Advances in Bioresorbable Triboelectric Nanogenerators

Abstract

With the growing demand for next-generation health care, the integration of electronic components into implantable medical devices (IMDs) has become a vital factor in achieving sophisticated healthcare functionalities such as electrophysiological monitoring and electroceuticals worldwide. However, these devices confront technological challenges concerning a noninvasive power supply and biosafe device removal. Addressing these challenges is crucial to ensure continuous operation and patient comfort and minimize the physical and economic burden on the patient and the healthcare system. This Review highlights the promising capabilities of bioresorbable triboelectric nanogenerators (B-TENGs) as temporary self-clearing power sources and self-powered IMDs. First, we present an overview of and progress in bioresorbable triboelectric energy harvesting devices, focusing on their working principles, materials development, and biodegradation mechanisms. Next, we examine the current state of on-demand transient implants and their biomedical applications. Finally, we address the current challenges and future perspectives of B-TENGs, aimed at expanding their technological scope and developing innovative solutions. This Review discusses advancements in materials science, chemistry, and microfabrication that can advance the scope of energy solutions available for IMDs. These innovations can potentially change the current health paradigm, contribute to enhanced longevity, and reshape the healthcare landscape soon.

1. Introduction

Numerous biological and metabolic processes are regulated by chemical or electrical signaling between cells and tissues,¹⁻⁵ providing opportunities to connect computers and the human body for healthcare purposes. The rise of implantable medical devices (IMDs) has enabled the quantification of various physiological signals that offer real-time monitoring for valuable information about the health state and the development of effective electrical impulse-based treatments for an array of disorders or diseases.⁶⁻¹⁰ The direct or close physical interfaces between IMDs and biological systems enable accurate and effective two-way interactions, and the improvement in surgical robots has relieved risks associated with surgical procedures.^6,9,11−16 Moreover, biocompatible polymers have been explored and applied to the development of IMDs with improved biosafety and clinical benefits, as they form stable, intimate tissue interfaces.^10,17,18 In short, many IMDs offer a fully integrated healthcare system based on the organic collaboration of sensing,¹⁹⁻²¹ data analysis,^22,23 communication,²⁴ and treatment.²⁵

Transient, bioresorbable IMDs, which are designed to degrade and be absorbed by the human body over time, have emerged as an attractive new class of biomedical devices.²⁶⁻³⁰ Their self-clearing capabilities reduce physical and economic burdens on the patients through periodic replacement to extract the device and allow us to circumvent the potential risks for infection and negative health consequences.^{17,19,31−39} Conventional semipermanent IMDs rarely offer a clinical functionality that takes a short time to manage and treat (e.g., neurostimulation for six months and wound healing for two weeks), since the threat of surgery is much greater than their benefits.^13,40−42 In this regard, the use of transient IMDs has broadened the scope of medical applications for IMDs to drug delivery,⁴³⁻⁴⁸ monitoring,^{10,17,27,33,49−52} diagnosis,^6,19 and therapeutic applications.^27,41,53,54 Further advances in materials science have led to the rise of on-demand transient IMDs whose degradation starts by events triggered at the intended time.^{44,45,55−61} While a majority of transient systems employ a passive operation that decays constituent materials gradually in vivo,^{28,37,48,62−67} this new class of bioresorbable IMDs with active operation is disintegrated quickly after the application of a corresponding stimulus, ensuring their stable function and avoiding the potential risk of residue materials remaining in the body.^{57,60,61,68−70} Continued advances in functional polymers, materials processing, and electronic components will drive progress in IMDs with biosafety and complex capabilities for a variety of biomedical applications.

Despite the noticeable progress of bioresorbable IMDs, short lifetimes and a limited power source remain technological challenges that should be addressed. Batteries are a traditional and reliable electrical power source of most IMDs integrated with power management integrated circuits (PMICs), however, they have finite energy capacity thus a limited lifespan without replacement or charging.^{7,11,12,32,36} Additionally, they are not bioresorbable, and device removal through surgical procedures is required. In addition, their large volumes complicate the miniaturization of bioresorbable IMDs crucially required for seamless implantation.^27,71 Wireless energy transmission (WET) technologies, such as near-field communication (NFC) and radiofrequency identification (RFID), enable the noninvasive transfer of large amounts of electrical energy via inductive coupling, with the capability of real-time adjustment.^{19−21,27,41,44,72,73} By incorporating WET technologies, implantable devices can be designed to be significantly smaller due to a lower reliance on battery capacity. This not only minimizes invasiveness but also eliminates the need for secondary surgeries linked with battery charging issues.^{27,41,74−78} Furthermore, these devices might not even require batteries, as they could be externally controlled through the targeted application of electromagnetic waves and specific inductive design. These advances paved the way for the development of fully bioresorbable IMDs.^{11,12,20,25,32} However, there are still several limitations that keep WET from serving as a fully bioresorbable power source for IMDs. First, the biosafety of WET technologies has not yet been fully proven, and their transmittable power is controlled by the intensity of the incident radiation, which is limited by the safety guidelines of the Food and Drug Administration (FDA) to prevent tissue heating.⁷⁹⁻⁸² Next, it is important to mention that the use of bioresorbable passive circuit elements, such as capacitors, diodes, and inductors, is necessary to achieve effective electromagnetic induction. However, this makes it difficult to predict or control the lifespan of the components under physiological conditions due to the varying degradation performance of the materials involved.^27,32,54 Additionally, diminished power transfer efficiency due to high attenuation through the skin and the strong dependence of the efficiency on the angle between the transmission and receiving antennas are another notable disadvantages.⁸³⁻⁸⁵

Energy harvesting technologies that convert incident to electrical energy through photovoltaic and thermoelectric effects have also been explored.^86,87 Photonic power transfer exploiting visible and near-infrared light has demonstrated high performance in energy harvesting, thanks to its high DC output characteristics.^85,88−90 Recent studies have explored the development of bioresorbable, implantable photovoltaic devices compatible with shallow skin.^87,91 The latest advancements in the comprehensive integration of microsized optoelectronic devices provide a viable foundation for power generation in extremely small wireless implants.^89,90 However, bioresorbable photovoltaic cells have exhibited low output power due to the low energy conversion efficiency of bioresorbable active materials,^87,91 and the light sources face high attenuation at the epidermal skin and underlying tissue, leading to a low-light-intensity environment and limited output power in deep tissue.⁹² Additionally, the output power is significantly affected by incident angles, which may limit device operation in a physiological environment.⁹³ Next, thermoelectric energy harvesters have shown promise as power sources in various research and commercial fields, including wearable devices and power plants.^86,94 These devices are desirable for their DC output characteristics and their ability to exploit an abundant energy source.⁸⁶ However, their performance is significantly impeded by the homotherm physiological environment within the body.⁹⁴ Finally, energy harvesters based on the piezoelectric effect have garnered significant attention due to their high output performance.^11,95−99 This performance has been achieved through the development of high-piezo response materials and composites,^100,101 ultrasound-mediated power transfer,^11,95 and advanced engineering techniques such as multistacking.^98,102 Despite the promise of piezoelectric energy harvesters, there are significant challenges to overcome. Most high-performance piezoelectric materials, such as lead zirconate titanate (PZT), possess inherent biological toxicity.^{100,103−106} Their brittle nature could also cause physical harm to surrounding tissue. Some studies have reported on flexible, biocompatible, and bioresorbable piezoelectric materials such as polyvinylidene fluoride (PVDF) and ZnO nanowires.^107−109 Recent research has even revealed the piezoelectric properties of viruses or living tissues. However, their output power is often significantly lower than required.^{101,110−113}

