Authors: Gagan K. Jalandhra (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Lauryn Srethbhakdi (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), James Davies (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Chi Cong Nguyen (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Phuoc Thien Phan (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Zachary Och (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Aditya Ashok (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia; School of Mechanical and Manufacturing Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Khoon S. Lim (School of Medical Sciences, University of Sydney, Sydney, NSW, 2006, Australia), Hoang‐Phuong Phan (School of Mechanical and Manufacturing Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Thanh Nho Do (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia), Nigel H. Lovell (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia; Tyree Institute of Health Engineering (IHealthE), University of New South Wales, Sydney, NSW, 2052, Australia), Jelena Rnjak‐Kovacina (Graduate School of Biomedical Engineering, University of New South Wales, Sydney, NSW, 2052, Australia; Tyree Institute of Health Engineering (IHealthE), University of New South Wales, Sydney, NSW, 2052, Australia)
Categories: Review, biomaterials, cardiovascular disease, devices, heart disease, tissue engineering
Source: Advanced Materials (Deerfield Beach, Fla.)
Materials Advances in Devices for Heart Disease Interventions
Authors: Gagan K. Jalandhra, Lauryn Srethbhakdi, James Davies, Chi Cong Nguyen, Phuoc Thien Phan, Zachary Och, Aditya Ashok, Khoon S. Lim, Hoang‐Phuong Phan, Thanh Nho Do, Nigel H. Lovell, Jelena Rnjak‐Kovacina
Abstract
Heart disease encompasses a range of conditions that affect the heart, including coronary artery disease, arrhythmias, congenital heart defects, heart valve disease, and conditions that affect the heart muscle. Intervention strategies can be categorized according to when they are administered and 1) Monitoring cardiac function using sensor technology to inform diagnosis and treatment, 2) Managing symptoms by restoring cardiac output, electrophysiology, and hemodynamics, and often serving as bridge‐to‐recovery or bridge‐to‐transplantation strategies, and 3) Repairing damaged tissue, including myocardium and heart valves, when management strategies are insufficient. Each intervention approach and technology require specific material properties to function optimally, relying on materials that support their action and interface with the body, with new technologies increasingly depending on advances in materials science and engineering. This review explores material properties and requirements driving innovation in advanced intervention strategies for heart disease and highlights key examples of recent progress in the field driven by advances in materials research.
Introduction
Cardiovascular disease (CVD) is the leading cause of death globally, representing one‐third of all deaths, with ischemic heart disease and stroke the major contributors to the disease burden.^[^
^1^
^]^ This alarming statistic underscores the urgent need for effective diagnosis, monitoring, and treatment strategies for cardiovascular disease.^[^
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Heart disease encompasses a range of conditions that affect the heart, including coronary artery disease, arrhythmias, congenital heart defects, heart valve disease, and conditions that affect the heart muscle. These conditions affect various regions of the heart and associated blood vessels, each with its unique anatomy, physiology, and function. Existing and emerging medical devices for the monitoring, treatment, and repair of cardiac function rely heavily on advanced materials that support device function and actively interface with the body. Consequently, clinical interventions require a deep understanding and precise tailoring of material properties for optimal device performance. The tuning of intrinsic material properties, such as chemical composition, surface features, stiffness, porosity, degradation, or endowing additional functionality, such as antifouling or antithrombotic properties can determine outcomes of medical device interventions.
This review will highlight advances in materials science and engineering that have been pivotal in the fight against heart disease. It is structured around three themes, 1) Monitoring, 2) Management, and 3) Repair (Figure
1 ). The Monitoring theme will explore technologies used in disease diagnosis and monitoring with the goal of timely intervention, using devices such as wearable and implantable sensors to obtain blood pressure, heart rate, and electrocardiography (ECG) information. The Management theme will discuss devices and technologies used in treating the symptoms of heart disease by assisting the heart in performing its function, typically by providing mechanical or electrical stimulus or support to the heart through devices like pacemakers, implantable defibrillators, cardiac patches, and stents for endovascular interventions. In the case of severe heart failure, interventions may include mechanical circulatory assist devices, such as left ventricular assist devices (LVADs), which were initially used as bridge therapies while patients were on a transplant list for a donor heart, but more recently, they are increasingly used as destination therapies. Finally, the Repair theme will explore technologies required when irreversible damage has rendered a structure unable to reprise its function via management strategies, and will specifically focus on strategies for the repair or replacement of damaged myocardial tissue and heart valves via tissue engineering.
This review will focus on the material advances that have revolutionized the diagnosis, monitoring, and treatment of heart disease, providing an overview of the current state of the art and future directions in this critical area of healthcare. We will first examine the requirements of devices for each form of intervention, then review notable trends in material innovations, highlighting pivotal works. This review aims to provide the reader with an understanding of the challenges associated with cardiovascular intervention, an appreciation of the diverse, interrelated strategies required to address them, and the important role of materials science and engineering in this field.
Monitoring
Monitoring represents the first line of CVD intervention. Sensor technologies provide information about key parameters of cardiac function. These assist in diagnosing conditions and symptoms that may lead to the onset of heart failure and in monitoring cardiac function to prescribe and assess the effectiveness of therapies. Effective and early detection of cardiovascular conditions that may develop into heart failure requires devices that enable longitudinal monitoring. These can take the form of less‐invasive wearable devices, which sense signals through the skin, or implantable devices, which directly interface with cardiac tissue. Both require electrodes that detect signals conveyed to onboard sensors and electronics via conductive interconnects. Output signals are then transmitted so that they can be interpreted to inform the state of cardiac function. Maintenance of conformal tissue‐device interfaces is paramount to reduce unwanted artefacts and impedance issues. Advances in soft bioelectronic materials have facilitated the development of sensing devices that provide critical insights into cardiac function for the diagnosis and monitoring of heart disease.
Wearable Sensors
The three most widely used sensing modalities to inform heart function are i) electrophysiology, ii) mechano‐acoustic sensing, and iii) photoplethysmography. These modalities provide information on electrical function, contraction function, and blood oxygen saturation, respectively.^[^
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^]^ Electrical activity of the heart is measured via sensing of biopotentials on the skin—known as electrocardiography (ECG). Electrodes placed on the skin convert the spatial and temporal summation of the biopotentials from the depolarization and repolarization of cardiac cells into voltage signals. These signals are used to diagnose numerous heart conditions, such as arrhythmias, ischemic heart disease, left ventricular hypertrophy, and coronary artery disease. Mechano‐acoustic sensors measure vibrations arising from within the chest cavity. These signals provide insight into the heart's mechanical contraction activity, i.e., ventricular contraction and valve opening/closing. Finally, a vast majority of commercial fitness trackers are based on photoplethysmography—where changes in light‐transmittance of tissue are used to measure blood oxygen saturation and heart rate. The following subsections present materials of choice, fabrication techniques, and sensing concepts in these three classes of wearable cardiac monitoring systems.
Flexible Epidermal Electrodes for ECG Monitoring
Holter monitors, which are typically used for 24–48 h for each measurement, have been the clinical standard for ECG monitoring to capture intermittent arrhythmias. These devices provide cardiologists with vital information to diagnose suspected arrhythmias and support the management of patients with cryptogenic stroke and paroxysmal atrial fibrillation.^[^
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^]^ A key burden for ambulatory use of Holter monitors is their bulky and tethered form factor, making them uncomfortable for many patients and more susceptible to losing conformal device‐skin contact, resulting in inaccurate and nonreliable readings.^[^
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^]^ As such, there is an emphasis on emerging diagnostic wearable devices with compact form factors that can maintain conformal contact with the skin for long periods, including physical activity. The Zio patch (Figure
2A) (iRhythm Technologies, United States) is an FDA‐approved lightweight adhesive patch applied to the chest that allows continuous monitoring of cardiac rhythm for up to 14 days via a single‐lead ECG.^[^
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^]^ The patch consists of electric components in a flexible housing adhered to the left pectoral via a perforated flexible adhesive tape. The perforated design maintains adhesion despite sweating. Newly detected atrial fibrillation or atrial flutter events lasting ≥30 s were recorded more frequently with 14‐day continuous monitoring using the patch compared with 24‐h Holter. Detection of nonsustained ventricular tachycardia was also more common with the Zio patch without any increased risks of adverse outcomes or excess resource utilization.^[^
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^]^ ECG monitoring is also becoming the mainstay in consumer products with the FDA 510(k) approval of Apple watches as a medical device for recording single channel ECG data and analyzing it for the presence of atrial fibrillation, sinus rhythm, and other arrhythmias including tachycardia. This was particularly useful for patient monitoring during the 2019 coronavirus pandemic.^[^
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^]^ Such approaches, however, are not continuous but episodic, relying on the user to routinely initiate a recording or if they had any concerns of irregularities in heart rhythm. Additionally, these devices are built on rigid materials, which leads to the presence of gaps in the device‐skin interface, causing areas that produce high impedances and result in inaccurate measurements with low signal‐to‐noise ratios.
![Figure 2: Wearable ECG electrodes on soft polymeric substrates for conformal adhesion to skin.A) The Zio patch is an FDA‐approved wearable device for single‐lead ECGs which allows continuous monitoring of cardiac function. Reproduced with permission.^[^
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^]^ Copyright, 2019, Springer Nature Limited. (Credit: iRhythm Technologies). B) Epidermal electronic system using a gold filamentary serpentine configuration. Reproduced with permission.^[^
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^]^ Copyright, 2013, WILEY‐VCH. C) Schematic illustrating the general process of flexible electrode formation coating a flexible polymer (e.g., PMMA) onto a rigid substrate (e.g., Si), evaporating and patterning (typically via lithography) of desired metal, and finally encapsulation of noncontacting areas. Flexible electrodes are typically characterized to evaluate their (i) electrochemical, (ii) electrical, and (iii) mechanical performance to ensure reliable operation in situ. D) Stretchable electrodes made using nanowires and PDMS. Reproduced with permission.^[^
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^]^ Copyright, 2018, RSC Publishing. E) Ultraconformable temporary tattoo devices based on inkjet‐printed PEDOT:PSS electrodes and gold interconnects (scale 2 cm) (i) Assembled circular electrode transferred on the arm; (ii) SEM images of a conformal contact between the E‐tattoo with a silicon substrate. (scale 400 µm). Reproduced with permission.^[^
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^]^ Copyright, 2018, WILEY‐VCH. F) SEM images of gecko‐inspired pillar structures to enhance adhesion between devices and skin (scale 20 µm). Reproduced with permission.^[^
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^]^ Copyright, 2016, American Chemical Society.](ADMA-37-2420114-g002.jpg)
In clinical settings, electrolytic gels—such as those based on silver/silver‐chloride—are used for device‐tissue coupling.^[^
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^]^ However, their propensity for drying and causing skin irritation has necessitated the development of dry electrodes from materials such as metals, conductive polymers, and nanocomposites. Despite the high conductivity of metals compared to conductive polymers or hydrogels, their higher stiffness leads to unwanted noise due to impedance arising from loss of conformal contact due to the stiffness mismatch at the device‐tissue interface.
One means of increasing the conformability of rigid conductive metals is by structurally engineering them into more compliant forms, such as filamentary serpentine configurations (visualized in Figure 2B) or nanowires supported by a soft polymer matrix. Filamentary serpentine design enables large strains (up to 100%) for otherwise rigid materials, such as gold, copper, and indium tin oxide, while maintaining contact between electronic components.^[^
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^]^ These structures are fabricated using lithography or cut‐and‐paste techniques. A general method for flexible electrode fabrication and characterization is illustrated in Figure 2C to aid readers’ understanding. Generally, fabrication involves four spin coating a flexible polymer onto a rigid substrate, evaporation, patterning of the desired metal, and finally, encapsulation of noncontacting areas. The resulting flexible electrodes are typically characterized based on electrochemical (i.e., cyclic voltammetry, electrochemical impedance spectroscopy, charge injection capacity), electrical (i.e., voltage‐current and noise analyses), and mechanical (i.e., flexion, bending, cyclic fatigue testing) performance (Figure 2C(i),(ii), and (iii)).
Nanowires are often preferred due to their high aspect ratios and lower percolation thresholds compared to nanoparticles. Nanowire‐based electrodes (e.g., gold, silver, indium tin oxide, graphene) show enhanced deformability and have shown promise in dry electrode fabrication.^[^
^14^ , ^15^
^]^ Silver nanowires (AgNWs), particularly, have shown great promise in dry electrode fabrication due to their high conductivity.^[^
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^]^ They have been incorporated into poly(ethylene oxide) (PEO) networks to form conductive inks for electrohydrodynamic printing of electrodes on polyethylene terephthalate (PET) and polydexamethasone (PDMS) substrates for stable ECG measurement and heating (Figure 2D).^[^
^18^
^]^ These nanowires can also be integrated into soft materials, such as polyvinyl butyral/hydrophilic polyurethane sponges or conductive polymer films for ECG signal detection.^[^
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^]^
Another strategy for maintaining conformal contact over extended periods is the use of self‐healing polymeric networks to fabricate stretchable electrodes that can repair themselves from damage. Self‐healing in polymers and electronic devices involves the recovery of mechanical, physical, and electrical performance after damage, through various covalent and noncovalent interactions and can occur autonomously or via external stimuli like light and heat. For example, a recently developed patch consisting of gold nanosheet electrodes and a self‐healing microporous foam allowed monitoring of multiple physiological signals including skin temperature, wrist pulse, and ECG.^[^
^21^
^]^ The patch consisted of a porous graphene foam coated with oxime‐carbamate bond‐based polyurethane and conductive thermoelectric polyaniline (PANI) for heat‐triggered self‐healing. The patch allowed stable physiological monitoring of multiple signals before and after self‐healing and in wet environments.
Similarly, gold nanowire networks were incorporated into polyurethane sponges to form deformable electrodes that provided conformal contact with skin while allowing air permeability.^[^
^22^ , ^23^
^]^ The electrodes were incorporated into an ECG module based on a flexible polyimide printed circuit board to collect ECG measurements during various everyday activities, including office work, driving, sleeping, and eating, demonstrating its capability for continuous measurement. Further, stable, high‐quality ECG and pulse signals were reported during intensive sporting activities, including walking, jogging, and running. Onboard wireless data transmission modules allowed real‐time continuous monitoring from anywhere via a smartphone. In the same study, gold nanowires were used to make electrodes for blood pressure sensing by dip‐coating gold nanowires onto tissue paper, followed by embedding in PDMS. The electrodes were integrated into a similar flexible pulse module and adhered to various locations via a sports adhesive tape to gather pulse waveforms at radial, temporal, carotid, and dorsalis pedis arteries. As a proof‐of‐concept, a machine learning algorithm was developed to estimate systolic and diastolic blood pressure based on simultaneous ECG and pulse measurements and trained on a 24‐person study. Validation using a cuff‐based sphygmomanometer showed results comparable to the FDA‐approved standard.
In addition to elastomeric materials mixed with metal nanowires, highly conductive materials, such as graphene have also been used in wearable ECG electrodes. In this context, graphene is biocompatible and enables scalable device fabrication through top‐down approaches, such as chemical vapor deposition. Submillimeter transparent graphene‐based electronic tattoos have been developed and can be transferred to the skin in a manner akin to temporary tattoos.^[^
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^]^
Developed in parallel with conductive elastomeric materials, a class of materials known as “epidermal electronics” or “electronic tattoos” has emerged to provide a promising means for epidermal biosensor attachment. In these systems, components such as electrodes, electronics, sensors, etc., are integrated into ultrathin, stretchable, skin‐like membranes, and adhered to the epidermis via van der Waals interactions.^[^
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^]^ Although currently not a long‐term monitoring solution, single‐ and multielectrode array tattoos provide an alternate approach toward seamless, intimate skin contact for continuous electrophysiological measurements (Figure 2E).^[^
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^]^ A challenge in electronic tattoo fabrication is the transfer of ultrathin electronics from one substrate to another. Tunable silk fibroin films have recently emerged as a viable solution for use as both the stamp and the receiver substrate.^[^
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^]^ Silk fibroin solution containing calcium ions was cast directly onto stretchable wave‐like Au conductors on a donor substrate to form films. The rigidity of the silk fibroin film was then increased by varying the humidity to above 100 MPa while peeling the metal conductors from the donor substrate to prevent damage due to excessive deformation. Following this, the silk was tuned to be soft (0.1–2 MPa) to apply the silk‐gold devices to human skin. The epidermal electrodes exhibit conformal contact and stable electrical and mechanical function under various skin deformation modes including compression, twisting and stretching.
