Evaluating and improving biocompatibility of conductive polymers for cardiac tissue engineering

Abstract

Conductive polymers (CPs) have gained increasing attention in cardiac tissue engineering (CTE) due to their ability to restore electrical conductivity, enhance cardiomyocyte (CM) function, and support tissue regeneration. Despite significant progress in the field, challenges related to variability in biocompatibility testing, including dopant-dependent cytotoxicity, poor reporting of biodegradability, unpredictable long-term stability, and regulatory uncertainty of CPs continue to hinder their applications. To address this, we reviewed the properties, applications, and biocompatibility of the three most studied CPs: polypyrrole (PPy), poly(3,4-ethylenedioxythiophene) (PEDOT), and polyaniline (PANI) in CTE, and contrast their advantages and safety challenges with inorganic electrodes and carbon-based materials. We critically assessed current methods for evaluating the biocompatibility of CPs, highlighting limitations in traditional in vitro and in vivo approaches. Our analysis revealed a significant gap in chronic implantation data beyond six months and provided dopant-centered assessment and toxicity risks across different CP platforms. A comprehensive roadmap was further suggested to guide the evaluation of the biocompatibility of CPs, including material characterization, in vitro cytotoxicity testing with particular emphasis on in vitro 3D human heart model testing platforms of human pluripotent stem cell (hPSC)-derived engineered heart tissues and cardiac organoids, and in vivo evaluation. Additionally, we discussed recent advances in improving the biocompatibility of CPs through hybrid scaffold development, molecular engineering, surface chemistry modifications, and the development of stimuli-responsive and targeted CP constructs. By establishing this standardized framework and highlighting critical regulatory requirements, this review aims to overcome current biocompatibility barriers and facilitate the improved implementation of CPs in CTE applications.

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Huaxiao Yang

Huaxiao Adam Yang, PhD, is an Assistant Professor at the University of North Texas (UNT) in Biomedical Engineering. He earned bachelor's degrees in chemical engineering from Huazhong University of Science and Technology and Biological Science from Wuhan University, MS in Macromolecular Physics and Chemistry from Fudan University, and a PhD in Bioengineering from Clemson University. He further completed postdoctoral training at Stanford University Cardiovascular Institute. His research focuses on applying human pluripotent stem cell-derived cardiovascular cells, bioengineering tools, and biomaterials in cardiovascular tissue engineering. He has published over 50 research articles in Science, Nature, Circulation, European Heart Journal, Cardiovascular Research, ACS Nano, etc. Recently, his team won the 2024 NIH Complement-ARIE Challenge Prize Award.

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1. Introduction

Heart disease is the most common cause of death globally.¹ It often affects the ability of heart muscles to conduct electrical signals, which is a critical function for maintaining coordinated and efficient blood circulation.^2,3 In conditions such as arrhythmias and cardiomyopathies, electrical communication between cardiomyocytes (CMs) becomes erratic or blocked, and this disrupts coordinated contraction, which ultimately leads to heart failure.^4,5 Current treatments like pacemakers, defibrillators, or left ventricular assist devices (LVADs) can manage electrical dysfunction, but they do not fully restore the heart's natural electrical pathways.⁶ To address these challenges, researchers considered the application of conductive polymers (CPs) in cardiac tissue engineering (CTE) to support and bridge the electrical conductivity of cardiac tissue. Additionally, the CPs have demonstrated tunable mechanical properties and the potential to support CM attachment, proliferation, and cell-to-cell electrical communication in engineered cardiac constructs.^7,8 As summarized in Fig. 1, intrinsic CPs including polypyrrole (PPy), poly(3,4-ethylenedioxythiophene) (PEDOT), and polyaniline (PANI) have been studied and applied in various tissue engineering fields including cardiac, bone, neural, skin, and skeletal muscle tissue, where they serve as conductive scaffolds that can deliver electrical stimulation to enhance tissue-specific cellular functions and development.^9–15 CPs can integrate directly into the native heart tissue and conduct electrical signals to enhance electrical communication between cells including CMs.¹⁶ Unlike external electrical stimulation, this approach can improve electrical conduction at the tissue level and help synchronize cardiac contraction, providing a more direct way to address certain conduction abnormalities.¹⁷ For example, CP-based scaffolds are being incorporated into engineered heart tissues (EHTs) and used as substrates for stem cell-derived CMs to promote electrical coupling, contraction synchronization, and cardiac maturation, with the goal of improving overall cardiac function.^18–21

Fig. 1 Diverse applications of CPs in tissue engineering. The diagram highlights their roles in CTE, including electrical stimulation and synchronized beating, as well as their potential for neural regeneration, bone tissue engineering, and wound healing.

The specific interest in CPs dates back to the late 1970s, when the first electrically conducting forms of polyacetylene were synthesized and reported.²² By the mid-1990s, Ingber and Langer had begun exploring their potential applications in cell cultivation, paving the way for their later use in CTE.^23,24 At the molecular level, CPs have a unique structure, which contains conjugated π-bonds.^16,25 These bonds allow them to have a wide range of electrical conductivity, which can be tuned to match the electrical conductivity of heart tissues. However, these polymers need to be doped to increase the charge carrier density and improve conductivity.^16,26 For instance, CPs have been doped with agents like polystyrene sulfonate (PSS) to enhance CP–cell interactions and stability in biological environments, making them more suitable for tissue engineering applications.^8,27 A notable example is PEDOT doped with PSS to achieve conductivities of up to 300 S cm⁻¹, and the doped PEDOT was found to enhance cell-to-cell communication between EHTs and native myocardium, making it one of the most widely studied CPs in CTE.^27–29

In contrast to inorganic electrode materials (e.g., gold, silver, metal oxides) and carbon nanomaterials like graphene and carbon nanotubes (CNTs), CPs provide unique advantages for cardiac applications. CPs can be readily processed into diverse formats, including hydrogels, electrospun fibers, printable inks, and thin, uniform surface coatings, often without the dispersion challenges associated with carbon nanomaterials.^30,31 Although carbon-based materials exhibit excellent conductivity, there are concerns about their long-term clearance and in vivo safety, and the incorporation of CNTs into composites can increase stiffness, potentially affecting biomechanical compatibility.^32–34

On the other hand, CPs can be chemically functionalized through the incorporation of peptides, extracellular matrix components, or bioactive dopants to promote CM adhesion, minimize leachable toxicants and deliver cellular signaling cues.³⁰

Despite these benefits that CPs offer, there are still a number of factors to consider, especially regarding their processability, stability, and potential cytotoxicity.⁸ To achieve electrical conductivity, all CPs must be doped, and the type of dopant used significantly influences both their conductivity and biocompatibility as summarized in Table 1. These dopants can range from small proton donors to large macromolecules. Commonly used dopants, such as PSS and camphor sulfonic acid (CSA), have been shown to negatively affect cell viability, depending on their concentration and release profile.^35–38 These challenges are more pronounced for the stimuli-responsive CPs, which rely on mobile ions or dopants that can leach over time, reducing conductivity and posing local toxicity risks for their dynamic behavior.^39,40 For instance, in PEDOT:PSS systems, PSS dopants undergo degradation in biological environments and release acidic degradation products that can induce tissue damage, cause inflammatory responses and local ionic imbalance.⁴⁰ This phenomenon extends to other doped systems where PANI materials must be chemically de-doped to move electrical conductivity from 10⁻³ S cm⁻¹ to 10⁻⁸ S cm⁻¹ through strong alkaline de-doping processes, showing inherent instability of dopant–polymer interactions under physiological conditions.³⁴ Additionally, in mechanically active tissues such as myocardium, chronic cyclic loading, enzymatic exposure, and repeated wet-dry cycling characteristic introduce additional failure pathways that accelerate loss of conductivity and compromise material integrity. However, the potential cytotoxic effects of these dopants are often inadequately addressed or reported in studies involving CPs.