Bioresorbable triboelectric nanogenerators (B-TENGs) are promising candidates to power transient IMDs due to their high output performance, low cost, and simple structure and the vast selection of materials available.^{49,51,52,64,114−124} B-TENGs convert various forms of abandoned mechanical energy in the human body (e.g., body motion such as walking, arm/leg swinging, organ movement, etc.)^118,120,124 and external mechanical energy (e.g., ultrasonic waves),^{58,117,123,125} into useful electrical power through the coupling of the triboelectric effect and electrostatic induction (Figure 1A). The typical working principle of B-TENGs is a contact-separation mode; they consist of pairs of triboelectric layers, underneath electrodes, and encapsulation layers and generate electrical output as deformation is applied and released.^{52,64,116,120,121,124,126} As such, a considerable number of studies have used B-TENGs to power batteries⁵⁸ or developed self-powered transient IMDs for biomedical applications including diagnostics,^{30,49,127,128} therapeutics,^123−125 rehabilitation,^129,130 and antibacterial activities^117,131 for future healthcare. Along with the development of on-demand transient materials, biodegradation behaviors can be classified as passive or active operations (Figure 1B).^{58,117,120,124} B-TENGs with passive operation lose their weight and function continuously after transplantation by predetermined biodegradation performance solely relying on the inherent properties and film dimensions.^{124,132−134} As biodegradation occurs, B-TENGs show a reduction in output performance. This is attributable to the deteriorating mechanical properties of the B-TENGs and contamination by physiological substances like water and proteins, as well as the interaction with the immune system.^132,133,135 The potential risk for negative health consequences due to prolonged residues and unreliable service quality and lifespan are challenges for clinical settings.¹³⁶ In contrast, despite remaining technological challenges, the active operation offers B-TENGs the ability for immediate clearing upon the application of stimuli such as heat, light,^122,137 and ultrasound,^58,117,125 eliminating concerns about prolonged residues and performance stability. The progress in high-performing bioresorbable triboelectric materials and an ultrasound-mediated power generation mechanism that utilizes biosafe medically approved ultrasound as incident mechanical energy has highly increased the level of output power of B-TENGs.^117,138Figure 1C provides an overview of the power consumption and implantation duration of conventional IMDs compared to the current state of output power of B-TENGs and nondegradable implantable TENGs. The power consumption of IMDs ranges from a few tens of microwatts to several tens of milliwatts, and they can be implanted from a few days up to several decades, depending on their specific therapeutic applications and functionalities.^{28,80,139,140} Transient implants can last up to two years, while semipermanent IMDs such as pacemakers and spinal cord stimulators are designed for long-term implantation. Presently, the power output of both B-TENGs and nondegradable TENGs has shown a dramatic increase, hitting 300 μW¹¹⁷ and 6 mW,¹³⁸ respectively. The significant improvements in output power have extended the applications of B-TENGs; thus, it seems they can serve soon as viable and practical power systems for transient, low-power IMDs, such as neurostimulators, infusion pumps, and electrical sterilizers. (Figure 1C).^14,141−143

Figure 1.

Overview of B-TENGs as an advanced energy solution for transient IMDs. (A) Schematic representation of B-TENGs in terms of their energy conversion, typical structure, and biomedical applications. (B) Profiles of the mass and performance of B-TENGs through passive and active operations for comparison of their working principles. (C) Conventional IMDs categorized according to power consumption and implant duration, illustrating the current state of powering capabilities of B-TENGs.

In this Review, we present an overview of B-TENGs for bioresorbable electronics and highlight recent progress in the field. To comprehensively introduce their fundamental science, chemistry, and applications, this Review covers the power generation and biodegradation mechanisms, mechanisms of passive/active operation, materials development strategies for high power, and biomedical applications. Most of the bioresorbable materials including on-demand transient materials have been employed in other research areas, such as transient electronics, drug delivery, stimuli-responsive materials, and chemistry. Considering the vast options of possible constituent materials for B-TENGs,^144−146 there are a lot of unexplored opportunities to develop triboelectric energy solutions with functional biodegradation performance as well as a large power supply. We also provide substantiated suggestions of candidate materials and insight for future works. Lastly, we outline the future perspectives of B-TENGs regarding their remaining challenges and milestones.

2. Working Principles and Materials for B-TENGs

2.1. Energy Harvesting and Transience Principles of B-TENGs

The triboelectric effect is a phenomenon where certain materials become electrically charged when they come into contact with another material and then separate from it (Figure 2A).¹⁴⁷ When two dissimilar materials are in contact, the electrons from one material can move to the other material, leaving one surface positively charged and the other negatively charged.^148−155 Despite ongoing debate, it is generally accepted that the electron transfer mechanism dominates the solid–solid triboelectric effect. Both the Fermi level model and the surface states model can adequately describe triboelectricity in metal–polymer and polymer–polymer contacts.^156,157 The overlapped electron cloud model was initially introduced to describe electron transitions by a triboelectric event. According to this model, an overlap of the electron cloud between two adjacent atoms at the interface of two materials under mechanical stress reduces the potential barrier between them, thus facilitating electron transfer from one atom to the other. This model also has been instrumental in explaining various phenomena, such as the plastic deformation of metals under stress and the bonding between different types of molecules. The number of transferred charges is dependent on several factors, including the material type, its surface roughness, and the environmental humidity.

Figure 2.

Power generation mechanism and target energy of B-TENGs. (A) Schematic of the triboelectric effect and overlapped electron cloud model. Reproduced with permission from ref (148). Copyright 2020 Wiley-VCH. (B) Working principle diagram of B-TENGs. (C) Ambient energy sources with a wide frequency range used for triboelectric energy harvesting. Reproduced with permission from refs (138), (165), (164), (154), and (139), respectively. Copyright 2019 American Association for the Advancement of Science. Copyright 2018 American Association for the Advancement of Science. Copyright 2018 Nature Publishing Group under the terms of the Creative Commons Attribution 4.0. (https://creativecommons.org/licenses/by/4.0/). Copyright 2021 American Association for the Advancement of Science. Copyright 2021 Nature Publishing Group under the terms of the Creative Commons Attribution 4.0 (http://creativecommons.org/licenses/by/4.0/).

The contact-separation mode, the typically employed configuration of B-TENGs, is used for a variety of applications. B-TENGs consist of two triboelectric layer–electrode pairs, a small gap between them, and encapsulation layers. A B-TENG generates electrical energy from mechanical motion through the coupling of the triboelectric effect and electrostatic induction (Figure 2B).^126,158 When two materials come into contact and then separate, two triboelectric layers are positively and negatively charged due to the triboelectric effect. In the contact-separation mode, the materials are in contact and then pulled apart repeatedly by an external force, such as mechanical vibration or deformation.^144,159,160 As the triboelectric layers move closer together and then apart, the triboelectric charges induce an electrical potential difference in the paired electrodes. This causes an electrical current to flow through an externally connected load between the two electrodes, thereby compensating for the potential difference. Single electrode mode TENGs comprised of one triboelectric layer–electrode pair have also been exploited in biomedical applications due to similar working principles and the simple structure.¹⁶¹ There are other modes of TENG, including sliding mode, freestanding mode, and other variations,^162,163 however, most of them have been rarely used in this area because the required motion is generally unavailable in the physiological environment.

Nonbioresorbable implantable TENGs (I-TENGs) and B-TENGs can be categorized according to the frequency of the mechanical motion involved (Figure 2C). Low-frequency mechanical energy refers to biomechanical energy (e.g., body motion, organ movement, and vascular dynamics),^{127,139,154,164} and high-frequency mechanical energy corresponds to audible sounds¹⁶⁵ and ultrasounds.^{58,117,123,125,138,166,167} Regardless of the frequency range, the working mechanisms of TENGs are the same, that is, based on the coupling of the triboelectric effect and electrostatic induction. However, the detailed mechanical dynamics of the triboelectric layers, device configurations, and output characteristics differ. Low-frequency TENGs employ conventional and standard models. Mechanical stress induced by biological activities deforms a B-TENG and leads to contact and separation of the polymeric triboelectric layers.^154,164,168 B-TENGs also can use inertial force derived from body motion such as walking, sitting, and running to vibrate the freestanding triboelectric layer in the device, generating power.¹³⁹ High-frequency TENGs operate based on the vibrations of polymeric membranes induced by audible sound or ultrasound. The sound waves are propagated through tissue and undergo constructive or destructive interference at the surface of the electrode, leading to high-frequency vibrations of the triboelectric layer and output generation.¹⁶⁵ In particular, the use of external ultrasound has gained considerable attention since it was first reported in 2019 owing to the superior performance and ability to transfer power to IMDs noninvasively.¹³⁸

Similar to other bioresorbable electronics, the transience of B-TENGs is caused by passive operation in many cases, and their lifetime is divided into two phases: “functioning time”, where the operating B-TENG maintains its performance, and “disappearance time”, where the bioresorbable component materials completely dissolve and are absorbed in the body (Figure 3A).^48,50 To be more specific, B-TENGs degrade over time once they are transplanted into the body. Weight loss occurs through a variety of degradation processes such as swelling, hydrolysis, oxidation, and enzymatic degradation.^{48,54,67,169−175} As the device loses mass over a certain quantity, physiological fluids begin to infiltrate the triboelectric layer or the electrodes, which leads to a cessation of its functioning, thereby rendering it unable to generate power.¹⁷⁶ This marks the end of its operational lifespan (functioning time), but the degradation process continues further until it is entirely absorbed by the body.⁶⁸ This phase is termed the disappearance time, after which there are no remnants of the B-TENGs left in the body. The biodegradation process of B-TENGs offers a major benefit regarding safety and clinical adaptability. This not only guarantees the device’s optimal performance throughout its operational lifespan but also mitigates concerns about the potential long-term impacts of implantable devices, as the device eventually disintegrates, leaving no residual materials behind.^55−57,61 Functioning time and disappearance time are determined by several parameters in passive operation, which are described in the subsequent chapters in this work: material properties (e.g., degradation mechanisms, chemical structures, water permeability, molecular weight, and crystallinity),^{173−175,177} formfactors and microstructures (e.g., thickness, surface morphology, and porosity),^178−180 and physiological conditions (e.g., temperature, pH, biofouling, and enzyme contents).^178,181,182

Figure 3.

