Nano Energy 76 (2020) 105076

Available online 21 June 20202211-2855/© 2020 Published by Elsevier Ltd.

Cardiac energy harvesting and sensing based on piezoelectric and triboelectric designs

Lin Dong 1, Congran Jin 1, Andrew B. Closson, Ian Trase, Haley R. Richards, Zi Chen *, John X. J. Zhang **

Thayer School of Engineering, Dartmouth College, Hanover, NH, 03755, USA

A R T I C L E I N F O

Keywords: Energy harvesting Sensing Biomedical devices Piezoelectric Triboelectric Cardiac implantable devices

A B S T R A C T

Cardiac implantable devices are effective both as methods of controlling irregular heartbeats in people with heart rhythm disorders and as interventional therapy for cardiac diseases. However, early failures and battery limi-tations of cardiac medical devices lead to periodic battery replacement surgeries for patients that are both risky and costly. Self-sustainable energy generation could significantly extend the lifetime and effectiveness of cardiac implantable devices and other implantable biomedical devices (IMDs). The human heart is a compelling in vivo energy source and is a natural battery to power IMDs. In this review, we discuss current trends of developing self- powered cardiac medical devices harvesting the energy from the heart. Based on key challenges and limitations, we propose design principles for cardiac energy harvesters and sensors. We further discuss advanced energy materials, structural and fabrication considerations, biosafety and biocompatibility, and comfortability and flexibility. Moreover, recent advances in cardiac energy harvesting and sensing devices in both in vivo and in vitro studies are reviewed and discussed. Such sustainable energy strategies based on piezoelectric and triboelectric designs provide a promising means to reduce the reliance on batteries for powering cardiac medical devices and other IMDs.

1. Introduction

Cardiac disease is the leading cause of death worldwide, accounting for nearly one third of all deaths [1]. The use of cardiac implantable electronic devices (CIEDs) remains among the most effective methods of interventional therapy for cardiac diseases [2,3]. Despite successfully offering supportive therapy to an unhealthy heart for numerous patients, these electronics face a critical limitation in that they need a sustainable power source to function properly. Currently, these implantable devices rely on built-in batteries to provide the needed power. A typical cardiac pacemaker lasts 5–10 years before replacement of the depleted batteries. A recent clinical study assessed implantable cardiac defibrillators (ICDs) longevity among 685 consecutive patients over the last 20 years, and 8% of devices encountered premature battery depletion at 3 years [4]. When the battery is depleted, a battery replacement surgery is inevitably needed to maintain the functioning of the device. These surgeries bring not only extra financial burdens, but also pain and potential infection to

the patient, especially for pediatric and geriatric patients. Engineering solutions are proposed to address these medical prob-

lems, and there are a variety of issues to resolve when designing and testing medical devices, especially when developing implantable medi-cal devices. For example, the devices should be small and lightweight to prevent additional mechanical loads on the heart and to make the im-plantation procedure as minimally invasive as possible. Biocompatible materials are commonly used to prevent an autoimmune response, but precautions should be taken during implantation procedures to mini-mize the risk of infection. The implanted device itself should guarantee many years of operation in the dynamic internal environment of the body without compromised performance. Time-dependent challenges in implantable device design need to be further considered, such as erosion of metal materials, tissue changes during wound healing, and deterio-ration of battery performance.

Energy harvesting (EH) and self-powered sensing strategies provide practical and promising power and diagnosis solutions for implantable

* Corresponding author. ** Corresponding author.

E-mail addresses: lin.dong@dartmouth.edu (L. Dong), zi.chen@dartmouth.edu (Z. Chen), john.zhang@dartmouth.edu (J.X.J. Zhang). 1 Equal contribution.

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https://doi.org/10.1016/j.nanoen.2020.105076 Received 27 April 2020; Received in revised form 28 May 2020; Accepted 8 June 2020

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cardiac devices [5–15]. An overview of design principles for cardiac medical applications is shown in Fig. 1. Materials, structures design, and fabrication approaches establish the foundation for device development. Piezoelectric and triboelectric materials are the two major materials that have been widely used for cardiac energy harvesters and sensors because of their excellent energy conversion capabilities. Examples of piezo-electric materials include polymers, ceramics, and single crystals, which exhibit the piezoelectric effect under stress or strain. Triboelectric ma-terials can be almost any material pair with a large difference in charge affinity. The structural design of the devices has multiple scales, ranging from the nanoscale and microscale to the macroscale. Nanowires, microporous structures, and multi-layer thin film structures are exam-ples of these three different scales. A variety of techniques have been employed to fabricate devices, including chemical synthesis, spin coating, and physical or chemical deposition [5,14]. Since energy har-vesting/sensing devices are usually either implanted inside the body or closely attached to the skin of the patient, it is critical that they are biocompatible and biosafe. Currently there are two strategies to ensure biosafety. The first one is to use biocompatible or biodegradable mate-rials to build the device. The second approach is to encapsulate the functional material within a coating of biocompatible material, with the most popular material being polydimethylsiloxane (PDMS). On the basis of an effective and biosafe device, researchers have been pursuing a better user experience. For example, the implanted or wearable device must not cause discomfort to the patient. Therefore, the device has to be flexible, soft, stretchable, and conformal to the implanted site. The de-vice could be fabricated by using flexible materials such as PDMS, sili-cone, and hydrogel; or it could utilize smart structures such as Kirigami patterns and helical structures to implement flexible and comfortable devices.

There have been a number of published reviews on energy harvesting and sensing devices with a focus on materials development [5–10], energy transducers and device designs [8,11–15], biomedical applica-tions [16–20], and micro-electromechanical systems (MEMS) or nano-scale devices [21–23]. By contrast, this review is focused on sustainable energy strategies for cardiac medical applications based on piezoelectric and triboelectric designs (Fig. 1). We begin by discussing the current

cardiac implantable electronic devices, key challenges, and limitations. Next, we carefully examine the design principles of piezoelectric and triboelectric based devices, which include the materials, structure and fabrication, considerations of biosafety and biocompatibility, and com-fortability and flexibility. We categorize designs into cardiac EH devices and sensing devices, and closely review recent progress with summaries of device performance in both in vivo and in vitro studies. Lastly, we conclude by providing our perspectives in the field.

2. Key challenges and design principles

Heart is among the most vital organs in the human body. However, it is also vulnerable to many diseases such as heart attack and arrhythmia. CIEDs have been developed to diagnose and support the normal function of hearts. In this section two common CIEDs, Cardiac Resynchronization Therapy (CRT) devices and Insertable Cardiac Monitors (ICMs), are discussed, with the device longevity emphasized. CRT is for heart failure patients whose heart chambers do not beat in unison. It was reported that in 2013, there were approximately 1,000,000 patients living with cardiac pacemakers and 400,000 with defibrillators according to physician claims, with the number of new implantations remaining relatively constant over the last decade [2]. ICMs are another common class of implantable cardiac devices. These devices keep track of cardiac behavior continuously, by measuring the electrical activity of the heart over extended periods of time. ICMs give key indications of heart failure, particularly in atrial fibrillation.

2.1. Cardiac medical devices and limitations

One of the most common CRT devices is the cardiac pacemaker, which provides electrical stimulation to the heart muscle to regulate its contractions and is implanted in over a million patients per year worldwide. A cardiac pacemaker consists of a subdermal pulse generator and pacemaker leads that extend into the chambers of the heart and are affixed to the heart wall in either the left or right ventricle (single lead pacing) or in both of the ventricles (dual lead). A pacemaker lead simultaneously senses the intrinsic electrical activity of the heart and

Fig. 1. Typical cardiac implantable electronic devices and a schematic of the design principles of cardiac energy harvesting and sensing devices including the materials, structures and fabrication, considerations of biosafety and biocompatibility, and comfortability and flexibility.

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provides a voltage pulse if it detects abnormal activity according to programmed thresholds for the intrinsic voltage and heart rate. Nowa-days, leadless pacemakers are a remarkable innovation in pacemaker technology. They are miniature self-contained devices affixed to the endocardium and implanted through a catheter. While leadless pace-makers offer a radically less invasive implantation process, they do not solve the fundamental burden of a limited power supply. Only two types of leadless pacemakers are currently available (Medtronic Micra and St. Jude Nanostim) and the high fabrication costs currently make leadless pacemakers prohibitive to most patients [24]. ICDs are the other type of CRT devices, and similarly apply an electric shock to restore the heart’s rhythm. In general, cardiac pacemakers and ICDs have similar stimu-lation mechanisms, device layout, and battery life, though defibrillators are used in events of cardiac arrest as opposed to routine pacing. A ventricular assist device (VAD) is an implanted mechanical pump in the chest used to support heart function and blood flow for weak or failing hearts. A VAD pumps blood from the lower chambers of the heart to the rest of the body. By relieving the heart’s workload, VAD technology will be used as a comparable alternative to heart transplantation for patients with heart failure [25].

ICMs are another common class of implantable cardiac devices, and they are leadless subcutaneous devices that continuously monitor the heart rhythm and record events over a timeframe measured in years. These devices allow for the diagnosis of infrequent rhythm abnormal-ities, which can be the cause of syncope, palpitations, and stroke [26]. Typical ICMs also rely on the limited power supply of lithium batteries, and the latest iteration of the device (Reveal LINQTM, Medtronic Inc.) has a lifespan of approximately 3 years [27]. ICMs are subdermal im-plants that continuously measure an ECG signal over a long period of time. Implantation and replacement procedures for ICMs are less inva-sive than those of CRT devices but are nevertheless performed in a clinic [28].

Although cardiac pacemakers and ICDs are effective tools for treat-ment of heart block and ventricular dysrhythmias, providing a sustain-able energy solution is the key challenge for CRT devices. Table 1 compares the longevity and power consumption of a variety of com-mercial CRT devices, assuming that device failures are caused by battery depletion. Lead-based cardiac pacemakers and defibrillators consume a similar amount of power despite differing stimulation mechanisms, while the leadless pacemaker consumes much less power but has a much smaller battery capacity due to size constraints. It is important to note that clinical longevity data is not available for all the devices, and the values in Table 1 are taken from product specification sheets [29–32]. Current pacemakers have limited battery life, and thus require invasive replacement surgeries every 7–10 years. While defibrillator power consumption and longevity is more difficult to predict, a cohort of 953 patients showed a median battery longevity of 4.9 years [33]. Notably, more ICDs are replaced every year than cardiac pacemakers, and ICDs comprise a larger percentage of the CRT device market than pacemakers worldwide. The periodic battery replacement surgeries inevitably in-crease the complications such as cardiac tissue infection and arterial puncture, and significantly hinder the quality of life of cardiac patients [34]. This work is an overview of engineering solutions for cardiac medical devices based on piezoelectric and triboelectric energy har-vesting and sensing strategies. Strategies in scalable manufacturing and device development are also taken into consideration.