Table 1 Common dopant-specific biocompatibility effects in conductive polymer systems

Dopant	Effects on biocompatibility	Ref.
Polystyrene sulfonate (PSS)	Cytotoxicity exhibited in indirect assays (MTT) due to residual toxic molecules; reduced cell viability, growth or adhesion at higher concentrations (e.g. 0.3% or above) potential leaching of toxic molecules into culture media; acidification; anionic (negative) charge	41 and 42
Dodecylbenzenesulfonic acid (DBSA)	Exhibits high cell viability (∼70%); demonstrates low cytotoxicity; supports cell growth and viability; maintains stability and resists delamination in aqueous environments	43
Perchlorate (ClO4^−)	Favorable biocompatibility-high cell viability (∼89%); supported cell proliferation; limited apoptosis or necrosis (∼7%)	44
Hexafluorophosphate (PF6^−)	Good cell growth and attachment (69%); enhanced electrochemical capacitance	44
Dodecylbenzene sulfonate (DBS)	Immobile anion; minimal leaching; enhances hydrolytic stability; minimal cell death	45
Heparin	Large immobile anion; bioactive polyelectrolyte; contains anticoagulant properties; minimal leaching; enhances hydrolytic stability; minimal cell death	46 and 47

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The biocompatibility of CPs themselves, such as PPy, PEDOT, and PANI, is also affected by factors like their chemical forms, concentrations, and polymer purity.⁴⁸ When compared to biodegradable and biocompatible nano-biomaterials such as polymeric nanoparticles, protein-based hydrogel and biodegradable elastomers, these materials offer predictable resorption and well-established safety profiles but generally lack inherent electrical conductivity like CPs.^49,50 For example, the non-biodegradable nature of PANI raises concerns about its potential long-term cytotoxicity.¹³ Additionally, it requires strong acids like hydrochloric acid (HCL) for activation and shows reduced conductivity under physiological conditions.⁵¹ All of these factors highlight how important it is to carefully design these materials and test their biocompatibility for CTE applications.⁵²

Currently, the most common methods for testing biocompatibility involve animal models and in vitro assays. The animal models assess how the material interacts with tissues, while in vitro cytotoxicity assays provide insights into cellular responses to CPs.^53,54 However, these methods have their limitations. Animal models often fail to accurately recapitulate human physiology, and cytotoxicity assays using human primary cells or cell lines in conventional culture (e.g., 2D culture in Petri dishes) do not capture the complex function and structure of tissues or organs for the long-term physiological impacts of CPs.^53,54 Additionally, ethical concerns associated with animal testing have led to increased scrutiny of these methods, and the invasive nature of histological analysis provides only endpoint data.,^53,54 Although several reviews have discussed CPs in cardiac and other tissue engineering fields,⁵⁵ none have provided a standardized roadmap for evaluating their biocompatibility, hence creating inconsistencies in biocompatibility evaluations among different studies. Existing reviews largely highlight material properties or broad biomedical applications, but they lack systematic frameworks for assessing degradation product toxicity or propose standardized testing protocols across different CP systems. This variability is largely because of the differences in how these polymers are synthesized, which affect their chemical composition and material properties. Other contributing factors include dopants and monomer selection. These issues become even more challenging because conventional in vitro and in vivo systems often fail to accurately replicate the physiological environment. As a result, researchers are turning to more physiologically relevant models to improve biocompatibility assessments in line with the FDA Modernization Act 2.0, which emphasizes the use of human-relevant testing methods that do not require animal models for testing biomedical safety and efficacy.⁵⁶

This review provides an overview of the properties and applications of CPs in CTE, and the advantages and disadvantages of the current techniques for determining their biocompatibility. The importance of incorporating in vitro three-dimensional (3D) human heart models for biocompatibility evaluation with attention to dopant-dependent cytotoxicity in line with recent regulatory trends is specifically emphasized. Lastly, this review seeks to provide a roadmap to guide the evaluation of CP biocompatibility for cardiac regenerative medicine.

2. Types of conductive polymers

To date, three major CPs, namely PPy, PEDOT, and PANI, have been greatly explored for their potential in general tissue engineering fields. Table 2 summarizes and compares their conduction properties, typical dopants, advantages, and limitations in tissue engineering.

Table 2 Material properties and characteristics of conductive polymers for tissue Engineering applications. Data compiled from various studies, including pure polymers, doped systems, and composite scaffolds used in tissue engineering applications

Polymer type	Typical dopants	Conductivity range	Physical properties	Advantages	Material limitations	Applications	Ref.
	PSS, DBSA, camphor sulfonic acid (CSA), hydrochloric acid (HCl), P-toluene sulfonic acid (PTSA),nitrate (NO3^−)	(10^−3–10^2 S cm^−1)	– Brittle when doped	– High conductivity.	Aggregation tendency, dopant leaching, non-biodegradable, brittle when highly doped	Cardiac patches, neural interfaces, bone tissue engineering	9, 48 and 57–62
			– Moderate flexibility	– Support cell growth.
	PSS, tosylate (TOS), ClO4^−	(10^−4 S cm^−1 to 10^3 S cm^−1)	Good flexibility, stable in aqueous environments	Exceptional stability, highest conductivity, good mechanical properties	Limited biodegradability, complex processing, PSS toxicity concerns	Nerve, cardiac, skin tissue engineering, neural electrodes	10, 11 and 63–66
	CSA, DBSA, HCL	(10^−2 S cm^−1 to 10^1 S cm^−1)	Brittle but flexible when doped, pH sensitivity	Tunable conductivity, low cost	Non-biodegradable, poor processability, pH instability, limited solubility, limited biodegradability	Cardiac tissue engineering, neural interfaces and muscle tissue engineering, biosensors	12–14, 20 and 67

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2.1. Polypyrrole (PPy)

In comparison to other CPs, PPy stands out as the most widely used in the field of biomedicine.^9,68,69 Specifically, in CTE, over 40 research articles have been reported between 2014 and 2024 (Fig. 2). It is because PPy offers a balance of strong electrical conductivity and stability and easy material processing. PPy also readily forms stable composites with both natural and synthetic materials and exhibits exceptional compatibility with living organisms.^57,70 The synthesis of PPy can be achieved through oxidative and microemulsion polymerizations.⁵⁷ Composite scaffolds and fiber approaches of PPy have demonstrated particular success in cardiac applications. For example, Liang et al. (2021) blended PPy in silk fibroin (SF) biomaterial to create a conductive composite nanofiber that exhibited electrical conductivity with a range of 0.05 to 3.6 mS cm⁻¹, demonstrating tunable electrical properties suitable for CTE applications.¹⁸ They found that the scaffolds containing PPy showed high cell viability and improved metabolic activity of CMs based on the live/dead and 3-[4,5-dimethylthiazol-2-yl]-2,5 diphenyl tetrazolium bromide (MTT) assays.¹⁸ The PPy/SF composite fibers supported the cellular elongation and sarcomere alignment of neonatal rat heart cells.¹⁸ Moreover, the cardiac function on PPy/SF composite fibers was evaluated with the hPSC-derived CMs, showing strong contraction. This study demonstrates PPy/SF composites provide both the physical structure and electrical environment to support the cardiac structure and function of CMs.¹⁸ Similarly, Tsui et al. (2018) reported great hPSC-derived CM survival, sarcomere organization, and gap junction formation on PPy-SF scaffolds over a 21-day culture period.¹⁹ PPy has also been blended with carbon nanotubes (PPy/CNT) and evaluated using neonatal rat ventricular myocytes (NRVMs) cultured over a 14-day period, which showed improved cell survival, organized gap junction formation, and cardiac functional development.⁷¹ PPy-derived nanoparticles (NPs) have also been added to the bacterial cellulose (BC) for the PPy NP/BC composites patch, which was evaluated with human cardiac fibroblasts and H9c2 cardiomyoblasts.⁷² Both cell types showed high cytocompatibility on the PPy NP/BC composite, with the H9c2 cells demonstrating improved cell viability, adhesion, proliferation and enhanced expression of cardiac troponin T (cTnT), alpha myosin heavy chain (αMHC), and Cx43.⁷² Similar results were reported by Spearman et al. (2015), PPy-polycaprolactone (PCL) interpenetrating networks supported CMs attachment and enhanced cell–cell communication through increased Cx43 formation in HL-1 CMs.⁷³

Fig. 2 Publication ranking of CPs in cardiac tissue engineering (2014–2024) – PubMed Database. Search terms used: (“cardiac tissue engineering” OR cardiac regeneration”) AND (“polypyrrole” OR “PPy” OR “PEDOT” OR “poly(3,4-ethylenedioxythiophene)” OR “polyaniline” OR PANI”).

Despite these promising findings, diverse reports highlight concerns regarding PPy cytotoxicity, particularly at high concentrations or when the polymer aggregates. Kai et al. (2011) reported that when the polymer is used in particularly high concentrations of 30% combined with PCL and gelatin, and they found hindered cell growth and metabolic activity after 8 days of culture.⁷⁴ Similarly, Humpolicek et al. (2018) showed that the decreasing biocompatibility of CPs, such as PPy can exhibit cytotoxic effects, potentially through mechanisms involving oxidative stress and apoptosis.⁴⁸ These varied findings suggest that PPy biocompatibility is highly influenced by polymer concentration, aggregation behavior, and how it is incorporated into composite biomaterials.