Service lifetime and biodegradation processes of B-TENGs. (A) Schematic illustration of the biodegradation mechanism; each step illustrates the mechanical/chemical dissolution processes along with the functioning and disappearance times after device transplantation. (B) Categorized diagram of bioresorbable materials along with their degradation rates. (C) Plot of the diffusion rate for degradation factors and the degradation rate for bioresorbable B-TENG materials. Encapsulation layers and inner active layers are illustrated to provide the required properties in the suggested methodologies.

The construction of B-TENGs involves the integration of multiple materials, including metallic electrodes, polymeric encapsulation or triboelectric layers, and potentially ceramics.^{51,118,120,121} Thus, the biodegradation rates and mechanisms of constituent materials are carefully considered in material selection and the design of B-TENGs for particular applications. Figure 3B and Tables 1 and 2 offer the biodegradation rates of various materials. These guide the selection of component materials to develop B-TENGs. Polymers and organic materials usually have high degradation rates over a relatively wide range,^{129,183−190} whereas inorganic materials have intermediate and low degradation rates with narrow tolerances.^{13,40,50,191,192} For example, poly(lactic-co-glycolic) acid (PLGA)¹⁸⁸ is mainly decomposed by hydrolysis with a degradation rate of approximately hundreds of nanometers per day, depending on chain composition, hydrophobicity, molecular weight, crystallinity, and other factors.¹⁹³

Table 1. Biodegradation Rate of Bioresorbable Polymers^{129,183−190}.

	biodegradation mechanism	biodegradation rate	test environment	ref
alginate	hydrolysis	20% mass loss after 40 d	phosphate buffered saline (PBS) at 37 °C	(183)
chitosan	hydrolysis	90% mass loss after 1 d	PBS (pH = 7.4) at 37 °C	(184)
gelatin	hydrolysis	68% mass after 15 d	PBS (pH = 7.4) at 37 °C	(185)
PLA	hydrolysis	100% mass loss after 49 wk	Hank’s balanced salt solution (HBSS) at 37 °C	(186)
PGA	hydrolysis	10% mass loss after 3 wk	PBS (pH = 7.4) at 37 °C	(187)
PLGA (PLA/PGA = 50:50)	hydrolysis	80% mass loss after 8 wk	PBS (pH = 7.4) at 37 °C	(188)
PCL	hydrolysis	11% mass loss after 4 wk	0.01 M NaOH solution at 37 °C	(189)
PVA	hydrolysis	90% mass loss after 90 d	PBS at 37 °C	(129)
PHBV	hydrolysis	4% mass loss after 15 wk	PBS (pH = 7.4) at 37 °C	(190)

Table 2. Biodegradation Rates of Bioresorbable Semiconductors^{13,40,50,191,192}.

	biodegradation mechanism	biodegradation rate	test environment	ref
Si	hydrolysis	4.5 nm·d–1	PBS (pH = 7.4) at 37 °C	(13)
Poly-Si	Hydrolysis	2.8 nm·d–1	PBS (pH = 7.4) at 37 °C	(50)
a-Si	hydrolysis	4.1 nm·d–1	PBS (pH = 7.4) at 37 °C	(50)
SiO2	hydrolysis	8.2 nm·d–1	artificial cerebrospinal fluid (aCSF) (pH = 7.4) at 37 °C.	(40)
Si3N4	hydrolysis	0.4 nm·d–1	PBS at 37 °C	(191)
SiGe	dissolution	0.1 nm·d–1	PBS (pH = 7.4) at 37 °C	(192)
Ge	hydrolysis	3.1 nm·d–1	PBS (pH = 7.4) at 37 °C	(192)

To ensure proper and precise control over their lifespan, a design strategy can be suggested; the encapsulation layer and inner active layers of a device should exhibit either slow or fast degradation rates and low or high diffusion rates for degradation factors, as shown in Figure 3C. This design strategy could potentially reduce the discrepancy between the device’s functional time and disappearance time, thereby minimizing the potential harms posed by prolonged residues before the full resolution of a B-TENG. As the encapsulation layer slowly decays from the surface with a low diffusion rate and finally allows the biofluid to get into the inner side of the devices, inner active layers dissolve quickly in a few hours or days due to fast swelling and a high decomposition rate, enabling almost simultaneous elimination.

2.2. Bioresorbable Polymers for B-TENGs

Polymers are a major constituent of the encapsulation and triboelectric layers of B-TENGs. Their surface charge transfer characteristics and the selection of triboelectric pairs are highly responsible for the surface charge density.^144,194 Thus, bioresorbable polymers significantly influence both the transience behavior and output performance. As shown in Figure 4, bioresorbable polymers can be categorized into nature-derived bioresorbable polymers (NBPs) and synthetic bioresorbable polymers (SBPs).^{18,37,38,48,169} NBPs are mainly of three types: polysaccharides, proteins (polypeptides), and bacterial polyesters. Among them, polysaccharides, such as alginate,^195,196 hyaluronic acid,⁶⁴ cellulose,^{63,66,145,197−202} chitosan,^203−205 and agar¹¹⁸ and proteins, such as silk fibroin,^206−209 collagen,^57,210 gelatin,¹⁸⁵ and elastin,²¹¹ are intensively employed in transient electronics and B-TENGs because of their superior biocompatibility and inherent enzymatic biodegradability in the human body.²¹² Polysaccharides are composed of different units of monosaccharide or disaccharide chains, and various polysaccharides can be obtained from distinct isomers and chemical bonds.^195,213 Cellulose and chitosan, which are present in plants and crustacean shells, represent structural polysaccharides,^63,66,197 whereas alginate, hyaluronic acid, and agar are storage polysaccharides.^64,195,196 Protein chemistry is tremendously diverse and heterogeneous.^210,214,215 A near-infinite set of proteins can be acquired from amino acids and their derivates with different primary, secondary, tertiary, and quaternary structures.^185,214,215 Despite their widespread utilization and the vast potential offered by diverse novel materials, the variability in material properties such as distribution, branching, composition, and molecular weight sequence arising from natural extraction processes can be a disadvantage when trying to predict the degradation performance.^216−218

Figure 4.

Representative bioresorbable polymers and their chemical structures. Bioresorbable polymers can be categorized into NBPs and SBPs.