In addition to limited battery life, there are other drawbacks of CIEDs

that inhibit the functionality of the devices and the quality of life of the patient. Many older pacemakers are incompatible with magnetic reso-nance imaging (MRI) given that the pacemaker lead may vibrate, in-crease temperature, or conduct induced electric current in the presence of a strong magnetic field [35]. In addition, there is a significant risk of infection associated with the implantation of CRT devices in the tissue proximal to both the pulse generator and the leads. These infections warrant immediate removal of the device entirely [36]. Mechanical lead failure is another common cause for device removal with an incidence of approximately 1% [37].

2.2. Materials, structures and fabrication approaches

2.2.1. Piezoelectric based designs One of the most common transduction mechanisms for mechanical

EH and sensing makes use of the direct piezoelectric effect. Piezoelectric based devices typically fall into one of three categories: bulk, thin film, and flexible. Bulk type consists of cantilevers, cymbals or diaphragms, and stack configurations. Thin films are typically used in MEMS based technologies. Flexible devices are of growing interest and many material configurations are being developed to better handle the required flexi-bility. All these devices are affected by a wide range of parameters associated with the input as well as the structural design, including ac-celeration, applied force, vibration frequency, impedance matching, mass, surrounding environment, and materials designs [8,11–15]. In this section we will focus on a variety of piezoelectric materials used for cardiac EH and sensing and their respective fabrication aspects.

In piezoelectric materials, mechanical stress of the material’s crys-talline structure induces an electric charge on the surface of the material. For applications in EH and sensing, devices make use of stresses on the materials to store or readout this charge buildup respectively. The piezoelectric effect is reversible, and in the reverse piezoelectric effect an applied electric field can induce a resulting mechanical strain in the material. This allows piezoelectric materials to be used in applications as actuators as well as for sensing and energy harvesting. Examples of this are with speakers and piezoelectric inkjet printing. Essentially, piezo-electricity occurs by breaking the central symmetry in a crystalline structure under external force, which deforms the crystal structure forming an electric dipole. This electric dipole induces opposite electric charges on the material surface and thus creates an electrical potential difference across the piezoelectric material. The direction of the applied stress in relation to the material’s polar axis determines electrical output performance. In ferroelectric materials such as lead zirconate titanate (PZT), lead magnesium niobate lead titanate (PMN-PT), Barium titanate (BaTiO3), and polyvinylidene fluoride (PVDF), poling direction de-termines the polar axis. For non-ferroelectric materials such as Zinc oxide (ZnO) and Aluminum Nitride (AlN) the polar axis is determined by the crystal orientation [14]. The crystalline makeup of piezoelectric materials can vary greatly from one material to another. For example, some natural materials such as DNA and bones have been shown to be piezoelectric, while the more commonly used piezoelectric materials are ceramics such as PZT, a Wurtzite structured ZnO, or even a semi-crystalline polymer of PVDF. Electromechanical coupling is the con-version between mechanical and electrical energy. In piezoelectric materials this conversion factor depends on both the piezoelectric con-stants and the elastic properties. Quantitatively, piezoelectricity is characterized by the strain-charge coupled equations [38]. Based on

Table 1 Comparison of reported typical CRT devices in the market.

Device Device type Battery Capacity (Ah) (Nominal)amplitude (V) Longevity /500 Ω lead (years)

Medtronic Adapta ADDR01 [29] Pacemaker Li/I2 1.3 3.5 6 Boston Scientific L300 [30] Pacemaker Li/Cfx 1.6 2.5 9.2 Medtronic Micra MC1VR01 [31] Leadless Pacemaker Li/Cfx-SVO 0.12 2.5 5.8

Boston Scientific INOGEN EL D142 [32] Defibrillator Li–MnO2 1.9 2.5 9

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these equations, the amount of electric charge converted for a given unit of mechanical stress can be enhanced by increasing the piezoelectric coefficient (which is material driven) or by increasing the relative strain on the material for a given stress. This gives researchers opportunities for improving mechanical to electrical energy conversion of piezoelec-tric materials for EH and sensing applications.

Piezoelectricity exists in many different materials, including ce-ramics, single crystals, and polymers. One of the most commonly used piezoelectric materials found in numerous commercial applications is PZT. PbZrxTi1-xO3 is a polycrystalline solid solution and one of many ceramic piezoelectric materials. Owing to its high piezoelectric constant, it is a popular choice among researchers for EH and sensing applications. PZT has a perovskite crystalline structure which exhibits both pyro-electric and ferroelectric properties. The latter allows for an electric poling process to orient its electric dipoles. PZT exhibits piezoelectric coefficients d33 of ~300–1000 pC/N [39] with 0 < x < 1, one of the higher coefficients of all piezoelectric materials. The most typical fabrication method for PZT is a sol-gel deposition method [40–42], in

which a PZT solution gel of desired Ti/Zr ratio is spin-coated onto a substrate followed by a baking and annealing process that allows the film to crystalize, forming the polycrystalline piezoelectric material. Aerosol methods [43,44] and screen printing methods [45–47] for depositing films have also been developed. Many of the MEMS based fabrication methods are reviewed by Lin et al. [14] (Fig. 2A).

Researchers have developed numerous microstructured morphol-ogies for improving properties such as flexibility and piezoelectric response. Qi et al. developed flexible PZT nanoribbons [48,49] depos-ited by an RF-sputtering process and using a transfer printing method, printed onto a PDMS rubber sheet with a high piezoelectric constant of d33 ¼ 101 p.m./V. Kwon et al. [42] also developed PZT nanoribbons using a sol-gel deposition method followed by etching and transfer to a flexible polyethylene terephthalate (PET) substrate with a PDMS stamp. The ribbon-based devices were able to produce a voltage of ~2V and a current density of ~2 μA/cm2. Utilizing electrospinning, Chen et al. [50] were able to deposit aligned PZT nanofibers over an interdigitated electrode on a silicon substrate, and these nanofibers produced a power

Fig. 2. Piezoelectric and triboelectric materials development, working mechanisms and device designs. (A) An overview of the materials and fabrication, configuration and mechanics, and applications for piezoelectric based devices [14]. (B) Examples of piezoelectric materials, their material types and piezoelectric coefficients [10]. (C) The four basic working mechanisms of TENG [112]. (D) Comparison of charge affinity for various materials used in triboelectric designs [124].

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output of 0.03 μW. Lee et al. [51] developed a high-performance flexible nanogenerator based on the chemically reinforced composite thin film, which contained a stable dispersion of PZT nanoparticles (PZT-NH2 NPs) within the polymer matrix to generate a high output voltage of 65 V and current of 1.6 μA. Seen in these examples is a movement towards fabrication methods to produce more flexible piezoelectric devices, due to improved methods and the inherent large piezoelectric responses of PZT based materials.

The relaxor ferroelectrics lead-magnesium-niobate lead-titanate Pb (Mg1/3 Nb2/3)O3–PbTiO3 (PMN-PT) and PB(Zn1/3Nb2/3)O3–PbTiO3 (PZN-PT) are single crystal solid solutions that have exceptional piezo-electric properties. PMN-PT can be grown in single crystals oriented along the <110> and <001> directions. Compared to PZT, a poly-crystalline ceramic material, this allows for piezoelectric constants to be many times greater [52]. Some studies have shown piezoelectric coupling coefficients (d33) of up to 2500 p.m./V [53]. However, one of the difficulties of dealing with these materials is fabrication into patterned devices. Bottom-up approaches, of growing single crystals in a patterned array have proven difficult, with many issues involving the purity of the crystalline structures. Chen et al. developed a top-down method for fabricating scalable, ordered, single crystalline PMN-PT nanobelt arrays, by etching away larger single crystal materials. These patterned nanobelts showed a massive piezoelectric coefficient of 677 p. m./V [54]. One of the major issues for PZT and PMN-PT involves their brittle nature, with a maximum strain of 0.2% [55]. For these materials, slight stretching can lead to material failure.

Another piezoelectric nanomaterial that has received immense in-terest in the past couple of decades has been Zinc Oxide (ZnO). ZnO has a wurtzite crystal structure, which lacks central symmetry, giving the material piezoelectric properties [56]. ZnO has been studied for many decades for its various semiconducting properties for optoelectronics, sensors, and actuators. It was brought to prominence in 2006 for its EH applications when Wang et al. first showed the ability of ZnO nanowire arrays to convert small amounts of mechanical energy into electrical current [57]. It was first found that ZnO nanowires have a piezoelectric constant d33 ¼ 14–26 p.m./V [58], and these nanostructures have spurred a large interest in the material for piezoelectric based energy harvesters and sensors. Aligned nanowires are commonly grown with a physical vapor deposition method with the use of substrates and catalyst particles or seeds [59]. Lower temperature, hydrothermal based chem-ical approaches have been developed without using ZnO seeds, utilizing zinc nitrate and hexamethylenetetramine (HMTA). These methods have been used to form various structures including vertically and laterally aligned nanowires as well as nanorings, nanosprings, nanobelts, and nanospirals [59]. In 2010, Wang et al. demonstrated the vertical and lateral integration of ZnO nanowires into arrays. A lateral integration of 700 rows of ZnO nanowires produced a peak voltage of 1.26V at a strain of 0.19%. A stacked 3-layer device produced a peak power density of 2.7 mW/cm3 and was able to power a nanowire pH sensor and nanowire UV sensor [60]. They further developed a flexible nanogenerator by trans-ferring vertically aligned ZnO nanowires to a flexible substrate, in which a single layer was able to produce a peak output power density of 11 mW/cm3 [61].

