Heart failure (HF) is the number one cause of morbidity worldwide [1] . Although HF could be treated with pharmacological administration or implantation of left ventricular assist devices, heart transplantation still remains the only viable approach for the treatment of HF [2] . However, the clinical success of heart transplanta-tion has been notably hindered due to the scarcity of donor hearts and immune rejection upon transplantation [3] . A single myocardial infarction (MI) can result in the loss of 25% of 2-4 billion cardiomyocytes (CMs) present in the left ventricle (LV) [4] . In contrast to the hearts of lower vertebrates such as adult zebrafish, neonatal mice, and neonatal swine, the adult human heart possesses a significantly limited regenerative capacity with a minimal annual replication rate of CMs ( < 1%) [5] .
After MI, the LV's healing process is mediated by a programmed cellular response that clears dead cells and matrix debris upon the development of the fibrotic scar tissue [6] . Numerous cellbased as well as cell-free strategies, have been proposed to date to promote regeneration of the injured host tissue upon MI [7] . Examples of cell-based therapies mainly include direct intramyocardial injection of non-CMs such as mesenchymal stem cells (MSCs) [ 8 , 9 ], cardiac stem cells (Sca-1 + cells) [ 10 , 11 ], as well as exogenous cells such as human pluripotent stem cell-derived CMs (hPSC-CMs) [12][13][14][15] , which often result in myocardial repair to some extent. For instance, using a non-human primate MI model, intramyocardially injected human embryonic stem cell-derived CMs (hESC-CMs) into infarcted hearts resulted in remuscularization and electromechanical synchronization [13] . In another study, delivery of the allogenic iPSC-CMs improved cardiac function three months post-transplantation [15] . However, in the majority of these studies, it has been reported that only about 1% of injected cells were grafted at one month post-injection into the host myocardium [16] . Furthermore, ventricular arrhythmias and potential teratoma formation are the other major drawbacks of cell-based strategies through bolus injection of cells [17] . Cell-free techniques include, but are not limited to delivery of microRNA, exosomes [18][19][20] , angiogenic growth factors (.g., VEGFA-B, FGF, PDGF-BB) [21][22][23][24][25][26][27] , as well as acellular injectable hydrogels [28][29][30][31][32] or patches [33] . However, despite improved cardiac function, reduced fibrosis, and increased vessel density [ 34 , 35 ] reported in these studies, full regeneration of the injured myocardium has not been reported yet, driving the need for developing new and innovative regenerative medicine-based strategies for the treatment of MI.
The engineering of therapeutic biomaterials has been an appealing strategy to enhance the efficacy of cell-based therapies or growth factor delivery. Numerous studies have shown that engineered biomaterials can improve retention, survival, and engraftment issues often seen with growth factors delivery or cell therapies [36][37][38][39][40] . Naturally-derived biomaterials such as fibrin [41][42][43][44] , collagen [45][46][47][48][49][50][51] , chitosan (CHI) [52][53][54][55][56] , gelatin [57][58][59] , alginate [ 44 , 60-64 ], hyaluronic acid [65][66][67] , and extracellular matrix (ECM)-derived matrices [ 32 , 41 , 68-72 ], have been widely utilized as scaffolding biomaterials for revascularization and regeneration of ischemic myocardium. Alternatively, commonly used synthetic matrices for treatment of ischemic myocardium include poly(ethylene glycol) (PEG)-based copolymers [ 22 , 25 , 26 , 73-81 ], polylactic-co-glycolic acid (PLGA) [22] , polycaprolactone (PCL) [ 74 , 82-84 ], poly(N-isopropyl acrylamide) (PNIPAAm) [ 24 , 27 , 85-93 ], etc.
Electrical conductivity is an inherent characteristic of the myocardium [94] . The conductive system within the heart comprises the internodal pathways, sinoatrial node (SAN), the bundle of Purkinje fibers, and the atrioventricular node (AVN). The action potential initiated in the SAN [95] propagate to the atrial, internodal pathway, and finally to the AVN. The generated action potential propagates to the ventricular myocardium through Purkinje fibers allowing the myocardial cells to experience these signals and contract synchronously [94] . Following MI, the remaining viable CMs are uncoupled and isolated in the non-conductive scar and fibrotic tissue, resulting in ventricular dysfunction and abnormal electrical signaling propagation [96] . However, delivery of non-conductive biomaterials in the form of injectable hydrogels or transplantable cardiac patches may inhibit the propagation of electrical signals across the infarcted myocardium and, preventing the synchronization of contracting CMs within the scar zone [97] . Admittedly, an electroconductive biomaterial may have the potential for mediating electrical signal propagation across the non-conductive fibrotic tissue. To that end, in the past few years, there has been a tremendous effort in the use of electrically conductive biomaterials in the regeneration of injured myocardium, a feature which was rarely considered in initial attempts of cell-based therapies or tissue engineering strategies.
In this review article, we will cover the application of electroconductive materials in the engineering of functional and biomimetic cardiac tissues for the treatment of MI (summarized in Fig. 1 ). We will cover the use of various types of nanomaterials, such as gold nanoparticles (GNPs), silicon-derived nanomaterials, carbon-based nanomaterials (CBNs) as well as electroconductive polymers (ECPs) in the development of mature and functional cardiac tissues in vitro . We will further broaden our discussion and will provide an in-depth overview of the in vivo use of electroconductive biomaterials and their delivery methods for cardiac repair after MI. Finally, we will discuss new avenues and our perspectives to establish better strategies for improved cardiac tissue engineering (cTE).
Tissue engineering has been shown as a promising strategy to recapitulate the topographical, mechanical, and ECM features of the native cardiac microenvironment [98] for regenerative medicine [7] as well as disease modeling applications [99][100][101] . As one of the key components, the proper design of biomaterials plays a crucial role in engineering functional and mature cardiac tissues [102] . The developed biomaterials for cTE must meet certain design criteria such as substrate stiffness and flexibility [103] , biocompatibility [104] , and appropriate pore size allowing cell and nutrient infiltration [105] . These features have been thoroughly reviewed and addressed in design of first-generation biomaterials such as decellularized matrix [ 106 , 107 ], collagen [108] , fibrin [41][42][43][44] , CHI [52][53][54][55][56] , gelatin [57][58][59] , alginate [ 44 , 60-64 ], etc. for engineering cardiac tissues. However, the poor electrical conductivity of previously reported biomaterials has been one of the major hurdles in achieving proper function in engineered cardiac tissues (ECTs). To make up this drawback, a new generation of electroconductive biomaterials has been explored by incorporation of conductive materials including gold nanoparticles (GNPs), silicon-derived nanomaterials, carbon-based nanomaterials (CBNs), and electroconductive polymers (ECPs). In this section, we will review the leading advances using these electroconductive nanomaterials or biomaterials for in vitro cTE applications.
Due to biocompatibility features, ease of fabrication process, and unique electrical characteristics, GNPs, including gold nanospheres (GNSs), gold nanorods (GNRs), and gold nanowires (GNWs), have been shown as promising nanomaterials in biomedical-related applications from imaging to diagnostics and notably in tissue engineering and regenerative medicine [109] . The specific emphasis on using GNPs in regenerative medicine has also emerged from their tunable physicochemical properties that could modulate tissue environment by providing cells with the necessary instructional cues for targeted biological responses such as proliferation, migration, and lineage-specific differentiation [110][111][112][113] . Numerous studies have demonstrated that the interactions of gold nanomaterials with CMs promote the electrophysiological response of ECTs, similar to the native myocardium ( i.e. , heart tissue). In this section, we will review the recent advancements in in vitro applications of GNPs for cTE applications.
Porous biomaterials have been widely used in cTE [36] , but one of their fundamental limitations is that their pore walls restrict cell-cell interaction and delay electrical signal propagation [114] . To address this issue, the pioneering work of Dvir et al. embedded GNWs within pore walls of microporous alginate scaffolds to enhance the matrix's electroconductivity ( Fig. 2 A) [115] . In an in vitro study, neonatal rat left ventricular cardiac cells (NRVCs) (composed of CMs and cardiac fibroblasts (cFBs)) were cultured on the conductive alginate hydrogel. Administration of stagewise electrical stimulation showed improved expression of connexin-43 (Cx-43), cardiac troponin I (cTnI) as well as sarcomeric α-actinin protein expression along with enhanced synchronous contractility of the cardiac tissues.