Using intrinsically soft materials (e.g., liquid metals) can improve signal‐to‐noise ratios by reducing impedance issues encountered by rigid materials. Embedding rigid electrical components in soft, elastomeric materials such as PDMS is a promising strategy for achieving conformal contact between skin and epidermal sensing devices. The soft nature of such polymeric materials enables intimate contact with curvilinear epidermal topographies and easy deformation throughout movements while effectively absorbing strains/stresses to mechanically isolate embedded electronics. This has been demonstrated with stretchable ECG patches containing a monitoring chip‐on‐board with microfluidic liquid‐metal interconnects embedded in an elastomeric PDMS matrix.^[^
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^]^ In addition, clever structural engineering strategies have been employed to enable conformal tissue‐device interfaces. Patterning of gecko‐inspired micropillars on conductive dry adhesive pads made from embedding 1D carbon nanotubes and 2D graphene nanopowder in a PDMS elastomer matrix showed excellent epidermal adhesion and highly accurate ECG signal recording (Figure 2F).^[^
^29^
^]^ The metal‐free dry‐electrode device maintained contact throughout extreme movements and when under water.
Wearable Mechano‐Acoustic Sensors
Wearable mechano‐acoustic sensors provide vital signals, such as gallops, murmurs of aortic valve stenosis, and pulse waves associated with cardiac activities, such as closure of heart valves and change in blood pressure. The use of mechano‐acoustic sensors compliments ECG measurements, aiding physicians in monitoring cardiovascular health and diagnosing heart disease.^[^
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^]^ For instance, mechano‐acoustic signals provide indicators of structural defects in heart valves, which are not always reflected in the ECG response. Currently, analog and digital stethoscopes are clinical standards for physiological auscultation, but their rigid and bulky form factors make it challenging to adapt them to outpatient monitoring applications. Recent advancements have seen the development of soft wearable acoustic sensors with several types of form factors, including strain sensing patches, micro‐electromechanical systems (MEMS), acoustic/inertial measurement unit (IMU) sensors, soft integrated circuits, and thin films of ultrasound transducer arrays.
Due to their simple operating principle, strain‐sensing patches are one of the most common platforms used in wearable acoustic sensors.^[^
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^]^ Sound waves generated from the heart's activities propagate through tissues and fluids to the skin surface, inducing mechanical strains in sensing elements embedded inside elastomers attached to the skin. The induced strain is commonly captured via piezoresistive, capacitive, and piezoelectric sensing mechanisms (Figure
3A). These strain‐sensing patches are produced similar to the flexible ECG electrode described above where highly conductive composites (nanowires or graphene mixed with elastomers, such as PDMS or Ecoflex) or metallic traces with smart mechanical designs (e.g., wavy structures, interdigitated electrodes) are employed to adapt to epidermal stretching and deformation. For instance, a dual‐mode fiber‐shaped capacitive sensor was developed, which can be attached to different body parts for skin strain measurement.^[^
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^]^ The device was fabricated using direct ink writing to create two helical silver electrodes on an elastic thermoplastic polyurethane (TPU) fiber (diameter 0.5—1.5 mm) and encapsulated within Ecoflex (Figure 3B). TPU was selected due to its excellent elasticity of over 400% strain. Skin deformation leads to expansion or contraction of the TPU fiber, changing the distance between the two electrodes and resulting in a change in their capacitance. Attaching the devices to different body parts, such as the neck, chest, and fingertip, enables the detection of multiple signals, such as vocal sounds, respiration, heartbeat, and pulse waves, providing useful information for physiological assessment.
![Figure 3: Wearable mechano‐acoustic sensors. A) Illustration of basic principles of the mechanisms behind commonly utilized mechano‐acoustic sensors. i) Piezoresistive: applied pressure on a diaphragm causes tensile strain in a piezoresistive material. The tensile strain is directly proportional to the change in resistance. ii) Piezoelectric: application of pressure to piezoelectric crystals or ceramics leads to charge generation across the face of the material. iii) Capacitive: applied pressure changes the distance between two parallel conductive plates which changes the capacitance of the circuit. B) A capacitive sensor using helical Ag electrodes printed on polymeric fiber. Right: SEM images of the Ag electrode printed on the elastic TPU fiber. Reproduced with permission.^[^
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^]^ Copyright, 2023, Chi Zhang et al. C) Integration of acceleration sensors, MEMS microphone and Bluetooth low‐energy (BLE) for wireless, wide‐band‐width acoustic sensors. Right: The device was attached and tested on a 15‐month neonate to monitor their body sounds. Reproduced with permission.^[^
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^]^ Copyright, 2023, Springer Nature Limited. D) Wearable ultrasound imaging device using a piezoelectric array and liquid metal interconnect. Reproduced with permission.^[^
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^]^ Copyright, 2023, Hongjie Hu et al.](ADMA-37-2420114-g020.jpg)
The bridge‐island design concept has also been employed to develop epidermal mechano‐acoustic sensors, where off‐the‐shelf accelerometers together with signal processing components (e.g., Op‐Amp, capacitance, and Bluetooth module) are embedded inside polymer (e.g., Ecoflex) and interconnected through serpentine copper traces. Here, the serpentine copper trace enables devices to stretch following skin deformation while minimizing strain induced at the interconnect between electronic components and soft circuitry. Representative examples of this approach include the work from Liu et al. that utilized a low‐power accelerometer with a bandwidth ranging from 0.5 to 550 Hz, a compact form factor with a total weight of 213 mg, a small thickness of 2 mm, together with a low Young's modulus of ≈31.8 kPa and thickness of 2 mm.^[^
^33^
^]^ These soft, conformal devices enable precise recordings of physiological signals. Experiments show these acceleration sensors can record seismocardiography (mechanical vibrations of the chest wall caused by the heart's activity) from the chest and heart murmurs from four auscultation sites, essential for cardiovascular monitoring. They also monitor ventricular assist devices (VADs), detecting vibrational acoustics changes that indicate thrombosis or mechanical failures, providing a noninvasive tool for monitoring heart failure patients. Beyond medical uses, these devices can be easily attached to different parts of the body (chest and neck) and demonstrate promise in human–machine interfaces like speech recognition and video game control, demonstrating their versatility in daily life and healthcare.^[^
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^]^
While acceleration sensors exhibit relatively low‐frequency bandwidth (several hundred Hz), MEMS microphones offer a wider frequency range (30 Hz to 20 kHz), enabling the detection of sound waves generated from different organs in the body (e.g., heart, lung, and gastrointestinal tract).^[^
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^]^ These soundwaves can be differentiated based on the frequency domain or with the help of artificial intelligence. For instance, a MEMS microphone was integrated into a lightweight design and a low‐power consumption circuit that can be attached to the wrist for continuous monitoring.^[^
^37^
^]^ The study compared the performance of the proposed algorithm with other photoplethysmography (PPG)‐based devices using data from 12 subjects. PPG is a noninvasive optical technique that measures blood volume changes in peripheral circulation. The results demonstrated high accuracy, with a mean absolute error of 0.28 bpm and a standard deviation of 0.86 bpm for heart rate monitoring at the wrist. The algorithm utilizes techniques, such as K‐means clustering, energy envelope construction, and artifact identification to achieve accurate heart rate estimation. This method can potentially complement or replace PPG devices for accurate heart monitoring at the wrist. Further investigation into other cardiac biomarkers is warranted to fully explore the potential of this method. Recent studies suggest an integration of both acceleration sensors and MEMS microphones into a single platform that covers most of the required measurement range for body sounds for multimodal sensing.^[^
^35^
^]^ Acceleration sensors detect low‐frequency sounds, while MEMS microphones track higher‐frequency sounds (Figure 3C). The mechanical flexibility and wireless communication features of these mechano‐acoustic sensors (embedded inside Ecoflex with low Young's modulus) have been utilized in recent clinical studies on neonatal monitoring. The systems simultaneously captured heart, respiration, and bowel sounds while minimizing the tethered interfaces typically associated with conventional monitoring devices. These measurements offer a powerful data set for clinical decision making and enhance user comfort, especially for patients with highly sensitive skin, such as neonates.
Epidermal electronics using bridge and island structures have enabled a new class of wearable ultrasound sensors. Ultrasound scanning devices have been a cornerstone of several clinical procedures for monitoring and visualizing the structure and dynamics of inner organs, including blood flow into and out of the heart.^[^
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^]^ However, their bulky design and operator requirements confine their applications to clinical settings and short measurements. The mechanical flexibility, lightweight nature, and wireless communication capabilities of epidermal ultrasonic devices show promise for continuous health monitoring. Representative examples of these devices include work from the Xu group, which utilizes piezoelectric composites interconnected by a network of eutectic gallium–indium liquid metal and encapsulated within two layers of triblock copolymer (Figure 3D).^[^
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^]^ The device has a low Young's modulus of 921 kPa, compatible with human skin, and offers stretchability of up to 110%, withstanding several deformation cycles. Ultrasound is generated from the piezoelectric electrode array, penetrating through tissue and biofluid, bouncing back from the targeted organ, and being detected by the wearable piezoelectric transducers. The device employs a wide‐beam compounding approach, obtaining higher image quality than plane‐wave and monofocus methods. The device successfully captured echocardiographic images, including the apical four‐chamber view, parasternal long‐axis view, and left ventricle, which are widely used to assess the blood delivery capabilities of the heart and its structure. Ultrasound images obtained from wearable mechano‐acoustic sensors can be used at rest and during motion, providing critical information related to cardiac activities such as myocardial displacement, stroke volume, ejection fraction, and cardiac output.
Wearable Devices for Measuring Oxygen Saturation
Blood oxygen saturation (SpO2)—the ratio of oxygenated hemoglobin to total hemoglobin concentration—measures oxygen ventilation and perfusion, critical during surgery and when monitoring patient health. SpO2 measurements—usually via photoplethysmography—can provide valuable insights into heart function for monitoring and diagnosis of chronic cardiovascular disease.^[^
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^]^ Clinically, oxygen saturation is measured using an oximeter. This rigid device is placed on the patient's index or middle finger, where photodetectors measure light from light emitting diodes (LEDs) passing through the finger. The difference in light absorption rates between oxyhemoglobin and hemoglobin gives oxygen saturation. The rigid nature of oximeters means that the light often scatters, leading to inaccurate readings.^[^
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^]^ Further, due to the presence of a lead wire, continuous measurement during activity is not possible. The growing popularity of wrist‐worn fitness trackers and smartwatches has allowed continuous measurement in a less invasive, more mobile, and convenient manner.^[^
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^]^ These devices function based on photoplethysmography, wherein photodetectors measure the intensity of transmitted or reflected light due to blood flow in the arteries of the wrist. Onboard LEDs act as the light source. However, the rigid nature of these devices, again, means conformal contact cannot be maintained through physical activity, leading to inaccurate continuous measurements due to motion artifacts.^[^
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^]^ Innovation in this space is driven by the need for flexible devices with a minimal footprint that can continue to provide accurate continuous oxygen saturation sensing during movement.
To address the rigidity of inorganic photodetectors used in photoplethysmography, organic photodetectors based on graphene sensitized with quantum dots have been developed.^[^
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^]^ Ultrathin (30 nm) and flexible photodetectors formed via assembly of chemical vapor deposited graphene and colloidal semiconducting PbS quantum dots are also semitransparent, allowing discrete, minimally invasive measurements. Devices were formed by incorporating the flexible graphene photodetectors onto a 125 µm thick flexible polyimide substrate containing Ti (2 nm) and Au (50 nm) electrical contacts and Al2O3 dielectric layer (100 nm). SpO2 and heart rate measurements could be simultaneously recorded and monitored wirelessly via a smartphone, enabled by a Bluetooth module. LEDs are commonly utilized as the light source is pulse oximetry. However, they are rigid, and their bulky form factor and need for a power source hamper their potential for miniaturization. This has led to the development of ambient light oximeters, which obviate the LED requirement by utilizing ambient light as the source. Spectrally selective organic photodetectors with light filters eliminate the need to control light source such that oximetry readings can be performed in sunlight and under fluorescent, LED, and incandescent light sources (Figure
4B).^[^
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^]^ The flexible device is placed on the index finger, where it collects measurements and relays them to a control board housed on the hand. The board processes these signals and wirelessly transmits them to a computer for visualization. Validation against a commercial oximeter showed less than 1% difference in SpO2 measurements.
![Figure 4: Oxygen saturation measurement using wearable optical sensors. A) The basic principle of photoplethysmography. Devices generally consist of a light‐source (LED) and a photodetector. Changes in light either reflected (configuration on the left) or transmitted (configuration on the right) correspond to changes in the volume of blood in vessels during cardiac cycles. B) Pulse Oximetry in Ambient Light using Organic Optoelectronics. Inset: image of the photodiode. Reproduced with permission.^[^
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^]^ Copyright, 2020, WILEY‐VCH. B) Wearable QD‐LED optical sensors using wavy, pretrained LED. Top SEM image of wavy structure. Top The LED retains stable performance under mechanical strain. Bottom, demonstration of pulse measurement. Reproduced with permission.^[^
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^]^ Copyright, 2017, American Chemical Society.](ADMA-37-2420114-g011.jpg)
Reliance on ambient light for measurements limits continuous long‐term monitoring, as measurements are not possible in low‐light environments or at night. This is a driving force for the development of light sources in wearable oximetry devices which can overcome the limitations of rigid and bulky LEDs. Quantum dot LEDs have emerged as a promising alternative (Figure 4C).^[^
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^]^ Recently, cadmium selenide/zinc sulfide quantum dots—which possess photoluminescence characteristics of ≈620 and 530 nm peak wavelengths—were incorporated into PET‐based 1D fibers which could be woven to form 3D structures.^[^
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^]^ The devices exhibit narrow full‐width half‐maximum characteristics and high selectivity and sensitivity, meaning they provide highly accurate SpO2 measurements. The flexibility of PET encapsulation, organic photodetectors, and the poly(3,4‐ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) transparent electrodes contribute to the utility of these fibers for the fabrication of advanced wearable technologies for monitoring.
Heart rate can easily be extracted by measuring the changes in transmittance during pulsatile flow of blood through arteries. It is desirable to measure multiple parameters simultaneously to provide a comprehensive window into cardiovascular health. As such, there is ongoing work to develop algorithms and machine learning approaches that can extrapolate parameters, such as blood pressure by analyzing photoplethysmography pulse wave data.^[^
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^]^ This may provide cuffless blood pressure measurements without the requirement of pressurization.
Implantable Sensors
While having many advantages around convenience and minimal invasiveness, wearable sensors are limited in precision and resolution due to their distance from the heart. This also makes them prone to motion artefacts and noise generated by the other structures inside the chest cavity including lungs and skeletal muscle. Due to their direct interface with the heart and large vessels, implantable sensors are powerful tools for precise and detailed measurement of vital parameters such as blood pressure, blood flow, blood volume, and mapping of electrophysiological activity. These platforms provide valuable insights into electrical and mechanical cardiac function to improve diagnostic capabilities and aid timely intervention. Advances in soft materials (for conformal contact with the soft heart tissue), microfabrication technologies (for installation at hard‐to‐reach and narrow sites), and battery‐free wireless communication (bypassing long wires in conventional tethered configurations) have enabled the development of several types of implanted cardiovascular sensors that can be installed onto the heart tissue or inside the pulmonary arteries. The following sections highlight recent developments of some key implanted sensors for heart failure monitoring, including i) electrode patches/arrays for cardiac physiology monitoring^[^
^51^
^]^ and ii) mechanical sensors for hemodynamic monitoring.^[^
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Bioelectronic Patches for Electrophysiology Monitoring
Implanted electronics enable continuous measurement or spatial mapping of cardiac electrophysiological monitoring, providing feedback signals for cardiac neuromodulation (e.g., used in pacemakers and implantable cardioverter‐defibrillators) and detection of dysfunctional cardiomyocytes. Several bioelectronic patches have been developed for endocardial (typically for acute treatment) and epicardial (generally aimed at chronic monitoring) applications.