To mitigate the cytotoxic effects associated with PPy aggregation and high concentrations. He et al. (2018) developed a novel approach crosslinking methyl acrylic anhydride-gelatin (GelMA) and PPy nanoparticles via oxidative polymerization.⁷⁵ In their study, biocompatibility assessment with live/dead cell viability assays demonstrated that CMs maintained excellent viability (>80%) even at high GelMA–PPy nanoparticle concentrations (50 mg mL⁻¹) after 7 days of culture.⁷⁵ Moreover, their conductive membrane exhibited conductivity of 0.3 S m⁻¹, suitable for cardiac electrical stimulation applications.⁷⁵ Similarly, the hyperbranched poly(amino ester) (HPAE)-PPy/gelatin hydrogel was synthesized to achieve comparable biocompatibility results of over 80% cell viability.⁷⁶ While the conductivity of CPs was traded off for improved biocompatibility, these studies suggest that embedding PPy NPs within GelMA matrix can potentially mitigate the cytotoxic effects associated with high PPy concentrations while maintaining desirable electrical properties.⁷⁵

Other studies combining PPy with natural biomaterials like alginate or silk fibroin (SF)^18,77 or decellularized cardiac extracellular matrix (CG)⁷⁸ have been shown to enhance biocompatibility while preserving electrical conductivity. In a study by Parchehbaf-Kashani et al. (2021), their PPy/CG scaffold improved electrical conductivity six times higher than CG alone and maintained appropriate mechanical properties.⁷⁸In vivo assessment in a rat myocardial infarction (MI) model demonstrated that PPy addition supported the restoration of heart function, retention of transplanted cells, and increased neovascularization.⁷⁸ However, the cross-linked scaffold yielded lower Young's modulus values (1.63 ± 0.27 kPa) than native myocardium (∼22–50 kPa).⁷⁸ PPy–chitosan hydrogel also significantly improved electrical signal propagation after implantation.⁷⁹ Injection of PPy–chitosan hydrogel one week after MI significantly improved cardiac electrical function, including reduced QRS intervals and increased conduction velocities, while showing enhanced electrical connectivity in the scar regions and border zones.⁷⁹

While these studies show promising biocompatibility in terms of in vitro cell viability and in vivo cardiac function, the host-immune responses were barely evaluated for potential inflammatory responses. Kim et al. (2023) investigated this aspect and developed a scaffold coated with PPy that can carefully control immune responses. This PPy coating directed macrophages toward a healing state rather than an inflammatory one, which is essential for helping tissues repair themselves and integrate properly with the native heart tissue.⁸⁰

2.2. Poly(3,4-ethylenedioxythiophene) (PEDOT)

PEDOT was also widely used for biomedical applications due to its excellent electrical conductivity (up to 10³ S cm⁻¹).^{10,63,81–84} Typically, it is combined with the dopant of PSS to create a polymer blend (PEDOT:PSS) to achieve high conductivity, making it particularly suitable for cardioelectrical environments.^26,85 CMs cultured on PEDOT exhibit good attachment and expression of α-actinin, cTnl, and Cx43.^64,86–88 Further incorporating PEDOT into hydrogels, such as polyethylene glycol (PEG)-based hydrogels, has confirmed their cytocompatibility with supportive H9c2 myocyte adhesion and proliferation.⁸⁹In vitro experiment of PEDOT-based hydrogel demonstrated excellent electrical coupling, with 77.7% of CMs responding to electrical pacing and over 62% showing sensitivity to stimulation.⁹⁰ The electrical conductivity of ∼10⁻³ S cm⁻¹ was comparable to native cardiac tissue, which enabled effective electrical signal transmission between cells.⁹⁰

To further enhance PEDOT's bioactivity/biocompatibility, biomacromolecules were successfully applied. Recent studies have demonstrated this approach through electroconductive photocurable polyethylene glycol diacrylate (PEGDA)–gelatin/PEDOT:PSS hydrogels, where gelatin served as both a bioactive component providing cell adhesion sequences and a co-initiator for photopolymerization.⁹¹ The incorporation of PEDOT:PSS decreased the time for photocrosslinking from 60 seconds to 10 seconds while imparting electrical conductivity.⁹¹ These hydrogels exhibited tunable mechanical properties within the range of native myocardium (8–28 kPa) and promoted human cardiac fibroblast adhesion, demonstrating that peptide gel CP composites combine the bioactivity of natural peptide components with the electrical functionality of CPs, potentially making them attractive for next generation CTE strategies. Similarly, Roshanbinfar et al. (2024) cultured hiPSC-CMs in collagen–PEDOT:PSS hydrogel for 21 days.⁹² The CMs formed more matured, well-organized sarcomeres and displayed improved calcium handling which was later translated into infarcted mouse myocardium to reduce ventricular tachycardia from ∼35% in controls to ∼25%.⁹²

However, it is worth noting that some reports include the possibility of PEDOT:PSS becoming cytotoxic at high concentrations. For example, GelMA-based photo-crosslinkable hydrogel containing 0.3% w/v PEDOT:PSS showed cytotoxic effects due to increased presence of anions.^41,65 The PSS dopant has the safety risk of leaching out into the tissue and disrupting ionic balance, triggering oxidative stress and cytotoxicity.

2.3. Polyaniline (PANI)

PANI is another popular type of CP due to its tunable electrical conductivity, which can be adjusted by altering the pH or the level of doping. PANI can be synthesized through electrochemical techniques and chemical methods, such as oxidative polymerization, conventional free-radical polymerization, and enzymatic synthesis.⁷⁷ Researchers have explored the potential of PANI by mixing it with other materials to create composites to study how they interact with tissues in a living organism. For instance, pure PANI films implanted under rat skin showed minimal inflammation responses and no significant neoplastic tissue development for two years.⁹³ In addition, the conductive salt form of PANI (i.e., emeraldine salt) and its non-conductive base form (emeraldine base) showed decent biocompatibility as they did not cause severe inflammatory responses in a rodent model and allowed for cellular attachment and proliferation.¹⁴ Another benefit of emeraldine PANI was its ability to maintain electrical conductivity in physiological conditions for extended periods (up to 100 hours).¹⁴ This sustained conductivity is important for supporting the electrical maturation and synchronization of heart cells.¹⁴

In addition to these promising biocompatibility findings, peptide-based gel CP composites demonstrated that the incorporation of PANI into simple self-assembling 9-fluorenylmethyloxycarbonyl-tryptophan (Fmoc-W) peptide scaffold could produce an injectable conductive hydrogel with improved stability and support for electrically sensitive cells.⁹⁴ Building on this, Chakraborty et al. (2021) also developed a supramolecular biomaterial based on nanoengineered peptides with antimicrobial properties that combined di-Fmoc peptides with PANI to create a semiconductive hydrogel that was mechanically stable and supported synchronized cardiac cell function.⁹⁵ The foundational work by these authors revealed the potential for creating self-repairing peptide hydrogel composites using only organic conductive materials. Such materials could serve dual functions as both pressure sensors and electrogenic cell soft substrate.⁹⁶ These hydrogels demonstrated excellent biocompatibility with CMs and the ability to maintain intrinsic conductivity through self-healing properties.

In addition to peptide-PANI hydrogels, other composite strategies have also been explored to enhance the biocompatibility and functionality performance of PANI. Wu et al. (2023) evaluated the biocompatibility of PANI/lignosulfonate (LS) nanorods using H9c2 CMs.⁹⁷ These nanorods showed uniform morphology (∼453 nm length, ∼80 nm diameter) and good dispersion stability in aqueous environments.⁹⁷ Furthermore, when these PANI/LS were embedded in alginate hydrogel and injected directly into the infarct border zone of MI rats, they maintained excellent biocompatibility over 28 days of implantation with minimal inflammatory response and no adverse systemic effects. Also, histological and electrophysiological assessments showed improved electrical coupling, reduced scar formation, and increased angiogenesis within the infarcted area.⁹⁷

Despite the benefits PANI offers, concerns about its use in biological applications revolve around its non-biodegradable nature, which can potentially lead to chronic inflammation in the case of long-term implantations.⁹⁸ Also, its conductivity significantly decreases at physiological pH levels, and it can cause possible harm during its manufacturing process when strong acids like HCL are used in the doping process.^99–101 Another major limitation of PANI is its poor mechanical stability in its pure form. Although some studies have combined PANI with polymers like PES to improve mechanical properties and conductivity,^102,103 PES is not widely used due to its poor biodegradability and limited ability to support cell attachment. Therefore, more biocompatible polymers such as PCL, gelatin, or SF are preferred to provide mechanical support and promote cell compatibility.

2.4. CPs as active regulators of electrophysiology beyond conductivity

The ability for CPs such as PEDOT, PANI, and PPy to influence cellular electrophysiology can occur in ways beyond their ability to conduct charge. CPs exert direct effects on cellular electrical behavior through modulation of ion channel activity, calcium handling, gap junction expression and membrane properties to improve electrical maturation in cardiac cells. They directly influence voltage-gated ion channels critical for cardiac excitability. More specifically, PEDOT:PSS microwires demonstrated the capacity to modulate CM action potential with high spatial precision, synchronizing contractions to externally applied voltage pulses without compromising membrane integrity.¹⁰⁴ Similarly, the incorporation of PEDOT:PSS into collagen hydrogel matrices has been shown to enhance calcium transients in hiPSC-CMs, improve excitation–contraction coupling, and upregulate calcium-handling genes such as calcium voltage-gated channel subunit alpha1C (CACNA1C), Ryanodine receptor type 2 forward primer (RYR2F), and ATPase sarcoplasmic/endoplasmic reticulum Ca²⁺ transporter 2 (ATP2A2), indicating that CPs can promote more physiological calcium dynamics.⁹² The enhanced calcium handling translates to improved contractile function, with CMs cultured on CPs showing increased contractile force generation and more synchronized calcium waves across cell population. These improvements likely result from both direct electrical coupling through the conductive matrix and indirect effects on calcium channel expression and localization.