To address some of the issues with more rigid piezoelectric materials, researchers have turned to piezoelectric polymers. PVDF and its co-polymers show the most potential and have been the most widely studied in the group of piezoelectric polymers. Along with piezoelec-tricity, PVDF has excellent chemical stability and biocompatibility [62–64]. PVDF and its co-polymers are semicrystalline with five different crystalline phases, the α, β, γ, δ, and ε-phases [65]. The β and γ-phases are the most electrically active, with the β-phase being the most widely studied and desired for EH and sensing applications [66,67]. The semicrystalline makeup of these polymers causes their desired proper-ties to be dependent on the polymers morphology [68]. This makes the fabrication of engineered nano and microstructures an important aspect of piezoelectric polymer device designs. The development of new

techniques for fabricating these structures have played a crucial role in the advancement of this technology. Many different fabrication tech-niques have been developed to increase the electrical output of PVDF. Thin films are the most common form of PVDF and are formed through a spin coating process. Physical patterning [69] and phase-separation techniques [70–72] have given researches the ability to create micro-structures on and within the PVDF thin films, giving rise to improved flexibility and electrical output. Micro/nano-fibers are another common form of PVDF, often formed through an electrospinning process [73,74]. These fibers are self-poled in their formation and exhibit a high crys-talline percentage of the β-phase. An interesting application of electro-spinning for pressure sensing is with the fabrication of core shell fibers by poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT: PSS) and PVDF-TrFE [75], which showed 4.5 times greater sensitivity than that of PVDF fibers. Due to the lower piezoelectric constant of PVDF compared to traditional piezoelectric ceramics such as PZT, PVDF is often doped with various other materials to improve the overall piezo-electric efficiency. Some of these include, multiwalled carbon nanotubes [76–79], nickel nanoparticles [80], gold nanoparticles [81], ZnO [82–85], BaTiO3 [86,87], as well as many others.

Fig. 2B summarizes a comparison of piezoelectric coefficients for different biocompatible piezoelectric materials [10]. More recently, piezoelectricity in biological materials from the basic building blocks (i. e., amino acids, peptides, proteins, etc.) to highly organized tissues (i.e., bones, skin, etc.) has been systematically reviewed [88]. The wide va-riety of piezoelectric material types allows multiple options, including ceramics, single crystals, and polymers. Higher electrical generation can be achieved when using some of the single crystalline and ceramic piezoelectric materials, while improved flexibility can be available with piezoelectric polymers (such as piezoelectric thin films and fibers). Different fabrication techniques through the engineered nano and mi-crostructures enable new piezoelectric polymers with high piezoelectric coefficients while exhibiting flexibility for EH and sensing applications. Various EH structures have been developed to implement the energy conversion in literature, such as cantilevers [89–93], beams [94–98], patches/membranes [85,98–106], compressed “8” shape [107], and helical structures [108–110]. By leveraging active piezoelectric mate-rials with smart structure configurations, EH devices can be designed for the desired functionality for applications. Therefore, piezoelectric ma-terials are on the frontier of new developments in biomedical EH and sensing applications.

2.2.2. Triboelectric based designs Triboelectric energy harvesting is an emerging technology that uses

the triboelectric effect to convert mechanical energy into electrical en-ergy. Triboelectric based devices have been successfully implanted to monitor cardiac signals such as heart rate, blood pressure, and endo-cardial pressure in real time. A triboelectric nanogenerator (TENG) has also been designed as a cardiac energy harvester that uses the otherwise wasted energy from the heart to power implantable electronics such as a pacemaker. This section focuses on the recent advances of triboelectric based devices, from working mechanism to material selection, structure designs, and surface morphology optimizations for biomedical applications.

In 2012, Dr. Wang and his coworkers designed the first triboelectric energy harvester that converted mechanical deformation to electrical output [111]. The basic working mechanism of the triboelectric gener-ator can be explained by the movement of electrical charges between two different materials [112]. When one material is in contact with another, the electrochemical potential will be balanced by the move-ment of charge between the two materials. When the two material start to separate, some atoms tend to give away extra electrons and others are inclined to keep them, creating an electrical potential difference. This electrical potential difference drives the triboelectric charges on the dielectric surface to produce a current until the difference disappears. In the process, the current I can be expressed by the equation I ¼ C ∂V=∂tþ

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V ∂C=∂t, where V is the voltage across the electrode, C is the capacitance of the system, and t is the time. In this expression, the first term repre-sents the rate of change of the voltage potential due to electrostatically induced charges and the second term describes the capacitive rate of change. All TENG are developed based on the fundamental mechanism of surface electrification and electrostatic induction. Though a comprehensive theoretical description of the working mechanism and performance of TENG is not within the scope of this review, systematic studies can be found in Ref. [113–115]. In general, there are four basic working modes of TENG devices: vertical contact-separation mode, contact-sliding mode, single electrode mode, and free-standing tribo-electric layer mode [112] (Fig. 2C). In the vertical contact-separation mode [116,117], surface charge transfer is induced by the contact be-tween two materials with different triboelectric polarity. The contact and change in proximity of the oppositely charged surfaces generates an electrical potential difference that drives the electrons to flow through an external load in response to the applied vertical movement on the device. In the contact-sliding mode [118,119], two surfaces with a large difference in charge affinity are in contact and sliding on each other. The generated current is a result of electron movement created by the coupled triboelectric and electrostatic effect during the relative displacement of the two contacting surfaces. In the single electrode mode, an object that is not a part of the TENG will play the role of the other frictional material. The principle underlying electrical generation of this kind is also based on a difference in potential between two different materials. The current flows between the electrode and ground through an external load. One example is the human skin based TENG developed by Yang et al. [120], where the human finger acts as a fric-tional layer and therefore this device can sense human skin touch. Recently, Wang et al. [121] designed a device with a frictional layer made of porous polytetrafluoroethylene (PTFE) and one electrode con-necting to ground. The device can convert the motion of a human hand with and without a latex glove into electrical signals with open-circuit voltages of 1.1 V and 6.9 V, respectively. Lastly, the free-standing mode involves an independent layer that is free to move. One electrode-free layer slides on top of two stationary bottom electrodes that connect through an external load [122]. Such a design does not have a firm contact between two frictional layers, and thus increases the durability of the device by decreasing wear and tear. One interesting example of a freestanding rotation based TENG was designed by Kwak [123], and the employed triboelectric material of butylated melamine formaldehyde (BMF) in that design demonstrated superior mechanical and triboelectric characteristics. It was found that the root-mean-square output of BMF based TENG was 210 V and 125 μA, which was higher than that of the one with PTFE.

One significant advantage of TENG is the wide spectrum of materials available: almost all materials, such as polymers, metals, and even paper can be used to fabricate triboelectric devices, since the tribo-electrification effect is nearly universal. The general rule to maximize performance is to pick material pairs that have large differences in electron affinity [124]. Fig. 2D summarizes the charge affinity of various common materials which can be used as a guideline for selecting ma-terials for TENG design [124]. A positive charge affinity number in-dicates that the material tends to lose electrons and a negative charge affinity number indicates a tendency to gain electrons. Following this guidance and some other designing criteria such as material’s mechan-ical and chemical properties, stability, availability, and cost, the following materials are commonly used. Organic materials such as nylon, polyethylene terephthalate (PET), PTFE, polyimide (PI), PDMS, and Kapton film are widely used, and inorganic materials such as gold (Au), silver (Ag) copper (Cu), aluminum (Al), silicone (Si), titanium (Ti), indium tin oxide (ITO), and titanium dioxide (TiO2) are employed in many TENG devices. The selection of the material heavily relies on the design purpose of the device. Lin et al. developed a TENG that can serve as both an energy harvester or a sensor for catechin detection using TiO2 and PTFE as the contact layers and Cu and Ti for the electrodes [67]. The

high sensitivity (5 μM) of the device owed itself to the interaction be-tween the catechin and the Ti atom in the TiO2. In order to sense tiny mechanical motion, a lightweight TENG device was made of carbon sponge, nylon, and PET or PVDF [125]. Nylon was chosen because it tended to lose electrons (a large positive charge affinity) when it was in contact with materials such as PET, which had a strong tendency accept electrons (a large negative charge affinity) (Fig. 2D). Gu et al. [126] developed an antibacterial composite film based TENG that could be used insole while effectively prevented the contamination of Escherichia coli and foot fungus. Ag-exchanged zeolite (Ag-zeolite) and poly-propylene (PP) composite films were fabricated to increase the dielectric permittivity and surface charge density of the composite film, resulting in a high voltage output of 300 V. In addition, Tang et al. used liquid-metals such as mercury (Hg) and gallium (Ga) with a frictional layer such Kapton, PTFE and PET. The device yielded a unprecedently large output charge density of 430 μC/m2 and demonstrated an instantaneous high energy conversion efficiency of 70.6% owing to the high contact surface area between the liquid metal and the frictional polymer layer [127]. In another example, thanks to the high surface-volume ratio between the liquid metal and the frictional poly-mer layer, Zhang et al. developed a device that showed a wide range of acceleration of 0–60 m/s2 and a high sensitivity of 0.26 V s/m2 [128].

From the structural design perspective, researchers have developed various mechanical designs for triboelectric devices. To generate higher energy, the basic two-layer structure of TENG is extended to triple layer or multi-layer structures. Chun et al. developed a sandwiched structure TENG device to boost the output performance [129]. The device yielded a sustainable energy output of 1.22 mA current and 46.8 mW/cm2

power density under a low frequency of 3 Hz with an energy conversion efficiency of 22.4%. Another example of a multi-layer TENG device was fabricated by Bai et al., and the device demonstrated its ability to harvest biomechanical energy from human walking by generating an open-circuit voltage of 215 V and a short-circuit current of 0.66 A with a power density of 10.24 mW/cm3 [130]. To harvest energy in full space (independent of the direction of motion), researchers have designed spherical 3D TENGs with single or multiple electrodes [131–133]. A triple-cantilever structured device was designed to harvest ambient energy, and the device produced an output voltage of 101 V and a current of 55.7 μA in open- and short-circuit, respectively [134]. Zhu et al. [135] designed a high performance TENG based on radial-arrayed rotary electrification by using micro-sized sectors in a radial array, and the device produced a power output of 1.5 W at an efficiency of 24%.

