Effects of electrically conductive nano-biomaterials on regulating cardiomyocyte behavior for cardiac repair and regeneration

Abstract

Myocardial infarction (MI) represents one of the most prevalent cardiovascular diseases, with a highly relevant and impactful role in public health. Despite the therapeutic advances of the last decades, MI still begets extensive death rates around the world. The pathophysiology of the disease correlates with cardiomyocyte necrosis, caused by an imbalance in the demand of oxygen to cardiac tissues, resulting from obstruction of the coronary flow. To alleviate the severe effects of MI, the use of various biomaterials exhibit vast potential in cardiac repair and regeneration, acting as native extracellular matrices. These hydrogels have been combined with nano sized or functional materials which possess unique electrical, mechanical, and topographical properties that play important roles in regulating phenotypes and the contractile function of cardiomyocytes even in adverse microenvironments. These nano-biomaterials’ differential properties have led to substantial healing on in vivo cardiac injury models by promoting fibrotic scar reduction, hemodynamic function preservation, and benign cardiac remodeling. In this review, we discuss the interplay of the unique physical properties of electrically conductive nano-biomaterials, are able to manipulate the phenotypes and the electrophysiological behavior of cardiomyocytes in vitro, and can enhance heart regeneration in vivo. Consequently, the understanding of the decisive roles of the nano-biomaterials discussed in this review could be useful for designing novel nano-biomaterials in future research for cardiac tissue engineering and regeneration.

Statement of significance

This study introduced and deciphered the understanding of the role of multimodal cues in recent advances of electrically conductive nano-biomaterials on cardiac tissue engineering. Compared with other review papers, which mainly describe these studies based on various types of electrically conductive nano-biomaterials, in this review paper we mainly discussed the interplay of the unique physical properties (electrical conductivity, mechanical properties, and topography) of electrically conductive nano-biomaterials, which would allow them to manipulate phenotypes and the electrophysiological behavior of cardiomyocytes in vitro and to enhance heart regeneration in vivo. Consequently, understanding the decisive roles of the nano-biomaterials discussed in the review could help design novel nano-biomaterials in future research for cardiac tissue engineering and regeneration.

1. Introduction

Cardiovascular diseases (CVDs) encompass all pathological changes that affect the heart and blood vessels, i.e., coronary heart disease, hypertension, and atherosclerosis are all global leading causes of death. Myocardial infarction (MI) and stroke are mainly caused by a blockage of blood flow to the heart or brain due to calcium (Ca²⁺), fat, or other deposits on the inner walls of the blood vessels, causing damage to the heart muscle and inducing the loss of cardiomyocytes (CMs) [1] (Fig. 1). MI initiates a progressive sequence of morphological changes, comprising typical ischemic necrosis as a result of blood coagulation, followed by inflammation and repair that resembles the injury response of other tissue types (Fig. 1) [2]. As a result, heart dilation reduces muscle relaxation capacity (diastolic dysfunction), which can progress to the loss of contraction capacity (systolic dysfunction) and reduced electrical conductivity. The injured tissue must be treated as quickly as possible to avoid persistent ischemia and heart failure, which could severely compromise quality of life [3]. The immune response to MI is initiated by the release of various inflammatory growth factors and cytokines. Transforming growth factor β (TGF-β) isoforms inflect CMs survival, proliferation and hypertrophic responses, and direct the immune response [4]. This process also results in the transdifferentiation of cardiac fibroblasts (CFs) into cardiac myofibroblasts, and is a prominent factor in the fibrotic healing response. These cells and other resident CFs can account for the excessive secretion of extracellular matrix (ECM), resulting in stiff, non-contractile and non-conducting fibrotic scar tissue, which has very different characteristics compared with healthy cardiac tissue (Fig. 1).

Fig. 1.

Schematic diagram of the pathophysiological changes that occur after MI. First, lack of blood supply to the surrounding tissue causes ischemia and hypoxia. The hypoxic state results in ROS generation causing necrosis and apoptosis. In the cardiac remodeling phase, the CMs die, repairing inflammatory tissue and the healing process is initiated by the release of various cytokines and growth factors, by infiltrating immune cells. Fibroblasts are responsible for excessive production of the ECM, obtaining very different tissue characteristics from healthy tissue, consisting of a network of CMs and fibroblasts interspersed within the stiffer ECM, rendering the tissue non-functional. Created with BioRender.com.

Recent MI treatments have involved revascularization surgery, angioplasty, and stenting for revascularization purposes, while pharmacological treatments are used simultaneously for remodeling from the heart [5]. Nevertheless, these efforts are not sufficient to overcome the loss of CMs and achieve revascularization for the treatment of large-scale cardiac injuries. There are only a few treatment options that exist, such as heart transplantation and electronic pacemaker implantation, which are long-term strategies for the treatment of large-scale cardiac injuries. However, the number of heart donors is limited, and only a small fraction of patients on transplant waiting lists receive a new heart. The abnormal conductive pathway can also still persist in the hearts of patients with pacemaker implants [6]. Thus, replacing the damaged tissue with functional and clinically relevant engineered three-dimensional (3D) cardiac tissue has emerged as an alternative solution [7-10].

CMs exhibit a minimal innate capacity for repair, so re-establishing matured CMs with contractibility is a high priority. Therefore, cardiac tissue engineering aims at assembling tissue constructs that can restore basic cardiac function by incorporating cellular components within scaffolds. These acellular or cell-laden scaffolds should be non-immunogenic, mimic cardiac tissue flexibility, allow for physiological electrical propagation, allow sufficient oxygen and nutrient delivery to cells, and have a degradation rate appropriate to the rate of native ECM replacement.

To address these challenges, hydrogels are an attractive source to achieve cardiac regeneration due to their broad spectrum of properties such as controllable swelling behavior, excellent body fluids absorption/retention capacity, high biocompatibility, tunable biodegradation, and the possibility to adapt their chemical structure by using different combinations of substrate materials and crosslinking networks [11]. Furthermore, based on these modifications hydrogels possess the ability to respond to light, pH, electrical, magnetic, thermal and chemical stimuli. These properties allow them to create mimetic scaffolds that can feasibly resemble the native microenvironment [11,12]. Moreover, hydrogels can provide mechanical support for weakened or damaged cardiac tissue to maintain ventricular wall integrity and can also work as a source of growth factors, drugs or cardiac cell lines [13,14]. In this sense, ECM based or naturally derived biomaterials have been widely used to incorporate cardiovascular cells into hydrogels, since the biomaterials closely mimic the natural ECM and show excellent biocompatibility. Furthermore, the biomaterials can be designed to conjugate soluble and insoluble biological factors to improve cellular behaviors [15-18]. However, some key limitations of conventional biomaterials are restricted cell-cell coupling and delayed electrical signal propagation due to the materials’ lack of electrical conductivity [19]. Furthermore, due to the insulating nature of conventional biomaterials, the propagation of electrical impulses between host tissue and implants might be impeded until the implant has formed an intercellular electrical coupling at gap junctions into host tissues [20]. To compensate for these limitations, various types of electrically conductive nanoparticles and materials, which possess unique physical and chemical properties, have been incorporated into conventional biomaterials to create electrically and physically tunable hybrid or composite biomaterials which allow for proper CM maturation and boost electrical signal propagation inside hydrogels [21]. In this review paper, we discuss the effect of the various electrically conductive materials, including carbon nanotubes (CNTs) [22], graphene derivatives [23], gold nanoparticles (AuNPs) [24], electroactive polymers (EAPs) [25], and bio-ionic liquids (Bio-IL) [26] to construct electrically and mechanically tunable engineered microenvironments, which play decisive roles in regulating CM behaviors (Fig. 2). Especially, this review aims to highlight an interplay between the electrical, mechanical, and topological cues from the electrically conductive materials incorporated into various types of acellular or cellular hydrogel scaffolds and their effects on the maturation, proliferation, differentiation, and electrophysiological properties of cardiac cell lines in vitro and in vivo for improving heart regeneration.

Fig. 2.