The slowed ventricular conduction velocity in cardiac arrhythmias has been associated with poor ventricular gap junctions' electrical coupling [116] . You et al. hypothesized that Cx-43 expression in CMs could be enhanced by using electroconductive scaffolds [117]. To that end, GNSs were homogeneously distributed in a tunable hydrogel scaffold synthesized by mixing thiol-2hydroxyethyl methacrylate (thiol-HEMA) with HEMA at different mass to volume (thiol-HEMA weight/HEMA volume, w/v) ratios to nanoengineer a conductive hydrogel (40% w/v thiol-HEMA/HEMA) with equivalent stiffness to the native myocardium during systole (60 0-160 0 kPa). Neonatal rat left ventricular myocytes (NRVMs) cultured on the conductive hydrogel showed improved Cx-43 expression with and without electrical stimulation compared with non-conductive scaffolds (0% w/v thiol-HEMA). Despite the promising outcomes, the conductive hydrogels were coated with fibronectin before cell seeding, which may mask the direct contact of cells and the conductive nanomaterials [117] .
GNWs and GNRs have also been successfully incorporated in photo crosslinkable gelatin methacrylate (GelMA) for cTE applications [ 118 , 119 ]. In a pioneer work, the Nikkhah group included GNRs into GelMA hydrogel to fabricate a hybrid and conductive UV-crosslink-able hydrogel ( Fig. 2 B) [119] . Electroconductivity as well as mechanical characteristics ( i.e. , stiffness) were significantly improved in the conductive GelMA compared with non-conductive GelMA. Furthermore, in vitro biological studies by seeding NRVMs on the nanoengineered hydrogels for seven days demonstrated significant enhanced cell adhesion, retention, and viability. Furthermore, the seeded cells exhibited improved expression of cardiacspecific markers ( i.e. , cTnI, Cx-43, and sarcomeric α-actinin) along with synchronous beating and enhanced calcium transient as a function of GNRs concentration [119] .
One of the critical challenges in cTE is fabricating a biocompatible scaffold without the risk of immune response and host rejection upon transplantation [32] . To address this issue and to better mimic the in vivo fibrous structure and architecture of myocardial tissue, Shevach et al. tested the efficacy of depositing a gold layer, with a thickness of 4 and 10 nm, on fibrous decellularized matrices derived from patients' omental tissues ( Fig. 2 Cad) [120] . NRVCs cultured on the conductive substrates for five days demonstrated enhanced expression of sarcomeric α-actinin and Cx-43 along with the formation of aligned and elongated cells ( Fig. 2 Ce-g). Notably, the engineered conductive tissues deposited with 4 and 10 nm gold layer exhibited stronger contraction force and faster calcium transient than the non-conductive pristine scaffolds. The same group further developed conductive nanocomposite coiled fiber scaffolds by electrospinning poly( εcaprolactone) polymer (PCL) followed by depositing gold to ensure the anisotropic electrical impulse propagation through the fabricated cardiac tissue [121] . After seven days, NRVCs cultured on non-conductive scaffolds showed rounded morphology with limited cell spreading. In contrast, cells cultured on the conductive scaffolds possessed an aligned and elongated morphology along with actinin striation.
Native heart tissue is organized intricately from macro-to nano-scale [122] . However, the majority of the previous work solely relied on a random mixture of GNPs within the scaffolds ( i.e. , only nano-scale cues) that lacked the necessary microscale mechanical cues. To address this need and integrated macro-to nano-scale cues within engineered tissues, in a work by Dvir's group, fibrous scaffolds were fabricated through elec- trospinning PCL-gelatin hybrid solution, followed by incorporating GNSs onto the surface of the scaffolds [123] . Compared with the non-conductive pristine scaffold, NRVCs cultured on the GNSincorporated scaffolds assembled into highly aligned and elongated cardiac tissues. In another recent study, Navaei et al. generated micro-topographical grooves (50 μm width and depth) on the surface of GelMA-GNR hybrid hydrogels to simultaneously introduce electrical and topographical cues into the engineered tissue ( Fig. 2 Da-c) [124] . Highly aligned cardiac tissues resembling the in vivo cardiac muscle bundle were formed upon culturing NRVCs on these engineered hydrogels. Enhanced expression of Cx-43 and sarcomeric α-actinin were also evident upon seven days of cell culture ( Fig. 2 Dd and e).
Cardiac tissues are primarily fabricated or engineered through the use of microfabrication techniques followed by culturing cardiac cells on the surface of the substrates. Although this method is straightforward and accessible, three-dimensional (3D) bioprinting is also a desirable technique since strides can be generated within highly organized tissues recapitulating the in vivo architecture of the myocardium [125] . In 2017, in a pioneer work, a printable conductive bioink was developed by incorporating GNRs in GelMA hydrogels for printing 3D cardiac tissues ( Fig. 2 Ea) [126] . The GNR concentration was optimized to generate a low viscous nanocomposite bioink allowing for the integration of cells into the bioink. NRVMs and cFBs were successfully encapsulated within the bioink and printed to engineer the organized cardiac tissue ( Fig. 2 Eb). The printed conductive cardiac tissue presented enhanced expression of Cx-43 and sarcomeric α-actinin, improved cell adhesion, and organization, along with enhanced synchronized contraction ( Fig. 2 Ec and d) [126] .
Although it has been hypothesized that GNPs might create conductive bridges across the biomaterials and enhance the electrical signal propagation [ 115 , 123 ], the mechanism by which these nanomaterials contribute to improved function and maturation of cTE has not been thoroughly and mechanistically investigated. To address this knowledge gap, recently, the Nikkhah group incorporated conductive GNRs and non-conductive silica nanoparticles with approximately similar sizes into GelMA hydrogel to examine the effect of conductivity of nanoparticles on functionalities and maturation of cardiac tissues. Surprisingly, in both nanoengineered hydrogels ( i.e. , conductive vs. non-conductive), enhanced cell retention and expression of Cx-43, sarcomeric α-actinin, and cTnI were observed, suggesting that conductivity of the scaffold per se is not the only parameter by which the GNRs contribute to the enhancement of functionalities of the ECTs. Nano-scale roughness as well as the mechanical properties of the scaffolds are the other key parameters which influence the biological function of cardiac cells [127] . Due to the undeniable influence of the mechanical properties of the scaffolds on cells, in another work, a biocompatible hybrid biomaterial was developed by incorporating GNRs into collagen hydrogel (GNR-Col) [128] . The GNR-Col scaffold with appropriate stiffness properties (7.25 ± 0.08 kPa) efficiently regulated intercalated disc assembly and formation in the cultured NRVMs. Further mechanistic studies showed that the improved formation of intercalated discs was partly associated with the ILK/p-AKT/GATA4 pathway mediated by β1 integrin. Admittedly, improved β1 integrin expression was also previously shown in the work of Navaei et al., demonstrating that GNRs could activate the mechanical pathway [119] . However, more studies are required in this area to delineate the detailed mechanism of improved functionalities of ECTs in the presence of GNPs. In addition, the majority of the previous studies in this regard have relied on the use of rat-derived CMs. To that end, more work is needed to assess the role of GNPs in biological response ( i.e. , protein and gene expression), maturation, and contractility of hPSC-CMs.
Silicon nanoparticles have been introduced for diverse applications in drug delivery and tissue engineering [129] . Their mesoporous structure allows pre-loading the nanoparticles with drugs, growth factors, and other molecules of interest that enhance their interaction with cells and engineered tissues [130] . However, one of the most attractive features of silicon nanomaterials has been the possibility of adjusting their electroconductive properties by introducing small impurities (dopants) in their lattice structure [131] . Particularly, due to their size, morphology, optical and electroconductive properties, silicon nanowires (SiNWs) have been of recent interest for the fabrication of ECTs. For instance, Rotenberg et al. created a hybrid system utilizing SiNWs internalized by neonatal rat myofibroblasts and CMs to introduce optical pacing and investigate cellular electrical coupling of the cells ( Fig. 3 Aa-c) [132] . When light pulses were applied to the hybrid myofibroblasts-SiNWs (MF-SiNWs), it was possible to indirectly stimulate and pace adjacent CMs in in vitro cocultures. Additionally, it was demonstrated that injection of MF-SiNWs in the native cardiac tissue was well tolerated, without significant signs of fibrotic encapsulation ( Fig. 3 Ad). With this, the authors demonstrated that it is possible to integrate minimally invasive optoelectrical silicon-based nanomaterials for cardiac tissue engineering applications.