CPs have also been associated with broader electrophysiological maturation, including increased expression of sodium voltage-gated channel alpha subunit 5 (SCN5A) and potassium voltage-gated channel subfamily H member 2 (KCNH2), and improved functional coupling between CMs, while dopant-dependent variations in surface roughness and chemistry further influenced progenitor cell viability and adhesion.¹⁰⁵ At the cellular level, CPs modify membrane capacitance and seal resistance to create more favorable conditions for electrophysiological recording and stimulation. The enhanced seal resistance at cell–substrate junctions facilitates more efficient electrical coupling while reducing current leakage.¹⁰⁶ Additionally, CP-based surfaces promote integrin clustering and focal adhesion formation, thereby supporting mechanosensitive pathways including calcium and mitogen-activated protein kinase (MAPK) dependent pathways that govern proliferation and electrophysiological maturation.^107,108 Collectively, these findings suggest that CPs function beyond acting as passive conductors but also act as active regulators of electrophysiological behavior. These approaches explain why CP-based cardiac tissue constructs often demonstrate high electrophysiological properties including improved action potential characteristics, enhanced calcium dynamics and better intercellular electrical coupling compared to non-conductive alternatives.^73,109

3. Material characterization of CPs for CTE

A comprehensive evaluation of CPs is important for CTE applications. This is because the physical, electrical, and chemical properties of CPs directly influence their electrical conductivity, mechanical performance, biocompatibility, and cell–material interactions.¹¹⁰ The material properties of CPs are typically assessed at three arms: morphology and mechanical properties, electrical conductivity, and surface chemistry and functionalization for CTE, as summarized in Fig. 3.

Fig. 3 Methods and analysis for material characterization of CPs. DMA: dynamic mechanical analysis; SEM: scanning electron microscopy; XPS: X-ray photoelectron spectroscopy (C1s, O1s, N1s refer to carbon, oxygen, nitrogen core – level spectra); FTIR-ATR: Fourier transform infrared spectroscopy-attenuated total reflectance; AFM: atomic force microscopy.

3.1. Physical characterization: morphology and mechanical properties

The physical structure of CP scaffolds is recommended have an appropriate porosity and pore size distribution to allow cell infiltration, nutrient diffusion, and cell–cell interaction. Different scaffold designs have different infiltration capabilities. Porous CP scaffolds allow cellular penetration through interconnected porous networks and are recommended to have an optimal thickness range of 100–300 µm to facilitate cell infiltration and nutrient diffusion, particularly in the absence of vascularization.^111,112 In contrast, electrospun fibrous CP scaffolds, such as PPy/PCL or PEDOT:PSS composites (including those with other polymers like poly(lactic-co-glycolic acid), poly(glycolic acid) (PGA), or collagen), often face challenges in promoting cell infiltration due to their dense fibrous structures that hinder cellular penetration.^113,114 Strategies such as salt leaching, cryogenic electrospinning, and sacrificial fiber incorporation have been shown to increase porosity and can potentially be adapted for CPs to enhance tissue integration.¹¹³

The tensile strength, Young's modulus, and stretchability of CPs are recommended to closely match the mechanical characteristics of native myocardium. Small variations in these properties can significantly affect cellular and tissue interactions with the material. For mechanical measurements, common approaches such as uniaxial tensile, biaxial, and cyclic stretch testing could be used to capture the mechanical behavior and anisotropic properties relevant to CTE applications. Additionally, to accurately mimic the in vivo heart conditions, dynamic mechanical analysis (DMA) is performed under physiological conditions at temperature (37 °C), frequency of 1–3 Hz (mimicking heart rate), strain range of 10–15% (typical cardiac strain), and in a physiological buffering system like phosphate-buffered saline (PBS).^115–118

3.2. Electrical properties assessment

Evaluation of electrical properties is conducted under conditions that accurately reflect the electrophysiology of the human heart. This requires testing in physiological culture medium (typically Dulbecco's modified Eagle medium or similar) at 37 °C with controlled pH (7.4 ± 0.1)⁴⁸ and ionic strength (150 mM NaCl equivalent).⁷ Moreover, conductivity measurements can be performed using a four-point probe²⁰ method with direct current to eliminate contact resistance effects. However, for cardiac applications, impedance spectroscopy using alternating current could be used to better understand how CPs behave electrically across different frequencies under physiological conditions.^119,120 Additional resistance measurements and electrochemical evaluations such as cyclic voltammetry could provide more information about CPs’ stability and charge transfer properties under physiological conditions. The conductivity is suggested to fall in a range of ∼0.3–0.6 S m⁻¹ to match heart tissue conductivity.^7,121 The long-term stability studies are also suggested up to 30 days to track any changes in electrical conductivity with cell culturing.

3.3. Surface chemistry analysis

Chemically modifying CP surfaces with bioactive molecules, peptides, or extracellular matrix (ECM) proteins can enhance the attachment, structure, and function of CMs.^122,123 With this, important parameters such as hydrophilicity, degradation, and doping process need to be carefully studied to ensure long-term stability, conductivity, and biocompatibility with CMs.

Surface analysis includes high-resolution X-ray photoelectron spectroscopy (XPS) to analyze key elements such as carbon (C1s), oxygen (O1s), and nitrogen (N1s), to enable the identification of specific functional groups present on the CP surface.¹²⁴ Special attention was given to functional groups containing oxygen, as these significantly influence protein adsorption.¹²⁴ In addition to XPS, Fourier transform infrared spectroscopy with attenuated total reflectance (FTIR-ATR) was used to identify functional groups and assess surface modifications of CP in relation to biocompatibility.¹²⁵ FTIR-ATR enables the detection of important bonding structures, such as hydroxyl (–OH), carbonyl (–C=O), and amine (–NH₂) groups, which play a critical role in cell–material interactions.¹²⁵ Additionally, atomic force microscopy (AFM)⁶⁵ were used to measure surface roughness and topography at various scales to better accommodate cell attachment, spreading, and interaction.¹²⁴ Furthermore, the surface energy using contact angle measurements with deionized water was applied to evaluate the hydrophilicity of CP's surface.¹²⁶ A proper range of hydrophilicity supports the cell attachment with increased biocompatibility.¹²⁷

4. Biocompatibility evaluations of CPs

Biocompatibility testing of CPs becomes more significant as these materials are increasingly utilized in CTE to support and boost cardiac electrophysiology in potential cardiac dysfunction and damage. Due to the increasing significance of CPs in CTE and the unique chemical and physical properties of CPs on biocompatibility, a systematic approach to guide the evaluation of CP biocompatibility is outlined in Fig. 4. This roadmap integrates physical, electrical, and chemical characterizations (discussed in Section 3), in vitro cytotoxicity, in vivo evaluation, and advanced biological evaluation methods, highlighting the need for human-relevant testing platforms in the assessment process.

Fig. 4 The roadmap of the systematic approach for biocompatibility assessment of CPs, highlighting material characterization, in vitro testing, and advanced biological evaluation emphasizing human-relevant testing models.

Generally, how to select the appropriate method to evaluate the biocompatibility of CPs requires understanding the properties of the materials and how they interact with cells, tissues, and organs.¹²⁸ Materials with good biocompatibility would function effectively in their intended application without causing adverse reactions such as toxicity or inflammation to host tissues.¹²⁹ The materials also need to maintain structural and functional stability over time when exposed to biological fluids and cellular activity.⁴⁸ Specifically, when designing biocompatibility evaluation assays for CPs, it is crucial to consider both the biocompatibility of CP itself and its doping process, as both can significantly impact the interactions with biological systems and the overall biocompatibility.^38,52

4.1 Roadmap for evaluating the biocompatibility of CPs for CTE applications

The roadmap outlined in Fig. 4 provides a general guideline on how to evaluate biocompatibility in CPs for CTE. The process begins with sample preparation in the desired format (films, scaffolds, or hydrogels) using techniques like electrospinning, 3D printing, or casting. These formats are selected based on the intended CTE application. This is followed by a thorough material characterization of these formulated constructs, which is important for understanding the physical and chemical properties of the CPs. This stage involves assessing the conductivity (e.g., through four-point probe or impedance spectroscopy), porosity (e.g., through gas adsorption), and mechanical strength (tensile test) of the polymer, alongside an analysis of its surface properties, including topography, roughness, and functional groups. The characterization also involves the polymer's chemical structure examination to find potential impurities that can compromise its biocompatibility. Both sample preparation and material characterization may require 1–3 weeks each, depending on the complexity of the formulation and the number of characterization techniques used. These timelines may overlap when preparing multiple batches or formulations. This is followed by in vitro testing using cytotoxicity assays, histological analyses, immunohistochemistry to evaluate biocompatibility of CPs at the cellular level.