Schematic diagram of the effects of considerations for biomaterial design for cardiac tissue engineering. It includes 1) Electrical properties, which should mimic the proper conductivity of the scaffold, which is mostly accomplished by the addition of electro-active components, such as CNTs, AuNPs, GO, or electro-active polymers. 2) Mechanical properties which mimic the ECM by material choice, stiffness, and elastic behavior of the biomaterial. 3) Size and shape of the biomaterial, mostly regarding the porosity and the shape of the incorporated particles, as this affects cell adherence and behavior. Created with BioRender.com.

2. Effects of material properties on regulation of CM behavior

The healthy heart requires non-stop conduction of electrical and mechanical signals due to continuous beating throughout one’s lifespan. Subsequent to MI, the heart tissue turns into a fibrotic scar, which has little to no electrical conductivity and has a mechanical stiffness that is not suitable compared to the native heart. Therefore, it is important that engineered scaffolds and grafts allow for adequate maturation, differentiation, and phenotype maintenance of cardiac cell lines to ensure cardiac tissue regeneration [3]. To do this, the electrical integration of these cells and grafts is critical to avoid detrimental side effects such as arrhythmia, and to improve their contractility. In order to enhance the graft’s functionality, the material’s properties, including mechanical and topographical properties, should be well designed as they are major factors that can affect the maturation of the cardiac cell lines. To evaluate these outcomes, previously published research has focused on quantifying the presence of maturation markers in various cell lines, which demonstrate the transition from early developed CMs into the nearest phenotype to primary adult CMs [27]. Within these markers, the most notable are the changes from a disarrayed sarcomere rich in the expression of N2BA titin isoform into highly organized structures with high expressions of N2B and TNNI3, isoforms of titin and troponin respectively [28 29]. Additionally, the presence of t-tubules, a regularly distributed mitochondrial network, and multinucleation represent signs of cell maturation [30 31]. Functionally, the formation of intercalated disks by gap junctions, generation of spontaneous beating and adequate conduction velocities demonstrate adequate maturation progression and thus a high potential for full regeneration [27 32]. To assess differentiation into CMs, the presence of transcription factors such as cardiac homeobox protein (Nkx2.5), GATA-4 and GATA-6 are used as early specific cardiac markers [33-35]. Furthermore, phenotypic differentiation into CMs can be verified by assessing the expression of myofibrillar proteins such as α-myosin heavy chain (α-MHC), β-myosin heavy chain (β-MHC), and cardiac isoforms of troponin-I (cTnI) and troponin-T (cTnT) [36]. Lastly, the assessment of the formation of intercalated disks with connexin-43 (Cx-43) subunits reveal a unique feature of CMs and thus represent an important differentiation marker [37]. In this sense, both differentiation and maturation markers are fundamental to comprehend the adaptation and behavior of cardiac cell lines when exposed to electrically conductive nanomaterials.

2.1. Phenotypical and functional characteristics of various cardiac cell lines employed in the evaluation of electroconductive nano-biomaterials

To evaluate the effects of the different electrically conductive nanomaterials, research groups have employed various cell lines with specific features, which could influence the impact of the functionalized nano-biomaterials. The most common cell lines used are H9c2, HL-1, and pluripotent stem cells (PSCs). In this regard, primary CMs from both healthy and ill sources can be excellent representatives of the heart tissue’s morphological and functional endeavors. However, these cells show donor to donor variability, limited proliferation, and rapid de-differentiation after most isolation methods [38]. To address these issues, immortalized cell lines have been widely used. H9c2 cells are immortalized cells derived from rat ventricular cardiomyoblasts and express several functional cardiac markers, such as transient K⁺ currents and l-type Ca²⁺ associated currents [39,40]. They have also demonstrated an exquisite response to both adrenergic and muscarinic agents mimicking the resident cardiac physiology and have allowed for the formation of multinucleated myotubes [41]. However, this cell line lacks some morphological properties of CMs as the expression of t-tubules or caveolae, nor can they beat spontaneously [40]. HL-1 represents another type of immortalized cell line; however, it is derived from the AT-1 mouse atrial CM tumor lineage. This cell line is highly employed since it retains the morphological, biochemical, and electrophysiological features of adult atrial CMs even after several passages [42]. HL-1 cells exhibit well-developed myofibrillar organization and intercalated discs, similar to those observed in early developing CMs. Additionally, these cells possess the ability to produce spontaneous beating [42,43] and maintain a high expression of specific cardiac markers such as atrial natriuretic peptide (ANP), sarcomeric α-actinin, α-MHC, and Cx-43. Functionally, they show adequate voltage currents similar to those found in primary CMs, such as transient K⁺ and Ca²⁺ currents [43]. A major disadvantage of these animal-derived cell lines is their translational potential to human disease [38]. Given these issues, PSCs have offered the advantage of being a direct source of human tissue potentially capable of making precise conclusions regarding human heart pathophysiology, since they can differentiate into derivatives of all three germinal layers and continuously self-renew [44]. They have been widely used to model several cardiac diseases since they demonstrated effective sarcomere formation, adequate Ca²⁺ and sodium (Na⁺) transients, and spontaneous beating similar to primary CMs [45]. However, these cells possess the disadvantage of maintaining several fetal characteristics in comparison to mature primary CMs. For example, they have shown few or absent t-tubules [46], irregular mitochondrial networks, and lower contractile force and conduction velocities [32,47,48]. The main differences discussed previously between these in vitro models should be considered in analyzing the impact of conductive nanomaterials on cell behavior. In this chapter, we discuss the decisive roles of the nano-biomaterial’s properties on the behavior of cardiac cell lines and, therefore, the suitability of those materials for cardiac tissue engineering, as summarized in Fig. 3.

Fig. 3.

Effects of nanofunctionalization on CM and biomaterial behavior. (A) Electrical conductivity effect on the electrophysiological behavior of CMs. i) Increased Cx-43 expression allows for increased action potential transduction. ii) CNT, GO, and rGO incorporation improves CM functionality. Reprinted (adapted) with permission from [62]. Copyright (2019) American Chemical Society. iii) Bio-IL improves CM functionality [88]. (B) Mechanical properties and topography effect on phenotype of CMs. i) Upregulation of sarcomeric α-actinin and cTnI and cTnT in the sarcomere as markers of cardiac maturation. ii) Improved mechanical properties [96]. Copyright (2015) National Academy of Sciences. iii) Increased vinculin expression through mechanotransduction. iv) Rho/ROCK mechanotransduction pathway for CM beating. v) Effects of topography on CM behavior. SA-CNTs improve alignment of CMs, in addition to improved electrophysiological behavior. Copyright 2017 Wiley. Used with permission from Yi-Gang Li et al. [120].

2.2. Effect of electrical conductivity on electrophysiological behavior of cardiac cell lines

Cardiac tissue engineering focuses on mimicking the myocardium’s cellular microenvironment via recapitulating the ECM as a scaffold to regenerate heart tissue after injuries. These advanced materials have the potential to improve the electrical integrity of cells with the host tissue, thus giving newly formed cells the ability to mimic the proper electrical conductivity of CMs and contributing to the repair of ischemic cardiac tissue. In order to understand the importance of electrically conductive biomaterials in general, we must first elaborate the biological basis of the heart, which possesses an inherent electrical conduction system. The heart contraction is a coordinated mechanism, which is controlled by various electrophysiological phenomena within different heart cells, such as CMs and cells that form an electrical conduction system similar to that of a pacemaker. It is especially important that the cell-cell interactions contain a tight junction to ensure proper signal propagation. The most important components of gap junctions in CMs are the Cx, particularly Cx-43, which is a key component in the electrical transmission system that leads to a synchronized propagation of action potentials by promoting the exchange of ions between the adjacent cells [49] (Fig. 3A-i). Abnormal expressions of Cx-43 can result in arrhythmia and can have lethal effects [50]. In addition to gap junctions, various ion channels control the external and internal levels of different ions, including Ca²⁺, potassium (K⁺), and Na⁺ [51]. The Ca²⁺ ion control is handled by the expression of sarcoplasmic/endoplasmic Ca²⁺ AT-Pase and the Ryanodine receptor, which are responsible for the sarcomeric Ca²⁺ release channel [52]. The sarcomere is a functional unit of the cardiac muscle, and contracts when it receives a increase in action potential [53]. The conduction of action potential signals are propagated through the tight cell-cell interactions and gap junctions.

