its side-chain carboxyl groups deprotonating into COO⁻ in the medium. The POG hydrogel also enhances CM resynchronization by participating in an efficient interaction with CMs via electrical communication. POG hydrogel has been shown in vivo to drastically minimize LV remodeling and recover myocardial performance (169). Another ionic strength-induced hydrogel was synthesized by dissolving gold nanorods (GNRs) in GelMA solution. When cultivated GelMA-GNR hydrogels, the ionic conductivity of the hybrid hydrogel was modified to facilitate rhythmic contraction of the CMs (219). The ionic strength-induced hydrogel was also prepared by ultrasonically distributing CNTs into a hydrogel precursor solution. The electrical connection between CMs was greatly enhanced via this hydrogel (220).

4.2.3 Redox-responsive hydrogels:

A redox state is critical for a variety of everyday biological processes in humans, including cell cycle progression (cell growth, maturation, and mortality), gene expression, Calcium ions modulation, energy metabolism, immunological response, and neurogenic activity. People with cancer, cardiovascular disease, and diabetic patients all have uncontrolled redox metabolism. In a normal person, a redox reaction, particularly glutathione/glutathione disulfide (GSH/GSSG), the archetypal cellular redox couple, works continually to maintain the redox state controlled by buffering the quantity of oxygen species reaction. Redox-induced hydrogels can be used in a wide range of biomedical applications. Redox-induced hydrogels can be used as biodegradable frameworks for organ and tissue healing in TE. Active species including growth factors may be placed into these hydrogels and their release is activated by redox stimulation to enhance cell penetration and proliferation. Redox-responsive hydrogels, particularly those formed by in situ gelling methods, can be exceedingly effective for topical medication delivery (221). The use of suitable polymers in the creation of redox-induced hydrogels is critical for achieving biocompatibility and desirable performance. Vong, L. B. et al., (222) have developed poly(arginine)-based nanoparticles injectable hydrogel (NO-RIG) to sustain nitric oxide (NO) release in myocardial tissue. NO-RIG hydrogels could scavenge ROS and so boost NO bioavailability. Following injection, using a Hylite-labeled polymer, the retention of NO-RIG was evaluated. NO-RIG is made up of two biofunctional polymers: PArg-PEG-PArg, which produces NO by activating macrophages in inflamed tissues, and PMNT-PEG-PMNT, which scavenges ROS to boost the activity of NO. The NO delivery and redox equilibrium balance in the infarcted tissues can be stably controlled by NO-RIG. The NO-RIG therapy considerably slows the course of MI and enhances cardiac functions, according to the findings. NO-RIG-treated animals showed a remarkable increase in angiogenesis. Researchers discovered that only the animals treated with NO-RIG demonstrated significant restoration of cardiac performances in both major and mild MI models (223). To identify reactive oxygen species (ROS) and prevent cells from oxidizing as a result, redox-responsive hydrogels were created. In the human body, ROS, which are highly reactive substances generated during cell metabolism, are prevalent. Furthermore, despite their powerful oxidation-induced damaging action on cellular fluids, nucleic acids, and proteins (224), they make a significant contribution to cell signaling. Hence, increased generation of ROS is frequently linked to various systemic diseases and disturbances in a number of body activities (225). ROS-responsive drug delivery devices have received much research in the treatment of cancer, immunotherapy, and gastrointestinal (GI) illnesses (226–229). However, the low ROS concentrations in the aforementioned microenvironments prevent additional ROS-triggered drug release applications (230). In contrast, Ischemia/reperfusion (I/R) heart damage causes a quick buildup and continual creation of ROS, which may be exploited as a trigger for medication release. In this regard, Zhenhua Li and co-workers (136) synthesized a basic fibroblast growth factor (bFGF)-loaded and ROS-responsive hydrogel (Gel-bFGF) and immediately injected it into the pericardial cavity as a method for heart repair as presented in Fig. 8A. The reasoning for this is that bFGF will be carried by ROS-sensitive crosslinked poly(vinyl alcohol) (PVA) hydrogel, which will dissolve under ROS to release bFGF into the heart "on-demand" (Fig. 8B). The influence of these growth factors was assessed in vitro using rat cardiomyocytes (NRCMs). The findings showed that bFGF had the maximum effect to support cell growth and proliferation indicated by the elevation of Ki67 positive cells (Figs. 8C and D). Moreover, they injected the gel into the pericardium and monitored the animals over 4 weeks (Figs. 8E and F). After being delivered, the hydrogel spread across the heart's surface and instantly formed an epicardiac patch. The myocardium might be penetrated by bFGF released from the gel (Figs. 8G, H, and I). With improved angiomyogenesis, such intervention preserves cardiac function and lessens fibrosis in the post-I/R heart. Additionally, both in pigs and human patients, it was shown that minimally invasive injection and access into the pericardial cavity were safe and feasible.