Mei's group utilized SiNWs for the fabrication of scaffold-free ECTs. Specifically, electroconductive SiNWs (eSiNWs) with a diameter of 100nm and length of 10 μm ( Fig. 3 Ba) were used along with hiPSC-CMs into the fabrication of cardiac spheroids [133] . To achieve this, a suspension of hiPSC-CMs and eSiNWs (1:1) was created and deposited in agarose microwells. The low-adhesive surface of the agarose promotes cellular aggregation, resulted in a composite microtissue, with eSiNWs located in the intercellular space ( Fig. 3 Bb). The electroconductive microenvironment created by the addition of the eSiNWs promoted enhanced cellular structures ( Fig. 3 Bc), denoted by increased expression of sarcomeric αactinin, Cx-43, and β-myosin heavy chain. Additionally, higher synchronization of spontaneous contraction and increased sarcomere alignment were attributed to the addition of eSiNWs. In a subsequent study of the authors, it was shown that the combination of eSiNWs and external chronic electrical stimulation (2.5V/cm, 1Hz, 5ms) resulted in the formation of more mature cardiac spheroids ( Fig. 3 Ca-c) [134] . Enhanced expression of the genes involved in the contractile machinery of hiPSC-CMs ( i.e. , MYH7, increased MYH7/MYH6 ratio, and MYL2) was observed, along with improved cell-cell junction formation, measured by expression of Cx-43 and N-cadherin ( Fig. 3 Cd). It was speculated that increased availability of cell-cell junction proteins would lead to better engraftment to the native myocardium. However, the size of these microtissues was limited by oxygen and nutrient diffusion [135] . Thus, delivery and treatment of largely affected injured areas could be a challenging task.
Despite the advantages and significance of SiNWs, their in vivo application and use need to be thoroughly investigated. For example, biodegradation and clearance pathways of silicon and siliconderived nanoparticles need to be investigated in detail. Additionally, morphology, size, mesostructure, surface chemistry, among several other features, need to be carefully optimized to avoid acute cytotoxicity and bioaccumulation in non-targeted organs [136] .
Nanosize carbon allotropes, including carbon nanotubes (CNTs), graphene-based nanosheets, carbon nanohorns, and carbon nanofibers, collectively known as carbon-based nanomaterials (CBNs), have been famous for their outstanding properties. Owing to their mechanical, topological, and electrical properties, this group of materials have gathered significant attention in tissue engineering applications [ 112 , 137-140 ]. For example, a recognized avenue to address the need for electroactive features in engineering tissues such as cardiac, muscle, and the nerve is incorporating CBNs in commonly used non-conductive biomaterials [ 139 , 141 , 142 ]. In other words, including CBNs into tissue-engineered scaffolds is a facile one-step approach to confer refined properties such as higher physical strength, biological activity, and conductivity to the bulk material, which ultimately can direct cells to form electrically conductive networks [ 138 , 143 ]. In this section, the recent advances in the application of CBNs to engineering scaffolds to mimic the myocardium microenvironment and ultimately replace infarcted heart tissue are discussed.
CNTs known as cylindrical hollow nanostructures with diameters ranging from below 1 to 100 nm [ 138 , 141 ] have shown great potential in the maturation of CMs [144][145][146] . In a work by Martinelli et al., the cultured NRVMs on multi-wall carbon nanotubes (MWCNTs) substrates showed enhanced viability, proliferation, and electrophysiological properties [144] . It was demonstrated that MWNTs significantly improved expression of sarcoplasmic reticulum Ca2 + ATPase 2a cardiac-specific markers ( i.e. , Cx-43 and sarcomeric α-actinin) in NRVMs and functional syncytia. These improvements endowed by CNTs could also safeguard the ECT against pathological hypertrophy, modeled by phenylephrine treatment [146] . One step further, it was demonstrated that the incorporation of CNTs in GelMA hydrogel formed fibril-like bridges between scaffold pores [ 143 , 145 ] that provided higher mechanical and electrical properties. In addition, the engineered CNT-embedded scaffolds imparted higher adhesion of NRVMs on the nanocomposite film, more uniform 2D cell distribution, higher cell retention and viability, and cell alignment index [ 145 , 147 ]. Notably, CNTs inclusion (1 mg/mL) also led to enhanced electrophysiological charac-teristics in terms of higher spontaneous beating rates, while the excitation threshold for NRVMs cultured on hybrid CNT-GelMA was 85% lower in comparison to the control condition (pristine GelMA). CNTs also affect the CMs phenotype demonstrating higher expression of sarcomeric α-actinin and cTnI and better-distributed Cx-43 enhancing contractile features of cultured tissue [ 145 , 147 ]. The work by Shin et al. demonstrated that incorporated CNTs preserved cardiac tissue function against heptanol (that disrupts gapjunctions and synchronized beating) up to 60 min by providing a non-cellular network for electrical signal propagation in the GelMA matrix [145] . It has also been reported that the aligned CNTs in GelMA hydrogel could induce cardiac differentiation of mouse stem cell-derived embryoid bodies that were further promoted by an external electrical stimulation [148] .
Another family member of CBNs, a powerful counterpart for CNTs in reconstructing electroactive tissues, particularly myocardium, is graphene-based nanosheets [ 149 , 150 ]. This group of nanomaterials, including graphene (G), graphene oxide (GO), and reduced form of graphene oxide (rGO) nanosheets, is formed of a single layer of carbon atoms connected in the form of 2D hexagonal honeycomb lattice [ 112 , 151 , 152 ]. Like CNT, the graphene family has been demonstrated to improve mechanical and electrical properties while supporting cellular function within cardiac tissues [153][154][155][156] . It has been demonstrated that the NRVMs cultured on rGO-GelMA hydrogel had enhanced cardiac function in comparison to GelMA and even GO-GelMA hydrogel sheets, proved by higher cardiac marker expression, higher spontaneous beating rate, and stronger contraction [143] . Notably, only the cardiac construct on rGO-GelMA showed a response to the external electrical stimulation. These observations have been attributed to the lower electrical resistance in hydrogels incorporated by rGO and the higher ability of rGO in protein absorption compared to GO, which enhances cellular adhesion [ 143 , 150 ]. In a recent study, the insertion of graphene, as the most conductive material among CBNs, in collagen resulted in enhanced expression of MF20 (sarcomeric myosin) and cTnT (cardiac troponin T) in murine embryonic stem cell-derived CMs and led to significantly higher metabolic activity [149] . To draw a comparison between these siblings, 2D cell seeding on GO-GelMA substrates displayed faster 3T3 fibroblast attachment and higher cell spreading with polygonal shapes than CNT-GelMA films. This observation was attributed to the 2D topological characteristics of GO and higher surface roughness, compared to the 1D shape of CNTs, which dictates the cells' focal adhesion points and sizes [157] . However, in a recently published study by Lee et al., it was revealed that NRVMs cultured on CNTand rGO-GelMA compared to GO-GelMA thin films showed higher elongation and more local alignment with closer morphology to cardiac muscle cells and enhanced expression of cardiac markers ( Fig. 4 Aa-c) [143] . Even though the authors have noted a few reasons for this observation, more detailed and mature investigations are required to uncover the underlying mechanisms affecting cell behaviors. In an interesting approach, the electrophysiological features of cardiac tissues developed on CBN nanocomposites were evaluated by measuring spontaneous action potential. The authors demonstrated that the NRVMs matured on CNT-GelMA films had expressed more ventricular phenotypes (assessed by depolarization and repolarization pattern) ( Fig. 4 Ad), which is of great demand as MI is usually followed by ventricular dysfunction. From a conductivity perspective, the sheet resistance of rGO biohybrid scaffolds ( ∼1 k Ω /sq.) was shown to be approximately 100 times lower than CNT counterparts ( ∼100 k Ω /sq.) [143] . However, compared to graphene, studies denoted that the percolation threshold of CNT nanocomposites is about 85% lower than that of the graphene-based nanocomposite, implying that CNTs provide much higher electrical conductivity than graphene nanosheets with the same loading content [158][159][160] . Nevertheless, the best conductiv-ity has been achieved when a combination of CNT and graphene nanosheets was applied [ 161 , 162 ]. This hybridization strategy also helps with higher mechanical properties and lower agglomeration [ 151 , 163 ]. Therefore, we anticipate that future studies will focus on engineering bio-constructs as cardiac patches based on hybridizing CNT and graphene-based nanosheets to provide higher mechanical and electrical properties.