Six major types of biopolymers are naturally synthesized in microorganisms: polynucleotides (such as DNA and RNA), polysaccharides (such as alginate, xanthan, and cellulose), polyhydroxyalkanoates (PHAs), polythioesters, inorganic polyanhydrides (such as polyphosphates), and polyamides (such as polypeptides, cyanophycin, ε-poly-l-lysine, and poly-γ-glutamate).^219,220 Among them, bacterial polyesters represented by PHAs are biodegradable, biocompatible, and naturally produced by many Gram-positive and Gram-negative bacteria, including Cupriavidus, Pseudomonas, Alcaligenes, Bacillus, and Aeromonas.^221−223 There are more than 90 monomers of PHAs, represented by 3-hydroxybutyrate, 3-hydroxyvalerate, and 3-hydroxyhexanoate, leading to polymers having various physical properties, from hard plastics to flexible elastomers. Some PHAs were commercialized earlier for packaging, agricultural, and medical applications owing to their mechanical toughness, enzymatic degradability, and high biocompatibility.²²⁴ In this context, a few studies have employed PHAs as encapsulation or triboelectric layers of bioresorbable electronics and B-TENGs.^58,124

Since the use of linear, aliphatic, and thermoplastic polyester polyglycolide (PGA) as a biodegradable suture in the 1960s, many SBPs have been developed and used in bioresorbable devices. SBPs can be divided into three main groups: polyester plastics, polyester elastomers, and polyurethane elastomers. Polyester plastics include poly(vinyl alcohol) (PVA),^52,225,226 poly(ethylene glycol),²²⁷ poly(caprolactone) (PCL),^133,189,228 poly(lactic acid), and PLGA.^37,169,229 Polyurethanes and cross-linked polyester elastomers include polyglycerol sebacate (PGS)⁴⁹ and poly(1,8-octane diol-citrate) (POC).²³⁰ SBPs can have uniform compositions and undergo hydrolysis and oxidation in a predictive manner. In particular, SBPs have advantages over NBPs such as high reproducibility, tunable properties, endless forms, and established structures.^213,231,232 Although SBPs often lack cell adhesion sites, requiring chemical modifications, their biocompatibility is well established, with some polyesters and elastomers, such as PLA and PLGA, being approved for medical applications by the FDA.^{37,173,175,213}

Several studies have focused on the development of B-TENGs using NBPs owing to their high triboelectric charges, low costs, and environmentally friendly extraction processes. Jiang et al. used several NBPs (cellulose, chitin, silk fibroin (SF), rice paper (RP), and egg white (EW)) as triboelectric layers of B-TENGs.²⁰⁸ The fully biodegradable TENGs were fabricated using a pair of NBP triboelectric layers, SF encapsulation layers, and magnesium electrodes (Figure 5A). The nature-derived materials used in this study degraded gradually, leading to a B-TENG with a functioning time of 42 d and a disappearance time of 84 d when immersed in PBS solution (pH = 7.4) at room temperature via rapid autocatalytic hydrolysis (Figure 5B). The in vivo degradation performance of these materials in rat models was also investigated using methanol-treated SF. Some gaps formed at the edges of the B-TENG at 21 d and the SF-encapsulated B-TENG disintegrated into two parts over time. This study first presented a triboelectric series of NBPs and provided a basis for triboelectric pair selection and device design. Interestingly, the triboelectric series of the five materials was EW > SF > chitin > cellulose > RP, from positive to negative (Figure 5C). The EW/Kapton pair marked the maximum transferred charge of 12 nC, and the as-fabricated B-TENG achieved a tunable output performance for open-circuit voltage (V_OC) of 8–55 V and a short-circuit current (I_SC) of 0.08–0.6 μA. While this work did not explain the factors defining this triboelectric series, one plausible cause could be that EW and SF have a high content of nitrogen-containing functional groups, which donate electrons to a counter material.¹⁴⁶ Conversely, the presence of hydroxyl groups could make polysaccharides like cellulose and RP negative.²³⁴

Figure 5.

Bioresorbable polymers for B-TENGs. (A) Schematic representation of B-TENGs composed of various NBPs, including field emission scanning electron microscope (FE-SEM) images and the atomic force microscopy (AFM) topology to demonstrate the nanostructured surface morphology of the NBP triboelectric layer (lower and upper scale bars are 5 and 1 μm, respectively). (B) In vitro biodegradation (scale bars are 5 mm) and (C) working principle and output power of B-TENGs using five different NBPs. (A–C) Reproduced with permission from ref (208). Copyright 2018 Wiley-VCH. (D) Structure of the B-TENG using calcium alginate films. (E) Weight loss of a calcium alginate film via in vitro biodegradation in water at room temperature. (D and E) Reproduced with permission from ref (195). Copyright 2018 Royal Society of Chemistry. (F) 3D printing process of CNTs@SF core–sheath fiber-based smart patterns to fabricate electronic textiles capable of triboelectric energy harvesting. Reproduced with permission from ref (233). Copyright 2019 Elsevier. (G) Illustration of a contact-separation mode B-TENG using nanostructured SBPs, including FE-SEM images and the AFM topology of the SBP triboelectric layer. (H) Output currents of B-TENGs with different SBP triboelectric pairs. (I) Triboelectric series of SBPs using polyimide (Kapton) as a reference. (J) In vitro biodegradation of a B-TENG encapsulated by PLGA in PBS (pH = 7.4, 37 °C, scale bars are 10 mm). (H–J) Reproduced with permission from ref (124). Copyright 2016 American Association for the Advancement of Science under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/). (K) Structure and dimensions of a breathable B-TENG based on electrospun PLGA, PLA, and Ag NWs. (L) FE-SEM images of the surface morphology of PLGA and PLA nanofibers (scale bars 10 and 2 μm, respectively). (M) Sequential photographs of the in vitro biodegradation of PVA, Ag NWs/PVA, and PLGA/Ag NWs/PVA nanofiber films in PBS at 37 °C. (K–M) Reproduced with permission from ref (52). Copyright 2020 American Association for the Advancement of Science under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/).

Subsequent studies further explored the use of NBPs as triboelectric materials for B-TENGs. Pang et al. discussed the development of a B-TENG based on calcium alginate film (Figure 5D and E),¹⁹⁵ explaining the development of a partially bioresorbable generator based on alginate that can harvest energy from water waves, with detailed descriptions and images of the device structure and its electrical output characteristics. Alginate is a kind of natural linear polysaccharide containing a linear chain of α-l-guluronic acid and β-d-mannuronic acid derived from brown sea algae.²³⁵ It possesses high biocompatibility, low toxicity, and biodegradability, rendering it an attractive polymer for use in medical and environmental applications.^236−239 This B-TENG employed a calcium alginate film and aluminum as the triboelectric pair, achieving a high output performance with a maximum voltage, current, and power of 33 V, 150 nA, and 9.5 mW in ex vivo conditions, respectively. Additionally, as shown in Figure 5E, calcium alginate rapidly dissolved in in vitro conditions, losing its entire weight in 72 h (deionized water).

A B-TENG was directly 3D printed in the form of an electronic textile composed of core–sheath fibers using SF and carbon nanotube (CNT) ink (Figure 5F).²³³ To fabricate the B-TENG, a CNTs@SF core–sheath fiber-based pattern was printed on a conventional fabric using polyethylene terephthalate/titanium dioxide (PET/TiO₂) as the contact material and electrode. The B-TENG is partially biodegradable and, owing to the highly triboelectric positive properties of SF, exhibited high output performance, generating an I_SC of 1.4 mA and a V_OC of 15 V at a displacement speed of 10 cm·s^–1, with a maximum power density of 18 mW·m^–2. In addition to plastic films, NBPs in the form of hydrogels, prepared by spontaneous gelation, are also used as triboelectric materials with potential superior biocompatibility given their high water content, hydrophilicity, and mechanical matching with living tissue.²⁴⁰ Kim et al. discussed the development of a simple and efficient method for preparing pure and cross-linked hyaluronic hydrogel films using solvent evaporation and 1,4-butanediol diglycidyl ether (BDDE) cross-linker.⁶⁴ The B-TENG illustrated in this study employed polytetrafluoroethylene (PTFE) and hyaluronic acid hydrogel (HA) as negative and positive triboelectric layers, respectively, indicating they are partially bioresorbable. Given the superior charge transfer between these layers, the B-TENG achieved a high output voltage of 20 V and a high output current of 0.4 μA (active area of 7.5 cm²), with a maximum output power of 5.6 mW·m^–2 at 60 MΩ external resistance. Notably, the output power and biodegradation performance of the HA film showed opposite trends at different cross-linking ratios of HA. Highly cross-linked HA with a high BDDE content exhibited a more aligned structure than pure HA. The aggregated cross-linked HA film (2 wt % BDDE) degraded slower and disappeared after 30 d. Meanwhile, the B-TENG based on the highly cross-linked HA films showed lower output power than pure HA. A fully biodegradable TENG was also fabricated employing pure chitosan (CS) and HA hydrogel films as triboelectric layers, along with Mg electrodes. The 2 wt % BDDE cross-linked HA hydrogel film exhibited the highest peak-to-peak V_OC of 1.2 V. Although some previously reported TENGs were not fully bioresorbable, partially using nondegradable materials, NBPs exhibited considerable potential for use as triboelectric and encapsulation layers for high-power B-TENGs.