In vitro cytotoxicity assays are typically the first step in biocompatibility evaluation. It involves the exposure of standard or primary cell lines to assays for cytotoxicity, cell viability, and initial immune response evaluation. Common assays such as MTT, 3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium (MTS), cell counting kit-8 (CCK-8), and lactate dehydrogenase (LDH) are widely used to evaluate cell viability and metabolic activity when CMs are cultured on the CP substrates.^19,72,75 These assays may be performed at multiple time points (days 1, 3, 7, 14, up to 30) to provide longitudinal monitoring of cellular responses and may take 2–4 weeks. Initial assessment can employ a range of cell types relevant to heart tissue, including cardiomyocytes (e.g., primary hiPSC-CMs or H9c2 rat cardiomyoblasts), cardiac fibroblasts (e.g., NIH/3T3 or mouse embryonic fibroblasts), and endothelial cells (e.g., human umbilical vein endothelial cells (HUVECs)). In some cases, co-culture systems are included to assess inflammatory responses by measuring cytokine release and immune cell activation. It is recommended that the testing protocol follow ISO 10993-5 guidelines with specific modifications for cardiac applications.¹³⁰ The acceptance criteria for cell compatibility is recommended to be ≥80% viability compared to control surfaces at all points and should be consistent with increasing metabolic activity over time. In addition, live/dead staining may be combined with confocal imaging to get spatial information about the distribution and viability (>90%) of the cells. Flow cytometry analysis using Annexin V/Propidium Iodide (PI) staining may be performed to distinguish between apoptotic and necrotic cell death.¹³¹ This analysis may be conducted at early (24 and 72 hours) and later (7 days, 14 days) time points to capture both acute and chronic effects of the polymer on cell viability.¹³¹

For accurate cytotoxicity and reproducible testing, cells need to be seeded at proper densities (typically in the range of 1–5 × 10⁵ cells per cm² for CMs) on CP surfaces as cell number directly affects measurement outcomes.^92,132 Cytotoxicity assessment can also extend beyond standard 24–72 hours to 14 days or longer to evaluate long-term cellular response.⁴⁸ Throughout this period, it is recommended that the cells are maintained under controlled conditions with a temperature of 37 °C, CO₂ of 5% ± 0.5%, humidity of 95% ± 2%, and medium exchange every 48–72 hours.⁴⁸ Additionally, cell morphology and functional tests may be carried out through techniques such as scanning electron microscopy (SEM), fluorescence microscopy and transmission electron microscopy (TEM). Functional assays like contractility measurements and calcium signaling tests may be performed to provide further insights into the CP's compatibility with cardiac-specific cellular processes. These analyses typically take 2–4 weeks. Table 3 summarizes in vitro biocompatibility studies and performance evaluation of CPs in CTE.

Table 3 In vitro biocompatibility studies and functional assessment of conductive polymers in cardiac tissue engineering

Conductive polymers	Cardiac assessment methods	Cell types used	Key findings	Limitations	Ref.
PPy	Cell viability assessment (live/dead assay and MTT, MTS, CCK-8 assay), western blot, immunofluorescence staining, four-point probe, contractility and calcium transient analysis	Cardiac fibroblasts, H9c2 cell, hiPSC-derived cardiomyocytes, NRCMs, HUVECs, (2D and 3D models)	Improved cell viability (>80%), adhesion, maturation, and proliferation, electrical conductance matching native myocardium, non-cytotoxic at optimized concentrations.	Short term studies (≤4 weeks), limited scope on gene expression analysis. Dopant leaching concerns, aggregation issues at high concentrations, lack of long-term biocompatibility data	19, 72, 75, 76 and 133
PEDOT	Cytotoxicity test, live/dead assay, MTT, SEM, FTIR, electrophysiological measurements	3T3 fibroblasts hiPSC-CMs, primary CMs (2D and 3D models)	Non-cytotoxic, improved hemocompatibility, enhanced electrophysiological maturation, stable conductivity under physiological conditions	In vitro testing to assess hemocompatibility. Short-term studies. PSS toxicity concerns in some formulations, insufficient long-term degradation studies	41 and 65
PANI	Live/dead staining, almar blue assay, cell proliferation assays (MTT, flow cytometry), immunofluorescence, western blot, FTIR, XPS, SEM, conductivity measurements	H9c2 CMs, neonatal primary CMs, cardiac fibroblasts (2D and 3D models)	Good biocompatibility, supported cell viability and proliferation, promote myotube formation and sarcomere organization. Good dispersion in water, maintain conductivity in physiological pH	Short term studies, pH sensitivity affect stability, lack of long-term biocompatibility data, insufficient 3D tissue integration studies	14, 97 and 134

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If in vitro screening results are promising, the next phase involves a more comprehensive biological assessment through two parallel approaches, including advanced in vitro human heart models and small/large animal models. The advanced in vitro human heart models include tissue-specific systems such as hPSC-derived CMs, EHTs, and cardiac organoids, and other physiologically relevant platforms that better mimic the native human cardiac tissue in terms of structure and function.^135–137 These advanced models provide more accurate predictions of human biological responses and can reduce reliance on animal testing.^135–137 Advanced in vitro human heart models typically require 4–8 weeks for functional maturation and assessment. In vivo testing involves the use of animal models, either small animals like rats and mice or large animals like pigs, depending on the specific experimental goals.^138,139 At this stage, key considerations include the selection of appropriate implantation sites, whether subcutaneous, intramuscular, or left ventricle, and corresponding biomedical assessments, such as histological analysis, inflammatory responses, tissue integration and regeneration, and cardiac functional readouts.^138,139 Additionally, long-term studies are conducted to evaluate degradation, chronic inflammation, and tissue responses over time, to ensure that the polymer's performance aligns with the intended therapeutic results.^138,139 Animal studies generally require 8–12 weeks to evaluate chronic tissue responses. Also, ethical guidelines for animal research should be strictly followed to ensure responsible experimentation.

Following these tests, data analysis may be conducted to compare results against established controls and standards.¹³⁰ Statistical analysis typically uses Student's t-test for pairwise comparisons or one-way/two-way ANOVA with post hoc tests (e.g., Tukey's or Bonferroni correction) for multiple group comparisons with significant levels set at p < 0.05, p < 0.01, and p < 0.001.¹³⁰ Quantitative image analysis is usually performed using software such as ImageJ or MATLAB.¹⁴⁰ This detailed analysis phase, including histological processing, immunostaining, and functional data interpretation, may take 1–2 weeks. Techniques such as mass spectrometry may be used to detect impurities and quantify doping agents in CPs, offering chemical insights that help interpret biocompatibility outcomes. The last step in this process is to document and report the outcomes. Testing procedures are to be reevaluated and improved based on these findings. The protocols may be adjusted to improve the biocompatibility of the CPs, ensuring that future tests yield more favorable results in CTE.

4.2. In vitro human heart model, a new model for evaluating CP's biocompatibility

A major bottleneck in CP's biocompatibility evaluation is how to recapitulate human cardiac physiology, which is largely struggled and nearly unaddressable by the conventional in vitro 2D cell culture and animal models.^141,142 The recent in vitro human 3D models, including hPSC-derived EHTs and cardiac organoids, offer potential promises.^20,143–145 These models provide more biologically relevant models in both cellular structure and function for better assessing the material–human tissue interactions compared to traditional cell culture or animal models. The in vitro human 3D heart models align with the recent FDA Modernization Act 2.0,^56,146 which emphasizes the use of human-relevant testing methods that do not require animal models for testing biomedical safety and efficacy.

4.2.1 Application of CPs in 3D culture systems. The integration of CPs into 3D scaffolds has gained significant attention due to their ability to support electrical conductivity and cell viability in tissue engineering. This process involves careful control of the fabrication processes to achieve optimal balance in conductivity, biocompatibility, and porosity.¹⁴⁷ 3D culture systems address limitations of traditional 2D toxicity assays by better replicating the complex tissue microenvironment. These platforms enable the assessment of the materials degradation products, diffusion, pH changes, and their effects on surrounding tissue.¹⁴⁷ Impedance spectroscopy and live-cell imaging within these models provide continuous monitoring of the interaction between CP and native heart tissue, offering an advantage over endpoint analyses typically performed in animal experiments.¹⁴⁸

Representative examples show the diverse strategies and capabilities of CPs in 3D cardiac culture systems. As previously mentioned in Section 2.1, in vitro characterization of bacterial nanocellulose–polypyrrole (BC–PPy) conductive scaffold showed a tunable platform for CTE applications with electrical conductivities ranging from 1.7 × 10⁻⁸ to 2.4 ± 1.5 S cm⁻¹ as demonstrated in Fig. 5(A).⁷² Biocompatibility assessment using H9c2 rat cardiomyoblasts and human adult cardiac fibroblasts revealed that BC–PPy scaffolds significantly improved cell viability, adhesion and proliferation compared to non-conductive controls with 5 mM formulations (BP5) achieving optimal cell coverage of ∼40% versus 7% for plain bacterial cellulose. Moreover, the BC–PPy scaffolds demonstrated the ability to promote H9c2 differentiation toward a CM-like phenotype, evidenced by increased expression of cardiac-specific markers including cardiac troponin T (cTnT), alpha myosin heavy chain (αMHC), and Cx43.⁷²

Fig. 5 Examples of in vitro platforms showcasing how conductive polymers support cardiac cell viability, maturation, and functional development. (A) BC–PPy scaffolds demonstrate biocompatibility with H9c2 cardiomyoblasts, showing maintained cell viability. Adapted from ref. 72 with permission from American Chemical Society, copyright 2023. (B) Conductive hydrogel systems (heparin/chondroitin sulfate/hyaluronic acid-based formulations with PEDOT modifications) support iPSC-CMs viability in both 2D and 3D culture configurations. Adapted from ref. 149 with permission from Wiley-VCH GmbH, copyright 2024. (C) Collagen–PEDOT:PSS hydrogels promote hiPSC-CM structural maturation with organized sarcomeric development, functional contractile activity, and preserved responsiveness to pharmacological and electrical stimulation. Adapted from ref. 92 with permission from Wiley-VCH GmbH, copyright 2025. Data: mean ± SD or SEM. Statistical significance: *p ≤ 0.05, **p ≤ 0.01, and ***p ≤ 0.001.