Recognizing the native myocardium as a flexible and stiff tissue with nanofibrous architecture has drawn great attention for the fabrication of electrospun fibrous mats recapitulating proper structural, mechanical, and electrophysiological characteristics of the native myocardium. In order to achieve these features, CBNs have been incorporated in a wide range of biocomposites including poly(glycerol-sebacate) gelatin (PGS-gelatin) [164] , CHI-poly(vinyl alcohol) [165] , silk [166][167][168] , poly(ester-amide)-CHI [169] , PCL-CHI-PPy [170] . The inclusion of 1.5% GelMA-coated CNTs within nanofibrous scaffold caused a significant improvement in mechanical toughness (7-fold) and finer fibers with an improved orientation that mimics the anisotropic architecture of the myocardium [164] . From a biological perspective, better CMs alignment and lower excitation threshold (3.5-fold), higher cell viability, and cardiac-specific markers ( i.e. , Cx-43 and sarcomeric α-actinin) expression were observed in the effect of CNT incorporation [ 164-166 , 168 ]. Regardless of these improvements, there was still a big difference between the diameter of fabricated nanofibrous scaffolds (ranging between 200 to 400 nm) and the diameter of fiberlike native myocardium (10-100 nm). Therefore, studies evaluating the effect of fiber diameter in the range closer to the native tissue are highly desired. Moreover, the thickness of the developed electrospun scaffolds was significantly lower than that of the infarcted tissue that needs to be replaced. Therefore, future studies are encouraged to focus on the layer-by-layer assembly strategies to stack the living electrospun mats together to get closer to the infarct thickness. In a similar context, inspired by the gradual transition of fiber alignment within the native myocardium (from endocardium to epicardium), Ma's group developed CNT-incorporated nanofiber networks within a GelMA hydrogel shell ( Fig. 4 Ba-c) [167] . They could ultimately achieve a two-layer network interwoven with an orthogonal orientation, with a multitude of cell layers having a gradual orientation transition ( Fig. 4 Bd and e). Further studies developing CBN-embedded scaffolds to mimic the native architecture and their in vivo function are still in high demand. To this end, biofabrication techniques such as 3D bioprinting with the potential to place cells and materials in a precise 3D high resolution with complex architecture can be of great interest.
In addition to incorporating conductive nanomaterials within scaffolds, several works that have relied on ECPs in the engineering of cardiac tissues. In this section, we will review the key ECPs which have been specifically used in engineering cardiac tissue constructs. Polypyrrole (PPy), Polyaniline (PANI), and poly(3,4ethylene dioxythiophene) polystyrene sulfonate (PEDOT:PSS) have been some of the most common ECPs which have been extensively used in several biomedical applications [171][172][173][174][175] . The structure of these ECPs contains conjugated π electron systems allowing for tuning their electrical properties. To that end, the ability to tune the electrical conductivity of scaffolds attracted tremendous attention, specifically in tissue engineering [ 110 , 175-178 ]. Their electroconductive nature makes them useful for neuronal and cardiac related studies [ 179 , 180 ]. Specifically, the addition of ECPs can better recapitulate the native cardiac tissue microenvironment by improving scaffold electrical properties and conductive cellular network formation [ 171 , 178 , 179 , 181 ]. The following sections focus on the recent progress in utilizing ECPs to enhance cellular ( i.e. , expres- sion of cardiac biomarkers, CM maturation) and tissue-level properties ( i.e. , beating, conduction velocities, Ca wave propagation) of scaffolds meant for cTE applications.
PANI is an attractive type of ECP for cTE due to its ease of synthesis, cost efficiency, and proper stability [182] . However, PANI suffers from limited cell-binding motifs and loss of conductivity in prolonged culture conditions [183][184][185] . Bidez et al. showed the efficacy of PANI film for adhesion and proliferation of cardiac H9c2 myoblast while maintaining its conductivity up to100 h in the culture medium [186] . However, integration of PANI into soft cardiac tissues is limited because of its biodegradability and inflexibility features [ 142 , 184 , 187 ]. To address this limitation, PANI has been blended with other polymers such as gelatin [185] , PCL [187] , polyurethane (PU) [188] , polylactic acid (PLA) [171] , poly(lacticco-glycolic acid) (PLGA) [189] , and PGS [ 190 , 191 ]. PANI/PLA scaf-folds made by either uniaxial and coaxial electrospinning not only promoted adhesion but could also modulate the shape of NRVCs [192] . Similarly, Wang et al. observed partially aligned and wellconnected sarcomeric structures on electrospun PANI/PLA scaffolds ( Fig. 5 A) [171] . Alternatively, blending PANI/PU/PGS was demonstrated to result in scaffolds with well-interconnected pores ( > 150 μm) with suitable compression moduli and mechanical strength for the culture of CMs [188] . Within this blended scaffold, even without electrical stimulation, NRVMs cultured on the conductive substrates expressed higher cardiac-specific genes ( i.e. , actin alpha 4 and troponin T-2) than non-conductive scaffolds and controls. Similarly, Hu et al. developed elastic and electroactive PGSco-aniline trimer films ( Fig. 5 Ba) [190] . These micropatterned ( i.e. , 50/50 μm groove/ridge dimension) films promoted synchronous calcium transients in cultured NRVMs. Enhanced expression of cardiac-specific markers (sarcomeric α-actinin and Cx-43) and improved cell alignment and elongation along the patterned surface were also observed that could support anisotropy of cultured CMs ( Fig. 5 Bc andd). In addition to blending, another viable strategy to improve the properties of PANI scaffold is by various polymer doping [ 189 , 193 , 194 ]. For instance, Hsiao et al. fabricated electrospun meshes of conductive scaffolds comprised of PANI and poly(lacticco-glycolic acid) (PLGA) followed by HCl doping ( Fig. 5 Ca) [189] . The doped meshes acquired positive charges, thereby electrostatically attracting negatively charged laminin and fibronectin proteins that in turn enhanced cell adhesion, supported synchronous beating of cultured NRVMs with an improved expression of Cx-43 and sarcomeric α-actinin ( Fig 5 Cb andc).
The other widely used ECPs in tissue engineering applications is PPy [ 98 , 181 , 195 ]. Previous studies have shown that PPy is amenable to surface modification ( i.e. , bioactive molecule functionalization), supports neuron, fibroblast, and endothelial cell adhesion and proliferation and exhibits biocompatibility both in vitro and in vivo [196][197][198][199] . These features have motivated utilizing PPy conductive scaffolds in cTE and myocardial repair applications [ 180 , 200 , 201 ]. However, the poor mechanical properties initially made it difficult to use PPy for scaffold fabrication [ 180 , 195 , 202 ]. To address these challenges, several research groups have tuned the concentration of PPy polymer blends to balance conductivity, biodegradability, and mechanical strength [ 142 , 181 , 200 , 201 ]. PPy composite scaffolds have been shown to improve conduction of ac-tion potentials, CM proliferation, cell-cell communication, as well as expression of Cx-43 gap junctions [ 181 , 196 , 203 ]. Similarly, incorporating PPy into other scaffolds such as PCL/gelatin [181] and PCL [204] has been shown to improve calcium wave propagation velocities and lower calcium transient duration in 2D CM cultures. Various strategies could be utilized to make composite polymers of PPy with improved properties [205][206][207] . For example, Tsui et al. developed electroconductive acid-modified silk fibroin-PPy (AMSF + PPy) substrates [207] . Nanosized grooves and ridges were patterned using composite biomaterials to recapitulate the ECM of the native myocardium. Human embryonic stem cell (hESC)derived CMs cultured on patterned AMSF + PPy substrates exhibited improved sarcomere development and intracellular organization. Furthermore, CMs cultured on AMSF + PPy substrates with the nanoscale topographical cues demonstrated a greater degree of Cx-43 expression. In contrast, improved troponin expression was not observed in the absence of PPy in the cells cultured on the substrates, even with topographical cues.