SBPs have several advantages over NBPs for the development of TENGs. One of their key advantages is their versatility, as different monomers can be mixed in various compositions to achieve a wide range of reliable physical, mechanical, and chemical properties.^37,213,241 Thus, the corresponding TENG can be easily tailored for a specific application. Zheng et al. used different SBPs to develop fully bioresorbable TENGs that can serve as implantable power sources for IMDs (Figure 5G).¹²⁴ The triboelectric series of several SBPs, including PLGA, PHB/V, PVA, and PCL, were validated by comparing the output current of their combined triboelectric pairs (Figure 5H). Using a polyimide film as a reference, the relative output current was ranked as PLGA > PHB/V > PVA > PCL, from positive to negative (Figure 5I). The triboelectric series of representative bioresorbable SBPs that can be used for biodegradable TENGs presented here provides a fundamental guideline for subsequent research. As shown in Figure 5J, the B-TENG encapsulated by PLGA degraded and fully dissolved in PBS after 90 d (pH = 7.4, 37 °C).

SBPs can be easily manufactured in various forms with good reproducibility. By utilizing advanced manufacturing methods, such as electrospinning and 3D printing, polyesters and elastomers that have high surface areas can be obtained, possibly increasing charge transfer performance and thus enhancing output power.^{120,242−244} Peng et al. used electrospun PLGA, PVA, and Ag nanowires (NWs) to develop a breathable, fully biodegradable, antibacterial, and self-powered electronic skin based on an all-nanofiber B-TENG (Figure 5K).⁵² A biodegradable nanofiber membrane was created from these SBPs that could generate large amounts of electricity through the triboelectric effect. The surface morphologies of the PLGA and PVA nanofibers were observed using scanning electron microscopy (SEM), and their average fiber diameters were 600 and 130 nm, respectively (Figure 5L). Their highly porous structure provides high gas, water, and air permeability that grants flexibility, durability, and a low pressure drop to the B-TENG. The maximum power density of the PLGA/Ag NWs/PVA B-TENG was 130 mW·m^–2 at an optimized resistance of 500 MΩ, which is four times larger than that of PVA/Ag NWs/PVA. As shown in Figure 5M, all component materials showed a high degradation ratio and were fully degraded after 40 d in PBS at 37 °C.

Chen et al. explored a single integrated 3D printing process to customize elastic B-TENGs for wearable electronics.⁴⁹ These TENGs employed fully bioresorbable PGS and CNTs as the two triboelectric components. The CNTs also served as electrodes, whereas the elastic PGS matrix enabled the TENGs to be intrinsically responsive to biomechanical motion and biodegradation, resulting in robust outputs. The hierarchical porous structure of the B-TENG contributes to a higher output than its conventional molded microporous TENG counterparts. Figure 6 summarizes the investigated NBPs and SBPs, classified according to their charge densities after contact with several tribo-negative reference materials and biodegradation rates.^{124,127,206,208,211,226,245−253} Although these characteristics can vary upon material modifications and manufacturing methods, the figure provides a reference point to procure materials candidates for the development of high-performing triboelectric materials and control the lifetime of B-TENGs.

Figure 6.

Triboelectric charge densities and degradation rates of NBPs and SBPs.

To date, as shown in Tables 3 and 4, the current level of output performance of B-TENGs remains lower than that of conventional nonbioresorbable I-TENGs, and triboelectric properties of bioresorbable polymers and metals are centered on this issue.^124,254,255 Different materials possess varying tendencies to either gain or lose electrons via triboelectric events, which are determined by their respective triboelectric polarities. Functional groups significantly impact a material’s triboelectric properties.^156,161,234 For example, functional groups abundant in fluorine atoms typically demonstrate strong electron-accepting tendencies, whereas those containing nitrogen atoms frequently exhibit electron-donating behavior.^146,256,257 With progress in the development of high-performing materials in the last few decades, conventional nondegradable TENGs have been typically fabricated using materials with functional groups that are highly efficient at either gaining or losing electrons.^157,160,258 On the other hand, B-TENGs are fabricated from biocompatible and biodegradable materials generally limited to certain types of polymers, proteins, or other organic materials for their safe degradation in the body over time.^116,120,254 In most instances, these materials possess predominantly neutral or slightly positive functional groups, resulting in their placement in proximity within the triboelectric series.^118,124 Furthermore, bioresorbable metals also exhibit neutral or weakly positive triboelectric properties. Consequently, for the time being, maximizing differences in triboelectric polarities using two distinct bioresorbable polymers or metals might be more challenging, a constraint not shared with conventional TENGs. Nonetheless, it is worth noting that despite the relatively short history of B-TENGs, the output power has been rapidly increased through rigorous research efforts aimed at finding or suitably modifying bioresorbable materials that can generate high output.^{52,120,123,254}

Table 3. Output Performance of Non-Bioresorbable TENGs^{115,243,258−267}.

			output power (peak to peak)
no.	triboelectric layers	active area (cm2)	voltage (V)	current	working mode	ref
1	PTFE, Al/Cu	100	7540	20.1 A	contact-separation	(259)
2	PTFE, Al	0.49	7000	37 μA	contact-separation	(260)
3	silicone rubber, nitrile rubber	78.5	5000	15.7 mA	single electrode	(261)
4	P(VDF-TrFE):BTO, Al	9	1130	1.5 mA	contact-separation	(258)
5	PFA, ZnO@MOP/PVP	4	534	26.8 μA	contact-separation	(262)
6	PDMS-CNT, Al	4	338.25	19.91 μA	contact-separation	(263)
7	PVDF-ZnO NWs, Nylon-ZnO NWs	10	330	10 μA	contact-separation	(243)
8	PTFE, PAN@ZIF-8	4	260	24.5 μA	contact-separation	(264)
9	PTFE yarns, PA66 yarns	3.25	232	6.8 μA	contact-separation	(265)
10	FEP, Ag nanoparticles	12	200	20 μA	contact-separation	(266)
11	PA(polyamide), Al	8.75	150	10 μA	single electrode	(115)
12	nanostructured PI, Cu	2	109	2.73 μA	contact-separation	(267)

Table 4. Output Performance of B-TENGs^{51,116,119−121,124,197,128,208}.

			output power (peak to peak)
no.	triboelectric layers	active area (cm2)	voltage (V)	current	working mode	ref
1	gelatin, electrospun PLA	16	500	17 μA	contact-separation	(51)
2	rice paper, PVC	9	244	6 μA	single electrode	(116)
3	crepe cellulose, nitrocellulose	6.25	196.8	31.5 μA	contact-separation	(197)
4	SF, Mg	8	60	1 μA	single electrode	(128)
5	egg white, rice paper	2	55	0.6 μA	contact-separation	(208)
6	starch/laver, PCL	6.25	50	1 μA	contact-separation	(119)
7	PLLA, chitosan	6.25	45	9 μA	contact-separation	(120)
8	PHBV, PCL	6	28	0.6 μA	contact-separation	(124)
9	PVA, PLGA	6	26	0.4 μA	contact-separation	(124)
10	rice paper, laver	4	23	315 nA	single electrode	(121)

2.3. Bioresorbable Metal Electrodes for B-TENGs

The rational selection of bioresorbable conductors is pivotal for optimizing the electrical output performance and clinical application of B-TENGs.^{51,118,122,124} Thus, a thorough understanding of the inherent properties of the materials is imperative in this decision-making process. Particular attention has been given to comprehending the complex mechanisms of corrosion and biodegradation in metallic biomaterials when they are subjected to physiological environments. These processes can be influenced by various factors such as biological fluids, pH variations, cellular dynamics, and the intricate chemical milieu within the human body.^268,269

The corrosion and biodegradation events encompass several phenomena. Galvanic coupling arises from dissimilar metal nobilities, resulting in the formation of anodic and cathodic regions characterized by accelerated metal ion release and reduction reactions, respectively.²⁷⁰ Uniform corrosion entails the comprehensive release of metal ions from the exposed material surface, driven by its high electrochemical activity, while localized corrosion strongly depends on specific chemical heterogeneity and local surface characteristics.²⁶⁸ For instance, certain metallic biomaterials may exhibit the preferential dissolution of specific components, leading to pitting corrosion or the creation of electrochemically active sites at distinctive features such as edges, kinks, and grain boundaries.