To regulate the electrophysiological behavior of cardiac cell lines, CNTs are one of the most attractive materials as they have been widely used in various biomedical applications, because of their unique properties such as extremely low density, high aspect ratio, large surface area, favorable thermal properties, strong mechanical resistance, and high electrical conductivity, etc. The electrical conductivity of CNTs is especially dependent on the material’s structure (single, double, and multi-wall) and size, as conductivity can be influenced by atomic arrangement, diameter, length, morphology, or nanostructure. As a result of pioneering research, Cellot et al. introduced an electronic mechanism to explain the effect that CNTs have on improving neuronal performance through the formation of tightly connected neuron cell membrane-nanotube networks, which allows for the creation of electrical shortcuts between compartments of the neuron [54]. This study provided insight into the design of intelligent biomaterials to control, and even enhance the electrical activity of excitable cells. Later, in the same group, Martinelli et al. also found that neonatal rat ventricular cardiomyocytes (NRVCMs) cultured on substrates coated with CNTs were shown to promote electrophysiological properties by forming physical tight interactions between CNTs and CMs [55,56]. Based on these findings, in order to create 2D or 3D functional cardiac tissue constructs with desired shapes and structures in the tissue engineering fields, CNTs have been integrated into various types of biomaterials such as chitosan [57], gelatin [58], methacrylate collagen [59,60], and gelatin methacryloyl (GelMA) [61,62]. etc., resulting in increased electrical conductivity along with other physical properties. An improvement in the electrophysiological activity of NRVCMs was shown when incorporating multi-walled CNTs into GelMA hydrogels, as the beating rate was improved threefold, and the excitation threshold and the voltage required for synchronous contraction were reduced by 85% compared with pure GelMA, in addition to the upregulated expression of Cx-43 and sarcomeric α-actinin [61]. Moreover, Shokraei et al. obtained a similar conductivity rate to the natural myocardium, ranging 1.6 × 10⁻⁵ S/cm to 5 × 10⁻⁵ S/cm longitudinally/transversally using MWCNTs on polyurethane (PU) nanofibers [63] and incorporating CNTs in collagen [64], leading to an improvement in the electrophysiological function of H9c2 cells and NRVCMs, respectively, and improved cellular attachment and alignment.

Graphene is another carbon nanoparticle with a unique shape and size, composed of carbon arranged in a tightly bound hexagonal honeycomb lattice. Graphene and its derivatives have tremendous biomedical potential due to their cytocompatibility and large surface area [65]. Reduced graphene oxide (rGO), made through a thermal, electrical, or chemical reduction of graphene oxide (GO), provides better conductivity compared to GO. Graphene-enabled hydrogels allow for the potential reinforcement of the mechanical structure of materials, and they also have an electrical conductivity and create synchronous beating behavior similar to that of cardiac cell lines functionalized with CNTs. However graphene-based hydrogels exhibit different electrophysiological properties than CNTs, as graphene and its derivatives have larger surface areas, enabling cell adhesion and promoting the transport of cytokines and action potentials, and are more readily dispersible due to the available oxygen molecules [66]. The exact molecular mechanism through which GO and rGO improve electrical signal transduction between cells remains largely unknown, however, the possible attachment of cells to a graphene sheet forms a hybrid network to facilitate signal transduction, similar to CNTs [67]. Jiang et al. [68] incorporated H9C2 cells in a chitosan scaffold with GO and showed GO concentration-dependent cell viability and enhanced expression of Cx-43 and cTnT, and native cardiac conductivity (13.4 × 10⁻⁴ S/cm). Moreover, a rGO embedded collagen scaffold showed low cytotoxicity when introduced to human umbilical vein endothelial cells (HUVECs) and NRVCMs, and induced conductivity (28 × 10⁻⁶ S/cm), leading to an upregulation of Cx-43, Actin4, and cTnT proteins in this last cell line [69]. rGOs were incorporated in poly(ester amide) and chitosan scaffolds, introducing a resistance of 10⁴ Ω/sq, prompting iPSC-derived mesenchymal stem cells to upregulate the expression of GATA-4 and Nkx2.5, suggesting a promoting effect of rGO on cardiac differentiation [70], highlighting the possibilities of rGOs in cardiac tissue engineering. Lee et al. [62] studied the effects of GelMA functionalized with CNTs, GO, and rGO on NRVCMs behavior (Fig. 3A-ii). The incorporation of CNTs and rGO improved electrical conductivity (sheet resistance of ~ 100 kΩ/sq for CNT and ~1 kΩ/sq for rGO), although GO-GelMA and pure GelMA hydrogels were non-conductive. High expressions of sarcomeric α-actinin, Cx-43, and cTnI were shown in the CNT functionalized and rGO functionalized GelMA, suggesting that CM function improved upon functionalization of rGO compared to normal GO, due to the inherent conductivity. Moreover, CNTs showed higher electrical conductivity, native ECM-like nanofibrous architecture with smaller pore sizes, and closer mechanical properties to native heart tissues compared with other nanoparticles, thus enabling CMs to have better electrically conductive networks in the scaffolds and to express higher levels of cardiac markers, resulting in ventricular-like myocyte phenotypes in vitro. Generating ventricular phenotypes is highly desirable because after MI, as the scar tissue often leads to ventricular dysfunction through the production of arrythmias in the infarcted regions. However, CMs on GO-GelMA hydrogels, which are electrically nonconductive and locally stiffer, may become functionaly similar to auricular-like myocytes.

Among other electrically conductive nanoparticles, AuNPs are highly desirable nanoparticles for cardiac tissue engineering due to the ease of producing particles with different sizes, shapes, and surface properties, allowing for tunable electrical and mechanical properties [71]. The main geometric structures described in cardiac tissue engineering are gold nanorods (GNRs) [72,73], gold nanospheres (GNPs) [74], and gold nanowires (GNWs) [75]. In this regard, a pioneer study conducted by Dvir et al. [76] demonstrated that the incorporation of GNWs into alginate scaffolds increased alginate electroconductivity, demonstrated with a current ranging from −12 to 12 μA for GNW-Alginate, compared to 0 μA for pristine alginate. This improved the alignment of NRVCMs, allowing for a strong expression of cardiac markers involved in excitation-contraction coupling such as Cx-43 and cTnI, and improving Ca²⁺ transients in CMs in comparison to alginate counterparts. Indeed, the GNWs were shown to interact directly with the CMs, which increased and synchronized electrical signal propagation. Other studies have shown that the incorporation of AuNPs facilitated Cx-43 expression, which lead to improved conductivity in NRVCMs [77]. More recently, Navaei et al. [78] incorporated GNRs and NRVCMs onto a 2D GelMA hydrogel, wherein a patch formation of interconnected cardiac tissue organization was observed by F-actin fibers and β-integrin. Moreover, an even distribution of sarcomeric α-actinin and Cx-43 provided stimulation to control the contractility of implanted engineered cardiac tissues for the treatment of MI, increasing the synchronous beating frequency up to 100 beats/min, compared to less than 20 beats/min for pristine GelMA. They observed increased cTnI expression upon incorporation of the GNRs compared to pristine GelMA. Moreover, Hosoyama et al. incorporated spherical AuNPs in a collagen hydrogel, forming a conductive hydrogel (8 × 10⁻⁵ S/cm), which increased Cx-43 expression. They compared AuNPs to silver nanoparticles (AgNPs) and noted that AuNPs facilitated better electrical conductivity, likely due to the partial oxidation of the AgNPs, which impeded signal propagation [79]. Lastly, Baei et al. showed that the incorporation of AuNPs in chitosan increased conductivity to 13 × 10⁻⁴ S/cm, facilitating the cardiac differentiation of mesenchymal stem cells (MSCs), as shown by increased Nkx2.5 expression [80].