4.3 Biological-responsive hydrogels:

4.3.1 Enzyme-responsive hydrogels:

Enzymes play a critical role in metabolic as well as biological functions. In cardiovascular diseases like atherosclerosis, MI, and thrombus, upregulated levels of certain enzymes were reported, which can be used as triggers for site-specific delivery and prompt release of contents at sites of action. Various enzyme-responsive nanocomposite hydrogel has been developed in this regard, which may be destroyed by certain enzymes to trigger or reveal bioactive components for later diagnosis or therapy. Proteases and esterases degrade short peptide fragments or esters, which are the major processes. The biological exclusivity and unique sensitivity of enzymes towards their substrates are key advantages of using enzymes to drive nanocarrier disintegration, allowing for targeted and biomaterials payload release (187). A range of enzyme-sensitive hydrogel was used to treat MI, taking into account overproduced enzymes beyond the first stage of MI. Carlini et al. created an enzyme-responsive hydrogel that flowed easily during injection and generated hydrogels in situ of MI in rats overexpressing matrix metalloproteinases (MMP-2) and elastase. In this scenario, cyclic self-assembling peptides were combined with a Proline-Leucine-Glycine-Leucine-Alanine-Glycine (PLG|LAG) polypeptide that may be broken by MI-related proteolytic enzymes. Due to the self-assembly of peptide sheets forming fibrils, the non-Newtonian solution of peptide solidified into viscous hydrogel when triggered by proteases in the infarcted cardiac disease (186). Purcell et al. demonstrated MMP-injectable hydrogels utilizing HA that may be infused into the MI region and liberate rTIMP-3 (a recombinant tissue inhibitor of MMPs) since the utilized MMP-cleavable peptide (GGRMSMPV) cross-linker can be decomposed by active MMPs. When aldehyde and hydrazide functionalized HA were combined instantaneously through a double-barrel syringe, MRI confirmed hydrogel formation. In a pig MI model, intramyocardial injection of these hydrogels significantly reduced MMP expression in the damaged region, although MMP activity remained unaffected (187). Further research revealed that the localized distribution of an MMP-responsive hydrogel discharging a recombinant TIMP is a viable method for preventing negative post-MI remodeling (231). In another research enzyme-responsive micelles hydrogel was created out of peptide-polymer amphiphiles which can be identified and destroyed by MMP-2, allowing for targeted deposition and long-term retention in MI cardiac muscle. In response to upregulated MMPs in MI cardiomyocytes, the produced nanoparticles experienced a morphological transformation from spherical substances to a network-like structure hydrogel after IV injection. The sick location can be covered with the in situ created scaffold for up to 28 days. Despite this, the therapeutic benefits of MMP-responsive and adaptable nanoparticles have yet to be proven (232). Recently, Fan, Caixia, et al. (188) created a recombinant protein glutathione-S-transferase (GST)-TIMP-bFGF by combining bFGF, an MMP-2/9-degradable peptide PLGLAG (TIMP), and GST, which was then entrapped in a GSH-modified collagen hydrogel to create MMP-sensitive hydrogels (GST-TIMP-bFGF/collagen-GSH) for discharging of growth factors in the MI region (Figs. 9A, B, and C). The quantity of GSTTIMP-bFGF loaded in the collagen-GSH hydrogel is greatly increased by the specific binding between GST and GSH. Moreover, they showed a substantial release of these growth factors from the hydrogel platforms responsive to the MMPs enzymes (Fig. 9D). During MI, the TIMP peptide that is sandwiched between GST and bFGF reacts to MMPs for on-demand release. Additionally, after a MI, the excessive breakdown of the myocardial matrix by MMPs is inhibited by the TIMP peptide, a competitive substrate of MMPs. After local injection of the hydrogel to the damaged location of MI rats, MMP-2/9 destroyed the precursor peptide TIMP, accompanied by bFGF emission, decreasing MMP activity, boosting angiogenesis, and encouraging MI healing by enhancing vascularization and relieving myocardial remodeling. According to the findings, simultaneously regulating binding and responsive release to encourage angiogenesis and lessen cardiac remodeling may be a potential strategy for treating ischemic heart disease.