PEDOT:PSS is another type of ECP used in cTE [208][209][210] . For example, Roshanbinfar et al. developed hydrogels made from collagen, alginate, and PEDOT:PSS mimicking ECM fibular structures of the native myocardium [210] . The engineered scaffold exhibited improved electroactivity and promoted the maturation of NRVMs. Specifically, cells cultured on the conductive composite hydrogels showed enhanced expression of Cx-43, improved synchronous beating, enhanced contraction amplitude. Also, the engineered hy-drogel improved the beating properties and cell maturation when used with hiPSC-CMs.
In general, it appears that ECPs hold great promise in cTE applications; however, several challenges still persist. For example, to better mimic the mechanical properties and flexibility of the native myocardial tissue, the mechanical features of ECPs, including stiffness, need to be well-tuned. Specifically, PANI is not a flexible polymer [ 142 , 180 , 187 , 191 ], and PPy is challenging to handle due to its mechanical properties and brittleness [ 142 , 180 , 181 , 195 , 202 ]. As discussed earlier, blending them with other polymers has been successfully shown to enhance their material characteristics for the development of engineered tissues. The clinical application of these polymers is also limited due to their solubility, toxicity, and limited host tissue integration [ 171 , 172 , 174 ]. Many studies using ECPs rely on evaluating the behavior of CM monolayers cultured on thin substrates or electrospun sheets [ 171 , 180 , 192 , 193 , 211 ]. Methods need to be developed to recapitulate the 3D microenvironment of the native myocardial tissue for successful clinical translation of the fabricated conductive tissue. Additionally, converting from a conductive to a non-conductive form upon contact with the physiological or culture medium is a significant obstacle towards the application of these ECPs in tissue engineering [182] . To address this issue, for instance, Mawad et al. immobilized PANI doped phytic acid on CHI films, generating a thin conductive patch that retained its electronic stability in the physiological medium for more than two weeks [212] . Therefore, new strategies are still required to promote the stability and electroactivity of ECP over the construct's lifetime for long-term applications in bioelectronic devices and tissue engineering scaffolds.
Over the past two decades, injectable/transplantable biomaterials or patches have been widely investigated for the treatment of MI. Injectable natural biomaterials such as fibrin [43] , collagen [46] , alginate [213] , and CHI [54] have resulted in improved cardiac function, increased LV wall thickness, neovascularization, reduced scar size, and decreased fibrosis [214] . Similarly, there has been also numerous synthetic polymers composed of PNIPAAm [ 24 , 90 , 93 , 215-217 ] and PEG [ 76 , 79 , 82 , 84 ]. Biomaterials should be engineered based on specific criteria, such as material selection ( i.e. , biocompatibility, biodegradability), mechanical properties, chemical properties, etc., to lead to the desired outcome. A detailed explanation of various types of biomaterials has been reviewed elsewhere [214] . However, electrical conductivity has often been underestimated in designing biomaterials specifically for in vivo treatment of MI. The application of conductive biomaterials for treating MI stems from the fact that they may enhance the propagation of electrical signals throughout the non-conductive scar tissues [ 218 , 219 ]. This section summarizes the leading studies on the use of electroconductive biomaterials and their roles in cardiac regeneration after MI in vivo (summarized in Table 1 ).
Similar to in vitro studies, PPy has been extensively used for designing conductive biomaterials for MI treatment. In one of the earlier works, a blend of PPy with alginate was used for in vivo myocardial regeneration. Intramyocardial injection of the synthesized conductive polymer blend into the infarcted zone of rat hearts induced arteriogenesis and increased the migration of myofibroblasts into the injured zone without any significant inflammation upon five weeks post-injection when compared with the untreated and PBS-treated groups [202] . However, despite these improvements, the infarct size was not significantly different across the experimental groups five weeks post-injection. These improvements were partially attributed to the electrical conductivity of PPy as it could trigger differentiation of the cells responsible for arteriole forma-tion and restore the conduction within the infarcted tissue [202] . Despite promising outcomes, the underlying mechanism of such functional improvements was not further studied.
Due to its biocompatibility [220] , CHI has been extensively used as the base polymer for developing conductive injectable/transplantable biomaterials. In 2015, Mihic et al. tested the efficacy of grafting pyrrole to CHI to generate a semi-conductive hydrogel (PPy:CHI) for in vitro and in vivo studies [219] . NRVMs cultured on PPy:CHI resulted in enhanced Ca 2 + signal conduction compared with pristine CHI. Cell-free injection of PPy:CHI into the border zone of injured tissue 1-week of post-infarct, improved cardiac contractile function and conduction velocity measured eight weeks after hydrogel injection. Compared with the saline control, the scar size was significantly reduced in biomaterial treated hearts; while, no significant difference between CHI and PPy:CHI groups was observed [219] . However, electrical impulse propagation through the damaged zone was not investigated. In 2018, in another study using a cryoinjury model, intramyocardial injection of PPy:CHI into the injured region of the heart, seven days postinjury, enhanced electrical signals propagation across the infarcted tissue and improved cardiac function without inducing arrhythmias at four weeks post-injection [200] . In these two studies, the efficacy of PPy:CHI hydrogel was successfully shown once injected around the border zone or into the scar area; however, the detailed mechanism of action of the proposed hydrogel requires a further look. In a follow-up study, the same group investigated the mechanism by which the conductive PPy:CHI improves cardiac function [218] . In an in vitro model, synchronous contractions were induced in isolated clusters of NRVMs cultured on PPy:CHI hydrogels. Furthermore, intramyocardial delivery of PPy:CHI into fibrotic tissue seven days post-injury induced angiogenesis, reduced scar size, and enhanced cardiac function at three months postinjection without causing any arrhythmias. These improvements were mainly attributed to the enhanced conduction velocity across the fibrotic scar tissue due to the delivery of PPy:CHI hydrogels, leading to cardiac synchronized contraction.
Wang et al. used a dopamine crosslinker to distribute PPy nanoparticles into GelMA and poly(ethylene glycol) diacrylate (PEGDA) hydrogels to form a cryogel ( Fig. 6 Aa and b) [221] . The incorporated PPy nanoparticles (DOPA-based MA-G/PEGDA/PPY cryogel) resulted in increased Cx-43 and sarcomeric α-actinin expression in cultured NRVMs and enhanced synchronous contracting of the conductive engineered cardiac patch in vitro ( Fig. 6 Ac andd). After being cultured with DiI-labeled NRVMs for eight days, cardiac patches were sutured onto the epicardium of infarcted rat hearts. Transplantation of the engineered patch resulted in enhanced cardiac function, reduced scar size, and improved electric conduction within the infarcted zone at four weeks post-implantation. Furthermore, immunofluorescence (IF) staining of sarcomeric α-actinin was performed to assess the cardiomyogenic effects of the implanted cardiac patches. As indicated ( Fig. 6 Ae-i), DOPA-based MA-G/PEGDA/PPy cryogel was able to retain a higher number of delivered NRVMs four weeks post-transplantation, confirmed by DiI + CMs. It was speculated that PPy nanoparticles might have migrated from the donor patch and fused into the surface of the native CMs, similar to what they observed within in vitro experiments [221] . The migration of PPy nanoparticles was similar to another study [222] , where CNTs transferred from the scaffold into the host tissue and CMs in vivo .
Cardiac patch transplantation and intramyocardial injection of injectable biomaterials could alleviate the adverse effects after MI. However, each approach has its unique benefits, and full regeneration may not be achieved through a single approach. To that end, in a work by Wu et al., a conductive injectable hydrogel along with a self-adhesive cardiac patch were simultaneously administered to the injured heart in a combinato- rial approach [223] . Specifically, the bioadhesive hydrogel contained a mixture of dopamine-modified polypyrrole (DA-PPy) and gelatin-dopamine (GelDA) (GelDA/DA-PPy). In contrast, the injectable hydrogel was synthesized by a reaction between hydrazided hyaluronic acid (HHA) and oxidized sodium hyaluronic acid (HA-CHO) (HA-CHO/HHA). Intramyocardial injection of HA-CHO/HHA hydrogel and painting the electroconductive GelDA/DA-PPy hydrogel on the damaged tissue surface resulted in improved expression of Cx-43 and sarcomeric α-actinin, improved cardiac function, thicker LV wall, and enhanced angiogenesis upon four weeks post-injection.