EAPs are inherently electroconductive materials and are synthesized as a blend of polymers solubilized in organic solvents. This produces a homogeneous scaffold in the desired shape or size, allowing for electrical signaling between the cells, cell adhesion, migration, proliferation, and differentiation [81]. The disadvantages of EAPs include fragile mechanical properties and a difficult synthesis process. He et al. [82] incorporated polypirrole (PPy) conductive nanoparticles into GelMA hydrogels (30 × 10⁻⁴ S/cm) and showed that these hydrogels mantained reduced cytotixicity, adequate cell adhesion, and improved the expression of cardiac markers such as sarcomeric α-actinin and Cx-43. Moreover, high conductivity and enhanced electromechanical properties were observed in NRVCMs, showing stable and synchronous contraction. Kai et al. [83] created PPy-PCL-gelatin electrospun fibers for cardiac tissue engineering, inducing a conductivity ranging 1.3 – 37 × 10⁻⁵ S/cm, leading to increased Cx-43, sarcomeric α-actinin, and cTnT expression in NRVCMs. Spearman et al. [84] used PPy incorporated in PCL and showed increased Cx-43 gene expression in HL-1 cells, good cell viability, and a similar electrical resistance (1.0 ± 0.4 kΩ cm) compared to native cardiac tissue. The nanofibrous scaffolds composed of polyaniline (PANi) and polyetersulfone (PES) produced by electrospinning successfully induced CM differentiation from induced pluripotent stem cells (iPSCs), shown by increased expression of Nkx2.5, GATA-4, and cTnT, as a result of induced conductivity of the aligned fibers (5.7 × 10⁻⁴ S/cm) or random fibers (3.8 × 10⁻⁵ S/cm), compared to pristine PES fibers at 6.58 × 10⁻¹⁶ S/cm [85]. However, the cytotoxicity was proportional to the increase in concentration of nanofibers, probably due to the release of PANi into the cell culture medium [50]. Moreover, NRVCMs seeded on top of the PANi-modified hydrogels showed cTnT expression, as well as sarcomeric α-actinin striations [86].

Lastly, Bio-IL, a group of organic salts with high solubility in water and impressive ionic conductivity and electrochemical stability, have been shown to increase the conductivity of biopolymers when conjugated with substrates such as GelMA or poly(ethylene glycol) diacrylate, with the advantage of preserving good biocompatibility and minimal cytotoxicity (Fig. 3A-iii) [87]. Choline-based Bio-IL conjugated GelMA hydrogels maintain elevated cell viability rates, improve the cell proliferation of 2D cultures and 3D encapsulation of NRVCMs, and enhance cellular attachment and spreading over five days of culture (Fig. 3A-iii-a). Additionally, the phenotype and function of native CMs were preserved, as seen with a maintained expression of sarcomeric-α actinin and the presence of transverse striations [26]. Similar results were seen in adhesive cardio-patches based on choline-based Bio-IL conjugated GelMA hydrogels. The beating behavior of cell-seeded scaffolds exhibited increasing spontaneous beating frequencies in 66% choline Bio-IL conjugated hydrogels compared with GelMA controls and hydrogels treated with 100% choline Bio-IL after seven days of culture, as well as increased expression of Cx-43 and sarcomeric α-actinin (Fig. 3A-iii-bc) [88]. Overall, these results show that Bio-IL conjugated hydrogels represent a promising area of opportunity to produce scaffolds with significant electroconductive properties and feasible adaptability for cells to maintain adequate morphological and functional capacities in vitro.

In conclusion, the electrically conductive properties of the nano-biomaterials incorporated in hydrogels have improved cardiac function, as shown by the sarcomeric α-actinin striations and improved electrical impulse propagation as a result of induced Cx-43 expression. The main reason for the improved signal propagation might be that carbon-based nanoparticles, AuNPs, Bio-IL, and EAPs form tight connections between the cell membranes and intrinsic electrical conductivity of the nanoparticles or polymer networks, forming a hybrid network, facilitating signal propagation and expressing specific cardiac markers. The various strategies to invoke an appropriate electroconductive response in various cardiac cell models during cardiac tissue engineering as summarized in Table 1.

Table 1.

Effects of electro-active components on regulating CM behavior.

Scaffold Materials	Nanoparticle	Cells source	Outcome	Ref
GelMA	Carboxyl acid group functionalized multiwalled	NR CMs	Enhanced electrophysiological and mechanical properties, biocompatibility	[55]
	CNT
	CNTs	Mouse embryoid bodies (EBs)	Enhanced the cardiac differentiation of the EBs and stimulation of stem cells for potential tissue regeneration and cell therapy applications	[121]
	CNTs	NR CMs	Hydrogels enhanced CM adhesion and maturation	[159]
	CNTs	NR CMs	Showed 3 times higher spontaneous synchronous beating rates and 85% lower excitation threshold	[61]
PVA and chitosan	CNTs	Rat MSCs	Gene expression of cardiac markers.	[115]
Poly(ε-caprolactone)(PCL)	Multiwall CNTs	Human MSCs	Promise to cardiomyogenic differentiation	[116]
	Carboxyl acid group functionalized multiwalled	H9C2 cells	Improved cell proliferation	[160]
	CNT
Poly(octamethylene maleate (anhydride) 1,2,4-butanetricarboxylate)	CNTs	NR CMs	Improved excitation threshold and suggested greater tissue maturity	[105]
Polydimethylsiloxane	Multiwall CNTs	NR CMs	Promote viability, proliferation, and functional maturation of cardiac myocytes	[161]
Methacrylated collagen and alginate	Carboxyl functionalized CNT	Human coronary artery endothelial cells and primary CMs	Showed significant cellular proliferation, migration, and differentiation over 10 days of incubation	[162]
Oligo(poly(ethylene glycol) fumarate) (OPF) hydrogels	GO	CMs from Sprague Dawley rats	Provided mechanical support, electric connection, and improved heart function after AMI	[156]
Nickel foam	Nanostructured graphene foams	HUVECs and neonatal CMs	It shows viability of the cells and adhesion and growth of neonatal CMs	[163]
Gelatin/polycaprolactone (Gt/PCL) nanofibers	Nanofibrous graphene	NR ventricular myocytes	Improved electrical conductivity	[164]
Poly(ester amide) (PEA) fibrous chitosan	rGO	10 T1/2 cells	Increased scaffold conductivity and cardiac differentiation	[113]
Collagen patch	GO	HUVECs and CMs	Upregulated expression of the cardiac genes and improve angiogenic properties	[123]
Poly(caprolactone)	Nanofibrous graphene	Stem cell-derived CMs	Cells adhered and contracted spontaneously and exhibited classical CM phenotype	[165]
GelMA	GO	NR CMs	Improved the electrical conductivity and mechanical properties	[166]
Injectable reverse thermal gel	AuNPs	NR CMs	Cell survival for up to 21 days	[74]
ECM from hearts of adult Sprague–Dawley rats	AuNPs	Adult rat CMs	Enhanced cell expression of cardiac-specific markers	[167]
GelMA	GNRs	NR CMs	Cellular retention and expression of cardiac-specific markers	[168]
	GNRs	NR CMs	Improved cell-to-cell coupling and promoted synchronized contraction of the bioprinted constructs	[169]
Chitosan	AuNPs	MSCs	Enhanced properties of myocardial constructs	[108]
Hybrid nanofibers	AuNPs	Bone-marrow derived MSCs	Enhanced cardiomyogenic differentiation and improve functional effects on infarcted myocardium regeneration	[170]
Chitosan and polycaprolactone	PPy	-	Improved electrical conductivity	[117]
Chitosan	PPy	NR CMs	Biocompatibility, improve cardiac-specific protein expression, and cell alignment	[153]
	PPy	NR CMs	Improved conductivity and mechanical properties	[152]
Silk fibroin	PPy	Bone marrow stem cells	Anisotropic morphology, biocompatible, and increased gene expression of cardiac markers	[171]
Polycaprolactone	PPy	HL-1	Differentiation of stem cell-derived CMs. Improved the density of cells	[84]
PPy/PVA	PPy	NIH/3T3	Excellent conductivity and low cytotoxicity	[172]
Polyvinylidene fluoride	PPy	CMs	Improved electrical contractility of CMs	[173]
Collagen, fibroin solution, hyaluronic acid, and polyvinylpyrrolidone	PANi	H9C2	Adequate physio mechanical structure,viability and proliferation of CMs	[118]
PLA or PLA/PEG	PANi	NRK, MCF-7, and MG-63	Increasing concentration of fibrous scaffold improved the cytotoxicity	[139]
Fumarate-co-PEG-co-sebacate	PANi	L929 and H9C2	Viability and proliferation of L929 and H9C2	[148]
Polyaniline/polyetersulfone	PANi	Adult human dermal fibroblast	CMs differentiation	[85]
Poly(glycerol-sebacate)	PANi	C2C12	Modulate cellular behavior via electrical stimulation, and biocompatible	[149]
Gelatin nanofibers	PANi	H9C2 rat cardiac myoblast	Suitable biocompatible scaffolds	[174]