To enhance energy generation performance and the conversion ef-ficiency of a triboelectric based device, the surface of the frictional layer is often treated to maximize the triboelectric and electrification effect. One method is to increase the frictional effect between two contacting layers. For example, Fan et al. fabricated PDMS into a pattern of micro- scale line, cubic and pyramid features by using a Si wafer mold, and the device showed an increase in electrical output due to the frictional surface pattern at the micro-scale [136]. Zhu et al. grew aligned nano-wires on polymer film (Kapton) using a dry-etch technique [137]. Their results showed that the device generated an open circuit-voltage of ~110V with an instantaneous power density of 31.2 mW/cm3. To further increase the contact area between the two frictional layers, Wang et al. [138] modified the morphology of both contacting surfaces of PDMS and Al with different patterns, and the overall device generated a significantly enhanced power density of 128 mW/cm3. Other surface morphologies such as nanowires, nanosheets, and microcylinders using different techniques and materials were further developed [116,139, 140]. Tang et al. [127] studied the surface morphology influence on TENG by comparing four different morphologies. They studied Al cyl-inder structures patterned on a Al sheet with the same cylinder height but different diameters and spacings. The results showed that higher output charges and current were generated from a larger specific sur-face. Triboelectric-based EH designs have been intensively explored to enhance overall performance, including materials selections, innovation

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in device structures and working modes, and morphology variations of material surfaces.

2.3. Biocompatible and biodegradable devices

Biomedical devices are commonly used for both wearable and implantable applications [136]. Wearable devices intimately interact with different parts of the human body such as the ankle, neck, wrist, knee, and foot, while implantable devices are embedded inside a human body and are directly and constantly in contact with organs and soft tissues. Both devices are required to be biocompatible to provide a safe user experience. Encapsulation by a biocompatible soft material such as PDMS [108,137], PI [95], and ecoflex [138] can effectively prevent the exposure of piezoelectric and triboelectric EH devices to body tissues and fluids. Dagdeviren et al. examined the adherence, growth, and viability of rat smooth muscle cells (SMCs) on the encapsulated PZT devices, and the results showed that the majority of SMCs are healthy [95]. Yao et al. packaged the TENG into a multilayer film of PDMS and ecoflex and implanted it on the stomach surface of a rat to generate electrical signals in response to stomach movement [138]. The device has been proven to be biosafe through a series of biocompatibility tests including a cell viability test, whole blood chemical analysis, and a pathological test on vital organs.

In some applications, however, we may not always require the de-vices to stay on our skin or within our body. Biodegradable devices are designed to dissolve within a period of time, after which the device is no longer needed and fully degraded. Water is widely used as the solvent to dissolve biodegradable devices, since it is the major component of the biofluid that is everywhere in human body. Liang et al. designed a fast soluble TENG that can be fully dissolved and degraded into environ-mentally benign products within minutes once triggered by water [139]. In their design, polyvinyl alcohol (PVA) and sodium alginate (SA) were used for the contact material, and Lithium (Li) and Al are used for the electrode in a buckled sandwich structure. The device showed a peak voltage of 1.5 V when a finger presses the device, while the entire device can also be completely dissolved in 10 min through a series of chemical reactions. Some devices are fabricated using biocompatible and biode-gradable materials that possess both properties. Zheng et al. developed a TENG based biomechanical energy harvester that can be degraded and resorbed in an animal body based on various polymers that are biocompatible, biodegradable and exhibit a large difference in tendency to lose and gain electrons [140]. These criteria lead to their material selection of poly(L-lactide-co-glycolide) (PLGA), poly (3-hydroxybutyric acid-co-3-hydroxyvaleric acid) (PHB/V), poly(caprolactone) (PCL), and PVA. Pan et al. presented a biodegradable TENG based on a gelatin film and electrospun PLA nanofiber membrane. The device generated a large output open-circuit voltage up to 500 V with a power density of 5 W/m2

[141]. In addition, Li et al. demonstrated that, through in vivo testing, the degradation rate of the biodegradable device can be manipulated by doping Au nanorods (AuNRs) into PLGA, PCL and PLA films which respond to near-infrared (NIR) light due to the AuNRs dopant [141]. However, compared to synthesized commercial polymers which are expensive and contain potentially harmful chemicals, natural bio-absorbable materials are of a growing interest for their low cost, excellent biocompatibility, and sustainability. Wang et al. used com-posite films that are based on Chitosan, a natural biomaterial obtained from chitin, to fabricate devices [142]. Jiang et al. used a variety of natural materials such as cellulose, chitin, silk fibroin, rice paper, and egg white to fabricate a bioabsorbable natural-material-based TENG [143]. It showed that the maximum open-circuit voltage from the device reached 55 V with a power density of 21.6 mW/m2. For biodegradable devices, in general, the electrical output decreases as the material de-grades. However, the degradation rate for different energy materials especially in cardiac application has not been studied yet. In future work, this topic will be further explored in designing degradable and implantable devices. Note that biodegradability is not applicable for

long-term energy harvesting designs, since we expect that the longevity of those EH devices should last for years or even the patient’s lifetime, and thus long-term in vivo studies would be necessary to evaluate the devices’ reliability.

Since the biocompatibility of devices is critical for implantable biomedical devices, encapsulation of piezoelectric and triboelectric based devices using biocompatible materials could effectively isolate devices from body fluidics and tissues in vivo. The additional mechanical changes that this creates, such as increased rigidity, as well as the long- term effectiveness of sealed layer still need to be further investigated.

2.4. Flexible and comfortable devices

Except for bones, almost all other human tissues and organs are soft, and some are stretchable, such as skin and muscle. Our body and its parts undergo constant movement generated both externally, from walking, running, talking, blinking, etc., or internally, from the heart beating, stomach deforming, and lungs inflating. These biomechanical motions not only provide abundant energy that can potentially power electronics but also require the devices themselves to be flexible and stretchable to couple with the motion [11]. Flexible and comfortable devices are developed by either using flexible materials or through smart structural designs.

Due to their outstanding flexibility, polymer-based materials have been broadly studied for biomedical systems [144]. Some interesting examples incorporate Poly(vinylidene fluoride-co-trifluoroethylene) P (VDF-TrFE) cantilevers or buckled beam arrays on pacemaker leads for cardiac EH applications [89,97]. Zhang et al. utilized PVDF to generate power from the pulsing of the ascending aorta [145]. Mesoporous PVDF was developed with a tunable mechanical modulus for in vivo EH ap-plications, such as blood pressure changes [146]. Another interesting device is that of a hybrid biomechanical/biochemical energy harvester utilizing PVDF nanofibers for mechanical EH coupled with an enzymatic biofuel cell for biochemical EH applications [147]. These two func-tionalities work simultaneously, giving the device improved outputs over standalone mechanical or chemical harvesters. In addition to polymer-based piezoelectric materials, researchers have integrated sin-gle crystalline materials into flexible devices. Xu et al. developed PMN-PT nanowires, fabricated using a hydrothermal method, in a flexible polymer (PDMS) composite [148]. The device was able to reach a peak output voltage of 7.8V and output current of 2.29 μA, demon-strating promising applications for these functional materials.

Some common flexible materials for TENG based devices include PDMS, silicone rubber, PTFE, polymers (VHB), and hydrogels. Fan et al. designed a transparent and flexible TENG using PDMS, PET, and ITO which generated a maximum output of 200 V and 7μA [149]. A flexible fiber-based TENG was developed by Wang et al. by using PTFE and carbon fiber coated in PDMS [150]. Yang et al. created a flexible and shape adaptive TENG for self-powered biomedical monitoring [151]. The device was fabricated by sandwiching a wavy-Kapton layer into a top and bottom PDMS layer which exhibited good conformability to human skin. Pu et al. later developed a transparent, ultra-stretchable triboelectric-based energy harvester [152]. The device had a sandwich structure in which the middle layer was Polyacrylamide (PAAm) hydrogel containing lithium chloride (LiCl) ionic hydrogel (PAAm-LiCl hydrogel) and the top and bottom layers were either PDMS or Very High Bond elastomer (VHB). The VHB-TENG demonstrated a mechanical strain of up to 1160% and an average transparency of 96.2% for visible light simultaneously, and its peak power density reached 36 mW/m2, enough to power an electronic watch. Other flexible devices include an ultrathin single-electrode TENG using flexible films of PET, PTFE, and fluorinated ethylene propylene (FEP) [153], and a self-powered tactile sensing device based on the soft materials of PDMS and iongel [154].

In addition, various structural strategies have been successfully in-tegrated into flexible EH and sensing devices for biomedical applica-tions. Buckling designs can sustain large levels of strain without

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breaking due to the unique wavy/wrinkle structures [48,49] with variation in the wavelength and amplitude [155]. Buckling beams are designed to transition from one state to the other state via a snap-through, which allows for the generation of rapid motion [156]. This contributes to large deformations of piezoelectric beams and cor-responding energy generation [96,157]. The serpentine design is another interesting structure which has been widely used in stretchable electronics [158,159]. The strategy utilizes serpentine interconnections, which have a much lower effective stiffness [160,161]. Origami/Kir-igami structures, the art of paper folding, have also been employed to achieve flexible device designs [162,163]. Smart structures can be leveraged to implement flexible EHs and sensors even by using rigid and brittle materials, such as ceramic, semiconductors, and metals.

For both implantable and wearable applications, the comfort of de-vices is an important aspect. A compact, lightweight device is more compatible with normal human activities, and thus enhances comfort. Especially for cardiac medical applications, most current cardiac EH devices are designed to be sutured on the heart. Comfortability over time for the patient needs more evaluation, since no human trials have been performed using those implanted devices [94,95,101]. An ideal comfortable wearable device is cloth-like and can be seamlessly incor-porated into the clothes. Researcher have developed multiple ways to use conventional clothing material such as nylon, polyester, and per-ylene textiles as well as unconventional materials such as PDMS and silicone rubber to fabricate energy harvesters and sensors that can be worn as a part of normal clothes [164–170].

3. Cardiac energy harvesting devices

Due to the unique integration of energy materials with structural designs, various energy harvesting strategies has been developed to convert energy from the beating heart into electrical energy to power biomedical devices. Two major cardiac EH devices are based on piezo-electric and triboelectric energy conversion mechanisms. Note that in this section, the discussion on cardiac EH applications focus on energy sources directly from the heart itself. Other methods of using biofuels and solar energy sources for powering cardiac medical devices, such as pacemakers, are not within the scope of this review.