In 2019, Song et al. developed a surgical suture using thermal plastic poly(glycolic acid) (PGA) coated with PPy, which could form an elastically conductive 3D network termed as biospring ( Fig. 6 Ba and b) [224] . In vitro culture of CMs on the biosprings resulted in increased sarcomeric α-actinin and Cx-43 expression and formation of elongated and aligned sarcomeres ( Fig. 6 Bc andd). Injection of biosprings (16 biosprings per injection) into the infarcted regions increased angiogenesis and reduced infarct size at four weeks post-injection. Compared with sham and non-conductive spring groups, injection of DOPA-coated PPy/PGA resulted in more sarcomeric α-actinin and Cx-43 positive CMs in the scar area ( Fig. 6 Be1,h2). It was speculated that the conductive 3D biosprings mediate the propagation of electrical signals from the healthy tissues into the infarcted tissue.
Although GNPs have been extensively used for cTE in vitro , they have rarely been used for in vivo MI treatment. In 2018, Hosoyama et al. tested the efficacy of incorporating GNSs into a collagenbased cardiac patch for in vitro biological studies and in vivo cardiac repair applications [225] . In vitro culture of NRVMs on the car-diac patch resulted in enhanced expression of Cx-43 under electrical stimulation. Transplantation of the patch into the infarcted mouse hearts via fibrin sealant seven days post MI resulted in reduced scar size, enhanced expression of Cx-43, and sarcomeric αactinin, as well as increased vasculogenesis within the infarcted area without inducing any ventricular tachycardia at 28 days posttransplantation.
In a different attem pt, a conductive injectable hydrogel was developed with equivalent stiffness to native myocardium by grafting poly-3-amino-4-methoxybenzoic acid (PAMB) onto non-conductive gelatin followed by crosslinking via carbodiimide (PAMB-G) [226] . Injection of the hydrogel into the scar area of rat hearts at 1week post-injury resulted in enhanced cardiac function, reduced scar size, and a thicker LV wall without inducing arrhythmias. It was speculated that injection of PAMB-G hydrogel enhanced propagation of electrical impulse across the scar tissue and helped activation of viable contracting regions within the scar. Recently, a conductive cardiac patch was also developed with equivalent conductivity to native myocardium comprised of a gelatin-based gelfoam grafted with PAMB for the repair of infarcted rat heart [227] . One week post-injury, the epicardial delivery of the cardiac patch loaded with NRVMs significantly improved cardiac function, enhanced propagation of electrical impulse, and reduced the arrhythmias upon four weeks of transplantation.
GO has also been extensively used to introduce electroconductivity features into the designed biomaterials for MI treatment. Paul et al. developed a soft GelMA-based injectable hydrogel to deliver functionalized GO sheets with vascular endothelial growth factor-165 (VEGF) pro-angiogenic gene (fGO VEGF /GelMA) for transfection of infarcted myocardium [228] . In a rat model of acute MI, injection of the proposed nano-complexed hydrogel in the periinfarct regions resulted in mitotic activities of endothelial cells, enhanced capillary density of myocardium, decreased scar size, and enhanced cardiac function at two weeks post-injection compared with the infarcted hearts injected with other experimental groups ( i.e. , sham, GelMA alone, fGO VEGF /GelMA). Although this study showed the efficacy of GO for myocardial gene delivery, the specific role of the GO conductivity was not assessed or isolated on overall improved cardiac function. To that end, despite the benefits of the proposed nano-complexed angiogenic hydrogel, the long-term impact of GO should be further evaluated in the host system [228] . Similarly, Zhou et al. designed an electroconductive injectable hydrogel by incorporating GO nanoparticles into oligo(poly(ethylene glycol) fumarate) (OPF) hydrogels [229] . The combination of GO and OPF hydrogels enhanced cell attachment in vitro . In a rat MI model, in addition to mechanical support, injection of the conductive OPF/GO hydrogels into the infarct zone resulted in improved Ca 2 + signal propagation, upregulation of Cx-43 expression, reduced infarct size, and enhanced neovascularization compared with OPF and PBS groups. In 2017, for the first time, Bao et al. developed a soft hydrogel by synthesizing a multi-armed PEGDA700-Melamine (PEG-MEL) using ππ conjugation ring [230] .
PEG-MEL was later crosslinked with thiol modified hyaluronic acid (HA-SH) (PEG-MEL/HA-SH). GO was further used to introduce elec-trical features to the designed hydrogel (PEG-MEL/HA-SH /GO). In a rat model of MI, intramyocardial injection of adipose tissuederived stromal cells (ADSCs), combined with the proposed conductive hydrogel, resulted in enhanced expression of Cx-43, improved heart function, thicker LV wall, and reduced scar at four weeks post-injection compared with sham, PBS, PEG-MEL/HA-SH, PEG-MEL/HA-SH/GO, and PEG-MEL/HA-SH/ADSCs groups. Despite the positive outcome of this study, GO nanoparticles' metabolism and biosafety are still a major concern upon injection into the damaged myocardium. To address this issue, the same group utilized tetraaniline (TA) in a recent study to partly enhance conductivity within the synthesized hydrogel while maintaining its biocompatibility [231] . Michael addition was used for the reaction of tetraaniline-polyethylene glycol diacrylate (TA-PEG) and thiolated hyaluronic acid (HA-SH) to generate a compliant hydrogel (TA-PEG/HA-SH) with conductivity resembling the native myocardium. They tested the efficacy of the proposed combinatorial approach towards heart repair by intramyocardial transplantation of the conductive hydrogel encapsulated with ADSCs and plasmid DNA encoding endothelial nitric oxide synthase (eNOs) nanocomplexes. The proposed approach promoted cardiac function, increased Cx-43 and sarcomeric α-actinin expression, reduced infarct size, and increased vessel density at four weeks post-transplantation. In addition, although the beneficial effects of the plasmid on neovas-cularization were demonstrated, the transgene's expression time should be controlled in future endeavors. In these two studies, including PEG-MEL/HA-SH/GO/ADSCs [230] and TA-PEG/HA-SH [231] , hydrogel treated hearts showed enhanced α-smooth muscle actin ( α-SMA) expression across the damaged zone. The authors speculated that the improved cardiac function has resulted from the increased population of fibroblasts within the infarcted region, promoting the transduction of mechanical signals within the myocardium. However, expansion of cFBs is not necessarily beneficial as they could differentiate into myofibroblasts and increase fibrosis within the scar tissue [232] .
In another study, a thermosensitive injectable hydrogel was developed by incorporating single-wall carbon nanotubes (SWCNTs) into PNIAPAAm hydrogel to improve bioactivities and adhesion of PNIAPAAm to encapsulated cells [85] . Preliminary in vitro studies demonstrated the PNIPAAm/SWCNTs hydrogel efficacy in promoting proliferation and adhesion of encapsulated brown adiposederived stem cells (BASCs). Furthermore, intramyocardial delivery of BASCs-laden PNIPAAm/SWCNTs enhanced cells' retention within the infarct region compared with PBS/BASCs group at one week post-injection. Four weeks post-transplantation, BASCs were differentiated to CMs, resulting in improved cardiac function and reduced scar size in PNIPAAm/SWCNTs/BASCs experimental group.