Similarly, the in vitro characterization of glycosaminoglycan–PEDOT (GAG–PEDOT) conductive polymers demonstrated a self-doping system with enhanced electrical properties for cardiac applications in Fig. 5(B). These hydrogels were synthesized by using GAG with varying sulfation degrees (heparin (HepK), chondroitin sulfate (CSK), hyaluronic acid (HAK)) as backbones for PEDOT side chains. The self-doping mechanism provided by sulfate groups in the GAG backbone eliminated external dopant requirements.¹⁴⁹ When tested with hiPSC-derived CMs, the GAG–PEDOT polymers demonstrated spontaneous beating rates and higher electrical performance by requiring significantly lower threshold voltages for electrical pacing compared to non-conductive controls. The CP system preserved normal calcium handling in CMs indicating that the system did not negatively impact cardiac electrophysiology.¹⁴⁹

Moreover, the incorporation of PEDOT:PSS into collagen hydrogels increased electrical conductivity from 40.59 ± 2.75 to 65.32 ± 4.76 mS cm⁻¹ and significantly enhanced hiPSC-CM maturation in vitro as shown in Fig. 5(C).⁹² The conductive environment enhanced contractile function, with CMs showing higher contraction amplitudes on days 16, 20, 24, and 36 compared to collagen-only controls. Furthermore, hiPSC-CMs in PEDOT:PSS hydrogels showed better responsiveness to both pharmacological stimulation (epinephrine and isoproterenol) and electrical pacing (1.5 Hz and 2 Hz) on day 42, indicating improved functional maturation. Additional RNA sequencing analysis confirmed that the CP environment promoted upregulation of genes associated with cardiac muscle development, structural proteins and calcium handling.

Building on these examples that demonstrate enhanced cell viability, maturation, and electrical coupling, EHTs show significant advantages over conventional biocompatibility methods using cell lines in monolayer culture. Unlike conventional cytotoxicity assays that provide only short-term data, EHTs enable long-term evaluation of CP–tissue interaction.¹⁵⁰ They allow for high-throughput screening of CP formulations with various surface chemistries, stiffnesses, and degradation profiles to optimize scaffold design. The ability to generate patient-specific cardiac cells from hPSC differentiation allows for more thorough assessment of inflammatory responses, immune reactions, and cellular stress markers specific to human heart tissue with various genetic backgrounds. For example, regarding PANI/PES scaffolds, EHTs not only revealed cell viability after electrical stimulation for 15 days but also captured complex cellular responses such as changes in gene expression profiles and metabolic activities over extended culture periods.²⁰ The number of cardiac Troponin T-positive cells significantly increased, showing the potential of CPs for enhancing CMs’ growth and maturation.²⁰

Another approach using crystallized PEDOT:PSS (c-PEDOT:PSS) demonstrated comparable biocompatibility advantages. Cell viability studies after 1- and 3-day cultures showed very high cell viability (>98%).¹⁵¹ In contrast, both PEDOT:PSS and cross-linked PEDOT:PSS had relatively lower cell viability values which decreased to 70% for PEDOT:PSS and 82% for cross-linked PEDOT:PSS after 3 days. The cardiac cells showed similar expression patterns of α-actinin and Cx43, suggesting that c-PEDOT:PSS properly support the viability, maturation and synchronized beating of primarily cultured CMs.¹⁵¹ Overall, these studies highlight that CP-based 3D systems not only preserve cell viability but also promote functional maturation compared to traditional 2D assays, making them effective platforms for CTE applications.

Various fabrication methods have been developed for creating 3D cardiac engineered constructs that could be adapted for better CP integration, including hydrogel encapsulation, electrospinning, decellularization, and 3D bioprinting.¹⁴⁵ Electrospinning technique is used to create fibrous scaffolds that provide structural guidance and facilitate electrical signal propagation when incorporated with CPs.¹⁴⁵ One approach to incorporate CPs into traditionally non-conductive scaffolds involves coating methods.¹⁵² Another significant advancement in 3D cell culture is the development of composite scaffolds that integrate CPs with natural ECM components such as collagen, laminin, and hyaluronic acid.¹⁵³ These composite scaffolds integrated the electrical properties of CPs and biological support of ECMs, to promote cell adhesion and growth.¹⁵³ Additionally, they can fine-tune the mechanical properties of the scaffold with Young's modulus ranging from 6 kPa to 45 kPa to realize the physiological stiffness of human heart.¹⁴⁸

In situ polymerization has emerged as a particularly effective method for these composite approaches. For example, other studies have demonstrated successful incorporation of PPy into chitosan-based matrices through chemical oxidative polymerization.¹⁵⁴ This involved the grafting of pyrrole monomers onto the chitosan backbone and polymerizing in situ using ferric chloride (FeCl₃). The hybrid scaffolds maintained interconnected porosity and enhanced surface roughness with conductivity levels (0.0062 ± 0.0048 S cm⁻¹), which is comparable to the native heart tissue.¹⁵⁴ Additionally, the scaffolds improved cell alignment, sarcomere organization, and cell–cell communication through increased Cx43 expression of CMs.¹⁵⁴ This modification of the CP supported the formation of 3D cardiac cell networks, making it particularly promising in CTE.

4.2.2. Incorporation of CPs in cardiac organoid models. Building upon the advancements in hPSC-derived EHTs and 3D cell culture systems, the field of CTE has increasingly integrated the use of more physiologically relevant 3D models called cardiac organoids for studying disease mechanisms, drug testing, and regenerative medicine.^155,156 Cardiac organoids belong to 3D cell cultures and resemble closer human heart in organ structure, heterogeneous composition, and cardiac function. They are created by allowing cells to self-organize, often using a combination of scaffold-free techniques and scaffold-based methods.^155,157 Lewis-Israeli et al. (2021) developed a novel protocol for generating human heart organoids with spontaneous chamber formation and vascularization that self-assemble through sequential Wnt signaling modulation of pluripotent stem cells and subsequently model pregestational diabetes-induced congenital heart defects.¹³⁵ Moreover, we also emphasized that the incorporation of multiple cell types, including CMs, endothelial cells, smooth muscle cells, and cardiac fibroblasts, is critical for replicating physiologically relevant cardiac function.¹⁴⁵ Specifically, we recently have developed a robust and reproducible vascularization protocol for hPSC-derived cardiac organoid formation with built-in lumen-like vasculatures and myocardium-like structure.¹³⁶ When designing CP scaffolds, this cellular heterogeneity must be considered to ensure proper electrical coupling, vascularization, and mechanical support within the organoid.¹⁴⁵

While cardiac organoids have shown promising potential as physiologically relevant platforms for CTE research, the direct integration of CPs within cardiac organoid systems remains largely unexplored. This represents a significant research gap, given the established benefits of CPs in enhancing electrical coupling, promoting CM maturation, and improving tissue functionality in EHTs and 2D cardiac cell cultures. Cardiac organoids offer an ideal system for CP biocompatibility assessment because they incorporate multiple cell types, self-organized tissue architecture, and similar cardiac electrophysiology that better mimics native cardiac tissue than the conventional 2D culture systems.

4.3. In vivo biocompatibility studies of CPs in CTE

In vivo studies represent the gold standard to assess the biocompatibility of materials in tissue engineering. It is typically performed through subcutaneous implantation, intramuscular injections, or implantation in specific organs like the heart. Recent studies have employed diverse approaches to evaluate CP biocompatibility, ranging from general tissue compatibility studies to targeted cardiac applications as summarized in Table 4. Zhang et al. (2023) demonstrated CP biocompatibility through subcutaneous implantation of a PANI-fibrin scaffold with embedded vascular endothelial growth factor (VEGF) in mice model.¹⁵⁸ After one-week, histological analysis showed normal tissue architecture with no significant signs of inflammation or damage. Additionally, there were minimal fluctuations in blood cell counts, which suggested negligible immune response.¹⁵⁸

Table 4 In Vivo Biocompatibility studies of Conductive Polymers in Cardiac Tissue Engineering