3.1. Piezoelectric based cardiac EH devices

Various piezoelectric materials have been investigated for cardiac EH strategies in vivo. A shaped polyethylene terephthalate (PET) skeleton-based EH device was recently designed by using a single crystal PMN-PT film [107]. The device shown in Fig. 3A directly powered a cardiac pacemaker by harvesting the natural energy of a heartbeat without using an external energy storage element. Lee’s group in 2017 utilized single crystalline PMN-PZT to fabricate a self-powered wireless transmission energy harvester that was implanted on the outer wall of a porcine heart [171] (Fig. 3B). The device produced a voltage of 17.8V and a maximum current of 1.74 μA, larger than similar piezoelectric devices for heart-based energy harvesting and showing the promise of these materials. Making use of the high piezoelectric constant (d33), the same group also transferred PMN-PT film to a flexible substrate using the inherent residual stress of a Nickle (Ni) film to pull the crystal from

Fig. 3. Reported piezoelectric based cardiac EH strategies. (A) A shaped PET skeleton based PMN-PT EH device directly powered a cardiac pacemaker by harvesting the energy of a heartbeat [107]. (B) A PMN-PZT energy harvester was fabricated and implanted in a porcine heart to drive wireless communication [171]. (C) A PMN-PT EH device was developed to provide artificial beating for a live rat [172]. (D) PZT ribbons were designed for cardiac and lung applications via the conversion of the mechanical motion from the organs [95]. (E) The PZT cardiac EH devices sutured on different positions of the heart were used to harvest the biomechanical energy of the heart [94]. (F) A helical-designed compact EH device integrated with an existing pacemaker lead was developed to convert pacemaker lead motion to electrical energy [108].

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the growth substrate. The device was able to turn on 50 light emitting diodes using a finger tapping input and was used to perform real-time electrical stimulation to provide artificial beating for a live rat [172] (Fig. 3C). Due to the high piezoelectric coefficients of PZT, it has also been widely explored as a material for energy harvesting in vivo. In 2014, Dagdeviren et al. utilized PZT ribbons to develop a mechanical energy harvester for cardiac and lung applications [95] (Fig. 3D). In vivo testing involved affixing the device on the epicardial sites of the right ventricle, left ventricle base, and free wall of the hearts of bovine. The devices were sutured at three points. Different suture orientations were tested and a peak to peak voltage of up to 8 V was found, but it was directly related to the animal hearts’ size, as the smaller ovine model showed a peak-to-peak voltage closer to 4 V [95]. A similar EH design (Fig. 3E) using PZT was tested in a swine model with a peak-to-peak voltage of 3 V by varying the fixture positions from the left ventricular apex to right ventricle [94]. Although PZT and PMN-PT/PZN-PT have been used in multiple in vivo based cardiac energy harvesters to great success, a major cause for concern with those materials and their de-rivatives is the high lead content. As in many industries, researchers are trying to move away from the use of lead in materials and thus, much work has been done on searching for the next candidate to replace these lead-based piezoelectric materials. Due to their extremely high piezo-electric constants, PZT and PMN-PT have remained relevant in many applications.

Moreover, Wang et al. showed the first example of using ZnO based devices to harvest energy from a biological source [173]. Using a connection of four single wire generators, the mechanical movement of a human finger and the body motion of a live hamster were converted to electrical energy. A maximum of 0.1 V peak output was achieved. In many cases, ZnO is seen as too rigid a material for interfacing with biology for EH and sensing applications and so it is often combined with other materials to form more flexible devices. Polymer-based piezo-electrical material of PVDF has been employed for cardiac EH devices due to its excellent biocompatibility [174] and processability with mi-crostructures [146]. The porous structures within the materials and electromechanical coupling efficiency were tuned via humidity control through a film solidification process [70,72]. Using those porous P (VDF-TrFE) materials, some recently reported EH strategies (one example shown in Fig. 3F) did not contact the heart but could efficiently convert the energy of the heart into electrical power without interfering with the cardiovascular function [97,108,110]. Those EH designs were integrated into part of the existing pacemaker lead and otherwise with no direct contact of heart. Long-term animal studies will be needed to better understand the performance of EH devices within the heart.

A comparison of the reported piezoelectric based cardiac EH devices is listed in Table 2. In vivo energy harvesting performance of the various

devices has been compared in terms of voltage/current generation. Most in vivo EH studies are performed in large animal models, such as porcine [94,107,110,171], bovine and ovine [95], since they have similar sizes to human hearts. The essential energy harvesting mechanism for those reported devices is to harness cardiac motion or pacemaker lead motion into strain within the piezoelectric materials to generate electrical charge. Long-term evaluation of EH devices has been intensively tested in vitro [95,107,110,171] over 104 cycles but with very few reported in vivo results [94].

3.2. Triboelectric based cardiac EH devices

In recent years, researchers have advanced the development of TENG based devices into the next stage that involves animal in vivo tests for cardiac energy harvesting. Those in vivo experiments, allowing for more clinical evaluation on the possibility for TENG to be implanted in human body, have demonstrated great potential for self-powered biomedical electronics and physiological signal monitors. In 2014, Zheng et al. demonstrated the harvesting of biomechanical energy in vivo from a TENG device for the first time [175] (Fig. 4A). The device consisted of a PDMS film with patterned pyramid arrays buried under the thoracic skin of a living rat, and it converts the periodic expansion and contraction of the thorax (relating to the respiratory movement) of the rat to electrical energy. Their results demonstrated that the energy collected by the TENG from the breathing movement of a rat could be stored in a capacitor large enough to power the pacemaker. Zheng et al. fabricated an implantable TENG consisting of a multi-layer structure for in vivo EH. The voltage generated from the implanted device which were attached to the heart of a living swine, reached up to 14 V in response to the heartbeat [100] (Fig. 4B). In vivo EH performance was evaluated for over 3 days while the electricity was generated continuously in the active animal model. A self-powered wireless transmission system was fabri-cated for real-time wireless cardiac monitoring via the electrical signal associated with the in vivo heartbeat [100]. The same group developed a biodegradable TENG as a power source for neuron-repairing applica-tions [140] (Fig. 4C). The biocompatibility of the device was verified via a cell viability test by culturing endotheliocytes (ECs) with their syn-thesized biocompatible films (i.e. PLGA and PCL films). The TENG was buried in the subdermal part of the backs of rats, and the voltage output was recorded while a small finger tapping force is applied at the implanted site. The output voltage decreased from 4 V to 1 V as the PLGA-coated device degraded and from 3 V to 0 V as the PVA-coated device degraded. More recently, a symbiotic cardiac pacemaker sys-tem (Fig. 4D) was developed and was driven by the harvested energy from the TENG device directly [101]. An in vivo study was performed in a porcine model, and the implanted TENG device generated up to 65.2 V

Table 2 Comparison of reported piezoelectric based cardiac EH strategies in vivo.

References and publication Year [107] 2019

[171] 2017

[94] 2015

[95] 2014

[172] 2014

[110] 2019

Materials PMN-PT PMN-PZT PZT PZT PMN-PT P(VDF-TrFE) Structures Compressed “8”

shape Patch Ribbon Ribbon Patch Helical ribbon

Mechanism Cardiac motion Cardiac motion Cardiac motion Cardiac motion Cardiac motion Pacemaker lead motion

Thoracotomy Yes Yes Yes Yes Yes No Max. in vivo open circuit voltage

(V) 20 17.8 3 1.2 μW/cm2 (stacks of 5

EH devices) 2.7 μJ (artificial stimulation)

2.1

Max. in vivo Short circuit current (μA)

15 1.74 – –

Long term performance (cycles) in vitro/in vivo

5*105 105 Drop at 56% at 2 h (in vivo)

2*107 3*104 104

Animal model Porcine Porcine Porcine Bovine, Ovine Rat Porcine Self-power pacemaker Yes No No No No No Other applications – Wireless

transmission – – Power a stopwatch –

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open-circuit voltage. The energy scavenged from each heart beating cycle was 0.495 μJ, which was large enough to directly power the pacemaker with an endocardial pacing threshold of 0.377 μJ.

Jiang et al. [143] proposed a bio-degradable TENG device using all-natural materials such as cellulose, chitin, silk fibroin, rice paper, and egg white that could be absorbed by rat tissue. The fully bioabsorbable device facilitated implantation recovery and was used as a power source for regulating the beating rates of dysfunctional cardiomyocytes, providing a possible solution to treat cardiac diseases such as arrhythmia and bradycardia. Shi et al. designed a packaged self-powered system based on a triboelectric and piezoelectric hybrid nanogenerator [176]. In the in vivo test, an electronic thermometer powered by the EH system detected subcutaneous temperature of a live rat through an implanted temperature transducer. Note that in that design, the source of the en-ergy was not from the rat directly, but from other motions applied to the hybrid nanogenerator. Recently, a transcutaneous ultrasound implant-able EH device using capacitive triboelectric technology was developed to power medical implants [177]. In that work, the ultrasound induced micrometer-scale displacement of a polymer thin membrane to generate electrical power through contact electrification (Fig. 4E). The device successfully recharged a lithium-ion battery at a rate of 166 micro-coulombs per second in water. Although, the proposed implantable triboelectric generator (VI-TEG) was not cardiac based device, such unit could harvest ultrasound in vivo for powering IMDs.

The reported triboelectric based cardiac EH devices demonstrate to generate sufficient power to self-power a cardiac device [100,101]. These devices have a similar multilayer patch structure and are based on the vertical contact separation mode of the triboelectric effect. The vertical contact separation mode of triboelectric devices is dominantly employed in cardiac EH applications for two reasons. The main reason is due to the contraction and relaxation of the heart motions, which are consistent with the relative vertical contact-separation movement, while the sliding movement in the contact-sliding and freestanding mode are less effective in energy conversion. In addition, it is advantageous over the single electrode mode (which also relies on the contact-separation

movement), since there is no need to be grounded, and thus is more suitable for implantable applications. Various materials are used for these devices, but they all share PDMS, PTFE, Al, and Kapton films. Sternotomy or mini-thoracotomy are used for the implantation surgery on porcine models. The maximum voltage and the current of these de-vices ranges from ~14 V to 65.2 V and 0.5 μA to 5 A, respectively, with a long term performance ranging from a few hours to 3 days [100,101].