2.3. Effect of mechanical properties on the maturation and differentiation of cardiac cell lines

Suitable mechanical properties and dynamic mechanical stimulations play a pivotal role in developing proper cytoskeletal tension, which affects the maturation and functionality of cardiac cell lines [89]. These developments by cardiac cell-ECM interactions can be detected at the cellular level by analyzing different gene or protein expressions. Mechanotransduction occurs when the ECM signals to the cells, and can occur as a result of the stiffness and elasticity of ECMs along with nanofibrous topography. Mechanotransduction is primarily mediated through integrin binding, receptor tyrosine kinase activation, and GTP-binding proteins (such as guanosine triphosphatase) activation. The signaling pathways involved in cardiac cell proliferation and differentiation include Rho/ROCK (Rho-associated protein kinase) (Fig. 3B-iv) [90], MAPK/ERK (Mitogen Activated Protein Kinase/Extracellular Signal-Regulated Kinase) [91], and stress-activated protein kinases (SAPKs) [92]. The RhoA and ROCK pathways are responsible for cell differentiation into the CM phenotype by virtue of being responsible for the formation of actin stress fibers that result in force generation and subsequently cause the beating of the heart [90]. The ECM plays an important role in signaling mechanisms, as it functions as an anchor for integrin-mediated cellular cascades, which are responsible for complex cell behavior, including differentiation and maturation. One of these markers for maturation is vinculin, which links the cytoskeleton to the sarcomeres, allowing for mechanotransductory cues (Fig. 3B-iii). Therefore, the stiffness and elasticity of scaffolds functionalized by nano-biomaterials should be able to provide sufficient elasticity and mechanical sup-port. Mechanotrasnduction pathways can and have been mimicked in vitro by several researchers who found an upregulation of cardiac maturation and differentiation markers sarcomeric α-actinin, cTnI and cTnT, MEF2c, GATA-4, Nkx2.5 and α-MHC upon continuously promoting the stretching of different cardiac cell lines [93-95]. Furthermore, the biologically and mechanically well-defined scaffolds have sufficient physiological stiffness (~10 kPa) and mechanical stimulation ranges (~15% stretch) and improved CM differentiation from human induced pluripotent stem cells (hiPSC-CMs), which have the potential to regenerate damaged heart tissue and create biomimetic in vitro models for drug screening (Fig. 3B-ii) [96].

However, it has been shown that stiffer biomaterials, such as polyacrylamide at 144 kPa, impede NRVCMs maturation, as seen by poor electrical excitability and reduced cTnI expression [97]. There is evidence that ECMs with lower substrate stiffness (7–27 kPa), mimicking the early stages of cardiac development, showed increased sarcomeric organization of primary rat CMs [98], while matrices of higher substrate stiffness (117–255 kPa), mimicking the stiffness of fibrotic cardiac tissue, showed increased proliferation, but impeded the differentiation potential of mice cardiac side population progenitor cells [98,99]. The elastic modulus of infarcted heart tissue increased (~50 kPa) compared with healthy heart tissue in rats (10–20 kPa) [100]. This observation suggests that using biomaterials with low elastic moduli are more beneficial in cardiac tissue engineering. Moreover, various research suggests that lower culture stiffness (~11 kPa) results in NRVCMs and MSC derived-CMs that are more likely to beat with a higher beating frequency [101,102]. Additionally, this same cell type cultured on less stiff substrates (10 kPa compared to higher modulus of 50 kPa) exhibit more defined sarcomeres and higher Ca²⁺ storage [100]. Neonatal mice primary CMs with higher contractile forces are expected to be found on substrates with lower elastic moduli (12 kPa compared to 39 kPa), since a softer substrate allows for signaling mediated by contractile strain wherein the contraction length is larger for the same force [103].

Naturally derived biomaterials such as fibrin, collagen, fibronectin or matrigel have shown excellent biocompatibility, however their poor mechanical properties in terms of elasticity and mechanical strength limit their use in cardiac tissue engineering [18]. Therefore, incorporating nanomaterials into biomaterials to improve stiffness and elasticity could stimulate mechanotransductory pathways which regulate cardiac cell behaviors, such as physiological and pathophysiological properties [104]. For example, higher concentrations of GO in collagen hydrogels led to drastically increased tensile strength (63 kPa for collagen, 75 kPa for 5 μg/mL GO to 162 kPa for 90 μg/mL) and the Young’s modulus (109 kPa for collagen, 160 kPa for 5 μg/mL GO to 750 kPa for 90 μg/mL), which increased the viability of HUVECs [69]. Moreover, the effects on the mechanical properties of CNT-, GO-, and rGO-GelMA hydrogels in NRVCMs maturation and function were evaluated by Lee et al. [62], showing that hydrogels with low rigidity and stiffness, such as CNT-GelMA hydrogels (Young’s modulus 16.8 kPa, compared to pristine GelMA at 4.5 kPa), can lead to a greater expression of transcripts (Vinculin, YAP / TAZ), inducing the assembly of actin stress fibers favoring the formation of intercalated disks, and increasing the rigidity, the electrical conductivity of the tissue and the maturation of the CMs. In this regard, carbon nanomaterials such as GO and rGO are characterized by excellent mechanical strength (Young’s modulus of 47.9 kPa for GO and 59.8 kPa for rGO), and higher rigidity and roughness. Although they offer significant benefits for electrical conductivity, they may be less suitable than a combination of hybrid hydrogels, such as CNT-GelMA, which offers acceptable mechanical strength with less roughness and higher adaptability. Further research has also demonstrated that CNT-GelMA hydrogels positively influence the differentiation of embryonic bodies from neonatal rats into CMs [105]. Additionally, CNT in combination with natural matrices such as collagen type I hydrogels can be an interesting alternative to improve the mechanical properties of these hydrogels (0.9 kPa for 0.2wt% CNT, 1.85 kPa for 2wt% CNT), allowing them to acquire a Young’s modulus closer to the neonatal heart stiffness (4–11.4 kPa), compared to pristine collagen (80 Pa). Concommitantly, the highest beating area of NRVCMs was found in the hydrogel with 2 wt% CNTs [106]. Similarly, Sun et al. also found 2 wt% of CNTs in collagen provided the optimal NRVCMs culturing conditions. However, they reported Young’s moduli of 13.2 kPa for 0.5wt% and 28.8 kPa for 2 wt%, close to the stiffness of the ventricular myocardium (20 kPa) [64].