Smart patches integrated with highly stretchable electrodes and sensors have been deployed in invasive endocardial tools such as balloon catheters. For instance, by employing a thin Au film coated on a flexible polyimide substrate, a highly integrated system of electrode arrays (Cr/Au, 10 nm/300 nm in thickness), temperature sensors (Cr/Au, 10 nm/100 nm in thickness), and pressure sensors (Au‐based 3D structures fabricated using mechanically buckling approach) was developed for endocardial procedures.^[^
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^]^ Pressure sensors ensure physical contact between the balloon and heart tissue, which is critically important for subsequent physiological mapping and ablation (Figure
5A). The 8 × 8 array of Au electrodes was employed to map cardiac electrograms and to remove dysfunctional and aberrant tissue conduction pathways using radiofrequency ablation electrodes. Au temperature sensors serve as a feedback loop for the ablation process. Integration of multimodal components (recording, stimulation, ablation) into a single platform provides a powerful tool for all‐in‐one monitoring and therapeutic procedures.^[^
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![Figure 5: Implanted electronic patches. A) Multifunctional sensors and electrodes for endocardial balloon catheter. Reproduced with permission.^[^
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^]^ Copyright, 2020, Springer Nature Limited. (scale (i),(ii) 500 µm, (iii) 500 µm (left), and 100 µm (right)). B) i) Design process for 3D multifunctional integumentary membrane (3D‐MIMS) including imaging and sectioning of cardiac tissue for topography matching. ii) Photograph of a device on Langendorff‐perfused rabbit heart. Rproduced with permission.^[^
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^]^ Copyright, 2014, Springer Nature Limited. (scale 6 mm). C) An epicardial bioelectronic device for electrocortical mapping based entirely on rubbery elements. Reproduced with permission.^[^
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^]^ Copyright, 2020, Springer Nature Limited. D) Simultaneous optical pacing and electrical recording using flexible printed circuits integrated with micro‐LEDs and Pt microelectrodes. Reproduced with permission.^[^
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^]^ Copyright, 2022, The American Association for the Advancement of Science.](ADMA-37-2420114-g007.jpg)
While flexible electrodes attached on deployable balloon catheters are mainly used in acute treatment, electronic patches with arrays of recording/stimulation electrodes and other functional components (e.g., strain sensors and energy harvesters) are employed on the epicardial surface for chronic, spatiotemporal monitoring of cardiac activity. Metallic materials (e.g., Au, Pt) are commonly used to fabricate recording and stimulation electrodes. They are structurally engineered in 2D serpentine shapes or 3D out‐of‐plane configurations to enable mechanical flexibility and stretchability to ensure compliant and intimate contact with the dynamic epicardial topography. For instance, one epicardial electronics system for cardiac activity mapping featured a 25 µm flexible polyimide (PI) substrate containing single‐crystalline silicon nanoribbons and metal (Cr/Au) interconnects.^[^
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^]^ An 18 × 16 array of gold electrodes—each connected to an associated amplifier and multiplexer—spaced 800 µm apart maintained contact with the curvilinear epicardial surface of an exposed porcine heart through cardiac motion and provided reliable electrophysiological mapping data. The 288 measurement points were integrated into 36 wires connected to external data acquisition and control units. In addition to metallic electrodes, Si nanomembranes are also utilized in epicardial patches, where the membranes are transferred onto a polymeric substrate using mechanical stamping.^[^
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^]^ In these devices, using serpentine structures or coiled bridges mitigates the strain induced into the Si material when the patches are subjected to mechanical contraction and extension.^[^
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^]^ Silicon nanomembrane‐based transistor arrays have been integrated with a thermally grown ultrathin silicon dioxide layer to fabricate multiplexed flexible epicardial patches.^[^
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^]^ The dielectric thermal oxide layer allows capacitive coupling to underlying tissue and protects integrated electronics from bio‐fluids to enable longer‐term physiological stability (120 days).
Although structural engineering of rigid materials (Si nanomembranes or metallic nanothin films) can improve mechanical flexibility and stretchability, the hard‐soft device‐tissue interfaces may induce local mechanical strains and interfere with natural deformations. To address this limitation, soft, stretchable substrates, such as those used in 3D multifunctional integumentary membranes (3D‐MIMs) (Figure 5B), have been developed. 3D‐MIMs employ a soft, stretchable layer of silicone cast around a mold, which is prepared using 3D geometric data from optical segmentation of the epicardial surface to match the tissue topography (Figure 5B(i)) and maintain the device‐tissue interface.^[^
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^]^ Embedded gold electrodes enable high‐precision spatiotemporal mapping of electrical activity. Furthermore, the incorporation of indium gallium nitride inorganic micro light emitting diodes (µLEDs) for optical mapping, iridium oxide pads for pH sensing, gold resistors for temperature sensing and heating, and silicon nanomembranes for strain gauging measurement provides a comprehensive mapping of cardiac activity markers on the anterior and posterior epicardial surfaces (Figure 5B(ii)). However, the device requires a cable connection to an external data acquisition, control, and power supply module. Further, efficacy was only demonstrated in an ex vivo rabbit model.
Another strategy for achieving and maintaining adhesion to wet curvilinear epicardial tissue is via thin photopatternable conductive‐polymer gel‐electrode arrays based on polyvinyl alcohol (PVA) dispersed in polyrotaxane.^[^
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^]^ The polymeric network can absorb stresses and maintain conformal contact for up to 3 h on the wet surface of a beating rat heart to provide stable ECG and strain measurements with high signal‐to‐noise ratios. More recently, epicardial bioelectronic patches made entirely of soft rubbery materials—containing rubbery transistor arrays, strain sensors, temperature sensor/thermal actuator, and mechanoelectrical transducer—have been developed (Figure 5C).^[^
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^]^ Semiconductors used herein are based on flexible poly(3‐hexylthiophene‐2,5‐diyl) nanofibrils in PDMS. Silver nanowires (AgNWs) embedded in PDMS serve as electrodes, and gold nanoparticles coated in AgNWs/PDMS are employed as strain sensors. The interconnections are made from a conductive rubbery paste instead of a serpentine or otherwise structured rigid metal. The device was adhered in vivo to a porcine heart via a surgical adhesive (BioGlue) where it allowed spatiotemporal electrophysiological mapping, and strain and temperature sensing. Currently, the device is powered and operated via a wired connection. However, incorporating a rubbery mechanoelectrical transducer showed the potential for a self‐powered implantable device by utilizing piezoelectric materials. As well as their intrinsically stretchable properties, several gel electrode materials exhibit excellent compatibility with 3D printing techniques, allowing for the development of complex structures that can match the curvilinear morphologies of the heart through additive manufacturing. Wang et al. developed high‐performance 3D printable bioelectronics based on PEDOT:PSS hydrogels. Sequential direct ink extrusion of substrate, conductive, and encapsulation hydrogel inks is used to fabricate devices which exhibit mechanical compliance and robust interfacing with Langendorff rat heart tissue ex vivo. Conformal adhesion—especially on beating hearts—reduced impedance and allowed high‐quality in vivo ECG monitoring in a murine model. Further, signals from 16 nodes were combined to allow spatiotemporal biopotential mapping. In addition to monitoring, the devices showed capability for restoring induced arrhythmia via electrical stimulation. This study highlights the potential for additive manufacturing—which is inherently scalable—in fabrication of multifunctional soft bioelectronics for cardiac monitoring and treatment.^[^
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In addition to physiological mapping, conductive epicardial bioelectronics can be leveraged for simultaneous monitoring and electrocoupling treatment of heart diseases. As such, in situ formation of compliant bioelectronic patches with dual functionality of bioelectronic monitoring and passive assistance in tissue repair via electrocoupling therapy have been developed.^[^
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^]^ A precursor solution based on functionalized polyaniline (f‐PANI) and PVA hydrogels is painted directly on the epicardial surface. Chemical and physical crosslinking anchor the adhesive hydrogel patches to the myocardium, which does not interact with surrounding tissue surfaces upon gelation. Introducing protonated imines (─NH+ = C─) into PANI creates more protons, resulting in a high conductivity (1.35 ± 0.22 S m^−1^) and high charge injection capacity (0.637 vs 0.126 mC cm^−2^ in PVA hydrogel), suitable for electrophysiological diagnostic and therapeutic purposes. The device employs changes in the resistance of PANI under mechanical strain to detect microdeformations from diastole and systole. It continuously monitors stroke amplitude and rhythm to provide insights into cardiac pulsation and pumping function. Through experiments in a rat model, the implanted patches demonstrated the potential to reduce the effects of myocardial infarction (MI) by providing a conductive path to compensate for the reduced electrical conduction in damaged post‐MI fibrous tissue. This enhanced electrical response, and propagation promotes vascular regeneration while inhibiting ventricular remodeling. The concept of real‐time monitoring and synchronous MI repair assistance using stretchable bioelectronic patches was further explored in a recent work employing ion‐conductive polyacrylamide hydrogels reinforced with core–shell curcumin‐polydopamine nanoparticles.^[^
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^]^ The nanoparticles provide stress‐dissipation and additional crosslinking to the highly mechanosensitive 3D hydrogel network. As a result, the bioelectronic patches exhibit high elasticity, and structural and mechanoelectrical sensing stability throughout cardiac deformations. When implanted in a post‐MI murine model, they enabled monitoring of the MI repair process by sensing strain and relative electrical resistance changes in myocardial tissue. Simultaneously, curcumin release promoted M2 macrophage polarization, reduced inflammatory cell infiltration, enhanced angiogenesis, and restricted adverse ventricular remodeling, thereby improving cardiac function and facilitating MI repair. The system currently requires wires connecting the sensing hydrogel to a portable monitoring device which restricts its long‐term use.
The above examples represent a range of engineering approaches for fabrication of soft electrodes arrays for electrophysiological recording. While electrical mapping is beneficial and widely used for monitoring, there are limitations in terms of electrode channel count, and it only captures one aspect of cardiac activity. For example, optical mapping (utilizing calcium‐sensitive dyes or fluorophores) complements electrical mapping to offer greater insights into cellular activities, such as calcium dynamics, metabolism, and mechanical deformations.^[^
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^]^ This provides a more comprehensive window into cardiac function for deeper understanding and better diagnostics. By employing a soft poly(lactic‐co‐glycolic acid) (PLGA) substrate and a molybdenum (Mo) nanogrid structure fabricated using electron beam lithography, highly transparent multielectrode arrays (MEAs) were developed for concurrent electrophysiological and optical mapping.^[^
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^]^ The small feature size of the Mo nanogrid (500 nm width; 6.75 µm pitch) allows optical transparency, enabling mapping of parameters including biopotentials, myocardial conduction and contraction, cell metabolism, and calcium homeostasis. The device enables single‐ and multi‐site pacing for the treatment of arrhythmias and concurrent electrical and optical mapping to enable cardiac monitoring and diagnostics. This obviates the requirement of gene transfection or separate light source and external electrical‐potential electrodes used in traditional optogenetic and optoelectronic approaches. Complete bioresorption of the implanted device within 6 weeks in vivo means it circumvents surgical extraction. Instead, it is intended for use in treating and detecting complications in the days to weeks after surgery or ischemic events, which account for over one‐third of postoperative deaths.^[^
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^]^ Bench testing showed that the MEA maintained moderate impedance and charge storage capacity during resorption, ensuring reliable performance for temporary applications.
In addition to recording electrodes and multimodal sensing elements, bioelectronic cardiac patches require power management and signal processing units. Several studies have demonstrated the feasibility of developing highly stretchable electronics and electrodes on polymeric substrates. However, these systems typically require tethered configuration for energy and data logging. A promising solution to bypass complex cabled electrical connections is the development of wireless, battery‐free devices with onboard computation. In particular, using a laser machining process, highly flexible printed circuits (FPC) integrated with μ‐LEDs and Pt microelectrodes were developed for simultaneous optical stimulation and electrical recording (Figure 5D).^[^
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^]^ The device utilizes a near‐field‐communication protocol at 13.56 MHz for wireless power transfer and signal transmission to an external reading circuit. Ex vivo results show successful optical pacing and electrical recording in mouse hearts, with heart rates closely matching set values and no significant deviations post device attachment. In vivo results reveal that the device allows seamless integration with the heart, enabling long‐term recording and stimulation without major postoperative complications, as evidenced by stable animal weight postsurgery. The device supports real‐time cardiac control and multisite stimulation, enhancing optogenetic modulation capabilities. The device facilitates social interactions in freely moving animals, promoting healing and enabling behavioral studies in naturalistic settings. This demonstrates its capability to underpin fundamental research into cardiovascular diseases and treatment on animal models. A combination of i) highly stretchable electrodes utilizing either intrinsically stretchable materials (e.g., hydrogel, conductive rubber) or engineered structures (such as serpentines for metallic and semiconductor nanomembranes) and ii) flexible power management systems (NFC or Bluetooth circuits fabricated on FPC) represent viable routes for fully implantable cardiac patches.
Vascular Implantable Sensors for Hemodynamic Measurements
While peripheral blood pressure measurement using wearable devices can provide predictive clinical values in some specific contexts, in general, central blood pressure measurement is recognized as a more accurate determinant of cardiovascular status and, consequently, for detecting abnormalities in function.^[^
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^]^ Several clinical studies suggest the effectiveness of implanted/invasive mechanical sensors for hemodynamic monitoring, which can reduce the mortality rate, reduce medical costs, and improve the quality of life for patients with heart failure.^[^
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^]^ The conventional standard for hemodynamic monitoring employs long pulmonary artery catheters connected to a pressure sensor through a distal port (e.g., Swan–Ganz pulmonary artery catheter from Edwards Lifesciences). This approach enables continuous measurement of the flow and pressure levels within the pulmonary artery to support physicians in heart failure assessment. However, these tethered devices with bulky designs fail to provide continuous monitoring, requiring repetitive monitoring periods and patient immobilization. Continuous hemodynamic monitoring using wireless, miniaturized, implanted sensors has been shown to improve patient outcomes.^[^
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The CardioMEMS (Abbott, USA) is one of the pioneering technologies for wireless, implanted blood pressure measurement (Figure
6A). To our knowledge, it is the only implantable device approved by the FDA for remote hemodynamic monitoring in clinical applications.^[^
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^]^ The implanted components (≈3.5 mm wide, 2 mm thick, and 15 mm long) comprise a silicon‐based capacitive pressure sensor and an inductive coil that form an LC (inductive‐capacitance) circuit. The device is implanted into the distal pulmonary artery via right heart catheterization and then fixed onto the vessel wall using polytetrafluoroethylene‐coated nitinol loops. An external antenna powers the implantable device through inductive coupling and reads the dynamic change in pulmonary artery pressures through the coupled frequency shift. The measured pressure levels are then gathered via an electronic unit and transmitted to a secure server for processing, allowing remote monitoring of patients. The device has already shown utility in monitoring heart failure patients—helping guide medication dosage and reducing hospitalization rates.^[^
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^]^ Further, it can assist in the early detection of aortic ineffective endocarditis and has shown potential in guiding LVAD implantation by assisting prior detection of post‐implantation complications.^[^
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^]^ Despite device‐related and overall complication rates of 1% and 2.8%, respectively, lack of education by physicians and patients has seen this device being severely underutilized.^[^
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![Figure 6: Implanted mechanical sensors for hemodynamic monitoring. A) A photograph of the CardioMEMS device used for wireless blood pressure measurement. Reproduced with permission.^[^
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^]^ Copyright, 2020, Jacob Abraham et al. B) 3D printed stent integrated with pressure sensor and inductive coupling antenna. Reproduced with permission.^[^
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^]^ Copyright, 2022, The American Association for the Advancement of Science. C) Multimodal implanted sensors with a BLE module, capable of measuring blood flow rate, blood pressure, and temperature. Reproduced with permission.^[^
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^]^ Copyright, 2023, Springer Nature Limited.](ADMA-37-2420114-g023.jpg)
Another approach to implanting mechanical sensors in the pulmonary artery involves the development of pressure and flow sensors integrated into smart stents using 3D printing techniques. The key advantages of this configuration include using stents as a backbone for sensor housing (instead of suturing onto the vessel wall) along with the common use of stents in heart failure treatment/prevention, with over 3 million devices implanted annually. An example of this approach is the recent work from the Yeo group, which applies aerosol jet‐printed soft pressure and flow sensors based on silver nanoparticle‐embedded PDMS (Figure 6B).^[^
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^]^ These functional components are integrated into flexible stents based on laser‐machined stainless steel coated in subsequent layers of polyimide, gold, and parylene, respectively. Like the method used in the CardioMEMS, inductive coupling allows wireless monitoring at distances of 5.5 and 3.5 cm in air and blood, respectively. Wireless performance was validated in a pulsatile silicone artery model while minimally invasive catheter implantation was evaluated in an in vivo rabbit model via iliac and carotid access.
The inductive coupling method allows for a short range of measurement (typically less than 4 cm) and requires complex circuits or equipment (e.g., Vector Network Analyzer) to interpret the shift in the coupled frequency resulting from blood pressure change. Bluetooth Low Energy (BLE) communications, on the other hand, enable a longer distance of communication (meters) and require simpler external circuitry for power transmission.^[^
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^]^ Utilizing the BLE module, Kwon et al. developed battery‐free, multimodal sensors capable of simultaneously measuring vascular pressure, flow rate, and temperature (Figure 6C).^[^
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^]^ The implanted device includes a sensing chip built on a silicon substrate (attached to the vessel wall) connected to a BLE circuit that was encapsulated inside a biocompatible layer and sutured to the adjacent tissues of the left ventricle. A Cu‐based receiver coil was designed to receive electrical power from an external, wearable transmitter coil (connected to a battery and a charging circuit) mounted on the chest. Different from the CardioMEMS devices and integrated stent‐based sensors, the device developed employed the piezoresistive effect for both pressure and flow sensing. Specifically, monocrystalline Si nanomembranes serve as the sensing element for pressure (using a diaphragm configuration), bidirectional flow rate (using an out‐of‐plane cantilever fabricated using mechanical buckling), and blood temperature (using the thermosensitive effect). Since a variation in blood temperature can influence the baseline of the piezoresistive sensing element, integrating a temperature sensor into the implanted system can potentially support the periodic calibration of the pressure and flow sensors. Demonstrations on large animals validated the performance of the devices, capable of continuously monitoring blood pressure level (measurement range between 0 and 160 mmHg) and flow rate (measurement range of ≈6.2–6.6 L min^−1^). The small footprint of the sensing chip (3 mm × 8 mm × 2.3 mm) allows for integration with other medical devices, such as stents and valves, as demonstrated with a metallic stent. While the developed system suggests using temperature sensors for periodic calibration, other engineering approaches, such as surface charge, geometry, and wettability, are recommended to overcome blood clotting issues for long‐term use.