4.3.2 Antigen/antibody-responsive hydrogels:

Chemical coupling or polymerization of the antigen or antibody's interacting portion with the hydrogel could be used to include antigens or antibodies in the hydrogel matrix. Hydrogels were commonly employed to contain these chemicals because of their exceptional characteristics, such as microporous, higher binding affinity, and significant stability (233). Artificial apoptotic cells (AACs) were created by modifying liposomes with phosphatidylserine (PS) exposed at the out layer, to replicate the activity of apoptotic cells and stimulate macrophage M2 polarization. The AACs and VEGF were enclosed in a sulfated glycosaminoglycan-like ECM-mimetic injectable hydrogel. The loaded AACs and VEGF could achieve faster and slower sustained release, which is controlled by different electrostatic interactions. On-demand management of inflammatory immune response and increased angiogenesis was achieved by the regulated spatiotemporal administration of AACs and VEGF. The local precise release of AACs decreased the inflammatory response and lowered cardiomyocyte death in a rat model of MI with the intramyocardial treatment of an injectable hydrogel carrying AACs and VEGF (189). As the VEGF signaling cascade was activated by the sustained release of VEGF, revascularization was made possible in the targeted infarcted area, ultimately leading to a considerable improvement in cardiac function and reduced pathological remodeling of the left ventricle. Magnetic basic structure nanoparticles (Fe₃O₄-SiO₂) hydrogels were produced and augmented with hydrazine hydrate and aldehyde-PEG to improve antibody conjugation (Fe₃O₄-SiO₂-PEG-CHO; GMNP). Exosomes were captured and damaged CMs were targeted using the nanoparticles that had been coupled with anti-CD63. Endogenous CD63-expressing exosomes were collected by anti-CD63 after IV injection, and the nanoparticles were accumulated near the region of heart damage using an exogenous magnetic field. Exosome bioavailability was increased, heart function was improved, and MI was decreased in animal models in vivo investigations (190). These strategies for fabricating an injectable smart hydrogel with a precisely tailored discharge of diverse growth factors to the targeted area of MI will provide outstanding platforms with great promises for the clinical treatment of ischemic heart problems.

To summarize, we have discussed the available options for CTE in the current review, which include the use of patches and cells, CTE approaches, scaffold-less procedures, cell assembly, ECM decellularization, and neovascularization. The employment of conventional hydrogels in CTE with the existing limitations of their use was highlighted. Then, we went over the various kind of SRHs that were exploited for CTE in great detail. A great deal of effort has gone into developing more accurate hydrogel-dependent scaffoldings for CTE applications during the last few decades. As a result, the planning, manufacturing methods, characteristics, and utility of SRHs have undergone a paradigm change. By changing the structure of hydrogels, researchers hope to increase their biological, chemical, and mechanical properties. As a result, they are among the most commonly utilized matrices in healthcare. Several attempts have been made to create SRHs for specific purposes by carefully assessing their surrounding microenvironments and leveraging these efficient improvements for future usage. The development of gels that respond to numerous stimuli and mimic the natural 3D framework is a topic that can be researched further in the future. The use of diverse polymers in the building of various hydrogels, as well as the design of new hydrogels sensitive to extrinsic stimuli, are examples of these scientific breakthroughs. Furthermore, the development of SRHs with crucial physicochemical properties that are both robust and efficient is still under investigation. Future research will focus on creating unique intelligent hydrogels with pre-programmed, even complex self-bending, twisting, and folding behaviors. Despite significant advances in hydrogel synthesis methods, biodegradation rate, microstructures, external hybridization, and immunological responses remain potential manufacturing roadblocks. As a result, more attention should be placed on designing gels that can control and decrease immune responses. The use of several solvents during the manufacture of SRHs raises the risk of toxicity. As a result, spontaneous polymer amalgamation during crosslinking is a crucial area for further research. Future study into the biodegradation of hydrogels, as well as the management of their levels of cell growth and adhesion, is necessary. Future research should concentrate on shrinking the gel matrix down to single cells, enabling cell-by-cell customization of sequencing data. This is required in order to have a better understanding of the biological subgroups in the sample. Furthermore, for smartly regulated drug delivery, changing the frequency and speed at which drugs and biomolecules are released from smart hydrogels is critical. The creation of enhanced electro-sensitive scaffolds with quicker reaction times should be the focus of future research. Platforms that respond to antigens or antibodies are great smart materials for biomedical diagnostics and CTE. Swelling-deswelling transitions are caused by either antibody or antigen. The way to fabricate smart hydrogels that react to antigen-antibody interactions has not yet been documented. As a result, further study on this subject appears to be intriguing. Furthermore, the utility of antibody-conjugated hydrogels has been demonstrated in several research. The production of SRHs with antibodies incorporated in their composition is yet a future research concern. As the design and function of SRHs improve, the cell-scaffold interactions will become more obvious. This will pave the way for new CTE goals to be fulfilled through the use of such enticing platforms. These bright hydrogels might provide a safer, more effective, and possibly feasible alternative for a variety of biological applications, including but not limited to early disease detection, prevention, and TE, including CTE. In addition, further research is needed before they may be used in therapeutic settings.