Next, the formation and subsequent stabilization of protein nanobiofilms on micro- or nanoparticles can expedite biodegradation or metal ion release through the detachment of metal-bound proteins, forming metal–protein complexes or conjugates.^214,268 Reactive oxygen species (ROS) play a significant role in compromising the chemical stability of bioresorbable metal conductors. Notably, Fenton reactions occur when released metal ions from iron/iron oxide interact chemically with ROS, generating oxidative agents (e.g., OH₂• and OH•) that contribute to enhanced degradation.^271,272

Next, the biodegradation of ceramic oxide nanoparticles within specific cellular compartments is mediated by the adsorption of apoferritin or ferritin protein molecules onto the nanoparticle surface, triggering the degradation process upon the formation of metal–protein conjugates and subsequent metal ion uptake.²⁷³

Finally, membrane protein assemblies can produce unstable radicals, either through spontaneous dismutation or via superoxide dismutase (SOD) reactions, leading to the production of reactive oxygen and hydrogen peroxide. The interaction between hydrogen peroxide and metals can induce metal oxidation, initiating the generation of ionic metallic ions.²⁷⁴

Although these corrosion and biodegradation mechanisms are of paramount importance when selecting suitable metal conductors for the B-TENGs, in-depth investigations are still required to identify underlying phenomena for each metallic element in an experimental manner. Table 5 describes representative experimental methods for identifying biodegradation processes of metallic materials.

Table 5. Experimental Methods for Identifying Biodegradation Mechanisms^{268,275−281}.

methods	measurement tools	purposes	ref
ex situ methods	scanning electron microscopy (SEM)	morphology changes	(268, 275−278)
	transmission electron microscopy (TEM)
	energy-dispersive X-ray spectroscopy (EDXS)	atomic percentage variations
chemical analyses	inductively coupled plasma atomic emission spectroscopy (ICP-AES)	metallic concentrations and metal ion release	(277, 279, 280)
	inductively coupled plasma mass spectrometry (ICP-MS)
electrochemical measurements	open circuit potential (OCP)	predicting how metallic materials will perform in electrochemical reactions	(268, 281)
	potentiodynamic polarization (PDP)
	electrochemical impedance spectroscopy (EIS)

Table 6 shows the dissolution rate for each metallic material. Since byproduct concentrations can cause acute immune reactions and affect the metabolism, both the degradation rate and the total mass of byproduct formed should be considered to minimize adverse health effects during device transience.^282−284 However, an excessively slow degradation process may harm the patient owing to prolonged exposure to residues within the body. For example, Mg and its alloys are widely used bioresorbable metals with a relatively fast transient rate (0.68 μm·d^–1).¹³⁴ When employing Mg in B-TENGs, it is crucial to predict the flux and content of its degradation products to avoid impacting the surrounding tissues and the metabolism. Mo is also a prevalent bioresorbable metal used in B-TENG electrodes. It is nontoxic and does not form oxide layers within biological systems, thereby maintaining a consistent electrical conductivity. However, its biodegradability is comparatively slow (0.005 μm·d^–1).^227,285 Consequently, careful consideration of the total Mo mass is required to compare the increased risk of prolonged residue exposure with the improved mechanical durability. Other bioresorbable metals used in B-TENG electrode layers, such as Fe, W, and Zn, are subject to similar considerations.

Table 6. Biodegradation Rates of Bioresorbable Conductors^{34,134,284,286−288}.

metals	hydrolytic reaction	biodegradation rate	test environment	ref
Mg	Mg + 2H2O → Mg(OH)2 + H2	0.68 μm·d–1	HBSS (pH = 7) at 37.5 °C	(134)
AZ31(Mg 96/Al 3/Zn 1)		1.64 μm·d–1	Earle’s balanced salt solution (EBSS) (pH = 7.5) at 37 °C	(286)
AM50 (Magnesium alloy)		5.81 μm·d–1	3.5 wt % NaCl solution (pH = 6.5) at 22 °C	(287)
Mo	Mo + 2H2O → MoO2 + 2H2	0.005 μm·d–1	PBS (pH = 7.4) at 37 °C	(288)
W	W + 3H2O → WO3 + 3H2	0.19 μm·d–1	PBS (pH = 7.4) at 37 °C	(288)
Fe	3Fe + 4H2O → Fe3O4 + 4H2	0.55 μm·d–1	HBSS (pH = 7.4) at 37 °C	(34)
Zn	Zn + 2H2O → Zn(OH)2 + H2	0.82 μm·d–1	HBSS at 37 °C	(284)

3. Biodegradation Mechanisms

3.1. Macroscopic Physicochemical Biodegradation Mechanisms

Precisely controlling the lifetime of B-TENGs is fundamental to ensuring their clinical safety and enabling their practical application.^91,117,254 Bioresorbable materials have different degradation performances owing to their distinct degradation mechanisms, kinetics, and environment. Thus, an in-depth understanding of the biodegradation mechanisms is essential for designing bioresorbable implants. At the macroscopic level, the degradation of a material is determined by the diffusion of degradation factors into the material and the cleavage of physical/chemical bonds.^135,169,289 In other words, most bioresorbable materials undergo decomposition or dissolution as a result of contact and multiplex reactions with the components of biofluids, including water, electrolytes, oxidants, and enzymes.^173,175

Biodegradation rates depend on the molecular-scale chemical decomposition kinetics and the macroscopic degradation modes, and the macroscopic degradation can be divided into two categories: bulk degradation and surface erosion (Figure 7). Bulk degradation involves the rapid diffusion of degradation factors inside the material compared to polymer bond cleavage reactivity, resulting in its simultaneous degradation and weight loss in both the core and surface of polymers (Figure 7A).¹⁸¹ Many organic bioresorbable materials, including polymers and hydrogels, undergo bulk degradation via hydrolysis owing to their high water diffusion coefficients. For example, polyesters such as PLGA^290,293,294 and PCL,^133,189 which have been widely utilized as TENGs components, predominantly undergo bulk degradation because of their high porosity (Figure 7B and C).²⁹⁵ Hydrogels, which contain large amounts of water, are more likely to rapidly lose mass via bulk degradation, leading to simultaneous and randomized hydrolysis throughout the matrix.^64,203 The accumulation of bulk degradation byproducts also affects the degradation rate, leading to autocatalysis and making it difficult to predict the functional and degradation times.^296,297 In contrast, materials with low diffusion coefficients and relatively high reaction rates for liquid water and oxidants, such as silicone-based elastomers, metals, ceramics, and PHAs, mainly experience surface erosion (Figure 7D).^298,299 These bioresorbable materials lose mass linearly from their exposed surfaces, as determined by the static concentration gradient of degradation factors (Figure 7E).²⁹¹ In this case, autocatalysis is not considered a relevant factor because the byproducts diffuse away toward the surrounding tissue. As a result, their functional and disappearance times are easier to predict.²⁸³

Figure 7.

Macroscopic degradation mechanisms. (A) Bulk degradation process with a diffusion coefficient higher than the reaction rate. (B) FE-SEM images of bioresorbable polyesters before and after bulk degradation: PLGA, 33 d in PBS (pH = 7.4) at 37 °C. Reproduced with permission from ref (290). Copyright 2016 Springer Nature. (C) FE-SEM images of bioresorbable polyesters before and after bulk degradation: PCL, 20 h in PBS (pH = 7.2) containing 18 U·mL^–1 lipase at 45 °C. Reproduced with permission from ref (133). Copyright 2020 Springer Nature. (D) Surface erosion process with a diffusion coefficient lower than the reaction rate. (E) FE-SEM images of bioresorbable Fe–Mn alloys before and after surface erosion (three months in Hank’s solution at 37 °C). Reproduced with permission from ref (291). Copyright 2010 Elsevier. Profiles of mass, molecular weight, and mechanical strength of polymers during (F) bulk degradation and (G) surface erosion. Reproduced with permission from ref (292). Copyright 2008 Elsevier. (H) Theoretical plot of the erosion number for the hydrolysis of polymers, ε, depending on water diffusivity inside the polymer, D_eff, the dimensions of the polymer matrix, L, and the polymer bond reactivity, λ. (I) Critical thickness, L_critical, the threshold that a polymer specimen must exceed to undergo surface erosion. (H and I) Reproduced with permission from ref (174). Copyright 2002 Elsevier. (J) Mass profiles of surface-eroding polymers with different volume-to-surface area ratios during degradation. Reproduced with permission from ref (178). Copyright 2020 MDPI under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/).