Management
This stage of intervention describes strategies employed to manage the symptoms that have arisen due to heart disease. Conditions such as MI and coronary artery disease impair the ability of the heart to function adequately, requiring clinical intervention. These can be lifestyle modifications and pharmacology, but in many instances, these are not effective, especially in the case of advanced heart failure. Transplantation remains the gold standard for late‐stage heart failure patients for whom management options prove inadequate.^[^
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^]^ However, the shortage of donors has necessitated the development of devices that assist the heart in achieving healthy hemodynamics and biomechanics to extend a patient's lifespan until they can receive a transplant. Some assistive technologies have been considered long‐term mechanical support devices, becoming a destination therapy rather than bridge‐to‐transplantation. These include devices that directly apply pressure to assist the ventricles during contraction (VADs), devices that provide electrical stimulation to restore function (pacemakers), implantable cardioverter defibrillators (ICDs) which revert life‐threatening rhythm disturbances, such as ventricular tachycardia and ventricular fibrillation, robotic endovascular catheter devices for minimally invasive procedures, and stents required to restore flow in atherosclerotic vessels.
Ventricular Assist Devices
Traditional VADs
VADs are the most common class of bridge‐to‐transplant devices employed in managing heart failure. They provide mechanical circulatory assistance to pump blood from one or both ventricles into the aorta and/or pulmonary artery. Traditionally, these were rigid, blood‐contacting, continuous‐flow devices implanted to support the left ventricle—known as LVADs (Figure
7A,B). These pumped the blood at high speeds, often causing hemolysis (damage to red blood cells), while the stiff foreign material contact led to thrombosis. Even third‐generation LVADs, which included electromagnetically suspended noncontacting centrifugal impellers to reduce blood contact and lysis, led to complications arising from lack of pulsatile flow and thrombosis (Figure 7C).^[^
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^]^ Ultimately, LVADs are a bridge‐to‐transplant, impractical as a long‐term solution as they render the patient reliant on anticoagulant therapy for the rest of their life due to the risk of stroke from thromboembolization.^[^
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^]^ Even the more physiologically mimetic pulsatile LVADs elicit higher rates of stroke, thromboembolic events, and device malfunction.^[^
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Strategies involving modification of contacting surfaces have been employed to reduce risks associated with conventional LVADs (Figure 7D). Passive surface coatings, such as diamond‐like carbon (DLC) and 2‐methacryloyloxyethyl phosphorylcholine (MPC) have been employed to reduce platelet activation and thrombosis.^[^
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^]^ DLC can be applied via physical or chemical vapor deposition methods.^[^
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^]^ These render LVADs more suitable for blood‐contacting applications as they form chemically inert, hydrophobic, and impermeable surfaces with low friction coefficients. MPC is a neutrally charged, hydrophobic polymer that mitigates surface protein attachment by forming clusters of water molecules. Both have been utilized in commercial EvaHeart LVADs (EvaHeart, Inc., USA). LVADs such as the HeartMate (Abbott, USA) incorporate titanium microspheres to achieve texturing of surfaces to reduce thrombus formation by providing pseudo‐biological surfaces for endothelial cell attachment and proliferation.^[^
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^]^ This also reduces the requirement of anticoagulation therapy. Heparin is the most commonly utilized anticoagulant in active surface coatings, but as it is readily cleared from the bloodstream, it must either be released slowly, i.e., via physical entrapment or ionic bonding, or be covalently bonded to the device surface.^[^
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Improvements in device design (e.g. magnetic levitation) and surface coatings have advanced this technology, but ultimately there is a need for innovation in softer materials that better resemble the biomechanics of the native heart tissue and facilitate physiologically mimetic hemodynamics. Risk of adverse thrombotic events, strict life‐long anticoagulant therapy, associated gastrointestinal bleeding, contraindications, such as allergies to anticoagulants, right ventricular failure, and anatomical abnormalities mean rigid LVADs are an inadequate long‐term strategy to support impaired cardiac function. Materials innovations in this field require the development of softer materials and devices that can minimize risks associated with thrombosis and hemolysis and provide more physiologically mimetic pulsatile flow. Another concern surrounds the percutaneous driveline used to power and control the device, whereby infections migrating along the driveline can pose a serious health risk to the patient.
Soft assistive devices that may facilitate pulsatile hemodynamics without high‐shear blood contact have emerged as a promising research endeavor. These devices are designed to wrap around the heart, prevent dilation, or provide external compression to enhance cardiac function. By only interfacing with the pericardium, these devices eliminate issues associated with high‐shear blood contact. Soft, elastic materials can stretch to accommodate for increased ventricular return or preload, store that energy, and then contract harder during subsequent systole.^[^
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^]^ This inherent ability to autoregulate changing cardiac loads without complex control algorithms mimics the natural Starling response of the heart. For single or biventricular support, this mechanism also helps to ensure the balance of stroke volume from each ventricle such that one ventricle does not become overloaded.^[^
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^]^ Right ventricular overload is a major concern for those implanted with conventional LVADs as they lack the required compliance and cannot modulate support based on changing needs.^[^
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Soft VADs
Passive soft VADs are fabricated from elastic materials and surround the pericardium in a sleeve‐like fashion. They are designed to reduce the effects of ventricular dilation characteristic of early stages of heart failure and help native myocardium remodel by reducing stress on the ventricular walls.^[^
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^]^ The HeartNet Ventricular Support System (Paracor Medical, USA) is a compliant band made from silicone‐coated nitinol wires. It conforms to the patient's heart without requiring adjustment during implantation and provides constant constrictive support through this intimate contact from diastole to systole.^[^
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^]^ The CorCap Cardiac Support Device (Acorn Cardiovascular, Inc, USA) utilizes a polyester net to restrict dilation and provides some reverse remodeling, whereby there was an improvement in cardiac function over the longer term due to the circulatory assistance offered by the device.^[^
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^]^ Others have employed polyester balloons surrounding the heart, which are inflated or deflated to adjust device compliance and therefore support level, depending on treatment needs.^[^
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While the HeartNet and CorCap devices are the most well‐studied passive soft VADs, these devices did not make it to market due to insignificant cardiac functional improvements and epicardial fibrosis (CorCap).^[^
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^]^ Since the abandonment of these devices, passive epicardial support devices have shifted into cardiac patches, which will be explored in detail in a later section. However, through the lens of mechanical support, such devices have utilized natural and synthetic acellular materials with properties similar to the native myocardium. This crucially includes anisotropic passive support. Given the native heart's anisotropic and highly ordered myocardial fiber arrangement, directional support has the potential to aid in specific functions while allowing freedom for others. Longitudinal reinforcement of MI regions of the heart has shown promise in maximizing stroke volume by improving filling during diastole. Slitted Dacron patches have been used to this effect, first demonstrated in a canine model,^[^
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^]^ and then explained with in silico models,^[^
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^]^ where it was found that longitudinal reinforcement reduces diastolic fiber stretch and systolic fiber stress in the myocardium surrounding the infarct area. Along with anisotropy, auxetic (negative Poisson's ratio) patches have been introduced for right ventricular support and monitoring.^[^
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^]^ Given the generalized cardiac motion of paired longitudinal and circumferential contraction and relaxation through the cardiac cycle, auxetic materials could be a good fit for passive support. Pirozzi's work on this utilized a polyurethane auxetic mesh fixed to an ex‐vivo porcine model using commercial tissue adhesives (Surgiseal, Adhezion Biomedical). Using this porcine model, they demonstrated increased passive ejection due to elastic recoil, while their in‐silico modeling shows promise for decreased ventricular dilation and healthy stroke volume.
In situations requiring greater therapeutic intervention, active VADs or direct cardiac compression devices can provide active contraction to assist the heart in pumping blood during systole. These devices are driven by advances in soft robotic biomedical engineering. While they do have the potential to provide pulsatile flow via epicardial pressure and match native hemodynamics, they must be optimally positioned, provide directional compressive force to match native contraction, provide co‐pulsation, and be able to account for comorbidities.
Devices, such as the HeartPatch and AdjuCor reBEAT (AdjucCor GmbH, Germany) consist of multiple inflatable pockets made from silicone and polyurethane, respectively.^[^
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^]^ The HeartPatch's inflatable silicone patches are porous, allowing biointegrated adhesion to the left and right ventricles and facilitating independent biventricular support. They leverage stiffness‐modulated actuation, where the surface adhered to the epicardium is more compliant than the outer surface, meaning pouch inflation is directed toward the epicardium. The reBEAT device (Figure
8A) has the same operating principle and houses three inflatable pouches—two for the left ventricle and one for the right—within an epicardial sleeve custom designed and fabricated to suit each patient's anatomy based on computed tomography scans to ensure intimate contact. The inner and outer surfaces are textured to facilitate long‐term biointegration. The implantation procedure has also been optimized to reduce invasiveness (lower partial sternotomy and pericardiotomy) and duration (<10 min).^[^
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^]^ This device uses continuous ECG signal detection and interpretation to identify R and T wave onset to initialize inflation/deflation of its pouches,^[^
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^]^ meaning it can react to the patient's needs. The reBEAT device has now reached its first in‐human clinical study with successful implantation.
![Figure 8: Soft Active Ventricular Assist Devices. A) The AudiCor reBEAT device consisting of inflatable balloons. A driveline connects the implanted device to an external portable drive unit. Reproduced with permission.^[^
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^]^ Copyright, 2022, The American Association for Thoracic Surgery. B) i) Soft VAD robotic sleeve based on McKibbon artificial muscle fibers arranged circumferentially and axially to allow compression and twisting motion. ii) Circumferential (top) and twisting (bottom) actuators on a porcine heart cadaver. Reproduced with permission.^[^
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^]^ Copyright, 2017, The American Association for the Advancement of Science. C) Electrothermally‐actuated artificial muscle fibers based on silver‐coated nylon. (scale 1 mm) Reproduced with permission.^[^
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^]^ Copyright, 2021, Wiley‐VCH. D) VAD sleeve based on hydraulically actuated artificial muscle fibers. ii) Arrangement of artificial muscle fibers to induce contraction and twisting motion. iii) Artificial pericardium deployed to ensure force transfer from the device to the underlying muscle. Reproduced with permission.^[^
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^]^ Copyright, 2023, Wiley‐VCH.](ADMA-37-2420114-g019.jpg)
Artificial muscles have emerged as a novel mechanism for active ventricular contraction through devices interfacing directly with the epicardium. In contrast to inflatable pouches, which offer pressure normal to the heart's surface, artificial muscles provide targeted directional strain to match and support the complex native myocardial muscle fiber motion. This is generally achieved by embedding soft, aligned contractile fibers in an elastic matrix. Several actuation mechanisms, including pneumatic, electric, thermal, and chemical, have been utilized to provide directional pressure to the ventricles in an attempt to restore cardiac output and hemodynamic parameters.^[^
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^]^ For detailed mechanics and history of direct cardiac compression devices and artificial muscles, readers are directed to the following reviews.^[^
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^]^ McKibben fibers, an extensively studied, soft, pneumatically actuated class of contractile fibers, have been utilized in a soft robotic epicardial sleeve.^[^
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^]^ The device consists of two independently controlled layers of McKibben artificial muscles embedded in silicone—one arranged circumferentially and the other axially—to provide radial pressure and induce twisting motion (Figure 8B). The combination of radial and helical motions attempts to mimic native heart muscle motion. Implementing a control system that simultaneously monitors physiological parameters allowed fine‐tuning of force generation to provide disease‐specific assistance synchronized with the native cardiac cycle.
Complications associated with the requirement of pneumatic‐power‐generation equipment, air leakage, and condensation, as well as the relatively bulky footprint of pneumatic artificial muscles, have led some to pursue electrically‐actuated fibers.^[^
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^]^ To overcome these challenges, electrothermal actuation, which is a mechanism that uses thermal expansion, generated by electrical heating, to produce motion or force in actuators, is used. For example, silver‐coated nylon fibers embedded in a silicone matrix yield an artificial muscle device that can produce pulsatile flow in a heart model (Figure 8C).^[^
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^]^ Such fibers will contract when heated due to electrical current. The twisting of the artificial muscle fibers in a helical fashion provides amplified contraction in response to thermal energy and electrical actuation. This is a significant development as a major constraint of electrothermally‐actuated soft robots is the elevated temperatures, around 65 °C, required for activation. This heating and cooling of the fibers is a slow process which inherently affects actuation speed, hence their efficacy in cardiac support. The supraphysiological actuation temperatures can also damage surrounding tissue. To mitigate temperature effects, hydraulically‐actuated artificial muscle fibers have been developed from silicone tubes wrapped in a constraining coil.^[^
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^]^ A single continuous fiber is arranged helically—for radial compression—and axially at an angle–for axial and twisting motion—around a conformal silicone hemispheroidal shell (Figure 8D). Depressurization via a small driveline connected to an external hydraulic supply provides contraction mechanics similar to native cardiac tissue. Force transduction from the device to the underlying tissue remains an important step in artificial contraction generation via such devices. In the device's current form, a dual‐layered artificial pericardial sac made of silicone and fabric, filled with water, transmits and evenly distributes pressure from the surrounding device to the tissue, avoiding local stress concentrations which may cause damage. This may assist in conforming to varying patient anatomies.
The requirement for external or internal battery sources still leaves these devices susceptible to driveline infections and cumbersome for long‐term use. While blood contact within the heart is avoided, integrating these devices into extracardiac tissue has not been well investigated, with most functional studies being done on ex vivo models. As ventricular pressure is inconsistent, advances in real‐time monitoring of filling volumes and pressures to tune applied contraction forces to maintain the Starling Law will be vital to mitigate further tissue damage. In addition, the durability of highly deformable soft materials in vivo needs to be characterized if they are to become a reliable bridge‐to‐transplant or destination treatment for heart failure.
Catheters for Endovascular Interventions
Soft/flexible robotic endovascular catheters (SRECs) have been widely used in standard and emerging endovascular interventions. They provide a minimally invasive approach to diagnosing and treating heart diseases. For example, endocardial electrophysical mapping and cardiac ablation use robotic catheters with electrodes to diagnose and treat arrhythmias.^[^
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^]^ Additionally, SRECs are being developed for hydrogel injections aimed at myocardial repair and reconstructive procedures.^[^
^106^ , ^107^
^]^ Use of SRECs can reduce the need for invasive procedures like open‐heart surgeries and cardiopulmonary bypass, which are linked to secondary organ malperfusion and injury, thereby minimizing recovery times and associated costs.^[^
^108^
^]^ However, catheter devices still have key limitations that would benefit from advances in materials engineering. Catheter surfaces may induce life‐threatening complications including catheter‐related bloodstream infections and thrombosis, and injuries associated with high friction, which result in reactive intimal proliferation or distal embolization.^[^
^109^ , ^110^ , ^111^
^]^ Platelet activation on the device surface can lead to thrombus formation, risking device malfunction, bacterial infection due to microbe anchoring on thrombi, and thromboembolic events.^[^
^112^
^]^ In addition, catheter navigation inside the dynamic and highly constrained environments of blood vessels and the beating heart presents a complex challenge.^[^
^107^ , ^113^ , ^114^
^]^ Thus, impactful material advancements lie in the fabrication of intravascular catheter materials that are antifouling and antithrombotic, provide a low Young's modulus and coefficient‐of‐friction surface for articulation with tissue, and are small and dexterous enough to navigate effectively. Additionally, actuation mechanisms to improve steering capabilities will assist devices in navigating complex intravascular environments.
Multifunctional Coatings for Improving Endovascular Catheter Function
One direction of materials innovation in catheter technology is the development of multifunctional coatings for device surfaces to prevent undesirable outcomes, such as infections and thrombosis. Hydrogel skins are made by interpenetrating hydrophilic polymers into the surface of polymer‐based devices with complex geometries.^[^
^115^
^]^ Poly‐dimethylacrylamide‐ (PDMAA) based hydrogel skins form ultrathin (10 µm) uniform layers which exhibit tissue‐like moduli (≈30 kPa) and provide low coefficient‐of‐friction surfaces for endovascular catheter articulation (Figure
9A).^[^
^111^
^]^ Application on medical device surfaces (polyvinyl chloride) resulted in a 95% reduction in bacterial (E. coli) adhesion and over 90% reduction in the formation and adhesion of blood clots to the catheters in vitro without adverse effects on blood's intrinsic clotting ability. This was verified in a porcine iliac artery bypass model where the hydrogel coating led to a 60% increase in occlusion time (time before blockage due to a clot) in the absence of systemic anticoagulants. Hydrophilic PDMAA hydrogel skins also demonstrated a 10‐fold friction coefficient reduction in soft ferromagnetic continuous robots developed for endovascular navigation.^[^
^116^
^]^ Aside from PDMAA, polyelectrolyte‐surfactant coatings—such as those based on P (SBMA‐co‐HEA‐co‐VS) (PSHV) polyelectrolyte and organosilicon quaternary ammonium surfactant (N^+^
Si)—have shown promise as multifunctional, durable catheter coatings (Figure 9B).^[^
^117^
^]^ Coatings remained stable in harsh conditions while exhibiting antibacterial, antiadhesive, and low coefficient of friction properties. Dynamic flow tests simulating real usage conditions demonstrated decreased biofilm formation on PSHV‐N^+^
Si by 99.6% compared to uncoated catheters.