3D: Three-dimensional; Alg.: Alginate; BASCs: Brown adipose-derived stem cells; CH: Chitosan; CMs: Cardiomyocytes; CNTs: Carbon nanotubes; CTE: Cardiac Tissue Engineering; CVDs: Cardiovascular diseases; dECM: Decellularized extracellular matrix; ECM: Extracellular matrix; EHT: Engineered heart tissue; ESCs: Embryonic stem cells; GelMA: Gelatin Methacrylate; Gt: Gelatin; HA: Hyaluronic acid; LV: Left Ventricular; MI: Myocardial Infarction; MSCs: Mesenchymal stem cells; PAA: Poly(acrylic acid); PAAm: Poly(acrylamide); PAni: Polyaniline; PBA-PGA: γ-Polyglutamic acid; PDGF‐BB: Platelet-derived growth factor‐BB; PEG: Polyethylene glycol; PNIPAAm: Poly (N-isopropylacrylamide); PPy: Polypyrrole; rGO: Reduced graphene oxide; SRHs: Stimuli-responsive Hydrogels; SWCNTs: Single-wall carbon nanotubes; TE: Tissue Engineering; US: Ultrasound; VEGF: Vascular endothelial growth factor.

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Authors' contributions:

E-H.H.M., and M.E.A. equally contributed to this work. E-H.H.M. conceived the idea and discussed it with other authors. E-H.H.M., M.E.A., and W.A.E-D. wrote the original draft. E-H.H.M., and M.E.A. prepared the figures. E-H.H.M., M.E.A., A.S.D., and R.T. revised the manuscript, supervised the work, and edited the draft. All authors have read the manuscript and agreed to the submission.

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Authors' information:

¹Laboratory of Veterinary Surgery, Department of Veterinary Medicine, Faculty of Agriculture, Tokyo University of Agriculture and Technology, 3-5-8 Saiwai Cho, Fuchu-shi, Tokyo 183-8509, Japan. ²Department of Surgery, Anesthesiology, and Radiology, Faculty of Veterinary Medicine, Benha University, Moshtohor, Toukh, Elqaliobiya,13736, Egypt.³Laboratory of Veterinary Physiology, Department of Veterinary Medicine, Faculty of Agriculture, Tokyo University of Agriculture and Technology, 3-5-8 Saiwai Cho, Fuchu-shi, Tokyo 183-8509, Japan. ⁴Department of Animal Hygiene, Behavior and Management, Faculty of Veterinary Medicine, Benha University, Moshtohor, Toukh, Elqaliobiya 13736, Egypt. ⁵Department of Pharmaceutics and Industrial Pharmacy, Faculty of Pharmacy, Badr University in Cairo (BUC), Badr City, Cairo 11829, Egypt. ⁶Department of Biochemistry, Faculty of Pharmacy, Badr University in Cairo (BUC), Badr City, Cairo, 11829, Egypt.

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Stimuli-responsive Hydrogels: Smart State of-the-art Platforms for Cardiac Tissue Engineering

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Abstract

Figures

1 Introduction:

2 Current Regeneration Strategies For Cardiac Tissue Engineering:

2.1 Scaffoldings and cells:

2.2 Engineering of the cardiac tissue:

2.3 Cell sheet technique (scaffold-less cell):

2.4 Cell Assembly:

2.5 Heart matrix decellularization:

2.6 Neo-vascularization:

2.6.1 In vitro neo-vascularization:

2.6.2 In vivo neo-vascularization:

3 Hydrogel Polymeric Scaffolds:

3.1 Classifications of hydrogels:

3.2 Limitations to using conventional hydrogels in cardiac tissue engineering:

4 Stimuli-responsive Hydrogels For Cardiac Tissue Engineering:

4.1 Physical stimuli-responsive hydrogels:

4.1.1 Temperature-responsive hydrogels:

4.1.2 Light/photo-responsive hydrogels:

4.1.3 Electro-responsive hydrogels:

4.1.4 Magnetic-responsive hydrogels:

4.1.5 Pressure/mechano-responsive hydrogels:

4.1.6 Ultrasound/acoustic-responsive hydrogels:

4.2 Chemical stimuli-responsive hydrogels:

4.2.1 pH-responsive hydrogels:

4.2.2 Ionic strength-responsive hydrogels:

4.2.3 Redox-responsive hydrogels:

4.3 Biological-responsive hydrogels:

4.3.1 Enzyme-responsive hydrogels:

4.3.2 Antigen/antibody-responsive hydrogels:

5 Conclusion And Future Prospects:

Abbreviations

Declarations

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References

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