Bulk degradation and surface erosion affect the diminishing trend of the mass, molecular weight, and mechanical strength of a polymer.^300,301 Bulk-degrading polymers exhibit nonlinear degradation dynamics (Figure 7F).²⁹² After a short induction period without significant degradation, the molecular weight first decreases significantly over time by chemical bond cleavage. Upon reaching a critically low molecular weight, the polymer undergoes a delayed degradation in mechanical strength and mass, finally leading to the mechanical failure of the implant.¹⁷⁸ In contrast, as shown in Figure 7G, ideal surface-eroding polymers lose mass, molecular weight, and mechanical strength almost linearly over time.³⁰¹ Therefore, the rate of mass loss and change in the dimensions of the polymer are determined by the surface area-to-volume ratio.

Determining the dominant macroscopic degradation mode of bioresorbable materials is critical for designing bioresorbable implants and predicting their operational and disintegration times under physiological conditions.^{190,228,295,302} This is due to the significantly different degradation behaviors of materials based on their macroscopic mechanisms, as mentioned above. However, many factors are interdependent, and erosion occurs in a complex and compositive manner, particularly in organic materials.^289,296,302 Burkersroda et al. reported a theoretical model for predicting the degradation mode based on the dimensions, reaction rates, and diffusion coefficient of a specimen.¹⁷⁴ In this model, the material dimensions and critical dimensions play a key role in determining the mode that dominates the erosion.^174,175,177 The erosion number, ε, is determined according to eq 1 as follows:

1

where L is the half-thickness of the material, λ is equivalent to the first-order rate constant employed in reaction kinetics and, therefore, directly related to the half-life of a polymer bond, D_eff is the effective diffusion coefficient, M_n is the number-average molecular weight, N_A is the Avogadro’s number, N is the degree of polymerization, and ρ is the polymer density. ε is dimensionless, with surface erosion dominating for ε > 1. L_critical is the value of L for ε = 1.^179,181Figure 7H provides a three-dimensional plot of ε with the aforementioned variables, indicating the determination of the macroscopic erosion mode. Thus, surface erosion occurs if half of the material thickness is larger than L_critical. Although most inorganic materials and metals have sufficiently low L_critical so that bulk erosion can be neglected, those of bioresorbable polymers normally are relatively high (above tens of micrometers), thus leading to bulk degradation except for several examples (Figure 7I).

Physical degradation, which occurs in response to severe mechanical stress, is a macroscopic physicochemical degradation mechanism that depends on a variety of physical elements, such as applied stress, shear stress, Young’s modulus, temperature, water content, intrinsic crack presence, intrinsic fracture energy, and fracture toughness.^18,173,303 High mechanical stress can break the bioresorbable system and halt the functions of the device. Figure 7J shows the influence of the surface area-to-volume ratio on the mass loss profile, indicating that a large surface area significantly increases the mass loss rate of a polymer film.¹⁷⁸ Mechanical disintegration also increases the surface area and diffusion flux of the degradation factors; thus, the fractured polymer possibly undergoes bulk degradation rather than surface erosion, which accelerates mass loss. Toughening bioresorbable materials is a key challenge in the development of implantable devices with reliable lifetimes and durability, as physical degradation can accelerate chemical degradation processes by increasing the surface area exposed to degradation factors. On the other hand, quick degradation upon mechanical disintegration can be exploited to realize on-demand transience.

3.2. Chemical Biodegradation Mechanisms

Although biodegradable polymers can experience various forms of degradation, such as photodegradation, thermal degradation, and mechanical degradation, from a molecular perspective, the chemical biodegradation mechanisms of bioresorbable materials can be categorized into hydrolytic,^175,228,295 oxidative,^176,289 and enzymatic degradation^57,172 (Figure 8). The majority of bioresorbable polymers have functional groups and bonds that readily react with water molecules in physiological conditions, leading to hydrolysis.^228,295 Carbonyl-derived functional groups, such as ester, amide, thioester, anhydride, imide, carbonate, urethane, and urea, are susceptible to hydrolysis.^175,304 Thus, many hydrolytic SBPs, such as polyesters, polyanhydrides, polyamides, polyethers, and polycarbonates, are hydrolyzable.^67,169,170 Hydrolysis leads to bond cleavage of the carbonyl moieties in these polymers, forming the corresponding bioabsorbable carboxyl acids and other byproducts. For example, the primary degradation mechanism of PLA, a widely used and commercialized SBP, involves the hydrolysis of ester bonds and decomposition to lactic acid monomers.¹³² Similarly, the ester groups of PCL hydrolyze in an aqueous environment to a caprolactone monomer, hydroxycarboxylic acids, and cyclic oligomers.¹³⁵ The imine functional group is also hydrolyzable, forming aldehydes and amines as byproducts. Polyimine undergoes similar degradation to diamine and dicarboxylic acid byproducts.

Figure 8.

Chemical biodegradation mechanisms, including hydrolysis, oxidation, and enzymatic processes.

Oxidation is a harsh degradation process that breaks polymer bonds through radical reactions with oxidants. The primary initiators of oxidative degradation in bioresorbable polymers under physiological conditions are peroxides, which are generated by immune cells such as neutrophils and macrophages.³⁰³ As these cells initiate phagocytosis, they simultaneously release oxidants, including peroxides, on the surface of foreign bodies to facilitate their clearance. The acute inflammatory response, which occurs from a few days to a few weeks, is usually short-lived. Depending on the chemical structures of the polymers, the oxidation process may continue once a sufficient radical concentration is achieved, as free radicals proliferate via a series of radical-forming chain reactions. The degradation rate and progression time largely depend on the number of polymer bonds susceptible to oxidation; thus, polyurethanes, polyolefins, and polyethers can be degraded by oxidation.³⁰³ The immune response-catalyzed oxidation may render the precise prediction of the functioning and disappearance times of bioresorbable devices difficult, as immune cell activity varies considerably among different individuals. Therefore, the degree of oxidation of the encapsulation polymer in bioresorbable devices should be considered.

The enzymatic degradation in bioresorbable implants is triggered by enzymes present in the extracellular matrix and body fluids.¹⁷² Several kinds of enzymes, including lipases, dehydrogenases, proteinases, diastases, and α/β-amylases, can participate in this enzymatic degradation under physiological conditions.^305−308 This degradation is influenced by several factors, including enzyme diffusion, reaction kinetics, and the diffusion of decomposition byproducts. Though various enzymatic reactions can occur, the majority involve hydrolytic reactions that can be effectively catalyzed by enzymes such as lipases, esterases, and proteases in both synthetic and natural polymers. Despite research efforts to measure the mass loss of bioresorbable polymers under in vitro conditions that mimic physiological media, such as water or aqueous solutions (e.g., bovine serum, artificial spinal fluid, and phosphate buffer solution), in vitro data are insufficient to predict in vivo lifetime of those polymers because of the complex enzyme content and interactions in biofluids and their variance among animal and human subjects.¹⁷² Therefore, if the employed polymer is susceptible to enzymatic hydrolysis in biofluids, it is recommended to perform a series of clinical tests on human subjects to obtain a database of ranges of functioning and disappearance times.

Material properties, such as monomer structure, molecular weight,³¹⁰ hydrophilicity,³¹¹ crystallinity,³¹² phase microstructure,²²⁸ and material processing, influence the polymer degradation rate and the macroscopic biodegradation mechanism (Figure 9A).¹⁷⁵ The crystallinity is determined by multiple variables, including material processing, phase separation, molecular weight, and monomer structure. In general, water diffusion is restricted within the crystalline lamellae of semicrystalline spherulites, whereas the amorphous region has a high diffusion coefficient (Figure 9B).³⁸ Thus, high crystallinity is largely responsible for the low degradation rate, leading to partial surface erosion.³¹¹ Considering water diffusion into hydrolytic polymers, increased hydrophobicity leads to slower degradation for polymers with the same molecular structure and composition (Figure 9C). In this context, a wide range of polymer architectures including copolymers, polymer blends, semi-interpenetrating networks (semi-IPNs), and nanocomposites have been used to attain high crystallinity and enhanced hydrophobicity (Figure 9D). These properties confer the long and predictable lifespan necessary for an encapsulation layer through surface erosion.

Figure 9.

Material-based factors that affect the rate of hydrolysis of polyesters. (A) Diagram of the material properties that influence polyester hydrolysis. (B) Illustration of restricted water diffusion due to high crystallinity. (C) Hydrophobicity decreases water diffusion and the degradation rate. (D) Polymer architectures to modulate biodegradation performance via crystallinity and hydrophobicity. (A–D) Reproduced with permission from ref¹⁷⁵. Copyright 2018 American Chemical Society. Sequential photographs showing (E) bulk biodegradation and (F) surface erosion of POC due to low and high cross-linking ratios, respectively (in PBS (pH = 7.4) at 37 °C). (E and F) Reproduced with permission from ref (230). Copyright 2022 American Chemical Society. FE-SEM images of the surface and core of (G) PHA and (H) a PHA–PLLA polymer blend to demonstrate the influence of hydrophobicity on biodegradation behavior. (G and H) Reproduced with permission from ref (309). Copyright 2006 Elsevier.