![Figure 9: Advances in biocompatible skins A–C) and actuation mechanisms D,E) for catheter‐based endovascular approaches. A) Ultrathin hydrogel coatings on medical catheter surfaces mitigating bacterial adhesion and thrombosis. Reproduced with permission.^[^
^111^
^]^ Copyright, 2020, Wiley‐VCH. B) PSHV‐N^+^
Si coatings on intravascular catheters’ surface for antibacterial, antiadhesion, and low‐friction functions wit h stability. Reproduced with permission.^[^
^133^
^]^ Copyright, 2022, Wiley‐VCH. C) Heparin coatings central venous catheters’ surface with robust, antibacterial, and antithrombotic properties. Reproduced with permission.^[^
^109^
^]^ Copyright, 2024, Springer Nature Limited. D) Schematic illustration of a ferromagnetic soft continuum robotic catheter with programmed magnetic polarities, resulting from NdFeB/PDMS composite, and a hydrogel skin (left). Demonstration of navigating through a 3D tortuous and narrow phantom (right). Reproduced with permission.^[^
^125^
^]^ Copyright, 2019, The American Association for the Advancement of Science. E) i) Design and working principle of a 900‐µm‐diameter end‐effector of a hydraulically steerable catheter. ii) Cross‐section images with the components annotated. Reproduced with permission.^[^
^114^
^]^ Copyright, 2021, The American Association for the Advancement of Science.](ADMA-37-2420114-g013.jpg)
Natural polymers are advantageous as they often demonstrate enhanced biocompatibility relative to synthetic polymers and are biodegradable. Multifunctional O‐carboxymethyl chitosan hydrogel coatings were created using eco‐friendly melding electron‐beam‐induced graft polymerization.^[^
^110^
^]^ The porous polysaccharide coating yields super‐hydrophilic surfaces, which significantly reduce the adhesion of E. coli and platelets. Despite their rough morphology, these chitosan hydrogels possess low coefficients of friction and were durable through 1 h of continuous friction under wet conditions when deposited on a PAAc‐g‐PEBA substrate. While promising, the clinical viability of all these coatings is contingent upon developing simple and efficient manufacturing processes and demonstrating their long‐term stability under physiological conditions.
Most coatings to date overlook the relationship between catheter‐related bacterial infections and thrombosis. One‐step HS‐DAC coatings based on antithrombic heparin (HS) and antibacterial dimethyloctadecyl[3‐(trimethoxysilyl)propyl]ammonium (DAC) aim to address these shortcomings (Figure 9C).^[^
^109^
^]^ Coatings were stable for up to 30 days under physiological conditions on meter‐long and thin catheters. Electrolyte‐mediated dissociation in physiological environments in vitro and in vivo canine models reduced thrombus by 60% on extrinsic and intrinsic catheter surfaces. In both models, more than 97% of bacteria were killed by a contact‐killing method.
Actuation of Soft Robotic Endovascular Catheters
Although advances in coating layers can help reduce friction when inserting catheters into blood vessels, navigating a robotic catheter to reach target tissue remains challenging due to the complex architecture, limited space, and dynamics of cardiac vessels and heart chambers. The operation of these devices relies heavily on the efficacy of the steerable catheters' end‐effectors (or proximal tips). To overcome these barriers, materials innovation in actuation mechanisms of soft robotic endovascular catheters is imperative to enhance their active steerability and safety. The delicate nature of endovascular environments has motivated research into soft robotic technologies which minimize stress concentrations while fitting into small‐aperture vessels.^[^
^118^
^]^ Innovation is also driven by the requirement for miniaturization of soft robotic endovascular catheters to enable greater access.
Tendon‐driven mechanisms—actuated via cables made of materials such as PTFE‐coated stainless‐steel wires, nylon, or ultrahigh‐molecular‐weight polyethylene—are a popular conventional method utilized in numerous commercial soft robotic endovascular catheters.^[^
^119^ , ^120^ , ^121^ , ^122^
^]^ Devices based on this mechanism, such as Sensei X and Magellan (Hansen Medical Auris Health Inc., USA) and CorPath GRX (Corindus, Inc., USA), support electrophysiological measurements and low‐force interventional procedures such as radiofrequency ablation.^[^
^107^ , ^123^ , ^124^
^]^ Soft polymeric materials or coatings for the tendons can partially reduce their intrinsic friction with the outer sheaths when the catheter is being navigated inside complex configurations of the blood vessels. Due to antagonistic pairs of wires, however, they are associated with buckling,^[^
^114^
^]^ losses in distal force transmission,^[^
^107^
^]^ and are difficult to miniaturize to submillimeter diameters for more tortuous vascular structures.^[^
^125^
^]^
Wireless actuation mechanisms have gained increasing attention to address the technical limitations of previous devices. As magnetic microparticles can be scaled to submillimeter sizes, magnetic actuation has become one of the most widely used actuation methods for robotic catheters. It offers wireless control of catheters with real‐time feedback and low friction. In addition, it may allow more straightforward fabrication of robots with greater control over their movement. For instance, self‐lubricating, submillimeter, soft continuum robots were fabricated via extrusion printing and injection molding, followed by magnetization of composite inks containing hard neodymium‐iron‐boron (NdFeB) ferromagnetic microparticles suspended in either PDMS or thermoplastic polyurethane matrices. Robots are endowed with omnidirectional steering and navigation capabilities via magnetic actuation (Figure 9D).^[^
^125^
^]^ Subsequent optimizations have further refined this technology by integrating theoretical modeling and the genetic algorithm to increase the workspace (from 0.13 to 0.27 of normalized half workspace). This improvement affords catheters the ability to traverse highly constrained environments of narrow and tortuous vasculature.^[^
^126^
^]^ As previously described, the integration of PDMAA hydrogel skins significantly reduced the coefficient of friction, thereby significantly improving device insertion and navigation.^[^
^115^
^]^ The use of this advanced magnetic composite in robotic catheters has found application across numerous studies, such as in an electromagnetically controllable microrobotic interventional system for targeted real‐time cardiovascular interventions^[^
^127^
^]^ and in ferromagnetic fiber robots for navigation, sensing, and modulation purposes.^[^
^104^
^]^
The introduction of variable stiffness within soft robotic endovascular catheters increases device dexterity and functionality required for puncturing or grasping tissue.^[^
^128^ , ^129^
^]^ This can be achieved by adding a nitinol core (80 µm) to existing devices to achieve regions of varying stiffness, but this approach comes at the cost of increased diameter. Thermoset shape memory polymers may offer a promising alternative while maintaining submillimeter dimensions. Soft robotic endovascular catheters based on composite materials of shape‐memory polymer NOA86H and ferromagnetic NdFeB microparticles (5 µm) could be selectively stiffened or softened based on the glass‐transition controlled via Joule heating of electrical contacts at either end.^[^
^130^
^]^ However, magnetic steering has been limited to low magnetic fields, preventing integration into medical systems operating at ultrahigh fields, such as magnetic resonance imaging (MRI) scanners. To address this, hard neodymium permanent magnets have been utilized to present magnetic guidewire design and demonstrate unique magnetic actuation opportunities at ultrahigh fields, such as in situ magnetization.^[^
^131^
^]^ While magnetic actuation affords the miniaturization of robotic catheter diameters and improves distal force generation, its implementation necessitates costly and intricate actuation systems for generating external magnetic fields, thus leading to potential complications.^[^
^107^ , ^132^
^]^
The complexity of soft robotic endovascular catheter actuation mechanisms is often associated with high costs. In an attempt to mitigate this limitation, a low‐cost fluidic‐driven actuation method was proposed to develop a millimeter‐scale soft robotic platform capable of deployment and self‐stabilization at the entrance to the heart and guide conventional interventional instruments towards target sites.^[^
^107^
^]^ The steerable tip of the apparatus was made of three stacks, each housing 20 balloons constructed from a 2D laminate of thermoplastic elastomer and PTFE, encasing 1‐mm‐diameter fluidic tubing. This configuration allows the steering tip to have sufficient maneuverability within the confines of the right atrium while generating high bending (up to ≈0.8 N) and extension (up to ≈1.8 N) forces when actuated at a modest pressure of 120 kPa. This device has been demonstrated to be suitable for coronary sinus cannulation and tricuspid valve annulus puncture. However, given the bulky structure of the robotic arm and utilization of large‐scale materials, the device is limited only to procedures in the right atrium, precluding its suitability for other procedures within narrower vasculature. By contrast, fluidic‐driven soft robotic microcatheters have a submillimeter diameter (900 µm), allowing smooth navigation within diminutive blood vessels (Figure 9E).^[^
^132^
^]^ The steerable tip of this robotic catheter was fabricated through a complicated injection molding process, necessitated by its submillimeter size, using Smooth‐Cast 327 for mold construction. Subsequently, the tip was coated with relatively stiff Dragon Skin 10 SLOW for both outer and inner layers, while the vacant regions were filled with Dragon Skin 10 SLOW mixed with Hexane in a 1 ratio. This device indicates the potential for microhydraulic soft robotic catheters in mitigating challenges associated with access, navigation, and treatment problems in endovascular interventions.
Cardiac Pacemakers
Cardiac pacemakers help to improve symptoms caused by arrhythmias and enable patients to have a more active lifestyle. They are small, implantable, battery‐powered devices that treat certain arrhythmias (slow or irregular heart rhythm) and heart failure. Pacemakers deliver low‐energy electrical pulses to control the rate and rhythm of the heartbeat. A conventional pacemaker is a transvenous device consisting of a subcutaneous pulse generator and battery module to create electrical pulses and one or more endocardial leads to carry the pulses to the heart's chambers (Figure
10A).^[^
^134^
^]^ Depending on the number of leads, a pacemaker can provide different pacing modalities such as single‐chamber, dual‐chamber, and biventricular (also called cardiac resynchronization therapy). Modern pacemakers can sense the heart's natural rhythm and pace it, if necessary, at a programmable baseline heart rate.^[^
^135^
^]^
The use of pacemakers along with implantation surgery may result in substantial complications, including device‐related infection, bleeding, deep vein thrombosis, and damage to blood vessels or nerves.^[^
^136^
^]^ Device‐related infection is relatively common (4.4%) and continues to be associated with high morbidity and mortality.^[^
^137^
^]^ Endocardial leads can disturb the movement of the tricuspid valve leaflets, causing tricuspid valve regurgitation.^[^
^138^
^]^ Another possible adverse event in the case of long‐term right ventricle pacing is pacing‐induced cardiomyopathy, which is a clinical syndrome defined by a reduction in left ventricular ejection fraction.^[^
^139^
^]^
Leadless pacemakers have been developed as an obvious technological advance to eliminate lead‐related complications.^[^
^140^
^]^ The most common leadless pacemaker architecture is a small capsule made of titanium encapsulating the pulse generator, battery, and electrodes for pacing and sensing (Figure 10B). A steerable catheter is required to deliver a leadless pacemaker into the right ventricle through the femoral vein, thus eliminating implantation surgery and its associated risks.^[^
^141^
^]^ The Micra transcatheter pacing system (Medtronic, USA) is a commercially dominant leadless pacemaker which received FDA approval in 2016. The Micra TPS has four nitinol tines as a fixation mechanism to the myocardium.^[^
^142^
^]^ Long‐term implantation of the Micra TPS showed a low rate of major complications (4.5%) through 60 months of follow‐up.^[^
^143^
^]^ In a systematic review and meta‐analysis of pooled observational data, leadless pacemakers were found to have a low incidence of complications and good electrical performance up to 1 year after implantation.^[^
^144^
^]^ In a retrospective review study of 198 pacemaker‐dependent patients over 5 years, the incidence of pacemaker‐induced cardiomyopathy was significantly lower with leadless pacemakers compared with transvenous pacemakers ‐ likely as a result of the lack of a subcutaneous pulse generator pocket and transvenous leads.^[^
^145^
^]^
Early generations of leadless pacemakers could provide only single‐chamber pacing of the right ventricle, thereby limiting their capacity for common cardiac pacing indications where atrial pacing was required. To address this limitation, a dual‐chamber leadless pacemaker system with one pacemaker implanted in the right ventricle and one in the right atrium was developed. Wireless bidirectional communication between the two leadless pacemakers was established to achieve synchronous, dual‐chamber pacing with a success rate of 99.2%.^[^
^146^ , ^147^
^]^ The proposed dual‐chamber leadless pacemaker system has the potential to treat a wider range of indications that require atrial pacing or consistent atrioventricular synchrony.
Current pacemakers are operated via batteries which have a limited lifetime, thus requiring the pulse generator to be replaced every ≈10 years and increasing the risk of procedure‐related complications.^[^
^135^
^]^ Self‐powered cardiac pacemakers using piezoelectric energy harvesters have been developed to eliminate the risks associated with battery replacement. Piezoelectric energy harvesters transform mechanical energy from the movements of internal organs, such as the heart, diaphragm, and lungs into electrical energy to power a pacemaker.^[^
^148^
^]^ An example of such materials advances are biocompatible, implantable, and flexible polymer‐based piezoelectric nanogenerator materials that utilize energy from the left ventricular motion to drive a self‐powered pacemaker.^[^
^149^
^]^ A piezoelectric nanogenerator was made from composite nanofibers of poly(vinylidene fluoride) and a hybrid nanofiller of zinc oxide and reduced graphene oxide. The device harvested 0.487 µJ from every heartbeat in a canine model. Liu et al. developed a self‐powered intracardiac pacemaker that harvests biomechanical energy from cardiac motion to provide endocardial pacing function following catheter‐based delivery in a porcine model.^[^
^150^
^]^ The device (1.75 g) utilized the back‐and‐forth rolling motion of polyformaldehyde pellets between gold electrodes deposited on a polytetrafluoroethylene film under heart beating to generate alternating current based on contact electrification and electrostatic induction. Another study introduced a high‐performance inertia‐driven triboelectric nanogenerator based on body motion and gravity.^[^
^151^
^]^ The device was made from amine‐functionalized poly(vinyl alcohol) and perfluoroalkoxy as triboelectric materials and a freestanding copper mass to harvest the inertial movement. The nanogenerator was implemented to create a proof‐of‐concept self‐rechargeable cardiac pacemaker system. Similarly, an implantable triboelectric nanogenerator harvested energy from a rat's breathing to drive a pacemaker prototype to regulate heart rate.^[^
^152^
^]^
Besides intracardiac pacemakers, epicardial pacemakers have been developed where electrodes are attached to the heart's surface rather than inside chambers (Figure 10C). A leadless, bioresorbable, battery‐free implantable epicardial pacemaker was developed for postoperative control of cardiac rate and rhythm.^[^
^153^
^]^ The device consisted of a wireless receiver connected to a contact pad to interface with myocardial tissue through a strip of bioresorbable tungsten‐coated magnesium electrodes. The wireless receiver, also made from bioresorbable materials, served as a power harvester and control interface. The entire device was encapsulated in bioresorbable poly(lactide‐co‐glycolide). The system provided effective pacing of murine, rabbit, and canine hearts followed by complete resorption within 3 months, providing a good basis for use in human cardiac patients. Conductive cardiac patches have also been developed based on trinity triboelectric nanogenerators to sense and repair infarcted myocardium via simultaneous triboelectric charge generation and electrical stimulation.^[^
^154^
^]^ The patches consisted of Ecoflex film, reduced graphene oxide (rGO) electrode, polyvinylidene fluoride film, and polydopamine modified rGO electrode, capable of harvesting and converting biomechanical energy into electrical energy for therapeutic stimulation. In addition, the electrical signals were wirelessly transmitted to a smartphone for remote monitoring and diagnosis. In another approach, an implantable cardiac mesh was developed for continuous electrophysiological monitoring, and electrical and thermal stimulation in an ischemic swine heart.^[^
^155^
^]^ The mesh was constructed from a highly conductive, stretchable network of gold‐coated silver nanowires dispersed in poly(styrene‐butadiene‐styrene) elastomer.