Conductive polymers	Method of evaluating	Animal models used	Key findings	Limitations	Ref.
PPy-GelMA nanoparticles	H&E staining, Masson's trichrome, echocardiography, neovascular density quantification.	Sprague-Dawley rats MI model, 4 weeks evaluation post epicardial transplantation	50% reduction in infarct area, 20% increase in LVFS, no significant inflammatory response at 50 mg mL^−1	Short-term studies. Limited inflammatory marker studies. Biodegradation studies limited	75
PPy-chitosan conjugate	Trypan blue/DAPI cell viability, electromyography, QRS interval measurement	Sprague-Dawley rats, cryoablation injury, intramyocardial injection, 28 days evaluation post injection	Improved electrical signal propagation, reduced QRS interval, enhanced cell synchronization, no arrhythmia or mortality observed	Short term studies, limited histological analysis, limited inflammatory evaluation	133
PPy-xanthan gum/gelatin	H&E staining, echocardiography, tissue resistivity measurement	Sprague-Dawley rats, MI model, intramyocardial injection, 28 days post-injection	Reduced infarct size, increased LV wall thickness, decreased inflammation, enhanced angiogenesis, improved conductivity	Limited immunological markers, biodegradable study limited	159
PPy/gelatin-HPAE hydrogel	H & E staining, Echocardiography	Sprague Dawley rats, patch painted on left ventricle of infarcted area. 4 weeks evaluation post implantation (painting)	Rapid bond with myocardium, hydrogel degradation in 4 weeks, improvement of cardiac function parameters improved, promoted mature neovascularization	Long term stability beyond 4 weeks missing. Immune response after 4 weeks not covered	76
PEDOT: PSS-collagen	Electrophysiological analysis of VT, quantitative scar area analysis, microscopic assessment	Female CD1 mice MI model, intramyocardial injection, 2-4 weeks evaluation post infarction	Decreased ventricular tachycardia occurrence, reduced scar area, improved collagen gel formation	Limited-long-term biocompatibility, minimal inflammation markers	92
PANI/LS nanorod	Histological analysis (Masson's stain, immunofluorescence) in vivo cardiac function, fibrosis degree, arrhythmia assessment	Rat MI model, 28 days study, intramyocardial injection	Enhanced electrical conductivity, promoted gap junction formation, reduced fibrosis, improved cardiac function, increased VEGF expression and improved myocardial tissue repair	Long-term biodegradation and immune responses not fully studied.	97
PANI-fibrin-VEGF	Histological analysis (H&E, immunofluorescence, blood test)	BALB/c mice, subcutaneous implantation, 1 week evaluation	Excellent biocompatibility, low toxicity, immunogenic response	Short-term assessment, long-term functional integration and degradation not evaluated	158
PEDOT-PVA/GA hydrogel	Calcium transient, fluorescence microscopy. Histology analysis	Healthy adults C57BL/6 mice, left ventricle implant	No significant changes in ECG/QRS after 2 weeks, no functional impairment, slight fibrotic tissue formation, no inflammation over 2-week period	Short-term assessment, long-term biocompatibility not assessed	90

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Other studies have used direct cardiac applications to assess biocompatibility.⁹⁷ Roshanbinfar et al. (2024) cultured hiPSC-CMs with collagen-PEDOT:PSS hydrogel for 21 days.⁹² The CMs formed more mature, well-organized sarcomeres and displayed improved calcium handling.⁹² They then applied the collagen-PEDOT:PSS hydrogel in in vivo applications by injecting it directly into infarcted mouse myocardium. Over 2–4 weeks of in vivo evaluation, the system reduced ventricular tachycardia occurrence from ∼35% in controls to ∼25%, demonstrating biocompatibility while electrical conduction velocities and function were enhanced as shown in Fig. 6(A).⁹² In another study, the in vivo evaluation of heparin with ketone crosslinker and PEDOT side chains (HepK-PEDOT) conductive hydrogels demonstrated excellent biocompatibility and controlled degradation following intramyocardial injection in healthy rats.¹⁴⁹ In Fig. 6(B), both 7.5% and 10% (w/v) concentrations were successfully injected and remained localized at the injection site without adverse effects throughout the 2-month observation period. Histological analysis revealed significant cell infiltration within the hydrogel matrix at all time points, indicating active tissue remodeling without encapsulation or excessive inflammation. Additionally, the biodegradation rate was dependent on the concentration, with 7.5% hydrogels showing reduced presence by 14 days and near complete clearance by 4 weeks. However, the 10% hydrogels remained longer with complete disappearance by 8 weeks. Photoacoustic imaging successfully tracked hydrogel degradation and showed significant decay within the first 2 weeks followed by slower clearance,¹⁴⁹ indicating minimally invasive CTE applications of these conductive hydrogels.

Fig. 6 Examples of In vivo performance of CP hydrogels demonstrate therapeutic efficacy and biocompatibility. (A) Collagen–PEDOT:PSS hydrogel provides cardiac protection in mouse myocardial infarction models, showing significant reduction in ventricular tachycardia incidence and reduced tissue damage at 2–4 weeks post-treatment. Adapted from ref. 92 with permission from Wiley-VCH GmbH, copyright 2025. (B) Self-doped GAG–PEDOT hydrogels exhibit excellent biocompatibility with progressive biodegradation over 8 weeks in healthy myocardium, demonstrating safe tissue integration without chronic inflammation and quantifiable clearance kinetics via photoacoustic monitoring. Data: mean ± SEM. Adapted from ref. 149 with permission from Wiley-VCH GmbH, copyright 2024.

Beyond injectable approaches, direct implantation studies have further validated CP's biocompatibility. Luque et al. (2024) provided a thorough biocompatibility study through direct cardiac patch implantation of a PEDOT-based conductive patch on the left ventricle of healthy adult mice.⁹⁰ After 2 weeks, histological analysis revealed minimal inflammatory responses and only a thin fibrotic layer forming around the patches without significant changes in heart rate or cardiac function.⁹⁰ A paintable HPAE-PPy/gelatin hydrogel patch was applied to the infarcted left ventricle and showed fast and seamless bonding with the beating myocardium without requiring external sutures or additional fixation methods.⁷⁶ H&E staining revealed excellent tissue integration with complete biodegradation of the hydrogel within 4 weeks after implantation.⁷⁶ There were no adverse inflammatory responses, and neovascularization at the implant site was promoted together with improved cardiac function, further supporting the biocompatibility of CP-based cardiac interventions.⁷⁶

While these studies demonstrated encouraging biocompatibility profiles, in vivo research on CPs remains limited compared to in vitro studies, particularly in the context of cardiac repair and regeneration.¹⁶⁰ Animal models are important for understanding the biological interactions of CPs. However, they have several limitations, including growing ethical concerns, limited scalability for high throughput screening, and challenges in translating findings to human physiology. These limitations urge us to establish advanced in vitro human heart models such as hPSC-derived EHTs and cardiac organoids, which are proven to provide more physiologically accurate and ethically sound models for the biocompatibility evaluation of CPs in CTE.

5. Improvement of CP’S biocompatibility for CTE

The biocompatibility of CPs remains a critical consideration for their application in CTE. Cytotoxicity in CPs often arises from residual compounds, such as unpolymerized oligomers or acidic byproducts used during polymerization.¹⁶¹ The dimensions and configurations of the conjugated polymers in the composites, along with their insolubility and hydrophobic nature, can cause these compounds to accumulate and disrupt cell function by eliciting unwanted immune reactions.¹⁶¹ These effects can lead to inflammation and other negative side effects,¹⁶¹ limiting the long-term safety and effectiveness of non-biodegradable CPs. Recent research has explored hybrid scaffolds, molecular modifications, and biodegradable CPs as means to reduce their potential cytotoxicity and enhance their integration with cardiac tissues.

5.1. Hybrid CP systems for enhanced biocompatibility

Recent efforts have focused on improving the biocompatibility and mechanical properties of CPs for CTE. To achieve this, researchers have been developing hybrid scaffolds that incorporate CPs with natural and/or synthetic biomaterials. For instance, a significant advancement was developed with hydrogel-based devices, which demonstrated enhanced conductive capabilities while being mechanically flexible.¹⁶² A combination of chemical, polymerization, and electro-polymerization of PEDOT and polyurethane (PU) was adapted to create PEDOT/PU hydrogel hybrids that exhibit electrical conductivity levels reaching 120 S cm⁻¹ at 100% elongation.¹⁶² Other studies have shown that incorporating gelatin into PPy scaffolds can improve cell adhesion and proliferation while retaining the conductive properties necessary for cardiac tissues.⁵⁸ These hybrids demonstrated excellent biocompatibility with both muscle and neural cells as well, making them promising candidates for applications in tissue and regenerative medicine.