3.3. Summary of cardiac EH devices

Engineers have developed various configurations using both piezo-electric materials and broad triboelectric materials to convert biome-chanical energy from organs such as the heart into electrical power for cardiac applications. Various engineering approaches have been explored for cardiac energy harvesting, and a summary of reported ap-proaches is listed in Table 3. There are five reported cardiac EH ap-proaches: piezoelectric, triboelectric, mass imbalance oscillation, electrostatic, and electromagnetic EH mechanisms. Piezoelectric and triboelectric EH strategies generate levels of electrical output high enough to meet the requirement of medical devices. As shown in Fig. 5A, the cardiac EH devices reach high energy generation in vivo [94,95,107, 108,110,171,172], especially, TENGs generate voltages above 10 V [100,101], which are much higher than typical medical devices opera-tion requirements of 2–3 V. For triboelectric EH approaches, the per-formance of the devices mainly relies on the contacting and separation between the layers, which limits the effectiveness of long-term power generation performance. Almost all reported implantable cardiac EH designs have not reached clinical translation, and no human trials using these implanted devices have been reported to date. Additional open chest surgeries for the placements of these EH devices are always required [94,95,101,107,171,172], and thus it may cause potential risks of these procedures to the patients. These suture-on-heart surgeries inevitably increase the risk of complications and infections. From a clinical perspective, an ideal cardiac energy harvesting strategy does not require any thoracotomies and is compatible with current clinical

Fig. 4. Reported triboelectric based EH devices. (A) A schematic of TENG drove a pacemaker prototype to regulate the heart rate of a rat from animal breathing [175]. (B). A triboelectric EH device attached to a pig’s heart outer wall converted heartbeat motion to electrical energy [100]. (C). An implantable, biodegradable TENG tested in a rat model for neuron-repairing application [140]. (D). A symbiotic pacemaker system was self-powered by TENG via the motions of a pig’s heart [101]. (E) A transcutaneous ultrasound based implantable EH device utilized capacitive triboelectric technology [177].

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procedures for patients. Recently, researchers propose to seamlessly couple the EHs on the pacemaker lead to convert the mechanical energy of the lead into electrical power [108,110,178]. Those energy harvesting strategies avoid the direct contact with the heart, and thus are promising for clinical translation in human trials.

In addition to piezoelectric and triboelectric EH methods, a me-chanical approach was developed by using mass imbalance oscillation. The device generated a power of 90.1 � 0.7 μW (basal) by fixing it onto a pig’s heart directly [179]. An electrostatic EH concept was proposed to use a honeycomb type variable capacitor with coil springs; however, the overall device was too large to fit into the thoracic cavity of the animals

[180]. An additional electromagnetic approach was developed by using neodymium permanent magnets with a flux density of 1.43 T and an array of copper coils. The results showed that the electromagnetic EH device generated an average power output of 0.78 μW at 84 beat per minutes (bpm). However, that electromagnetic device could only pro-vide short-term protection from body fluids. From clinical side, the induced magnetic field may interfere with the functions of the IMDs and thus cause potential safety risks for patients.

In addition, the cardiac EH device itself need to be biosafe in the body and patients need to be comfortable. The biocompatibility of EH devices has been intensively tested and evaluated in the literature [100,

Table 3 Comparison of different cardiac EH methods in vivo (adpated from Ref. [108,110]).

Mechanisms Materials and components Thoracotomy and devices anchoring approaches Animal models Power (voltage/current) or power density

Piezoelectric P(VDF-TrFE) [110] Porcine 4.5 μW

PZT [94,95] Swine [94] Bovine [95]

0.18 (μWcm� 2) [95]

PMN-PT [107,172] Porcine [107] Rat [172]

12 V, 15 μA [107] 2.7 μJ [172]

PMN-PZT [171] Porcine 17.8 V, 1.74 μA ZnO NWs [103] Rat 2 mV, 4 pA

Triboelectric Broad material options [100,101] Porcine 14 V, 5 μA [100] 65.2 V, 0.495 μJ [101]

Mass Imbalance Oscillation

A clockwork and an electrical microgenerator [179]

Porcine 82.0 � 4.4 μW (apical) 90.1 � 0.7 μW(basal)

Electrostatic Honeycomb-type variable capacitor with coil springs [180]

None (concept design)

36 μW (in vitro)

Electromagnetic Permanent magnets and an array of coils [181]

Porcine 0.78 μW (84 bpm) 1.7 μW (160 bpm)

Fig. 5. Summary of cardiac EH performance, stability and biocompatibility evaluation. (A) Comparison of reported cardiac EH devices performance in terms of in vivo voltage generation and stability evaluation of EH devices. (B) Fluorescence images of L929 cells on encapsulation layers of EH devices and normalized viability of L929s for 3 days [101]. (C) Fluorescence images of neonatal rat cardiac myocytes on encapsulation EH device and beating frequencies of cardiac myocytes from day 2 to day 5 [107].

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101,107,110,143]. Cytotoxicity and biocompatibility of EH devices were evaluated by growing the mouse fibroblasts (L929s) on the encapsulation layer of materials and the cell culture dish as a control group [101]. The results demonstrated similar cells spreading and cellular structures for the device group and the control group, and no significant difference in cells viability was observed [101] (Fig. 5B). Neonatal rat cardiac myocytes and human pericardial fibroblasts were further used as cellular models for evaluation of biocompatibility. It was found that the cardiac myocytes started beating on day 2, and no sig-nificant differences in the beating rates were shown in the two groups during the following days observation [107] (Fig. 5C).

The long-term performance of devices is critical in biological sys-tems. However, most reports demonstrate short time operation (a few hours) in live animal models [94,95,100,110,172]. A comparison of the stability evaluation of reported implantable cardiac EH devices in recent years is also demonstrated in Fig. 5A. Most reported devices were tested over 104 cycles in vitro [95,107,110,171], and there were very few re-ported in vivo long-term performance results [94,100]. One of the re-ports showed that as the time elapses the peak-to-peak voltage generated by the implanted EH devices decreased by 10% at 40 min and drops by 56% in magnitude at 2 h [94]. Zheng et al. [100] tested the TENG after 24 h of implantation. However, an obvious decreased voltage output was also found, and they concluded that might due to weakened cardiac function related to surgery. Due to the triboelectric mechanism, almost all TENGs in cardiac applications are directly contacted with organs, such as hearts, therefore, long-term animal studies will be needed to determine the effectiveness of such contact. As for piezoelectric based cardiac EH devices, the fatigue behavior of the piezoelectric and ferro-electric materials and the decreasing polarization upon years’ cycling still needs verification through long-term animal studies.

4. Cardiac sensing devices

In addition to harvesting energy from the motion of the heart and internal organs, implantable piezoelectric and triboelectric devices can function as self-powered pressure sensors that monitor important physiological and pathological characteristics. Reported implantable piezoelectric and triboelectric sensors utilize similar materials and integrate with flexible structures [182–186]. Due to similar transduction mechanisms, a single device could carry out both sensing and EH functions. Cardiac sensors are normally required to be sensitive, biocompatible, lightweight, and long lasting. Implantable pressure sensors are a subset of a broad and rapidly growing field of flexible cardiac sensors. In vivo studies have reported piezoelectric and tribo-electric pressure sensors that monitor heart rate, endocardial pressure, and arterial blood pressure. The point-measurement signal generated by these sensors can be used to detect heart rate, blood flow rate, and cardiac pathologies such as atrial fibrillation and ventricular premature contraction. Related types of cardiac sensors include wearable pressure sensors, which detect pulse rate and blood pressure by interfacing with the skin in regions of the body such as the neck, wrist, and chest, and epicardial mapping devices, which generate spatiotemporal data by interfacing with large regions of the epicardium.

4.1. Implantable sensors

Implantable cardiac sensors generate valuable information about the heart’s behavior, such as heart rate, endocardial and arterial blood pressure, and pulse wave velocity. These sensors detect changes in pressure by directly interfacing with the heart wall or artery, either attached with sutures or laminated on the cardiac tissue using conformal elastomeric substrates such as PDMS [182], Poly-L-lactide (PLLA) [183] or POMaC [184]. Liu et al. recently measured endocardial pressure in the left ventricle and atrium of the pig’s heart using a small nPFTE TENG

Fig. 6. Reported implantable pressure sensors. (A) A triboelectric sensor was implanted in the left ventricle of a pig’s heart with a heparin-coated polyvinyl chloride (PVC) stretchable catheter to detect ventricular pressure [182]. (B) A PLLA sensor was used for in vivo force measurement within a mouse abdominal cavity [183]. (C) A biodegradable sensor monitored post anastomosis blood flow in the femoral artery with a bilayer coil structure for wireless data transmission [184]. (D). A PZT piezoelectric sensor laminated on explanted bovine organs was developed to measure the modulus of tissue samples [185]. (E). A schematic of working mechanism of an nPFTE triboelectric active sensor used to detect ventricular endocardial pressure variations and heart rate [186].

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sensor implanted through a catheter [182] (Fig. 6A). The measured pressure showed a good sensitivity of 1.195 mV/mmHg. The proposed sensor could also detect arrhythmias of the heart such as ventricular fibrillation and premature contraction. These studies have established clinical significance of conformal piezoelectric systems for rapid and non-invasive characterization of skin mechanical properties. Curry et al. calculated respiration rate using a biodegradable PLLA piezoelectric sensor sutured to the abdomen that measured the pressure of dia-phragmic contraction [183] (Fig. 6B). Implantable sensors were also reported to sense blood pressure in arterial veins as well as regions of the heart. Boutry et al. detected the change in diameter of the femoral vein of a rat using a biodegradable capacitive poly(glycerol sebacate) (PGS) sensor and determined arterial occlusion and pulse rate [184] (Fig. 6C). A sensor comprised of an array of PZT sensors and actuators detected changes in the stiffness of bovine and ovine explanted epicardial tissue [185] (Fig. 6D). An nPFTE triboelectric sensor in Fig. 6E was sutured onto the pericardium to detect heart rate and changes in ventricular endocardial pressure indicative of atrial fibrillation and premature ventricular contraction [186].