Most clinically used cardiac implants risk causing tissue damage and bleeding due to reliance on surgical suturing, glues, or direct insertion of electrodes into cardiac tissue. To address this, a bioadhesive pacing lead was developed which can directly adhere to cardiac tissue via physical and covalent cross‐linking to provide continuous and reliable monitoring and pacing.^[^
^156^
^]^ The device was fabricated via a multimaterial 3D printing process, which included sequential deposition of a nonconductive bioadhesive ink based on hydrophilic polyurethane modified with polyacrylic acid and a conductive ink based on PEDOT:PSS mixed with the nonconductive ink. The polyurethane was also modified with a monomer, which enabled on‐demand, gentle detachment of the device. The system's efficacy was demonstrated in rodent and porcine hearts for 2 weeks. In a similar vein, a soft bioelectronic patch was created that could adhere instantaneously and conformably to heart tissue for continuous cardiac monitoring.^[^
^157^
^]^ The patch was constructed from 3 an ionically conductive tissue‐adhesive hydrogel, a self‐healing conductive composite based on eutectic gallium–indium liquid metal, and an electrospun self‐healing polymer fiber network structure. Together, these components enable robust conformal tissue adhesion owing to efficient strain energy dissipation while maintaining effective electrical performance for long‐term sensing and stimulation. Simultaneous ECG recording and pacing capabilities were demonstrated in a murine model.
Coronary Artery Stents
Coronary artery disease is a common heart disease in which coronary arteries struggle to supply enough oxygen‐rich blood to the heart muscle due to atherosclerosis—arterial narrowing caused by gradual plaque buildup—which can partially or completely block blood flow. Symptoms of coronary artery disease include angina (chest pain), shortness of breath, and even MI if complete blockage occurs. Depending on the disease progression, treatments may include lifestyle changes, medications, percutaneous coronary intervention, and coronary artery bypass graft surgery. Percutaneous coronary intervention is a two‐step, minimally invasive procedure involving coronary angioplasty and stent placement. Coronary angioplasty is performed to widen a clogged artery to improve blood flow using a balloon catheter inserted into the heart and coronary arteries through the femoral or radial artery. Stent placement involves the insertion of a small, expandable wire mesh tube into the artery to prevent it from closing. The stent implantation after coronary angioplasty seals dissected tissues and prevents arterial recoil, resulting in better treatment outcomes than angioplasty alone.^[^
^158^
^]^
Bare metal stents were among the earliest stents developed and are made from a wide range of metals and alloys, including stainless steel, cobalt, chromium, cobalt‐chromium alloys, and platinum‐chromium alloys.^[^
^158^ , ^159^
^]^ Material advances for bare metal stents aim to achieve high flexibility and thin struts while maintaining good radial strength and radio‐opacity. However, the critical disadvantage of bare metal stents is the high rate of intracoronary stent restenosis, at around 15% of cases after a follow‐up to 6 months.^[^
^160^
^]^ Drug‐eluting stents—bare metal stents coated with drug‐eluting polymers—have been developed to decrease restenosis rates by impeding neointimal hyperplasia. Polymers such as polyethene‐co‐vinyl acetate (PEVA), poly‐n‐butyl methacrylate (PBMA), polylactic acid (PLA), and polyvinyl pyrrolidone (PVP) are used as drug carriers that allow sustained, controlled release of antiproliferative drugs, such as sirolimus and paclitaxel. However, while drugs inhibit smooth muscle proliferation to reduce intimal hyperplasia, they also affect stent re‐endothelialization, thus lead to increased rates of stent thrombosis in long‐term implantation. New drugs, including zotarolimus and everolimus, were shown to result in good endothelial coverage in animal studies and are used in commercial stents, though a series of clinical trials show mixed results in terms of adverse events.^[^
^158^
^]^
While drug‐eluting polymers show good control over drug release duration, polymers may cause inflammation, slow vascular healing, and contribute to stent thrombosis. Polymer‐free drug‐eluting stents are produced by directly coating the drug onto the stent surface by chemical bonding, thus eliminating polymers and their potential risks. The stent surface may be treated to have a microporous structure to improve drug elution duration.^[^
^159^
^]^ The BioFreedom stent (Biosensors, Singapore) combines a microstructured abluminal surface and a highly lipophilic antirestenotic drug (Biolimus A9) used for patients with high bleeding risk. The BioFreedom stent showed better outcomes compared to a bare metal stent after 1 month of dual antiplatelet therapy.^[^
^161^
^]^
The long‐term presence of stents in coronary arteries poses potential complications, such as neo‐atherosclerosis, chronic inflammation, and stent thrombosis. Therefore, biodegradable polymer and metal stents have been developed to provide initial support to prevent arterial recoil until vascular recovery, and then they gradually degrade to minimize the rate of stent thrombosis. Biodegradable polymer stents made from materials, such as PLLA, PLGA, and poly(glycolic acid) (PGA) generally have better biocompatibility than those made from biodegradable metals, such as magnesium, iron, or zinc but have lower mechanical strength, requiring thicker struts to achieve sufficient radial support. Biodegradable polymer stent development also benefits from a wide range of biodegradable polymer options and well‐established manufacturing modalities, including injection molding, laser cutting, and 3D printing.^[^
^162^
^]^ The DESolve scaffold system (Elixir Medical, USA), a PLLA‐based bioresorbable stent coated with novolimus, has demonstrated safety and efficacy at a 2‐year follow‐up.^[^
^163^
^]^ The DREAM 2G (or Magmaris) scaffold (Biotronik AG, Switzerland) is a magnesium‐based scaffold coated with PLLA and sirolimus. Clinical trial (BIOSOLVE‐II) of the DREAM 2G showed favorable safety outcomes at 3 years.^[^
^164^
^]^ In some pioneering approaches, biosensors have been incorporated into stents to obtain real‐time vascular conditions.^[^
^165^
^]^ An example is the implantable, wireless vascular electronic system described in Section 2.2 consisting of an inductive stent and printed soft sensors that can monitor arterial pressure, heart rate, and flow in real‐time.^[^
^166^
^]^
Repair
Once irreversible damage to cardiac tissue occurs, treatment options become limited. In such cases, devices designed for tissue repair, replacement, or regeneration are necessary. Examples include replacement valves, commonly used to treat heart valve disease, synthetic vessels for coronary bypass surgery, and various scaffold patches used in cardiovascular repair procedures. Tissue engineering plays a crucial role in long‐term cardiovascular repair and regeneration.^[^
^167^ , ^168^ , ^169^
^]^ It aims to restore native tissue function by either growing functional human tissue in the laboratory or developing biomaterials that guide the body's natural repair and regeneration processes while the biomaterial gradually degrades.
Myocardial Tissue Engineering
MI causes irreversible loss of cardiomyocytes, scar formation, thinning and weakening of ventricular walls, and abnormal rhythm. MI is generally triggered by prolonged ischemia due to occlusion of a coronary artery. MI leads to mass necrosis of cardiomyocytes, triggering the generation of reactive oxygen species (ROS), and the release of intracellular components.^[^
^170^ , ^171^
^]^ These events trigger proinflammatory cytokines and recruit immune cells. This destructive milieu impairs cellular function via calcium homeostasis disruption, contributes to further cell apoptosis, and aggravates oxidative stress and inflammation. Eventually, this results in the formation of a fibrotic area and adverse ventricular remodeling, wherein the ventricular wall becomes dilated and unable to meet the hemodynamic demands of the body. Current treatments, such as pharmacological therapies or reperfusion surgeries, only address the symptoms with little effect on healthy tissue regeneration.^[^
^172^
^]^ Pharmacological treatment strategies include pain management, antithrombotic therapy, reduction of oxygen demand via nitroglycerin and β‐blockers, lowering LDL cholesterol, and attempts to improve myocardial remodeling via ACE inhibitors and aldosterone antagonists.^[^
^171^ , ^173^
^]^ Clinically, coronary reperfusion surgery—which often involves catheter‐assisted restoration of blood flow in the occluded artery—is used. However, myocardial reperfusion injury introduces additional damage while preserving the infarcted myocardial tissue with its destructive milieu.^[^
^174^ , ^175^
^]^ As adult cardiac tissue possesses limited regeneration capabilities (≈1% per year in adults, decreasing to 0.45% by the age of 75), cellular therapies emerged as a promising approach.^[^
^176^
^]^ However, issues associated with challenging delivery, poor integration, and poor cell stability and survival in the hostile myocardial environment have necessitated the integration of biomaterial‐assisted tissue engineering approaches.^[^
^177^
^]^ To our knowledge, bone marrow‐derived stem cells, myoblasts, cardiac progenitor cells, and adipose‐derived stem cells have been used in clinical trials, with mixed results. Cardiac tissue engineering is a multidisciplinary approach combining biomaterials, cells, and bioactive molecules to repair and regenerate heart tissue and consequently restore proper function. Various natural (decellularized ECM, fibrin, gelatin, cellulose, alginate, chitosan, collagen, hyaluronic acid, silk)^[^
^178^ , ^179^ , ^180^
^]^ and artificial (PGA, PLA, PLGA, PU) hydrogel biomaterials are explored as both delivery platforms for therapeutic biochemicals (such as growth factors and extracellular vesicles) and as support matrices for cell proliferation and/or differentiation in the formation of cardiac tissue in vitro. Currently, the most common biomaterial platforms for cell and biochemical delivery include preformed scaffolds or injectable hydrogels.^[^
^171^
^]^
Biomaterials for cardiac tissue regeneration must be biocompatible, biodegradable, support cell infiltration and integration with native tissue, match native mechanical properties, and provide appropriate electrical conductivity. An electroconductive environment can facilitate intercellular communication to promote synchronous cardiomyocyte contraction and anisotropic electrical wave propagation.^[^
^181^ , ^182^ , ^183^
^]^ This is often done via integration of materials, such as electrically conductive polymers (PANi, polypyrrole, polythiophene) or nanomaterials (CNTs, gold, or silver nanoparticles and nanowires).^[^
^62^ , ^179^ , ^182^ , ^184^ , ^185^
^]^ Further, it is desirable that scaffolds for myocardial tissue engineering mitigate the hostility of the post‐MI environment, while enabling controlled release of any bioactive factors.^[^
^186^
^]^ As the myocardium experiences nonuniform 3D deformations, suitable biomaterials must be able to withstand appropriate tensile, compression, and shear stresses.
Cardiac Patches
Application of a cardiac patch to the infarcted ventricular surface to provide local mechanical support and stimulate regeneration via cell or biomolecule delivery has been extensively explored for post‐MI repair. Cardiac patch delivery to the myocardial surface is based on two main 1) delivery of a preformed patch, and 2) in situ gelation of injected biomaterials.^[^
^187^
^]^ Preformed patches are generally adhered to the surface via electrostatic interactions, surgical sutures, or adhesives, whereas injectable hydrogels can be deposited into the pericardial cavity.^[^
^188^
^]^ Ideal biomaterials should adhere to the heart surface, possess mechanical strength to provide structural support but be elastic enough to integrate with the host tissue throughout movement. Further, biomaterials must possess appropriate electrical conductivity to promote synchronous cardiomyocyte contraction. Biomaterials’ degradation rates must complement new tissue formation to provide adequate support while avoiding adverse effects associated with long‐term presence. Decellularized cardiac ECM has been a popular choice of matrix for cardiac patches as it contains a complex composition and architecture replete with proteins and proteoglycans required to provide adequate mechanical support and signaling to cells.^[^
^189^ , ^190^ , ^191^ , ^192^
^]^ In fact, an injectable hydrogel based on decellularized porcine myocardial ECM—VentriGel—was recently tested in first‐in‐man clinical trials in post‐MI patients.^[^
^193^
^]^ Improved left ventricular remodeling was observed in patients who were more than 1‐year post‐MI. Delivery of decellularized ECM was also recently demonstrated by painting directly on the surface of a beating rat heart. Decellularized porcine cardiac ECM was combined with tyramine‐modified hyaluronic acid to obtain a viscous hydrogel which could be painted on the cardiac surface for in situ patch formation (Figure
11A).^[^
^192^
^]^ Increased wall thickness, reduced fibrous area, and angiogenesis were observed in the murine MI model after 28 days. Although in situ patch formation via painting is a scalable approach in terms of size, it requires the cardiac surface to be exposed. This means highly invasive open‐heart surgery is still necessary for human administration. Meanwhile, issues associated with autologous tissue scarcity, host responses, and pathogen transmission have prompted innovation in cardiac patches from other natural or synthetic materials.
![Figure 11: Advances in cardiac patch biomaterials for myocardial tissue engineering. A) Paintable hydrogel based on decellularized porcine cardiac ECM combined with hyaluronic acid modified with tyramine for the formation of cardiac patches in situ. Reproduced with permission.^[^
^192^
^]^ Copyright, 2024, Wiley‐VCH. B) i) Visualization of the fiber orientation and changes in angle (left) in the left ventricle from the epicardium to the endocardium. iii) 3D stretchable architecture provides dynamic stretchability with the deforming heart throughout systole and diastole. Reproduced with permission.^[^
^200^
^]^ Copyright, 2020, The American Association for the Advancement of Science. C) Intrapericardial delivery of biomaterial‐assisted exosome delivery for MI repair in a mouse model. Reproduced with permission.^[^
^188^
^]^ Copyright, 2021, Dashuai Zhu et al. D) Curcumin nanoparticle‐loaded gelatin nanoparticles (GelB‐Cur NPs) and recombinant human collagen III (rhCol III) are loaded into a carboxymethyl chitosan and oxydextrin scaffold for rapid curcumin and sustained rhCol III release to treat MI. Reproduced with permission.^[^
^201^
^]^ Copyright, 2023, Elsevier B.V.](ADMA-37-2420114-g012.jpg)
Electrospinning is a popular fabrication method for cardiac patches as it allows tuning of porosity via control of fiber thickness and deposition density. The highly porous nature and high specific surface area of electrospun scaffolds assist cell infiltration and attachment, in addition to endowing them with enhanced deformability.^[^
^194^ , ^195^
^]^ Further, fiber alignment can be controlled to enable cellular alignment to match native myocardial architecture.^[^
^190^ , ^196^
^]^ Another key advantage of electrospinning is its compatibility with a wide range of materials, allowing integration of enhanced functionality for improved therapeutic outcomes. For example, nitrous oxide‐ (NO*─*) functionalized PCL was processed into a fibrous electrospun mat for the formation of acellular cardiac patches capable of local delivery of therapeutic NO.^[^
^197^
^]^ Implantation in rat and porcine MI models showed improvements in cardiac function, reduced infarct size, and attenuated adverse remodeling. Electrospinning was also utilized to fabricate piezoelectric, mechanoresponsive cardiac patches based on poly(vinylidene fluoride)‐trifluoroethylene fibers.^[^
^198^
^]^ The scaffolds supported the differentiation and maturation of seeded primary and induced pluripotent stem cell‐ (iPSC) derived cardiomyocytes in vitro. It was hypothesized that the piezoelectric material generated a voltage in response to cardiomyocyte‐induced deformation, further stimulating cardiomyocyte maturation. Similarly, conductive biodegradable nanofiber scaffolds were fabricated based on a conductive polypyrrole‐chitosan‐collagen composite.^[^
^196^
^]^ However, functionality was only tested in vitro with fibroblast cells. Electrospinning was also utilized for the formation of layered cardiac patches based on silk fibroin and PCL mats containing carbon nanotubes (CNTs), and brown adipose derived stem cell sheets.^[^
^199^
^]^ The CNTs enabled high conductivity which enhanced cardiac remodeling capabilities. In a murine post‐MI model, the cardiac patches demonstrated enhanced myocardial remodeling (including cardiomyocyte gap junction repair), reduction in fibrotic tissue, and neovascularization.
3D printing represents another versatile and scalable method for fabricating scaffolds with controllable architectures.^[^
^202^
^]^ It can allow reproducible fabrication of scaffolds with spatiotemporal control of cells and biomaterials.^[^
^191^
^]^ Multilayer deposition printing of hollow microchannel filaments based on a combination of PCL, poly(glycerol sebacate), and gelatin resulted in flexible patches with hierarchical microporous networks.^[^
^203^
^]^ The device is administered via thoracoscopy and fixed to the pericardium with barbed needles. The interconnected porosity and soft construction materials enable stable contact throughout beating without generating local stress concentrations while simultaneously enabling cell infiltration for improved integration with host tissue and angiogenesis. In addition, reduced infarct size in a murine post‐MI model was observed after 28 days. Recently, beam scanning stereolithography printing was utilized for the microfabrication of cardiac patches based on GelMA and PEGDA, which contained a highly stretchable mesh‐like microfiber structure that mimicked the anisotropic myocardial fiber orientation (Figure 11B,i)).^[^
^200^
^]^ The microstructure reversibly changed shape between mesh and wavy pattern in accordance with diastole and systole to maintain conformal contact with the changing curvilinear heart surface (Figure 11B,ii)).