Despite innovative approaches and promising outcomes, shown in previous studies, with the use of injectable conductive biomaterials or transplantation of patches for the treatment of MI, the mechanism of action of these materials needs to be further studied. Multiple mechanisms have been speculated for the effect of conductive biomaterials in improved cardiac function and tissue regeneration in vivo . Some studies proposed that the improvements were partially attributed to the conductivity feature of the conductive biomaterials, which may promote electrical propagation across the scar regions as it can trigger the migration of cells into the infarct region and restoring the conduction of the infarcted tissue [ 200 , 218 , 223 , 226 , 227 , 231 ]. Electrical signal propagation within the cardiac muscle undergoes significant changes after MI. Specifically, after MI, the formation of fibrotic tissue causes disconnection of functioning and beating CMs resulting in delayed electrical signal propagation and the loss of tissue synchronicity [218] . To better investigate this mechanism, recently, He et al. designed an experiment to explain how electroconductive biomaterial ( i.e. , CHI-PPy) could act as a wire to restore electrical signal propagation within non-conductive and fibrotic scar tissue after MI [218] . They specifically assessed the energy loss and latency time of the generated electrical signals by heart tissue passing through an equivalent circuit in vitro . It was found that the latency time was 5-fold shorter for electroconductive CHI-PPY groups compared to the non-conductive gelatin groups confirming higher conduction velocity within the conductive biomaterials. In addition, electrical signal energy upon passing through the conductive CHI-PPY biomaterial was more than 2 times higher than the non-conductive gelatin. Injection of CHI-PPY into scar tissue resulted in reduced impedance and enhanced electrical conductivity. Specifically, higher field potential amplitude was observed in the scar tissue of the CHI-PPy injected hearts compared to the damaged hearts injected with nonconductive CHI. Moreover, microelectrode array analyses indicated that current failed to re-enter the scar region in non-conductive CHI-injected hearts, while in the conductive CHI-PPy injected hearts, the electroconductive signals well propagated throughout the scar region. Compared to CHI injected hearts, faster conduction velocity and reduced resistivity of the PPy-CHI injected hearts were also confirmed. Overall, the generated current by myocardium faced less resistance, and the energy loss was significantly lower in CHI-PPY injected hearts compared to the non-conductive injected group [218] .
In another study, Zhou et al. constructed three-dimensional conductive ECTs (c-ECT) by incorporating SWCNTs into gelatin hydrogel scaffolds to provide a conducive microenvironment for cTE in vitro [222] . Cell-laden c-ECTs were sutured onto the rat hearts' epicardial surface at two weeks post-infarct. Four weeks after implantation, the c-ECTs were integrated into the host myocardium, which led to the migration of transplanted SWCNTs and CMs into the scar regions. The formation of well-aligned cell bundles, and enhanced expression of N-cadherin and Cx-43 compared with sham and gelatin groups, were evident. The authors hypothesized that SWCNTs might have helped the fusion of c-ECTs into the host tissues while triggering the native heart cells to migrate into the infarcted areas. Additionally, the authors demonstrated infiltration of vasculature from the host tissue into the transplanted constructs one week after implantation. From a mechanistic point of view, the authors found that the expression levels of β1-integrin, ILK, p-AKT, and β-catenin were higher within the c-ECT groups compared with sham and non-conductive groups suggesting the possible activation of ILK/Akt-catenin pathway in the infarcted zone. Four weeks post-transplantation, CD68 positive macrophages were also shown to accumulate in the myocardium's scar area. Therefore, it deserves further investigation to delineate whether infiltration of macrophages contributed to the positive outcome of the engineered c-ECT.
An essential criterion of biomaterial-based heart regeneration after MI is to develop a less invasive approach for the delivery of the engineered biomaterials [214] . Conductive materials have been applied to the infarcted hearts primarily through direct intramyocardial injection or suturing of patches onto the epicardial surface. Through intramyocardial injection, the amount of injected biomaterial could be controlled in a minimally invasive approach. However, in preclinical studies using large animals ( e.g. , porcine), multiple injections are required to evenly distribute the biomaterial into the infarcted tissue [37] . Some biomaterials could also be delivered into the epicardial surface in the form of cardiac patches. Cardiac patches are advantageous because they could deliver more cells and therapeutic agents and mechanically support the LV wall [233] . However, cardiac patch delivery often requires invasive surgical procedures ( i.e. , suturing) that could increase the likelihood of inflammatory responses [234] . Multiple innovative strategies have been successfully developed recently to address these issues. For instance, in 2016, a suture-free technique was created to transplant cardiac patches through light illumination (photoadhesion) [212] . Specifically, a bioadhesive cardiac patch was fabricated by adding a photoactivated dye ( i.e. , Rose Bengal) into a CHI film ( Fig. 7 Aa-d). The bioadhesive cardiac patch was then successfully transplanted onto the epicardium surface after exposure to green laser ( λ = 532 nm) without any need for suturing. Two weeks post-transplantation, the cardiac patches remained firmly adhered to the heart tissues, confirming the viability and efficacy of this method for delivery of cardiac patches in a minimally invasive fashion ( Fig. 7 Ae andf). In 2018, Tal Dvir's group developed another suture-free technique to transplant cardiac patches into the heart tissues [235] . In this study, a cardiac patch composed of albumin electrospun fibers and GNRs was developed, loaded with cardiac cells ( Fig. 7 Ba-c). The cardiac patches were positioned on the rat epicardium surface and exposed to a near-infrared (IR) laser (808 nm). It was suggested that the GNRs converted the absorbed light into thermal energy, changing the molecular structure of the fibrous scaffold, which firmly adhered it to the wall of the heart ( Fig. 7 Bd-f). Although this approach was promising and innovative, the safety of the procedure for clinical cardiac therapy of patients could be further studied.
In another attempt, Liang et al. developed an adhesive paintable conductive hydrogel that could be painted on the surface of the epicardium in the form of a patch, avoiding the need for suturing or light exposure ( Fig. 7 C) [234] . This paintable conductive biomaterial was designed by two-step Michael addition reactions to sequentially introduce pyrrole groups and dopamine. First, the reaction of pentaerythritol triacrylate (PETA), dopamine hydrochloride, and PEGDA resulted in a hyperbranched poly(amino ester) polymer (HPAE). The second Michael addition step was performed to react the pyrrole monomer with acrylate groups of HPAE resulted in a conductive bioadhesive polymer (HPAE-Py). To this hyperbranched polymer, gelatin was further added to enhance the biocompatibility of the precursor solution. Fe 3 + as a multifunctional curing agent was then used to initiate the polymerization of pyrrole groups. The dopamine-Fe 3 + introduced bioadhesive feature while the PPy formed in situ introduced conductivity to the biomaterial. The delivered hydrogel (HPAE-Py/Gelatin) to the infarcted hearts infiltrated and fused within the host and firmly bonded to the beating heart within four weeks post-delivery. Compared to sham and HPAE/Gelatin groups, the conductive adhesive patch (HPAE-Py/Gelatin) resulted in reduced scar size and fibrosis, thicker LV wall, improved expression of Cx-43, and increased generation of new vessels within the scar tissue.
Electroconductive biomaterials have widely been used in various tissue engineering disciplines from neural [236] , skeletal [237] , to cardiac [142] . As reviewed above, the use of electroconductive elements such as GNPs, CBNs, or ECPs has been the key strategy to promote electrical signal propagation and conductivity within engineered tissues. The synthesized electroconductive scaffolds need to fulfill specific requirements, such as cytocompatibility and acceptable biodegradation rate, to successfully support the functionalities of engineered tissue. However, there are certain advantages or disadvantages associated with the use of each approach which will be briefly discussed below.
GNPs have been favorable for the use in cTE due to their acceptable biocompatibility, controllable size, and shape fabrication [238] . As an advantage, GNPs are biocompatible and could be synthesized in different geometries and sizes, such as nanospheres, nanorods, and nanowires. However, the effect of GNPs geometry and size on functionalities of ECTs remained to be mechanistically investigated. The electroconductivity features of the resultant scaffolds also depend on the type of GNPs and the base polymer. For instance, GNRs (average length, ∼60 nm) when incorporated into alginate hydrogels did not impart a significant improvement in terms of electroconductivity compared to the pristine alginate [115] . On the other hand, when GNWs (average length, ∼1.0 μm) were incorporated into the same hydrogel, a significant increase in electrical current was detected [115] . In contrast, the electroconductivity of GelMA hydrogel was significantly promoted when incorporated with GNRs (average length, ∼50 nm) [ 119 , 124 , 127 ] as functional of nanomaterial concentration. Although cytocompatibility of GNPs has been confirmed in several publications, their cytotoxicity has been shown to be dose-dependent [239] which poses a limitation in tuning the electroconductivity of the final synthesized scaffolds. Furthermore, incorporation of GNPs results in scaffolds with low mechanical strength compared to CBNs, where the lack of homogenous dispersion within the base hydrogel material may also limit specific applications that require high concentrations of these nanomaterials [238] .