Previous studies have also demonstrated that AuNPs can be suitable nanomaterials to enhance the mechanical properties of synthetic hydrogel matrices, and can have positive impacts on cell growth and development. Turner et al. demonstrated that by incorporating GNRs into double hydrogel networks of alginate and acrylamide, the hydrogel composites increased their mechanical strength (2.79 ± 0.5 kPa and 2.6 ± 0.3 kPa with concentrations of GNRs at 1.0 and 2.0 nM, respectively) and elasticity, and could be stretched to more than 3500% of their initial size without fracturing [107]. Similarly, incorporation of GNWs increased the compressive modulus of alginate from 1.6 kPa to 3.5 kPa [76]. Moreover, the addition of AuNPs to chitosan has been shown to improve the gelation and conductivity of these materials, and their effects have also been seen to influence the cell behavior of MSCs, which expressed higher cardiac markers such as α-MHC or Nkx2.5, suggesting that AuNPs may positively influence differentiation into cardiac tissue [108]. Furthermore, AuNPs incorporated into GelMA hydrogels showed increased Young’s modulus (up to 1.3 kPa with GNRs at concentration of 1 mg/mL in comparison to 450 Pa for pure GelMA), and reduced impedance, swelling ratios and porosity. In this same study, this strategy allowed for adequate cell adhesion, phenotype and beating behavior of NRVCMs, showing that by using this approach, improvements in mechanical properties can lead to better cell viability and functional cardiac behavior [109].

Interestingly, incorporation of Bio-IL into hydrogels decreases the compressive and elastic moduli of the hydrogels in a concentration-dependent manner: a 15 w/v% pure GelMA hydrogel has a compressive modulus of 22.10 kPa, yet a 20/80 GelMA/Bio-IL ratio obtains a compressive modulus of 0.60 kPa, and 50/50 exhibits a compressive modulus of 5.53 kPa. Similarly, the eleastic modulus for pure GelMA is 89.30 kPa, 5.40 kPa for 20/80, and 21.30 kPa for 50/50 GelMA/Bio-lL. Logically, incorporation of a larger amount of liquid, i.e. Bio-IL, into the hydrogel leads to a decrease in polymer concentration, demonstrating the mechanical tunability.

The electrical properties of the heart also influences its mechanical resistance, because the beating of the heart is induced through electrical conduction. Aligned cardiac cells are more likely to have an increased expression of Cx-43, allowing for enhanced electrical conductivity. Conversely, the mechanical properties affect the electrical conductivity because the strain of the cardiac cells affects the gene expression of K⁺ channels and currents, affecting the heartbeat’s action potential in turn [110,111]. The static and dynamic stress induce and upregulate Cx-43 expression as well [112].

Overall, the mechanical properties are highly tunable and the stiffness can either be increased through incorporation of various nanoparticles (CNT, graphene, and AuNPs, etc.) or decreased by Bio-IL incorporation, which allows for the formation of the most optimal scaffold for cardiac differentiation and maturation. These desirable mechanical properties along with electrical conductivity concomitantly lead to improved functionality, as exhibited by larger beating areas and increased expression of specific cardiac markers.

2.4. Effect of topography on the phenotype of cardiac cell lines

The shape, size, and aspect ratio of nano-biomaterials can all change the topography of substrates or scaffolds, and influence the phenotype and behavior of cardiac cell lines modeling the role of CMs (Fig. 3B-v). As mentioned above, naturally derived biomaterials are preferred, as there is no risk of toxic degradation. Moreover, the formation of porous structures is highly desirable, as it increases the interconnectivity of the hydrogel and facilitates cell-cell interactions, and can be achieved in different ways [113]. For example, the incorporation of CNT onto PU fibers through electro-spraying reduces the fiber size and improves porosity, as observed by scanning electron microscopy (SEM) [63]. Another study was conducted using chitosan and the incorporation of CNT, resulting in high porosity as well [114]. Lastly, a blend of chitosan/polyvinyl alcohol (PVA) with CNTs [115], as well as poly(ε-caprolactone) (PCL) with CNT additions [116], both demonstrated increased porosity compared to pristine hydrogels. PPy, chitosan, and PCL can each be used to obtain thin films for cardiac regeneration. The addition of the PPy reduces the overall fiber diameter, resulting in the formation of highly porous and rough surfaces [117]. Moreover, the use of positively-charged PPy intercalated negatively-charged fibroin has exhibited anisotropic morphology and highly porous interconnectivity [118]. Tsui et al. [119] used silk fibroin conjugated with PPy to evaluate CMs differentiation from human embryonic stem cells (hESC). The biomaterial exhibited anisotropic morphology, enhanced biocompatibility and increased viability of hESC-derived CMs within 21 days. Upon incorporation of PPy and a patterned surface, there was increased cellular alignment, shown by sarcomeric α-actinin staining, as well as an increased expression of Cx-43.

In addition to the topography and porosity of the material, the size and shape of the incorporated nanoparticles are very important parameters to be considered. The incorporation of CNTs into polymer scaffolds and their effects would thereby depend on structure and size, as CNTs are influenced by atomic arrangement, diameter, length, morphology, and nanostructure. Ren and colleagues [120] observed a sheet of super-aligned carbon-nanotube (SA-CNTs) to recapitulate the aligned structures of the heart, which was shown to improve the electrical transmission of the NRVCMs (Fig. 3B-v). The incorporation of CNTs in GelMA improved the porous interconnectivity of the hydrogel, thereby improving the electrophysiological activity of NRVCMs, however, it did not influence the pore size [61]. In accordance, Ahadian et al. [121] functionalized CNTs with carboxyl groups and showed improved cellular alignment without changes in the hydrogel’s porosity. Lastly, Sun et al. engineered a patch, made from a GelMA bottom layer, CNT middle layer, and a microneedle-shaped top. The top structure induced the directional alignment of iPSC-derived CMs, whereas the conductive element of the CNTs improved cell-cell interactions [122].

The interaction between carboxyl groups of GO and amine groups (NH₂) of a collagen hydrogel produced a 3D structure with homogeneously distributed interconnected pores [123]. The incorporation of rGO in hydrogels reduces its toxic aggregation and facilitates homogenous distribution. Moreover, it was demonstrated that rGO sheets were smaller than GO, resulting in a decreased in the pore size in GelMA [124]. rGO also was incorporated onto CS scaffolds (rGO/PEA)-CS, but homogenous fiber morphology was not observed. However, the presence of rGO increased the size and uniformity of the pores, creating a more suitable environment for iPSCs-derived CMs maintenance [113]. Additionally, the nanoroughness of graphene has been shown to be key in cardiomyogenic differentiation. The researchers showed increased expression of Nkx2.5, MEF2c, Cx-43, cTnT, and MLC2a as markers of cardiac differentiation in human embryonal stem cells, as the nanoroughness of graphene facilitates increased cell adhesion and activates ERK signaling [125]. Corroborating these findings, Park et al. showed increased cardiac differentiation markers (GATA4-, Nkx2.5, Cx-43, MLC2a, sarcomeric α-actinin) of graphene compared to a coverslip control, and a decreased expression of early cardiac differentiation marker GATA-4 between weeks 2 and 3 of culture, indicating cardiac maturation of the MSCs due to the surface topography of graphene [126].

Despite the inertness of gold, some studies suggest that size-dependent cytotoxicity of AuNPs exists in human organs such as reproductive systems, lungs, kidneys, and the brain [61,72,127-129]. GNRs coated with cetyltrimethylammonium bromide (CTAB) were homogeneously incorporated in a GelMA hydrogel. A microscopic examination showed a rough topography, favoring the interaction between NRVCMs and GNRs, and improving the electrical conductivity. Compared with AuNPs with spherical morphologies, the GNRs showed reduced toxicity [130]. GNWs stabilized with CTAB incorporated in an alginate scaffold demonstrated a homogenous distribution of a star-shaped forms inside of porous material, but no interconnections among pores were observed [76]. Pores provide the diffusion of essential nutrients and oxygen to the cell, and, more importantly, allow for cell-cell interactions, Cx-43 upregulation, and the propagation of an action potential. Therefore, it can be concluded that highly porous hydrogels with interconnected pores are desirable for cardiac tissue engineering.

In conclusion, the incorporation of nano-biomaterials improves the nano-topography of conventional hydrogels, which, in turn, facilitates cell-matrix interactions, resulting in increased mechanotransduction signaling. Lastly, nanopatterning of hydrogel patches induces cellular alignment, which favors the formation of α-actinin striations. As a result, all major characteristics of nano-biomaterials play an important individual or combinatorial role in facilitating cardiac function, such as improving signal propagation, as well as upregulating cardiac specific markers such as Cx-43, sarcomeric α-actinin and cTnT. The effects of nanofunctionalization, with the electrical, mechanical, and topographical properties on cardiac cell behavior in this section are summarized schematically in Fig. 4.