Delivery of Therapeutic Cells and Biomolecules
Poor retention rates of therapeutic cells and biochemicals for endogenous repair can be improved via biomaterial‐assisted delivery. It is desirable to seed scaffolds for cardiac regeneration with cells (i.e., cardiomyocytes, cardiac stromal cells (CSCs), mesenchymal stem cells (MSCs), iPSC‐derived cardiomyocytes or cardiac progenitor cells) as they secrete paracrine factors and directly interact with host cells to improve therapeutic outcomes.^[^
^204^ , ^205^ , ^206^ , ^207^
^]^ While endogenous repair factors, such as growth factors and extracellular vesicles have demonstrated therapeutic potential, without biomaterial‐assisted delivery, they are unstable in vivo, exhibit low retention rates, and produce off‐target effects.^[^
^199^ , ^208^ , ^209^ , ^210^
^]^ Promoting angiogenesis in the post‐MI myocardium is believed to be an effective therapeutic strategy. Administration of proangiogenic factors, such as vascular endothelial growth factor (VEGF), fibroblast growth factor (FGF), and hepatocyte growth factor (HGF), has shown significant improvements in blood vessel formation and myocardial function.^[^
^178^
^]^ Extracellular vesicles are cell‐secreted nanoparticles that mediate intercellular communication through the delivery of bioactive cargo consisting of components including proteins, lipids, and genetic material (i.e., microRNAs).^[^
^211^ , ^212^ , ^213^
^]^ They play key roles in regulating healthy cardiac function, including tissue repair. Exogenous delivery to injured hearts in small and large animal models has yielded improved cardiac function.^[^
^214^ , ^215^
^]^
Incorporation of iPSC‐derived cardiac progenitor cells and MSC‐derived exosomes into injectable hydrogels based on decellularized porcine ECM has shown promise for the formation of a cardiac patch in situ (Figure 11C).^[^
^188^
^]^ Cardiac progenitor cells differentiated into cardiomyocytes, smooth muscle cells, and endothelial cells. Coupled with the associated paracrine effect, this promoted angiogenesis and reduced infarct size. MSC‐derived exosomes were delivered in a sustainable manner via encapsulation in methacrylated‐hyaluronic acid followed by incorporation into the porcine ECM injectable hydrogel (Figure 11C). In an alternative approach, therapeutic payloads of cardiac stromal cell‐secreted factors were encapsulated in biodegradable PLGA microparticles for sustained release.^[^
^189^
^]^ The microparticles were named “synthetic cardiac stromal cells” as they provide a more shelf‐stable proxy for live cells. Microparticles were embedded in a decellularized ECM matrix, providing mechanical support and mimetic structure and composition for easier integration with host tissue.
The post‐MI tissue microenvironment exhibits high levels of oxidative stress. Tissue engineering strategies for myocardial repair aim to mitigate this by incorporating ROS‐scavenging materials into the hydrogel matrix. Self‐healing, injectable recombinant human collagen III hydrogels containing natural curcumin nanoparticles have shown efficacy in reducing ROS levels owing to the antioxidant and anti‐inflammatory capabilities of the nanoparticles (Figure 11D).^[^
^201^ , ^216^
^]^ Further, since collagen III is the most abundant ECM protein in the myocardium, its use as the scaffold enhanced cell migration, proliferation, and angiogenesis in vitro and in vivo in a murine model. In another use of natural antioxidant materials, melanin nanoparticles derived from cuttle‐fish ink incorporated into seaweed‐derived alginate hydrogels were shown to regulate oxidative stress via ROS‐scavenging in vivo in a murine model.^[^
^180^
^]^ The composite inhibits cardiomyocyte apoptosis, downregulates proinflammatory M1 macrophages while upregulating regenerative M2 macrophages, and promotes angiogenesis while being biodegradable.
The desire to localize the release of payloads to only the target tissue has prompted the development of microneedle‐based cardiac patches that can integrate with the tissue without causing adverse injury. Polymeric microneedle patches based on micromolded poly(vinyl alcohol) microneedles and cardiac stromal cells encapsulated in a fibrin gel have shown promise in this regard.^[^
^217^
^]^ Microneedles act as communication channels between the cells and the myocardium, allowing delivery of paracrine factors and uptake of nutrients. Validation in vivo in a post‐MI rat model showed attenuation of adverse remodeling, promotion of cardiac repair, and improved left ventricular ejection fraction after 3 weeks. Similarly, microneedle‐patches based on gelatin methacryloyl have been utilized to load and sustainably release galunisertib, a transforming growth factor‐beta (TGF‐β) inhibitor for over 2 weeks.^[^
^218^
^]^
There is growing emphasis on reducing the invasive nature of cardiac patch delivery. In parallel to injectable biomaterials for intrapericardial injection followed by in situ formation of a mechanically stable, biodegradable structure, development of highly deformable patches which can be crimped and delivered via catheters aims to meet this requirement. Shape‐memory cardiac patches for post‐infarction repair based on highly conductive and injectable elastin‐gelatin‐carbon nanotubes have been delivered via catheter‐based thoracoscopy in vivo in murine and porcine models.^[^
^219^
^]^ The patches can be crimped to allow minimally invasive delivery and deployed at the target site via temperature‐mediated shape recovery. The porous hierarchical structure of these scaffolds—achieved by ice templating—allows cell infiltration and growth. Pre‐seeding with cardiomyocytes accelerated repair. Interwoven carbon nanotubes endow conductivity, which supports cardiomyocyte growth and action potential firing. Similarly, biodegradable poly(octamethylene maleate (anhydride) citrate) patches have been deployed via injection through orifices as small as 1 mm.^[^
^220^
^]^ Patches show full shape recovery without affecting the viability and function of the cardiomyocyte in the patch.
Valvular Tissue Engineering
Heart valves are living, dynamic connective tissue structures that continuously open and close to ensure sufficient and unidirectional blood flow from one heart chamber to the next.^[^
^221^
^]^ Their tissue‐level mechanics are highly anisotropic and reflect the bidirectional organization of the ECM, with circumferentially aligned collagen for strength and stiffness and radially aligned elastin for flexibility and recoil.^[^
^222^
^]^ Disruption of this aligned fibrous organization and composition leads to valvular heart disease, the third leading contributor to CVD worldwide.^[^
^223^
^]^ Diseased valve leaflets lose their mechanical integrity and pliability, leading to outflow obstruction (stenosis) or backflow (regurgitation).^[^
^224^
^]^ While heart valve replacement is often the only solution for severe valvular disease, existing mechanical or biological valve prostheses are associated with notable limitations. Mechanical valves offer excellent long‐term durability but are prone to infection, inflammation, pannus formation, and leakage. Their non‐physiological geometries and rigid materials cause thrombosis, necessitating lifelong anticoagulation therapy and increasing risk of bleeding complications and stroke.^[^
^225^
^]^ Biological valves—derived from glutaraldehyde‐fixed allogenic or xenogeneic valve roots or pericardial tissue—emerged to address thrombogenicity issues associated with mechanical valves, but they are susceptible to immunogenicity and calcification‐induced structural degeneration,^[^
^226^
^]^ especially in younger patients.^[^
^227^
^]^ This leads to leaflet thickening and stiffening, and subsequently, insufficient valve closure and leakage, requiring reoperation after 10‐15 years.^[^
^228^ , ^229^
^]^ Over the past 20 years, a paradigm shift in heart valve replacement has seen the emergence of transcatheter valve replacement (TVR), wherein a crimped biological valve is deployed via a minimally invasive catheter‐based procedure instead of conventional open‐heart surgery.^[^
^169^ , ^230^
^]^ However, current TVR‐compatible leaflets are still based on glutaraldehyde‐fixed xenogeneic and allogeneic materials, inheriting the limitations of surgical bioprosthetic valves.^[^
^231^ , ^232^
^]^ Additionally, there are concerns regarding accelerated structural valve degradation due to the stresses induced during crimping and deployment.^[^
^233^
^]^ The need for alternative valve materials is driving innovations in materials that are strong, durable, and resist calcification, as well as those that guide native tissue repair.
Nonregenerative Polymeric Valves
Early studies explored synthetic polymers, including silicone,^[^
^234^
^]^ PTFE^[^
^235^ , ^236^ , ^237^
^]^ and polyurethane^[^
^238^ , ^239^ , ^240^
^]^, due to their favorable mechanical properties and chemically defined composition. However, clinical translation was hindered by premature structural degradation, thrombosis, and calcification in preclinical studies.^[^
^241^ , ^242^
^]^ Recent advances in polymer science and the potential for crimpable valves via transcatheter delivery have driven development toward thin, durable, biostable, and elastic polymeric leaflets with more favorable thrombogenic and calcification profiles.^[^
^243^ , ^244^
^]^ However, these polymers still exhibit fatigue failure during cyclic loading and limited hemocompatibility. As such, research efforts have been driven toward fabrication techniques that best replicate the anisotropic architecture and mechanics of native valve leaflets, allowing them to withstand high transvalvular pressures with low flexural stiffness.^[^
^245^
^]^ Multilayered scaffolds, such as those consisting of a porous polycarbonate urethane (PCU) core—mimicking the spongy intermediate layer—stacked between highly oriented films made of electrospun PCL fibers in solution‐cast PCU—mimicking collagen and elastin‐rich outer layers—have demonstrated structural and mechanical anisotropy resembling native leaflets (Figure
12A).^[^
^246^
^]^ Leaflets show excellent biostability in accelerated oxidation environments and resistance to protein adsorption and calcification both in vitro and in vivo. However, concerns remain over the durability and scalability of the lengthy fabrication process.
![Figure 12: Material and biofabrication advances for polymeric heart valves. A) Biomimetic, tri‐layered valves using polycarbonate‐based polyurethane (PCU) porous foam covered by electrospun polycaprolactone (PCL)‐enhanced PCU films. Leaflet‐substitute material demonstrates improved biostability, durability, flexibility, and anticalcification potential compared to most commercial patches. Reproduced with permission.^[^
^246^
^]^ Copyright, 2022, Elsevier Ltd. B) Bioinspired silicone heart valves using direct ink writing. A heart‐valve‐shaped mandrel is spray‐coated with different stiffness silicone to create bilayered leaflets. Biomimetic fiber supports, inter‐leaflet edges and a stent‐like frame are printed onto the leaflets. Finite element analysis and accelerated wear testing demonstrate superior material durability and high fatigue resistance of the fiber‐reinforced leaflets. Reproduced with permission.^[^
^247^
^]^ Copyright, 2019, Elsevier Inc. C) Stereolithography (SLA) printed heterogenous polymeric heart valve from a polyacrylamide‐polyacrylic acid (PAAm‐PAA) hydrogel and strengthened with carboxyl‐Fe^3+^ complexes. Printed hydrogel skeleton is injected molded with a soft hydrogel PAAm matrix to form a composite to allow the polymer skeleton and matrix to entangle topologically. Hydrogel valve remains intact after 10 000 cycles in a hemodynamic test system. Reproduced with permission.^[^
^248^
^]^ Copyright, 2021, Elsevier Inc. D) Whole heart valve via SLA printing of urea‐based poly(N‐acryloylsemicarbazide‐co‐acrylamide) (P(NASC‐co‐Aam)) hydrogel ink and surface functionalization with heparin‐like sodium polystyrene sulfonate (PSS) chains via reversible addition‐fragmentation chain transfer (RAFT) polymerization to reduce thrombogenicity and hemolysis. Reproduced with permission.^[^
^249^
^]^ Copyright, 2022, American Chemical Society.](ADMA-37-2420114-g005.jpg)
Additive manufacturing strategies alleviate limitations associated with lengthy fabrication processes. For instance, multiaxis 3D printing has been utilized for fabrication of novel silicone heart valves.^[^
^247^
^]^ A heart‐valve‐shaped mandrel was spray‐coated with firmer silicone to mimic the collagen‐rich layer and a subsequent softer silicone layer to mirror the stretchable elastin layer in native leaflets (Figure 12B). Verified using finite element analysis to reduce peak stress, stiff silicone structural support fibers were printed onto the two‐layered valve leaflets. A hard silicone was then used to construct the inter‐leaflet edges and an auxetic patterned crimpable mesh stent, which conformed to a patient‐specific valve root geometry, effectively preventing paravalvular leakage. In vitro assessment indicated that the silicone valve fulfilled the minimum ISO 5840 standard for transcatheter valve prostheses in terms of the minimum effective orifice area and regurgitation fraction and withstood up to 40 million cardiac cycles (equivalent to 1 year in vivo) under accelerated wear testing.
Stereolithography printing has been used to fabricate anatomical, heterogenous polymeric valves from polyacrylamide‐polyacrylic acid (PAAm‐PAA) hydrogels (Figure 12C).^[^
^248^
^]^ The PAAm‐PAA hydrogel skeleton strengthened by carboxyl‐Fe^3+^ complexes, was filled with a softer PAAm matrix to absorb and distribute applied stress. While achieving significantly higher fatigue resistance compared to a homogenous control valve, stress decay under cyclic loading due to breakage of carboxyl‐Fe^3+^ coordination bonds resulted in significantly lower fatigue resistance than the 200 million cycles required by ISO 5840. This highlights the need for mechanical reinforcement of purely hydrogel‐based valves. Recently, patient‐specific, biomechanically compatible, hydrogel aortic heart valves were fabricated by coupling reversible addition‐fragmentation chain transfer (RAFT) polymerization with conventional SLA 3D printing^[^
^249^
^]^ (Figure 12D). The key component is a novel urea‐based polymer ink, poly(N‐acryloylsemicarbazide‐*co‐*acrylamide) (P(NASC‐co‐Aam)), which forms a tough hydrogel. This construct was then surface‐modified with a heparin‐like polymer through RAFT polymerization, resulting in low hemolysis, anticoagulation, and an acceptable inflammatory response. The extensive and robust hydrogen‐bonding network within the hydrogel enabled excellent fatigue resistance (exhibiting minimal damage even after 1.8×10^5^ cycles), with sufficient regurgitation fraction and transvalvular pressure gradient.