Similar to GNPs, CBNs have also been extensively used in cTE for both in vitro and in vivo applications. However, CBNs pose potential issues such as hydrophobicity that may result in aggregation once incorporated into the base hydrogel materials [240] .
CNBs cytotoxicity yet remains to be another challenge that may lead to host immune reaction upon in vivo transplantation of the engineering tissues [142] . CBNs also appear in various shapes, sizes, and geometries that impact their toxicity. For example, the biocompatibility of graphene and its derivatives has been reported to depend on their concentration, shapes, and number of layers [ 241 , 242 ]. To that end, more investigations are required to unveil their side effects upon in vivo applications. In addition, the biodistribution of these nanomaterials (GNPs, CBNs) and accumulation within other organs upon injection have led to arguments over their safe use for clinical applications.
Apart from the key nanomaterials (GNRs, CBNs), ECPs such as PANI, PPy, and PEDOT:PSS have been widely used in cTE. In general, ECPs have been shown to have poor biocompatibility [243] , while their conductivity characteristics also decline over time within the physiological environment [182] . According to previous studies, the biocompatibility of ECPs is also dose-dependent. For instance, 30% PPy in PCL and gelatin was shown to exhibit cytotoxic effects [181] . Although there are certain benefits with the use of ECPs for engineering functional cardiac tissues, the longterm degradation rate, cytotoxicity, and carcinogenic effects may limit their clinical application for the treatment of injured myocardium [238] . In addition, the current knowledge regarding the potential immune response of the implanted ECPs in vivo is limited. In-depth mechanistic investigations are required to address the remained questions in this regard.
Owing to the inherent biocompatibility and degradation issues of the abovementioned electroconductive nanomaterials and polymers, recently, more attention has been devoted towards the development of new generation of biocompatible and electroconductive materials for cTE. For instance, in a recent study, He et al. synthesized a naturally aligned self-conductive cellulosebased scaffold from sea squirt, named tunic hydrogel, for cTE application ( Fig. 8 Aa) [244] . The self-conductive hydrogel was further doped with PPy (PTC hydrogel) to enhance its electroconductivity ( Fig. 8 Aa and b). The pristine tunic hydrogel, tunic hydrogel, and PTC hydrogel showed uniform current distribution; however, the current and conductivity within the tunic and PTC hydrogels were stronger than the pristine tunic hydrogel ( Fig. 8 Ac). NRVMs, after cultured for 7 days on the tunic and PTC hydrogels, showed improved orientation, enhanced Cx-43 and sarcomeric αactinin expression compared to the pristine tunic control hydrogel ( Fig. 8 Ad). Transplantation of the cell-laden cardiac patches fabricated from the tunic and PTC hydrogels into the rat infarcted hearts resulted in enhanced cardiac function, increased ventricle wall thickness and reduced infarct size, and angiogenesis. In addition, enhanced expression of Cx-43 and more organized sarcomeric α-actinin were observed within the patch transplantation site. Although PPy doping was shown to enhance the electroconductivity of self-conductive tunic hydrogel, it may still pose safety issues due to its non-biodegradability and side effects upon in vivo implantation. However, the tunic hydrogel itself with suitable conductivity could be a promising candidate for MI repair with minimal side effects.
In another recent work, Song et al. incorporated hydrophilic ionic polymer, namely polyacrylic acid (PAA), into oxidized alginate (OA)/gelatin (Geln) (OG) hydrogel to develop a porous ionic conductive hydrogel, named POG 1 , with appropriate biocompatibility and electrical conductivity (35.36 ± 7.72 × 10 -3 S/cm) equivalent to the native human heart tissues [240] . The ionic conductive hydrogel could be advantageous over traditional electroconductive components for cTE as they maintain electroconductivity under strains condition, avoiding the disconnection of electroconductive paths. Furthermore, this new class of the synthesized hydrogel introduced a homogenous electroconductivity within the scaffold without aggregated conductive phases, as opposed to the use of other traditional electroconductive materials ( i.e. , PPy, rGO, and CNT) ( Fig. 8 Ba). Compared to the non-conductive OG and CNT-OG with inhomogeneous conductivity, NRVMs cultured on the POG 1 hydrogel exhibited minimal cytotoxicity, significantly oriented and elongated sarcomeric α-actinin structure, enhanced Cx-43 expression ( Fig. 8 Bb), and improved synchronization. Notably, in a rat model of MI, implantation of cell-laden POG 1 engineered cardiac patches resulted in suppression of cardiac fibrosis ( Fig. 8 Bc), increased LV wall thickness, and improved cardiac function compared to control MI and OG groups at four weeks post-transplantation. The formation of new blood vessels is also a crucial factor for oxygen and nutrients supply within the infarct region for treatment of MI [245] . Several in vivo studies have reported the formation of angiogenesis within the infarct zone after injection of conductive hydrogels [ 202 , 218 , 224 , 229 , 231 ] or transplantation of conductive cardiac patches [ 212 , 222 , 225 ]. Similar to previous studies, in this work, transplantation of engineered POG 1 cardiac patches into the infarcted rat hearts resulted in angiogenesis within the scar tissue confirmed by immunostaining for vWF (micro-vessels) and α-SMA (arterioles) markers ( Fig. 8 Bd) [240] .
Engineered scaffolds or biomaterials that incorporate electrical cues may provide the cells with the necessary conducive microenvironment, resembling the native myocardium, for the formation of functional cardiac tissues in vitro . Over the past decade, various conductive materials, including GNPs, silicon-derived nanomaterials, CBNs, and ECPs, have been successfully utilized to endow electroconductivity features to semi-and non-conductive biocompatible materials (either natural or synthetic) for enhanced cTE and cardiac regeneration after MI. The idea of injecting conductive biomaterials into the scar area stems from the fact that the electroconductive biomaterial could electrically connect the isolated CMs across the scar tissue. However, the remaining viable CMs in the scar area might not be sufficient for full regeneration of the infarcted hearts. One possible solution could be the delivery of exogenous CMs via injectable conductive biomaterials. Although cell delivery has been shown to have beneficial effects to some extent, the maturation level of the injected CMs, the risk of immune rejection, and tumor formation are still a big concern [17] . To that end, an appropriately designed acellular biomaterial could provide the cells in the infarcted tissue with a new microenvironment to promote regeneration and electrical signal propagation, avoiding the risks associated with cell-based therapies [246] . Furthermore, ECM protein engineering approaches [247] could be practiced to develop conductive ECM-based biomaterials for the effective delivery of CM proliferation cues such as agrin, which has been recently found as a stimulator for CM proliferation [248] . Therefore, a combined strategy for developing conductive acellular injectable or implantable scaffolds for the proliferation of resident CMs could be a new approach for enhanced cardiac regeneration. Additionally, there exist some gaps in this field that are worthy of further investigation. Despite the advancements of the field and extensive reports on the beneficial impact of the hybrid conductive materials on cultured CMs, further mechanistic studies are needed to unravel the underlying molecular mechanisms that potentially influence the native-like structures and functionalities of the engineered tissues as well as molecular signatures of CMs such as protein and gene expression. Apart from mechanistic biological knowledge gaps, the fate and biodistribution of the conductive nanomaterials in vivo should be further unveiled. Moreover, the electroconductivity properties of the designed tissue constructs diminish in the physiologic environment [182] . Although previous studies have successfully investigated the electroconductivity up to two weeks [212] , further attempts are required to design electroconductive biomaterials with long-term effects for potential in vivo applications with enhanced efficacy. In the native tissue, cells are located in an ECM with a complex composition; however, most of the designed cardiac tissues incorporate only one or two ECM components. It has been shown that a multicomponent ECM scaffold could efficiently differentiate stem cells towards cardiomyogenic fate [249] . One possible route could be endowing electrical features to the cardiac tissues generated from an optimized combination of ECM proteins. Therefore, a conductive biomaterial with appropriate electrical and ECM cues ( i.e. , protein and biochemical composition) could enhance the functionalities of the cTE toward translational applications.
There are no conflicts to declare.