Fig. 4.

Schematic representation of incorporation of nanomaterial in hydrogel and its effect on matrix properties, conductivity, and cardiac functionality. Created with BioRender.com.

3. Biocompatibility of nano-biomaterials

Polymers and nanomaterials which incorporate metallic-based nanoparticles are considered promising candidates to improve the electrical and mechanical properties of heart tissue, however, biocompatibility and biodegradation remain major challenges [131]. Moreover, the size and shape of the incorporated nanomaterials affects the cytotoxicity and reactive oxygen species (ROS) generation. Therefore, these should be carefully assessed before in vivo applications. Biocompatibility entails that the biomaterial should not provoke a significant foreign body response, yet should initiate proper electrophysiological and mechanical behavior without cytotoxicity. Moreover, the material should guide the proper immune response, ideally initiating a regenerative behavior of macrophages and other immune cells.

As discussed in the above section, CNTs are immensely helpful for the improvement of the electrical properties of biomaterials, however, they are not biodegradable so they might accumulate in specific organs and tissues, and are likely to induce potential cytotoxicity. For example, inflammatory responses have been reported for CNTs incorporated in biomaterials in the lung, where the CNTs could not be cleared by the immune system [51]. CNTs incorporated in polydimethylsiloxane (PDMS) scaffolds have been shown to be biocompatible, however PDMS is limited to in vitro use and cannot be used in vivo [56]. Additionally, chitosan can be used to incorporate the CNTs, as it is a biocompatible and biodegradable stabilizing agent. Research has shown good biocompatibility of gelatin-chitosan scaffolds with incorporated CNTs [132]. Other studies have demonstrated the cytotoxicity of CNTs to be dependent on their size, shape, and concentration, while nanofibers were shown to be more cytotoxic than nanotubes [51]. Moreover, the incorporation of CNTs could lead to enhanced contractility of the heart tissue, possibly leading to arrhythmia [133].

Graphene and its derivatives have high biomedical potential, however, the irregular corners and protrusions along the edges of graphene sheets can puncture cell membranes and initiate internalization in the cell. These characteristics threaten the biocompatibility of graphene, as it would disturb normal cell functions and could possibly lead to cell death [134]. Recently, graphene biocompatibility was evaluated using a culture of human fibroblasts, with graphene nanoparticles showing good cell viability. However, cytotoxic effects, including the generation of ROS, were found to be concentration-dependent. [135].

Gold has been shown to be a very biocompatible material due to its non-reactivity. Shukla et al. showed the biocompatibility of AuNPs with macrophages [136]. Spherical AuNPs were shown to be highly biocompatible, as no cytotoxicity was observed. AuNPs coated with polyethylene glycol (PEG) exhibited high cell viability over the course of 7 days, however, more studies are required to assess their long-term cytotoxicity [45]. Despite the inertness of gold, some studies suggest the cytotoxicity and biocompatibility of AuNPs depend on their size [74] and surface properties, as the AuNPs might interact with certain proteins [137]. Therefore, the biocompatibility of gold-based materials, including AuNPs, should be thoroughly assessed, especially considering the shape of the NPs.

EAPs are biocompatible and biodegradable materials [138,139]. The most commonly used EAPs for cardiac tissue engineering are PPy and PANi. PPy is biocompatible and has been adjusted to a conductivity level that resembles the physiological environment in cardiac regeneration [119]. Talebi et al. [117] used PPy, chitosan, and PCL to obtain films, which favored the formation of porous and rough surfaces, and the incorporation of PPy showed improved electrical conductivity and low cell cytotoxicity. PANi exhibited low solubility [140], questionable cytotoxicity, [139] and reports of chronic inflammation in tissue [141], however, its electrical conducitvity properties are highly desirable for cardiac tissue engineering. Therefore, chemical modifications are required to render the material biocompatible and suitable for in vivo use.

Bio-IL offer a highly biocompatible and biodegradable alternative for electroactive scaffolds for cardiac tissue engineering, since they are comprised of naturally derived compounds, which can degraded physiologically and environmentally into smaller chain molecules [26].

To improve the biocompatibility and minimize the cytotoxicity of various nanobiomaterials, chemical modifications of materials can alter the degradation speed and biocompatibility of the material [142]. Functionalization of CNTs can include carboxyl and alcohol groups, positively affecting cell growth [143]. Moreover, the addition of a PEGylated coat or functionalization with diethylene triamine pentaacetic acid have been shown to improve the biocompatibility of CNTs [144]. CNT functionalization improves solubility and avoids cell internalization, as serum proteins would otherwise adsorb on the CNT surface, resulting in toxic agglomeration, which could lead to cell death, phagocytosis, and inflammation [145,146]. Modifications of rGO to improve biocompatibility include PEGylation of the graphene, amine-functionalized graphene quantum dots, and graphene-PCL composites [147], in addition to oxidizing and reducing modifications [51]. Chemical modifications of PANi require mixing and complex formations with other materials. Alves et al. [118] stabilized PANi with polyvinylpyrrolidone, increasing its solubility and formation. The hydrated EAP showed high electrical conductivity, interconnected pores, and high cell viability of the H9C2 line. Komeri and Muthu [148] developed an injectable PANi hydrogel conjugated to fumarate-co-PEG-co-sebacate, showing the cytocompatibility of L929 and H9C2 cells, and an electrical conductivity similar to that of the native myocardium (0.550 ± 0.016 × 10⁻³ S/cm). Similar biocompatibility was found by Qazi et al. [149] using blended PANi with poly(glycerol-sebacate) and C2C12 cells. Lastly, coating with PLA polymer using coaxial electrospinning showed a decrease in the PANi release to the cells and an increase in the biocompatibility of the nanofibers [139].

In conclusion, biocompatibility and biodegradability are critical aspects for in vivo applications. Most nano-biomaterials showed good biocompatibility, but many nanoparticles lack biodegradation properties which might induce potential cytotoxicity in the body. Therefore, nano-bio materials should be chemically altered to possess biodegradation abilities or excreted from the body, which allows them to be used for in vivo applications.

4. Effect of electrically conductive materials on in vivo cardiac tissue regeneration

Several strategies for developing electroconductive materials for cardiac tissue engineering have also demonstrated appropriate in vivo biocompatibility and functionality, as summarized in Fig. 5. Walker et al. [88] developed naturally-derived Bio-IL conjugated cardiopatches and proved their suitability in an in vivo model of MI by luring the anterior descending coronary artery. Their results showed that after three weeks of the given injury, both GelMA and GelMA Bio-IL conjugated hydrogels showed strong adhesion to the injured cardiac tissue. Animals receiving GelMA and GelMA Bio-IL patches showed significant reductions in ventricular thinning compared to sham counterparts, which also presented aneurismatic malformations at the apex of the left ventricle’s free wall. Animals receiving the Bio-IL conjugated cardiopatches showed higher expressions of both Cx-43 and sarcomeric α-actinin, which are well-known markers of myocardial viability and phenotypic preservation of native tissue (Fig. 5A). In terms of degradability, both hydrogels showed significant degradation after three weeks of implantation. Like these results, choline-based Bio-IL hydrogels have exhibited efficient enzymatic degradation after 21 days of implantation in healthy rats [26]. Also, these hydrogels have shown minimal inflammatory infiltration and immune response after 28 days of implantation, seen in only a slight increase of CD3 and CD68 markers of lymphocyte and macrophage infiltration at the site of the implant [26]. However, the delivery of electrical currents by the potential of conductive scaffolds remains a concern since these hydrogels can be pro-arrhythmic [150]. Although this is a possible disadvantage, this parameter, in conjunction with cardiovascular function panels, has been scarcely evaluated.

Fig. 5.