Various implantable sensing devices have also been developed to detect and monitor the heart rate. Sato et al. developed a simultaneous monitoring system for the heart rate and breathing rate of anesthetized mice [187]. They used custom-designed analogue circuitry and a microprocessor program for high reliability sensing, leading the way to applications in monitoring of animals in laboratory settings [187]. In another example, Tseng et al. used a sol-gel process to deposit PZT thin films onto a flexible stainless steel substrate. The tactile sensor was used to measure human pulses including carotid, brachial, finger, ankle, radial artery, and the apical region with a sensitivity of 0.798 mV/g [188]. This device could lead to the ability to diagnosis hypertension

and cardiac failure in patients. A TENG-based self-powered wireless cardiac monitor developed by Zheng et al. could successfully generate voltage in response to the respiratory cycle and heartbeat in an in vivo test [100].

4.2. Wearable and epicardial sensors

Wearable cardiac sensors are the largest and most frequently re-ported group of flexible cardiac pressure sensors. The piezoelectric and triboelectric materials in wearable sensors are similar to those in implantable sensors and generate electrical signals that used to detect cutaneous pressure [189], pulse [190,191] and heart rate [192,193]. Device designs can be modified to sense particular properties and filter out certain types of environmental noise [17]. Many wearable sensors monitor the heartbeat by sensing blood pressure through the surface of the skin, and measure blood flow with a similar sensitivity to implant-able cardiac sensors despite lower voltage outputs. Dagdeviren et al. described a PZT piezoelectric sensor that detected transient changes in blood pressure through the wrist due to arm cuffing [189] (Fig. 7A). A bending-sensitive single electrode PDMS triboelectric sensor using a trench structure was reported to detect incident and reflected pulse waves through contact with the wrist [190]. Another triboelectric pulse sensor shown in Fig. 7B was reported to detect cardiac disease indicators through the contours of the pulse wave, such as irregular R wave in-tervals, heightened blood pressure in the right ventricle, and abnormal pulse wave velocity through the peak delay of two wearable sensors [191]. The outputs from wearable sensors matched heart rate mea-surements obtained using commercial pressure electrocardiogram (ECG) and photoplethysmography (PPG) sensors [191]. In that work, the sensor transmitted the data to a mobile device via a Bluetooth chip

Fig. 7. Reported wearable and epicardial sensors. (A) A square array of piezoelectric thin-film transducers based on thin films of PZT was used for cutaneous pressure monitoring [189]. (B) A schematic of multilayer triboelectric sensor detected cardiac disease and the real-time signal outputs when the devices were placed over the finger and radial artery [191]. (C) A schematic of triboelectric nanogenerator by body sensor network for heart rate monitoring via cell phone through a Bluetooth module [192]. (D) Photograph of wireless pressure sensor for heart rate monitoring by using ZnO/PVDF hybrid film [83]. (E) A 3D multifunctional integumentary membrane integrated on a rabbit heart with various sensing function elements in the system [194].

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on the thin film device. The sensitivity of wearable sensors could be tuned for specific applications. Similar, Lin et al. developed a wireless wearable heart rate monitor system powered by human biomechanical energy [192]. The system could convert the inertia energy of walking to electrical output with a maximum power of 2.28 mW and a 57.9% conversion efficiency. The acquired continuous heart-rate signal could be sent to a cell phone through a Bluetooth module and accessed by the user easily [192] (Fig. 7C). Jiang et al. reports a PVDF piezoelectric sensor that simultaneously detects subtle arterial pulse using a more sensitive bending unit and respiratory rate through a less sensitive unit [193]. Other materials, such as ZnO nanoneedles, were fabricated with a PVDF film to form a hybrid pressure sensor. Researchers used this sensor to detect heart rate from the carotid and radial arteries, with a detectable pressure as low as 4 Pa [83] (Fig. 7D). These flexible wearable devices represent a promising trend in short-term, low cost, cardiac monitoring sensors.

A spatial map of sensor data throughout the epicardium generates important information used to categorize cardiac function and disease experimentally and clinically. Many implantable cardiac sensors capture three-dimensional dynamics using an integrated mesh of sensors covering a large portion of the surface area of the heart. The strain sensing network described by Xu et al. incorporated ECG, pH, temper-ature and optical sensors in integumentary membrane structures. The device obtained three dimensional pixelated measurements, and addi-tionally provided structural support to the heart wall [194] (Fig. 7E). The majority of epicardial sensors measured electrical potential as opposed to mechanical stress, using piezoresistive materials such as silicon [195,196] and silver nanowires [197]. Fang et al. measured the electrical potential of regions of the epicardium using a small thin-film network of 96 silicon nanomembrane transistors [195]. Kim et al. re-ported an electrical sensor integrated on the surface of a balloon catheter that detected electrical potential using conductive PDMS and conductive silicon transistors [196]. While piezoelectric and triboelectric materials were not applicable for electrical potential sensing, the combination of this epicardial ECG data with piezoelectric and triboelectric pressure sensors could better determine the root cause of cardiac abnormalities [198].

Advantages of wearable sensors include low fabrication cost, ease of wireless data transmission [192,199,200] and noninvasive application via elastomeric lamination instead of sutures [189–191]. Multiple wearable sensors across the body can be used to generate flow mea-surements, particularly pulse wave velocity [189,191]. Since wearable sensors do not monitor internal organs, they do not generate a conclu-sive picture of how the heart moves and are subject to unpredictable environmental wear over time. In future work, an integrated network of piezoelectric or triboelectric sensors encapsulating the epicardium could potentially generate dynamic flow measurements and would be an im-pactful development in the field.

4.3. Summary of cardiac sensing devices

Implantable cardiac piezoelectric and triboelectric sensors generate abundant biological information as self-powered heart rate monitors. Those implantable sensors are sustainably self-powered from internal biological motion, and do not require the high power output of energy harvesters used to charge batteries and power small scale devices. Re-ported sensors can detect subtle cardiac events and could potentially be integrated in a network to spatially map blood flow in regions of the heart and circulatory system. Reported biocompatibility experiments demonstrate that these implantable cardiac sensors could last for years in the body, however, in vivo studies should further categorize the long term wear of the device, interfacing of heart functions, and the and ability of elastomeric substrates to remain attached to the heart wall [17].

The structures and sensing mechanisms of wearable sensors are similar to implantable devices. Wearable cardiac sensors are comprised

of flexible structures that interface directly with the skin instead of the heart wall. Wearable sensors are limited to arterial pressure measure-ments, though avoid invasive implantation methods that restrict the clinical viability of implantable sensors, and frequently report seamless wireless data transmission and data analysis. Wearable sensors are thereby effective as short term peripheral blood pressure monitors, especially in vulnerable patient populations such as elderly patients and neonates that are not eligible for implantation surgery. In both wearable and implantable sensors, triboelectric mechanisms demonstrate superior sensitivity and pulse wave consistency, though may be better suited for short term purposes given the frictional wear of contacting surfaces in the triboelectric mechanism. Electrical sensors and epicardial mapping devices generate flow measurements (pressure, heat, electrical poten-tial, pH) to detect heart rate and indications of cardiac disease. While reported epicardial mapping devices are not comprised of piezoelectric and triboelectric sensors, a multiplexed network of cardiac sensing de-vices could give dynamic 3D information of the heart’s motion and blood flow.

Sensitivity is an important metric of implantable pressure sensor performance and is formally defined as the magnitude of the voltage response due to a change in pressure [182]. However, sensitivity metrics for piezoelectric and triboelectric cardiac sensing are not standardized; some groups characterized sensitivity using a force voltage calibration [183,184] while others measured in vivo peak signals for heartbeat, blood pressure, and respiration events [182,186,201]. Longevity and biocompatibility are necessary conditions for clinical viability and are a more standardized point of comparison for reported implantable cardiac sensors. A summary of these parameters for both implantable and wearable cardiac sensors is listed in Table 4. The considerable power generated by some implantable sensors [183,186] is within the reported range of EH devices [95,101,172] and suggests that a single device can generate energy to charge battery powered implantable devices while monitoring heart and respiratory rate. These studies report promising in vitro longevity, and retained function over days in vivo, though do not confirm long-term clinical longevity. Triboelectric based devices generate higher power and sensitivity than piezoelectric based devices across this group of implantable sensors, however, the longevity of triboelectric in body systems has been called into question given the frictional wear of the triboelectric mechanism [198]. Biocompatibility requirements vary depending on the tissues and fluids in contact with each sensor, though immunohistochemical staining, Masson’s trichrome stain, and Hematoxylin� eosin staining remain standard tests in the field. Long-term in vivo studies would better characterize the biocom-patibility result of device wear on surrounding tissue. Recently, biode-gradable arterial and cardiac sensors have emerged as a method of sensitive blood pressure monitoring that avoids removal surgery and the biocompatibility challenges of long term implantation [184]. The sensing application of piezoelectric and triboelectric materials could be integrated with EH devices without reducing the electrical power generated, though spatiotemporal blood flow measurements remain a challenge. Other new directions in cardiac sensing include wireless data transmission from inside the body, and in-sensor analytics to detect disease characteristics [202]. Another recent work reported piezoelec-tric polymer microsystems for applications in EH and multifunctional sensors for robotic prosthetic interfaces with improved responsivity, and in bio-integrated devices [203]. The application of piezoelectric and triboelectric materials to energy harvesting and sensing enables a promising range of opportunities for biomedical device innovation [204, 205].

5. Conclusions and perspectives

Cardiac implantable devices are some of the most effective methods for helping control irregular heartbeats in people with heart rhythm disorders and for interventional therapy for cardiac disease. However, two key limitations of these cardiac devices and many other IMDs are

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device failure and longevity. In general, IMD failure often comes in one of two forms. In the first form, the cardiac device is either broken before surgery, during surgery, or soon after surgery, which is indicative of a reliability problem with the device, either in the manufacturing process or the implantation procedure. Manufacturing reliability can be improved by following best practices, such as design-for-manufacture (DFM) or design-for-assembly (DFA) [206]. All devices should go through both finite element analysis and rigorous real-world testing to determine possible failure modes and to improve device design if necessary. Other failure events can be caught by following Failure Mode Effects Analysis, especially failures not associated with mechanical is-sues. The other form of failure is long-term failure, in which the device fails over time, or works for a significant time and then starts to fail. This is often because of material deterioration, as the body is often a chal-lenging and corrosive environment. Materials that held up to significant stress testing outside the body may be slowly degraded by a combination of chemical and mechanical stress. In other cases, devices may be tested at a certain number of mechanical cycles but will be in the body for far longer. It is impractical to perform mechanical tests for the full lifetime of the device, so approximations must be made. The best way to limit these failures is to stick with materials and designs that have been previously shown to be effective. Failing that, first-principles analysis of chemical, thermal, and mechanical durability can help to limit failure potential. Devices which function for the majority of people may fail for a select few – in this case, it is important to analyze those patients with failures to attempt to determine a unifying cause.