Regenerative Tissue‐Engineered Heart Valves
Tissue‐engineered heart valves utilize biodegradable polymers or decellularized allogeneic/xenogeneic scaffolds— sometimes seeded with cells—and are designed to be gradually replaced by endogenous tissue formation, leaving behind autologous, highly‐organized functional valve leaflets to provide life‐long performance.^[^
^169^
^]^ This eliminates the requirement of multiple replacement surgeries and anticoagulation therapy. Allograft scarcity and xenograft immunogenicity risks have shifted innovation toward bioresorbable natural (e.g., collagen, fibrin, elastin, hyaluronic acid, chitosan, etc.) and synthetic (e.g., PGA, PCL, PLA, etc.) polymers.^[^
^250^ , ^251^
^]^ To date, these materials have demonstrated promising results from in vitro and preclinical in vivo studies in terms of early functionality, tissue remodeling, endothelialization, and suitability for transcatheter delivery.^[^
^169^
^]^ Recent findings offer encouraging evidence for the potential of tissue engineering to address the longstanding challenge of late‐term calcification.^[^
^252^
^]^ Among these materials, silk has been extensively explored in vascular grafts and cardiac patches due to its hemocompatibility, endothelialization, minimal inflammatory response, resistance to calcification, and favorable degradation and tissue remodeling properties.^[^
^253^
^]^ While its application in valve replacement has been limited,^[^
^254^ , ^255^
^]^ recent advancements in silk engineering, including the development of high‐strength fibers produced by transgenic worms,^[^
^256^
^]^ are paving the way for its application in this field. Traditionally, tissue‐engineered heart valves have relied on cell‐based manufacturing to construct living tissues in vitro. Relevant autologous or allogenic cells (e.g., valvular cells, bone marrow stem cells, progenitor cells from peripheral blood, or amniotic fluid) are seeded onto scaffolds and conditioned in bioreactors that mimic physiological conditions to induce ECM deposition and spatial organization mimetic of native leaflets.^[^
^257^
^]^ However, this process is logistically challenging and costly to implement due to the need for patient‐specific cells and the lengthy in vitro culturing phase.^[^
^169^
^]^
In situ tissue engineering has emerged as a promising approach for heart valve replacement. It involves directly implanting a bioresorbable elastomeric scaffold into a patient, eliminating in vitro culturing and providing “off‐the‐shelf” valve replacements. Materials are designed to immediately replace valvular function upon implantation and maintain functionality throughout remodeling in vivo. As such, the balance between tissue formation and graft resorption is critical. Rapid resorption may lead to premature graft failure, while excessive tissue deposition may cause leaflet thickening and retraction.^[^
^258^
^]^ Therefore, innovation is focused on the bioinspired fabrication of mechanically suitable scaffolds with adequate porosity to enable cell infiltration while integrating key elements of the heart valve structure–function relationships to guide in situ tissue regeneration. Most materials for this approach have used electrospinning to create fibrillar scaffolds that closely mimic the distribution and orientation of the long, thin collagen fibrils found in native leaflets. Preclinical^[^
^259^ , ^260^ , ^261^
^]^ and clinical studies^[^
^262^ , ^263^
^]^ of fibrous bioresorbable elastomeric heart valves in the pulmonary position have generally demonstrated extensive in situ cellularization and tissue remodeling, progressing from the leaflet hinge toward the tip, strongly implying that these initially infiltrating cells originate from adjacent arterial tissue. However, some studies report impaired valve functionality associated with leaflet thickening, retraction, and degradation,^[^
^263^
^]^ as well as intervalve and interleaflet variabilities.^[^
^261^
^]^ In a significant study, researchers developed a slow‐degrading, porous, microfibrous electrospun bis‐urea‐modified polycarbonate (PC‐BU) tube that was sutured onto a polyether ether ketone (PEEK) supportive frame and coated with non‐thrombogenic fibrin (Figure
13A).^[^
^261^
^]^ Surgical implantation of this pulmonary valve in sheep demonstrated sustained functionality over 1‐year, with extensive in situ cellularization and tissue formation. This provided the first long‐term preclinical evidence for sustainable in situ formation of living valvular tissue using solely bioresorbable synthetic material. However, subsequent investigations revealed that predefined fiber alignment in the PC‐BU electrospun tissue‐engineered valve did not induce native‐like oriented collagen fibers in the circumferential direction.^[^
^264^
^]^ Instead, unwanted deposition of ECM predominantly occurred on the scaffold surface, resulting in leaflet thickening, stenosis, and loss of cell‐fiber interactions. The limited control over fiber architecture in conventional electrospinning likely contributed to the dense fibrous microstructure, compromising porosity and limiting cell infiltration.^[^
^265^
^]^ Recent advancements in electrospun valves have modified the design of collectors for polymer deposition, enhancing fiber alignment, morphology and porosity to better facilitate cell infiltration.^[^
^266^ , ^267^
^]^
![Figure 13: Material and biofabrication advances for in situ tissue‐engineered heart valves. A) (Left) Bis‐urea‐modified polycarbonate (PC‐BU) elastomer electrospun onto polyetheretherketone (PEEK) frame to form a porous, microfibrous heart valve. (Center) Implanted acellular scaffold triggers sequential recruitment of inflammatory, progenitor, and ECM‐producing cells, driving host tissue remodeling and regeneration to form a living, autologous heart valve. (Right) Explanted after 12‐months of implantation in the pulmonary position in sheep shows neo‐matrix formation in leaflets. Reproduced with permission.^[^
^261^
^]^ Copyright, 2017, Elsevier Ltd. B) Melt‐electro‐writing (MEW) produces serpentine architecture poly(caprolactone) (PCL) tubular scaffold. Resulting MEW‐PCL macroporous scaffold injection‐molded with elastin‐like recombinamer (ELR) hydrogel for favorable cellular infiltration and hemocompatibility. Reproduced with permission.^[^
^268^
^]^ Copyright, 2022, Wiley‐VCH. C) Soft and elastic PCL fibrous heart valve (FibraValve) fabricated using focused rotary jet printing (PRJS) manufactured in less than 10 min by increasing the rate of fiber deposition on a rotating mandrel. Biomimetic fibrous valve is interfaced with expandable stent for transcatheter delivery, readily supporting initial cellular adhesion and infiltration, both in vitro and in large‐animal model. Reproduced with permission.^[^
^270^
^]^ Copyright, 2023 Elsevier Inc. D) Cell‐free direct ink writing printed aligned PCL fiber scaffold, replicating the orientation and structural rigidity in the fibrosa layer of native leaflets. The fiber‐reinforcing PCL layer is molded with a cell‐laden 3D‐printed gelatin methacrylate and polyethylene glycol diacrylate (GelMA/PEGDA) bioink layer, encapsulating valvular interstitial‐like cells, to develop a multilayered leaflet capable of repair and remodeling. Reproduced with permission (left).^[^
^271^
^]^ Copyright, 2020, Elsevier Ltd. Reproduced with permission (center and right).^[^
^273^
^]^ Copyright, 2024, Wiley‐VCH GmbH.](ADMA-37-2420114-g001.jpg)
Melt electro‐writing (MEW) was recently introduced as a promising additive manufacturing technique to achieve highly accurate layer‐by‐layer deposition of fibers at submicrometer resolution. PCL was processed using MEW to accurately engineer a wavy‐like, continuous, fibrous, macroporous tubular scaffold that mimicked the native collagen arrangement (Figure 13B).^[^
^268^
^]^ This macroporous PCL MEW scaffold was then injection molded with a soft, microporous elastin‐like recombinamer (ELR) hydrogel network to support rapid cell infiltration, endothelialization, and hemocompatibility. The resulting MEW/ELR heart valve satisfied ISO 5840 hemodynamic requirements under both aortic and pulmonary conditions but showed higher‐than‐ideal leakage (regurgitation fraction). This was likely attributed to the leaflets’ temporary porosity, which allowed some backflow through the leaflets in vitro. However, this does not necessarily reflect in vivo functionality, as it is expected that the scaffold would be instantaneously filled with fibrin from the blood. In vitro optimization of the scaffold design, followed by an in vivo study, is required to evaluate the capability of the heterogenous architecture to remodel tissue and guide ECM formation.^[^
^268^
^]^ Another study employed MEW to fabricate trilayered anisotropic nanofibrous PCL scaffolds resembling the native collagen framework. VICs were encapsulated within a soft microporous gelatin‐methacrylate (GelMA) hydrogel containing bioactive chondroitin sulfate methacrylate (ChsMA), which was infused into the MEW‐PCL scaffold. In vitro and in vivo evaluation revealed the scaffolds’ ability to promote VIC growth, ECM remodeling, hemocompatibility and endothelialization, while minimizing proinflammatory cell infiltration and calcification after 4 weeks.^[^
^269^
^]^
Recent work introduced the use of Focused Rotary Jet Spinning (FRJS) to fabricate 3D hierarchical micro‐ and nanofiber tri‐leaflet FibraValve using streams of focused air to quickly and accurately extrude circumferentially oriented PLA and PCL (PLCL) copolymer fibers onto a spinning heart‐valve‐shaped mandrel, replicating the load‐bearing fibrosa layer found in native leaflets (Figure 13C).^[^
^270^
^]^ The entire manufacturing process takes less than 10 min, a significant improvement over conventional electrospinning, which can take several hours. The enhanced fiber production rate and accurate micro‐ and nanoscale resolution enables rapid manufacturing and design iteration capabilities, which is essential to efficiently screen multiple valve scaffold designs, prior to more costly in vivo applications. Acute performance in the pulmonary position of an adult sheep using a transcatheter delivery system, demonstrated sufficient leaflet motion and coaptation area, minimal regurgitation, and absent stenosis during a 1 h implantation. Initial cellular adhesion and infiltration into the bulk scaffold without loss of macroscopic structure were also observed, providing promising results for functional tissue remodeling and replacement with the use of living tissue. However, to truly assess the long‐term viability of the synthetic PLCL scaffold, chronic in vivo studies are necessary to evaluate the degradation rate, calcification, and potential for leaflet contraction and native remodeling in the unique, high‐shear environment of the heart.
Another approach to in situ tissue engineered heart valves involves a one‐step pre‐seeding procedure, allowing more control over the initial host response. For example a multilayered aortic valve leaflet was developed by layering 3D printed circumferentially aligned PCL, mimicking the load‐bearing fibrosa layer, with a gelatin‐methacrylate/poly(ethylene glycol) diacrylate (GelMA/PEGDA) hydrogel containing autologous human induced mesenchymal stem cells (iMSC), mimicking the spongiosa/ventricularis layers (Figure 13D).^[^
^271^
^]^ Under the correct microenvironment, iMSCs are believed to have the potential to differentiate into valvular interstitial cells (VICs)—the most abundant primary cells in native valve tissue, responsible for mediating tissue remodeling.^[^
^272^
^]^ The resulting PCL‐reinforced GelMA/PEGDA hydrogel scaffold achieved an elastic modulus within physiological range and demonstrated superior cell adhesion and proliferation compared to hydrogel‐free PCL scaffolds.^[^
^271^
^]^ Notably, the encapsulated iMSCs facilitated cellular remodeling for VIC‐like cells, promoting ECM production for up to 14 days under physiologically relevant dynamic shear and stretching conditions in vitro. This improved cellular microenvironment minimized the expression of a myofibroblast‐like phenotype, which is typically observed with stiff PCL scaffolds. The scaffold successfully replicated the biomechanical properties and structure of the native trilayered structure, maintaining structural stability while facilitating the valvular remodeling in vitro. Despite yielding a promising leaflet replacement with improved properties for cellular integration and biological function, further studies are required to assess the long‐term durability, degradation, and hemodynamics of the multilayered leaflet inside a flow loop.^[^
^272^
^]^
While a complete understanding of in situ neo‐tissue formation and the associated long‐term function and growth of tissue‐engineered heart valves is still missing,^[^
^169^
^]^ these exciting advancements in biofabrication to produce materials that mimic the complex, heterogenous microarchitecture of native valve leaflets offer a promising outlook toward the production of effective regenerative valve replacements.
Future Directions
This review presented an overview of the drivers behind notable recent and innovative approaches at each stage of cardiac intervention, including diagnostics, management, and repair. As we look to the future, several promising directions emerge that could revolutionize cardiac care (Figure
14 ).
One significant advancement is the development of in situ power generation to eliminate the need for batteries in implantable medical devices. This innovation reduces the risks associated with multiple surgeries and battery malfunctions and aids in miniaturization by removing a large component from the device. Creating more reliable and efficient piezoelectric or triboelectric materials capable of harvesting or generating power is crucial to this progress. Additionally, development of materials which can generate power by harnessing biomechanical or biochemical energy, such as those based on triboelectric nanogenerators^[^
^274^
^]^ or thermoelectric nanomaterials^[^
^275^ , ^276^
^]^ can enable long‐term battery‐free continuous monitoring of wearable devices.
Miniaturization is a major trend in almost all areas where a device is not directly replacing a tissue structure. In transcatheter robotics, a smaller footprint means devices are more dexterous, create fewer local stresses on vessel walls—reducing the chance of injury—and can facilitate access to more remote areas of the cardiovascular system for interventions. In wearable sensor devices, miniaturization allows for seamless integration with the wearer's daily activities, making them more comfortable and likely to maintain conformal contact through various activities.
In ventricular assist devices, there is a shift toward developing nonblood contacting devices to eliminate adverse effects of device geometry, composition, or blood interactions. Advances now aim to mimic native cardiac muscle movements more closely to match natural motion and hemodynamics. More intelligent control and increased efficiency of artificial muscle actuation are expected to drive innovation and implementation of these devices in the clinic. Additionally, implementing battery‐free power generation and improved distances of real‐time wireless monitoring will enhance practicality of such devices.
Translating the effectiveness of cardiac devices from small animals to human applications presents various challenges due to differences in physiology between species. While devices being developed show great promise, many have only been tested in small animal models. It is unclear whether the currently demonstrated capabilities will extend to human physiology. Bioelectronics intended for tissue actuation will require higher power input and voltage outputs over greater tissue surface areas. Similarly, monitoring devices in larger anatomies will require integration of a greater number of signals over larger surface areas. Thus, large animal models are essential for more accurate representation and testing.
The application of AI in heart disease diagnosis, therapeutics, and management is a rapidly growing field. Despite significant concerns and limitations around the validation of deep learning algorithms, the provenance of training data, and the explainability of AI, the potential of AI to revolutionize cardiac care is immense. As these issues are resolved, we expect to see increased integration of AI into clinical workflows, mitigating bottlenecks associated with processing and analyzing large volumes of data.^[^
^277^
^]^ FDA approval of AI‐based algorithms for medical diagnostics, such as AliveCor, which supports early detection of atrial fibrillation via a six‐lead smartphone ECG, shows great promise.^[^
^278^
^]^ Meanwhile, an AI‐powered clinical decision support tool is being developed which showed promise during a recent pragmatic, randomized clinical trial (22 641 adults) for early diagnosis of low EF during routine 12‐lead ECG monitoring.^[^
^279^
^]^ Interestingly, the randomized trial found that abnormality detection was more pronounced in outpatient settings—where ECG testing is less frequent—compared to in‐hospital. This indicates its potential for detection of a readily treatable condition where diagnoses may be missed or delayed due to infrequent testing and/or scarcity of resources. Future work seeks to increase cost‐effectiveness to enable greater accessibility and extend diagnoses to other conditions that may be detectable via simple, relatively inexpensive ECGs.
The implementation of machine learning to create training databases can improve the robustness of extracted information, assisting in the improved, patient‐specific early detection of symptoms associated with abnormal heart function. ECG waveforms contain tens of thousands of data points, and subtle variations in waveforms may enable AI to inform cardiovascular disease diagnoses. Other common measurement techniques, such as MRI and CT also provide information which can be easily used to train deep learning/neural networks.^[^
^280^
^]^ Deep learning algorithms are being developed, which allow AI to extract relevant information from raw ECG waveforms from patients to develop screening tools for cardiovascular conditions, such as aortic stenosis,^[^
^281^ , ^282^
^]^ atrial fibrillation,^[^
^283^
^]^ and coronary artery disease.^[^
^284^
^]^ Interestingly, Lin et al. recently developed and validated a deep learning model which detects coronary artery disease based on facial features using patient photos.^[^
^285^
^]^ Considerable work is being done to develop algorithms that can more accurately extrapolate blood pressure from data measured by devices, such as wearable ECG and photoplethysmography‐based sensors.^[^
^286^ , ^287^
^]^ Improvements in AI and machine learning technologies will enhance the accuracy of these approaches, enabling more comprehensive noninvasive health monitoring for individuals and easy integration into clinical practice for improved monitoring, management, and treatment. In treatments requiring stem cell therapy, AI has the potential to improve the determination of stem cell safety, viability, and bioefficacy and inform improved patient matching for optimized patient response. It can improve patient outcomes and increase accessibility by driving down the cost of such treatments.^[^
^288^
^]^
From a materials development perspective, using computational modeling and AI can accelerate advances in material development by reducing time and resource‐intensive design, synthesis, and testing steps. For example, modeling the assembly of polymers and peptide structures can reduce the time associated with experimental trial‐and‐error, informing decisions regarding materials and therapeutic factor compositions. Techniques such as high throughput virtual screening have been utilized to determine ideal candidate materials based on application‐specific properties by subjecting them to simulations, thus drastically accelerating materials development.^[^
^289^
^]^ Smart materials designed to inform ideal parameters, such as unwanted mechanical and biological reactions, while being cost‐effective, have already been demonstrated in the design of stents for coronary artery disease interventions.^[^
^290^
^]^ In cardiac tissue engineering, often the most expensive and cumbersome experiments involve lengthy biological assays. Recently, an AI‐powered software was developed which aims to alleviate some of these challenges. MLATE (Machine Learning Application in Tissue Engineering) is an open‐source cardiac tissue engineering scaffold database developed by mining data from literature based on specifications, such as materials, cell lines, and fabrication materials, then rating scaffolds based on cell viability, growth, proliferation, and cell differentiation.^[^
^291^
^]^ Twenty‐eight machine learning algorithms were then trained on the data to determine the most effective one for predicting cell behavior based on parameters such as scaffold composition and fabrication method. This demonstrates the potential of machine learning algorithms to predict cell behavior to guide tissue engineering strategies for cardiovascular treatment.
Additive manufacturing offers an automated process for fabricating cardiovascular medical devices, spanning wearable and implantable sensors to cell‐laden constructs for tissue repair and regeneration.^[^
^292^ , ^293^ , ^294^
^]^ However, it is becoming evident that the various subsets of additive manufacturing technologies, such as inkjet, lithography, and extrusion, are limited in isolation.^[^
^295^
^]^ Newer technologies combining different modalities are of interest, driven by material adaptability where often only a subset of materials is compatible with specific printing techniques, making the resultant functional outcomes somewhat restricted. Therefore, the convergence of different printing technologies allows multimaterial integration into a device, enabling multifunctionality and potentially higher performance.^[^
^296^
^]^ A recent study combined volumetric printing with MEW to fabricate tubular hydrogel‐based composites with enhanced mechanical behavior.^[^
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^]^ By using this converged approach, hybrid polymer‐hydrogel tubular structures were successfully engineered, showing tri‐layered cell‐laden vessels and features (valves, branches, fenestrations). Imaging acquisition, which underlines the additive manufacturing process, is another aspect to consider. Given advancements in imaging modalities such as CT and MRI, 3D models of higher precision can be constructed for prints of higher resolution.^[^
^298^ , ^299^ , ^300^
^]^ This can lead to personalized devices, such as implantable sensors, cardiac patches, and valves, specifically designed for patients for optimal outcomes. The expansion of additive manufacturing into additive biomanufacturing now allows the incorporation of cells and biological molecules into these devices for tissue regeneration, with precise spatial organization and localization.^[^
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^]^ However, this technology is still in its infancy, as preserving cell viability and function during the biomanufacturing process remains challenging, and patient‐to‐patient variation also contributes to unpredictable results. Given the recent interest in machine learning, creating training databases from clinical data curation might digitalize additive biomanufacturing in the future, better predicting clinical outcomes and preventing complications.^[^
^302^ , ^303^
^]^
Conclusion
In conclusion, heart disease interventions require a multitude of approaches, all of which present their own unique challenges. Improved understanding of cardiac function and its pathophysiology, along with material innovations to tailor approaches, accordingly, hold the key to successful intervention. A critical enabler of advances in this area and their successful translation from laboratory to clinic is collaboration between multiple disciplines, codesigning appropriate biomaterials and device technologies, and critically adequate device testing to address unmet health needs.
Conflict of Interest
The authors declare no conflict of interest.