Effects on cardiac function and biocompatibility of different electroconductive hydrogels on in vivo models of acute MI. (A) Effects of GelMA and GelMA/Bio-IL hydrogels on the development of fibrotic tissue in histological samples stained with Trichrome-Masson (i); expression of the cytoskeleton protein sarcomeric α-actinin (ii), expression of the gap-junction protein Cx-43. (B-i) Effects of HPAE-Py/Geln hydrogels in reduction of fibrotic scar after MI in histological samples stained with Trichrome-Masson. (B-ii) Electrocardiographic signals produced by implantation of HPAE-Py/Geln hydrogels in injured myocardial tissue. (B-iii) Biodegradation of HPAE-Py/Geln hydrogels in injured myocardial tissue after 28 days of implantation. (C-i) Effects of OPF/GO hydrogels on expression of macrophage marker CD68 after implantation by immunostaining. (C-ii) Effects of OPF/GO hydrogels on expression of Cx-43 after one week of implantation. (C-iii) Echocardiographic functional tests of animals implanted with OPF/GO hydrogels. Images reproduced with permission [88,151,156]. Copyright Wiley (2018), Elsevier (2019) and Ivyspring International Publisher (2018).

Other types of strategies have also shown the positive impact of electroconductive hydrogels on in vivo acute MI (AMI) models focusing on cardiovascular function. In this regard, Liang et al. [151], developed PPy conjugated with hyperbranched poly-amino ester (HPAE-Py) adhesive hydrogels and implanted them on rats with induced AMI (Fig. 5B). Results showed that two weeks after implantation, hydrogels exhibited good penetration and adaptability to host injured tissue (Fig. 5B-i-ii). Echocardiography showed increased ejection fraction (EF) and fractional shortening (FS) in animals treated with HPAE-Py hydrogels compared to controls. Additionally, these animals exhibited reduced end-diastolic and end-systolic volumes in comparison with controls. These results demonstrate that animals treated with these scaffolds had preserved cardiac contractility and hemodynamic responses to injury. Additionally, electrophysiological studies showed that animals treated with HPAE-Py hydrogels had fewer deeper Q waves on an electrocardiogram and maintained an average QRS interval duration (Fig. 5B-iii). These results suggest that treatment with this strategy can reduce the detrimental effects of necrosis associated with AMI on the electrical depolarization behavior of the injury site. Interestingly, HPAE-Py hydrogels also had positive effects on reducing the infarct size, fibrotic area, and increased angiogenesis (Fig. 5B) [151]. Similar results have been shown by Shell-PPy cardiac patches, which ameliorated the AMI area, reduced fibrotic scar, increased angiogenesis, and increased EF and FS compared with controls [152].

In vivo models have exhibited that PPy material has good biocompatibility [153]. In this regard, positive effects on the velocity conduction of electrical impulse are critical for maintaining an adequate excitation-contraction coupling on in vivo models. In this sense, injectable hydrogels composed of PPy and chitosan exhibited biocompatibility and presented longitudinal conduction velocities of the cells at 74.3 ± 4.7 cm/s, comparable to uninjured tissue at 81.3 ± 0.59 cm/s after being injected in AMI rat models [153]. However, Costa et al. [32] reported the adverse effects of PPy in zebrafish embryos/larvae. There was no observation of cell death; however, morphological alterations in body length, and ocular and locomotor behavioral changes suggested the toxicity to be dose-dependent. To overcome biocompatibility issues, shell-derived chitosan hydrogels were crosslinked with PPy, which formed an adhesive sheet. The crosslinked sheet showed a desirable biocompatibility, promoted cell-cell coupling, and contraction of NRVCMs in vitro. In vivo heart function was significantly increased, showing no cytotoxic effects. Alongside these results, PPy hydrogels have been shown to produce only slight immune responses after cardiac implantation in animal models, suggesting that these composites’ adverse effects on cell biocompatibility are minimal [82].

McCauley et al. evaluated CNTs to re-establish myocardial conduction. In these conditions, CNT fibers maintained the electrical conduction for 30 days without an inflammatory response [154]. Similarly, in vivo assays using a rat model of MI improved the Ca²⁺ signal conduction of native CMs after injecting the hydrogel with GO in the injured tissue. Therefore, the use of graphene poses an exciting potential for manipulating cardiac phenotype purposes. These advanced materials can provide the electrical integrity of stem cells with the host tissue [155]. Similarly, Oligo-polyethylene-glycol-fumarate/GO (OPF/GO) injectable hydrogels have shown positive AMI improvements, resulting in less extensive scar tissue area formation and preservation of left ventricular thickness (Fig. 5C). Additionally, in this same study, treatment with these types of hydrogels favored the preservation of EF and FS, and significantly reduced left ventricular end-systolic volume compared to controls and presented acceptable biocompatibility (Fig. 5C-i to 5C-iii). Overall, these results show that these hydrogels promoted hemodynamic recovery after AMI [156].

Although several electrically conductive nano-biomaterials for tissue regeneration have significantly and positively affected histocompatibility and morphological and functional adaptations in small in vivo models, electrically conductive nano-biomaterials in large animal models have not been tried yet. Thus, much research, such as long-term immune response and electrophysiological studies, still needs to be done to significantly decrease the risk of failure for future large animal studies. Therefore, the safety of large animal models represents an area of opportunity for upcoming research.

5. Future perspectives and conclusion

To increase the regenerative potential after MI, emerging strategies using functionalized biomaterials offer new perspectives for the patients suffering from CVDs, however, researchers still face many challenges in developing an optimal therapeutic strategy. The incorporation of electro-active or electrically conductive nano-biomaterials into biocompatible and biodegradable hydrogels not only improve the electrical or electrochemical properties, they also tone the mechanical properties of hydrogels and bring topographycal cues into hydrogels that resemble native ECMs. This allows for improvements in cell differentiation, maturation and the contractile functionality to resemble that of native CMs, which are essential for the development of an ideal scaffold. Additionally, the electrically conductive networks in the hydrogels, as a result of the incorporation of nanomaterials, leads to improvements in the electrophysiological behavior of cardiac cell lines and to transfers in action potential amongs these cells, or between host tissues and implants.

Although restoring heart function and cardiac muscle regeneration without a severe immune response were not reported after implants, various electrically conductive scaffolds were reported in many articles. However, the incorporation of nanomaterials still possesses several limitations in clinical application. For example, long-term biocompatibility remains questionable due to cellular internalization or travel/transfer to other organs such as the liver or kidney, in the case of no degradation of the electrically conductive nano-bio materials. This limitation could be overcome through chemical modifications and coatings on the nano-bio materials to eliminate the immune response after implantation for a long time, or through expelling them from the body. In terms of polymeric-based micro or nano-sized materials, they can be conjugated with biodegradable segments such as monomers, oligomers, proteins, or enzymes, which can help to degrade or break the nano-bio materials at the nano-scale, as long as the electrical conductivity of the material remains unaffected. Also, during the synthesis or polymerization process of nano-bio materials, biodegradable bonds such as ester, carbonate, anhydride, amide, orthoester, and urethan could be added. Later, these bonds could help to degrade or break materials to small sizes [157].

Recreating the hierarchical architectures and blood pumping function of the native heart, which has four rooms surrounded by aligned myocardium and vessel networks, is still a significant challenge. The electrically conductive nano-bio materials could be an excellent candidate for bioinks, allowing researchers to build a contractile vascularized myocardium or full-size human heart constructs fabricated by recent advanced 3D bioprinting techniques. Moreover, this research could be translated towards heart-on-a-chip models to test the suitability of the scaffold, newly developed drugs, or the cardiotoxicity of anticancer drugs [158]. After improving the therapeutic efficacy, these electrically conductive nano-bio material strategies could be coupled with other approaches, such as gene therapy, miRNA therapeutics, and growth factor delivery into the heart tissue. Using the unique physical properties of nano-bio materials, which can respond to external stimuli such as electrical, magnetic, light, temperature, pH, etc., an intelligent delivery function could be fashioned on the scaffolds that allows for the release of specific growth factors, miRNAs, or drugs that could achieve a synergistic therapeutic effect in clinical trials. In conclusion, the formation of electrically conductive and mechanically tunable scaffolds for cardiac tissue engineering requires a multidisciplinary approach in order to substantially improve the quality of life for patients.