Longevity is a critical factor in measuring cardiac device efficacy. Even a successful device may only be intended to work for a certain number of years. If it functions effectively throughout its rated lifetime and then stops working, most will not consider it a failure. Pushing the longevity of IMDs higher requires many of the same analysis steps as preventing long-term failure but requires additional electrical design and analysis. An ideal IMD will last significantly longer than the ex-pected lifetime of the patient – the goal is to prevent or limit re- implantation or replacement surgeries, as these expose patients to a great deal of medical risk as well as costs. Most longevity issues for powered devices result from a lack of battery capacity. Either the battery itself is small, or the device draws too much power. Improving overall energy usage can take one of several forms. The battery size or density can be increased, which allows for a direct increase in longevity. In general, this approach has several issues – larger batteries will often not fit or will place undue stress on the body, and more efficient batteries are naturally more dangerous and generate more heat. Lowering the power requirements of the device is always an attractive option but is limited by the technology available and the base requirements of the device. For example, a sensor needs enough power to transmit, record, and amplify a signal, while something like a pacemaker needs to be able to generate a signal of a specific voltage and current. The third method of increasing longevity involves attempting to either recharge the battery or provide an alternative source of power. This power source can come from har-vesting chemical or mechanical processes in the body, or through wireless power transmission from an external source.

Since the heart is a compelling in vivo energy source and is a natural battery to power IMDs, self-sustainable energy generation could signif-icantly extend the lifetime and effectiveness of IMDs. Piezoelectric and triboelectric EH and sensing technology can be leveraged to infinitely power or replace life-saving commercial cardiac devices. With recent advances in energy materials, researchers have developed sustainable energy strategies for implantable cardiac devices, among which the most prominent two are piezoelectric and triboelectric based energy har-vesters and sensors. These devices can harvest biomechanical energy from the body and convert into electrical energy that can be used to power IMDs. Piezoelectric energy harvesters are built on piezoelectric materials which exhibit the piezoelectric effect, such that the material will generate electricity when subject to mechanical stress or strain due to electric polarization in the material. The most widely used piezo-electric materials for IMDs include ceramics, single crystals, and poly-mers. Triboelectric based devices are another excellent candidate for EH and sensing applications due to their high energy conversion capability and abundant choice of materials. Almost any pair of materials with a difference in charge affinity results in a combination of tribo-electrification and electrostatic induction between two contacting materials.

Strategies of cardiac energy harvesting and sensing have been intensively developed in alternative power and diagnostic solutions for IMDs. However, there are still certain challenges to address for optimi-zation of the device’s performance for clinical translation ultimately. Materials and devices biocompatibility are critical issues for EH and sensing devices design. Certain materials have either short or long-term toxicity effects on the body and must be avoided when designing implantable devices. This can sometimes significantly restrict the design space of the device, though current research is pushing the boundaries of what can be done with biocompatible materials. Non-biocompatible materials can be used in IMDs, but only if they are securely enclosed within a protective barrier, and designs such as these would be open to increased regulatory scrutiny. A common issue with devices is that metal wires, electrodes, and integrated circuits need to be fully isolated from the body both for biocompatibility and for function. An increasingly common method of encapsulation uses the polymer of PDMS, which has been shown to be safe for use in IMDs and can provide a robust hydro-phobic seal.

In order to be clinically translatable, a cardiac EH device must be able to operate in a highly-reactive environment for many years, all while constantly being deformed/undeformed or contacted/uncon-tacted. Mechanical biocompatibility must be considered – certain de-signs will place undue strain on the bodily functions and can negatively impact organ or musculoskeletal functions. This is often an important consideration for heart-adjacent IMDs especially and tests where the device is unpowered will often be conducted on animal models to verify that the device does not impact function. Further research will also be needed to evaluate the level of degradation of the electrodes, reactions between the materials in the body and the materials of the EH that could affect performance, and any possible adhesion or delamination. Addi-tional considerations for cardiac EH and sensing devices also ensure that

Table 4 A summary of reported cardiac implantable and wearable sensors.

References /publication Year [182] 2019

[183] 2018

[184] 2019

[186] 2016

[189] 2014

[191] 2017

Key material n-PTFE PLLA PGS n-PTFE PZT Cu and Kapton

Mechanism Triboelectric Piezoelectric Capacitive Triboelectric Piezoelectric Triboelectric Reported sensitivity 1.195 mV/mmHg 0–18 kPa – 17.8 mV/mmHg ~0.005 Pa 45 dBa

Long term performance (cycles) in vitro/in vivo

108 104 3000 3:5 *105(in vivo) 103 107

Application Epicardial blood pressure

Respiratory rate, pressure

Arterial pressure

Heart rate, blood pressure

Arterial pressure

Arterial pulse

a The peak signal to noise ratio.

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the device operates well at a range of driving frequencies, including repeated shifts between lower and higher frequencies.

A fully functional energy harvester for cardiac pacing or any other biomedical device consist of three primary components: the energy harvesting devices; the power management/storage electronic and the load, for example the pacing component. Many energy harvesting studies focus on the energy harvesting materials and devices, which we have primarily focused on in this review, however, the power manage-ment and storage aspects are critical for the implementation of these devices in practical applications. In an in vivo environment, the con-verted energy from these devices tends to be unstable and thus difficult to use as a direct power source for electronic devices, so it must be stored, typically in rechargeable batteries or supercapacitors. Piezo-electric and triboelectric generated power have alternating current (AC) behavior, high voltages, low currents, and extremely high internal im-pedances, all of which can complicate the efficient storage of the power.

A standard power management circuit consists of rectification of the AC output signal, followed by a direct current (DC) conversion to bring the voltage within the charging specifications of the storage component. Standard rectification circuitry is unable to optimize the energy utili-zation efficiency and coupled with a mismatch in impedance of the energy harvester and the storage component, leads to large losses in power. Others have thoroughly reviewed various circuit topologies for energy harvesters and energy storage components to better improve the voltage regulation and to optimize power extraction [207]. For biomedical applications, these storage components will have similarly desired material properties to the energy harvesters, such as flexibility and a small size. The development and the integration of these energy harvesting and storage devices have also been reviewed [208,209]. Some of the interesting approaches include, integrating piezoelectric energy harvesters and lithium ion battery energy storage in a single coin-type cell [210] and the integration of fabric-like triboelectric har-vesting devices directly with flexible batteries for a more seamless design for wearable electronics [166].

In addition to energy harvesting, piezoelectric and triboelectric based devices can also be integrated with self-powered cardiac sensors or monitors that can directly measure vital physiological signals such as heart rate, blood pressure, and respirator rate. Moving forward, long- term animal studies will also be needed, in which the animal recovers and is allowed to undergo normal physical activity over the course of weeks, months, and years to understand the impact of daily life on the device’s stability and energy storage. Biomedical self-sustainable energy generation represents a new frontier to greatly extend the lifetime and effectiveness of IMDs. Eventually the characterization of these cardiac EH and sensing devices through human trials will be needed, and it is promising that these devices will be clinically employed in the future to reduce the reliance on batteries for powering cardiac medical devices and any other IMDs.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

Acknowledgements

The authors acknowledge financial support from the National Insti-tute of Health (NIH) Director’s Transformative Research Award (R01HL137157, PI: X.J.Z.), the National Science Foundation award (ECCS1509369, PI: X.J.Z.), and the startup fund from the Thayer School of Engineering at Dartmouth. L.D. acknowledges the Arthur L. Irving Institute for Energy and Society Award at Dartmouth. Z.C. also appre-ciates the support from the Branco Weiss-Society in Science fellowship, administered by ETH Zürich.

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Lin Dong is a Research Associate at Thayer School of Engi-neering at Dartmouth College. She received her PhD from Department of Mechanical Engineering at Stevens Institute of Technology, where she was awarded Innovation and Entre-preneurship Doctoral Fellowship. Her research interests include advanced materials for implantable and wearable energy de-vices as well as actuating and sensing devices using soft materials.

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Congran Jin is currently a Ph.D candidate at Thayer School of Engineering at Dartmouth College, USA. He obtained his B.S. and M.S. degrees in Mechanical Engineering from Rensselaer Polytechnic Institute (RPI), USA in 2016. His research interests include flexible energy harvesters, sensors, nanomaterial fabrication and soft robots.

Andrew Closson is a Ph.D. student under the supervision of Professor John XJ Zhang at Thayer School of Engineering at Dartmouth College. He received his B⋅Sc. degree in Bioengi-neering from the University of Maine in 2016. His research focuses on the development of materials and devices for energy harvesting applications in biomedical devices.

Ian Trase received his Bachelor’s degree in Mechanical Engi-neering from Princeton University in 2014 with certificates in Materials Science and Engineering Physics, and recently grad-uated from the Thayer School of Engineering at Dartmouth as a Ph.D. Innovation Fellow. Ian’s research experiences include internships at the Princeton Electric Propulsion Laboratory and MIT CSAIL. His research focuses on flexible electronics and wearable haptics with an interest in translating technologies out of the lab to commercial applications.

Haley Richards received her bachelor’s degree in Engineering Sciences Dartmouth College in 2020 and competed a senior honors thesis under the supervision of Professor John XJ Zhang. Her research focuses on the design and analysis of self-powered wearable cardiac sensors.

Zi Chen is an Assistant Professor at Thayer School of Engi-neering at Dartmouth College. He received his PhD from Department of Mechanical and Aerospace Engineering at Princeton University and was a Research Scientist at Wash-ington University in St. Louis. His research interests include mechanical instabilities, multistable structures, energy har-vesting, origami/kirigami structures, robotics, and tissue/cell biomechanics.

John X.J. Zhang is a Professor at Dartmouth College. He received a Ph.D. from Stanford University, and was a Research Scientist at MIT. He is a Fellow of American Institute for Med-ical and Biological Engineering, and a recipient of NIH Di-rector’s Transformative Research Award, NSF CAREER Award, DARPA Young Faculty Award, Facebook SARA award and Sony Faculty Innovation Award. His research focuses on exploring nanomaterials, biophysics, and nanofabrication technology for biosensing, energy harvesting and wearables.

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