To address long-term biocompatibility challenges, recent reviews by Jadoun et al. (2021) and Kurowiak et al. (2023) provide insights into the integration of biodegradable polymers such as poly-D,L-lactic acid (PDLLA), PGA, polylactic acid (PLA), PCL, and certain PUs into CP composites for biomedical applications. When these biodegradable materials were mixed with CPs like PPy, PEDOT, and PANI, hybrid materials were synthesized with an ideal electrical conductivity and controlled biodegradation.¹⁶ The synthesis of these materials typically involves either creating block copolymers, where conductive oligomers connect with degradable ester bonds, or copolymerizing modified biodegradable components with electroactive elements.¹⁶ These materials can be fabricated through various techniques such as sol–gel methods, electrospinning, and 3D bioprinting, depending on their applications.¹⁶³ The resulting hybrid materials have shown promise in areas such as CTE, controlled drug delivery,^164,165 and antibacterial treatments.^166,167

5.2. Molecular engineering and surface chemistry modifications of CPs to improve conductivity and biocompatibility

Besides hybrid scaffolds, molecular engineering approaches, such as copolymerization and advanced doping techniques, have allowed for the fine-tuning of CP conductivity and biocompatibility.^2,168 By modifying the polymer backbone and side chains, researchers have been able to tailor the electronic properties of CPs. Innovations in doping, such as using biocompatible macromolecular dopants (e.g., heparin) and non-covalent dopants(e.g., ionic liquids) have resulted in materials that not only have higher conductivity but also retain their electrical properties under varying environmental conditions.^169,170 A promising approach to overcoming conventional CP limitations with dopant leaching is the development of self-doped CPs. For example, an innovative dopant-free conductive polyurethane elastomer (DCPU) was created by integrating biodegradable PCL, a conductive aniline trimer component, and dimethylolpropionic acid dopant molecules into a single polymer chain.¹⁷¹ This design eliminated dopant leaching concerns and maintained electrical conductivity under physiological conditions. Notably, the DCPU outperformed traditional doped systems by increasing conductivity to 264% of initial values after 150 hours in cell culture medium, making it potentially suitable for cardiac applications.¹⁷¹

Regarding surface chemistry modification of CPs, techniques such as grafting bioactive molecules, peptides, or growth factors onto the polymer surface have been shown to improve interactions between the material and biological tissues. For example, PPy and PEDOT have been functionalized with ECM proteins to promote cardiac and neural cell attachment and growth, making them more suitable for applications in cardiac and neural tissue engineering.^172,173 Hosseinzadeh et al. (2018) developed a conductive gel using aniline polymerization that mimics the native myocardium and promotes Cx43 expression in neonatal rat CMs.¹⁷⁴

6. Limitations

With growing ethical concerns over animal testing, there is a clear need to evaluate CP's biocompatibility in in vitro human-relevant heart models. Advances in hPSC-derived cardiac cells, and 3D culture models, such as EHTs, cardiac organoids, and organ/organoid-on-a-chip platforms, may provide promising alternatives to provide a more human cardiac microenvironment for biocompatibility evaluation of CPs.

Although CPs offer promising electrical conductivity and biocompatibility to enhance the outcome of CTE, their mechanical properties could be compromised along with cytotoxic by-products. A critical concern that remains inadequately addressed is the role of dopants in CP biocompatibility, which can negatively affect cell viability depending on their concentration and release profile. The systematic underreporting of negative dopant effects in literature also presents another significant limitation in the assessment of current CP biocompatibility. While our analysis identified several dopants with favorable biocompatibility profiles, toxicity data including leaching rates, dose–response relationships, and chronic inflammatory response are lacking for many commonly used dopants. Advances in surface functionalization and the development of CP composites have shown improvements, but there remains a significant gap between in vitro findings and in vivo performance, supporting the need for further research in this area. Additionally, while hPSC-derived EHTs have been used in CP biocompatibility studies, the integration of CPs into cardiac organoid platforms remains largely underexplored, representing a significant research gap.

Another critical consideration is biocompatibility evaluation under ISO 10993 framework, which provides internationally recognized standards for assessing cytotoxicity, sensitization, irritation, systemic toxicity, genotoxicity, hemocompatibility, and implantation responses.¹⁷⁵ Despite these standards, most CP studies remain short-term, with efforts to examine chronic effects limited and evidence on long-term responses scarce. Comparative work on PANI and PPy highlighted cytotoxicity and impurity-related limitations that must be addressed prior to their safe use in regenerative applications.¹⁷⁶ More advanced constructs, such as printable PEDOT-based cardiac patches and PEDOT:PSS-modified microelectrodes for 3D CM spheroids have demonstrated functional integration in vitro and in vivo, over the course of several weeks (2–8 weeks), leaving the understanding of long-term effects and safety unresolved.¹⁷⁷ This gap is consequential because chronic or progressive dopant leaching, material fatigue under cyclic cardiac loading, fibrotic encapsulation leading to electrical insulation and chronic immune activation are not predictable by short-term studies. Although CP platforms are beginning to be tested under ISO 10993 frameworks, systematic evaluation of dopant release, degradation by-products, scalability and chronic performance still remain important with ISO 10993-6 recognizing 12 weeks as the starting point for long-term implantation studies.

7. Future perspective

Moving forward, several key areas could be prioritized to improve the applicability of CPs in CTE and regenerative medicine. First, biocompatibility testing protocols need to be standardized, including long-term biocompatibility evaluation, and introducing in vitro human heart models to better predict outcomes that are relevant clinically. The suggestive roadmap in Fig. 4 provides a starting point. Advances in computational modeling and high-throughput screening could also accelerate the discovery of optimized CP formulations for specific cardiac therapy.

Future research is suggested to prioritize systematic long-term studies beyond six months with advanced in vitro and in vivo models. Additionally, researchers would employ potential accelerated degradation and aging models to anticipate chronic responses relevant to clinical translation. All polymer components including dopants and residual monomers and manufacturing processes (e.g. crosslinking, sterilization, batch to batch consistency of polymer synthesis on conductivity and mechanical properties) should be validated in preclinical testing following ISO 10993 standards with specific modifications for cardiac applications.¹³⁰ In-line quality controls, such as conductivity measurements, surface chemistry checks and mechanical testing, are highly recommended to be integrated with additional scalable fabrication strategies under current good manufacturing practice (cGMP) conditions.

Several strategies can potentially enhance CP stability and performance for cardiac applications. Bioactive stabilizer approaches, such as hyaluronic acid-stabilized PEDOT nanoparticles, have demonstrated improved conductivity over conventional PEDOT:PSS due to excellent NIH 3T3 fibroblast biocompatibility over 72-hour exposure periods.^40,178 Chemical crosslinking and additive strategies can potentially yield water-resistant, free-standing CPs, such as PEDOT:PSS films with enhanced conductivity and aqueous durability, addressing a key requirement for chronic wet-state implants or patch constructs.^179,180 Similarly, incorporating bio-functional polymers is an attractive approach. Dopamine-functionalized hyaluronic acid was paired with conductive polymers and produced adhesive, highly stretchable conductive hydrogels that retained appreciable conductivity in wet conditions, making them promising candidates for soft cardiac patches or injectable scaffolds.¹⁸¹

To overcome the limitations of unstable conductivity, new strategies aim to remove or immobilize mobile dopants and to decouple conductive functionality from chemically unstable polymer regions. Building on these advances, stimuli-responsive and targeted CPs strategies including PEGylation, surface functionalization, antibody conjugation, and incorporation into eutectogels¹⁸² have shown promise in enhancing circulation stability, enable targeted delivery, allow temporal control of electrical or mechanical signals, and mitigate immune responses. However, these approaches require systematic validation under regulatory frameworks.

Further efforts could explore the incorporation of therapeutic agents, such as growth factors (e.g., VEGF) or anti-inflammatory drugs, into CP-based scaffolds to promote angiogenesis and reduce inflammation at the implantation site. These bioactive-loaded CP systems could provide controlled release while maintaining electrical functionality.¹⁸² Additionally, the incorporation of antioxidants and cardioprotective molecules may further enhance long-term biocompatibility and support cardiac regeneration processes.^183,184 Emerging CP families such as polycarbazole-based materials and polythiophene derivatives offer chemically tunable backbones and side chains that can be engineered to introduce reactive handles for covalent bioconjugation.^15,185,186 However, while these chemical capabilities are well established in polymer and materials chemistry, direct demonstration of therapeutic growth factor conjugation or drug-loaded polycarbazole/polythiophene scaffolds applied specifically to CTE are relatively scarce.^185,187

Moreover, the emergence of 3D printed CP via additive manufacturing could enable the fabrication of intricate, functional structures with precisely controlled properties that would be impossible through conventional methods and can be adapted to create patient-specific cardiac scaffolds for precision medicine.¹⁸⁸ Importantly, the integration of these strategies with in vitro human heart models, direct in vivo cardiac injections, and large animal studies can potentially bridge the gap between experimental findings and clinical applications of CPs in CTE.

8. Conclusion

CPs including PPy, PEDOT, and PANI offer a unique combination of electrical conductivity and mechanical flexibility, which makes them suitable for supporting electrical coupling and conduction in cardiac tissue. Despite recent advances in hybrid scaffolds, molecular engineering, and surface modifications that have improved biocompatibility, these materials remain underutilized compared to conventional biomaterials. Current challenges center on standardizing biocompatibility testing protocols and bridging the gap between in vitro findings and in vivo performance. Hence, a systematic approach for evaluating CP biocompatibility that particularly focus on the use of human-relevant 3D heart platforms, such as hPSC-derived EHTs and cardiac organoids, will provide a more physiologically accurate system for evaluating CP–tissue interactions. These developments will further improve CP's performance and applications in CTE.

Author contributions

JA, HXY, and YH: original manuscript concept and draft; JA and ME: graphic design and illustration; HXY, JG, and YH: manuscript review and editing.

Conflicts of interest

There are no conflicts to declare.

Data availability

It is a review article, so there is no experimental data generated and available to share.

Acknowledgements

This work was supported by NIH R56HL174856 (HXY), Harry S. Moss Heart Trust (HXY), NSF CAREER DMR#1554835 (YH), and NIH R01HL175960 (YH).

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