The sequential images of POC shown in Figure 9E and F illustrate its contrasting biodegradation behaviors through bulk degradation or surface erosion, which are contingent on the cross-linking ratio.²³⁰ At low cross-linking ratios, POC had a high diffusion coefficient for liquid water, leading to simultaneous hydrolysis in the bulk sample. Conversely, POC with a high cross-linking ratio had a lower diffusion coefficient, leading to predominantly surface erosion. Figure 9G and H provide another example of how the characteristics of materials can influence macroscopic degradation. PHAs, which typically undergo bulk degradation, can exhibit surface erosion when combined with hydrophobic PLLA.³⁰⁹ Upon degradation (45 d, 37 °C, water), pure PHAs were found to exhibit porosity in both the edge and core regions, owing to their hydrophilicity and high water diffusion. In contrast, the PHAs-PLLA polymer blend showed only morphological changes at the edge, with the core appearing to have densified due to its low glass transition temperature (41 °C) and recrystallization. This indicates a degradation process that is primarily limited to the surface and implies that these polyesters possess considerable hydrophobicity, a necessary condition for achieving surface erosion/degradation behavior.

In addition to material-originated properties, environmental factors are largely responsible for the biodegradation performance of polymers. As shown in Figure 10A, in the human body, there are highly acidic environments (e.g., intragastric, pH = 1–2.5) as well as weak basic environments (e.g., distal small intestine, pH = 6.8–7.88).^178,182,186 In an acidic environment (pH < 7.4), the same carbonyl group becomes protonated, rendering it receptive to nucleophilic attack by water (Figure 10B), whereas in a basic environment (pH > 7.4) the carbonyl group of the compound is vulnerable to hydroxide ion attack (Figure 10C).³⁰⁴ Due to different catalyzed hydrolysis mechanisms and responsible moieties, the degradation rates of certain materials can vary significantly between basic and acidic environments. Sailema-Palate et al. investigated the degradation of PCL under highly acidic (pH = 1) and highly basic (pH = 13) conditions and found that PCL samples degraded more rapidly at pH = 13 but exhibited less water uptake.³¹⁴ Additionally, PCL degraded more slowly than PDLA in highly basic media, which was attributed in part to the lower electrophilicity of PCL’s carbonyl carbon atoms.¹⁸² Conversely, PCL degraded faster than PDLA in highly acidic media due to the higher nucleophilicity of PCL’s carbonyl oxygen atoms.

Figure 10.

Environment-based factors and surface properties for the biodegradation of polymers. (A) pH circumstances in gastrointestinal organs of the human body. Reproduced with permission from ref (178). Copyright 2020 MDPI under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/). (B) Acid- and (C) base-catalyzed hydrolysis of polyesters. (D) Illustration of the influence of surface morphology on biodegradation performance. (E) FE-SEM images of a PLA solid film and a PLA porous scaffold before and after degradation in water at 60 °C for 14 d. (F) Mass profile of a PLA film vs hydrolysis time depending on the pore size. (E and F) Reproduced with permission from ref (179). Copyright 2011 American Chemical Society. (G) Schematic of how surface coating suppresses water diffusion and the adhesion of proteins and microsomes. (H) Schematic representation of nonmodified and zwitterionic polymer-coated beads, demonstrating antifouling methods against biotinylated serum proteins. Reproduced with permission from ref (180). Copyright 2018 American Chemical Society. (I) Super hydrophobicity was achieved by incorporating GOgODA nanosheets (NSs) into PLA aimed at a decrease in weight loss rate and moisture permeability. Reproduced with permission from ref (313). Copyright 2017 Royal Society of Chemistry.

The precise control of the lifetime of B-TENGs can be achieved by modulating the surface morphology and using an appropriate barrier layer that undergoes surface erosion.^179,180 Surface morphology influences diffusion by affecting the surface area available for diffusion, as well as the surface energy and reactivity of the materials (Figure 10D). A material with a rough surface has a large effective surface area that offers more pathways for diffusion, thus increasing the rate of diffusion. On the other hand, a material with a smooth surface may have lower surface energy and be less reactive, which can decrease the rate of diffusion. Additionally, surface features such as pores can act as diffusion pathways, allowing atoms or molecules to move through the material more easily. For example, Odelius et al. investigated the impact of porosity and pore size on the degradation profiles of PLA.¹⁷⁹ The PLA solid film initially had a smooth and homogeneous cross-section, but after 2 weeks of degradation in deionized water at 37 °C crystalline structures were formed throughout the bulk due to faster degradation of amorphous regions (Figure 10E). However, the structure close to and at the surface differed from the bulk, supporting the theory of an autocatalytic process contributing to faster degradation. The porous scaffolds had regular structures with homogeneously distributed pores, and pore size affected the degradation rate. Sharp pore edges were observed after 28 d of hydrolysis, indicating that degradation primarily occurred in the amorphous regions (Figure 10E). All of the materials experienced a decrease in remaining mass with hydrolysis time, as expected (Figure 10F). Initially, higher mass loss was observed for materials with larger pore sizes, while the solid films exhibited the fastest mass loss rate. These differences decreased with increasing degradation time, eventually leading to almost no remaining number-average molar mass of the samples. The faster mass loss and lower average molar mass of the nonporous films compared to the porous samples are attributed to an autocatalytic effect, where degradation products catalyze further hydrolysis in the solid films. Increased pore size also resulted in faster mass loss due to the greater thickness of the pore walls. This autocatalytic effect has been observed even for very thin films in previous studies.

Figure 10G summarizes the main effects of surface coating on biodegradable polymers. A zwitterionic polymer contains both positive and negative charges in its structure. Thus, using it as an encapsulation layer reduces biofouling and immune response and allows for improved system integration with the surrounding tissue (Figure 10H).¹⁸⁰ Zwitterionic polymers possess high hydrophilicity owing to their inherent polarity from the presence of both positive and negative charges.¹⁶ While the exclusive use of zwitterionic polymers as encapsulation layers might result in undesired high water diffusion and hydrolysis, it confers several benefits when the polymers are coated onto bioresorbable systems. In particular , the hydrophilic nature of zwitterionic polymers reduces the risk of radical attack by immune-response-originated reactive oxygen species (ROS) and the interaction with various proteins and microorganisms by inhibiting their adhesion.^315,316

Employing a highly hydrophobic surface is a promising strategy for inhibiting the bulk degradation of encapsulation polymers. Zhou et al. proposed a novel approach to enhance the durability and reliability of encapsulation polymers through the development of a superhydrophobic 3D network architecture,³¹³ incorporating interconnected graphene oxide-grafted octadecylamine (GOgODA) NSs into PLA (Figure 10I). To prepare uniform-sized PLA microspheres and superhydrophobic GOgODA, the building blocks were decorated using a vapor barrier-postmolding assembly. The resulting nanocomposite PLA films displayed excellent water resistance, with a water permeability coefficient 6.5× lower than that of pure PLA films, leading to a remarkable decrease in the weight loss rate. The surface modification technique utilized by the authors effectively limited water infiltration into the polymer matrix, which is a key factor in the degradation process. Thus, the lifetime of the encapsulation system was extended by reducing the rate of hydrolysis and other degradation mechanisms. A hydrophobic surface can also enhance the stability and mechanical properties of the polymer matrix, ultimately leading to improved performance and longer lifespan.¹¹⁹

In the context of B-TENGs, bioresorbable metals, such as Mg, Zn, Fe, Mo, and W, and their alloys serve as electrodes and conductors in the form of thin deposited films or plates.^62,65 In addition, biodegradable metals can provide mechanical support for biodegradable devices within biological tissues by forming metal–polymer layer-by-layer composites with relatively strong mechanical toughness.³¹⁹ As shown in Figure 11A, metals undergo oxidation in contact with body fluids, generating electrons that are consumed in cathodic reactions.³⁵ This produces hydrogen gas and hydroxide, forming a protective metal oxide layer on the surface.³²⁰ The protective oxide layer of magnesium alloys can be weakened by the high concentration of chloride ions found in body fluids, accelerating the degradation process. The presence of calcium and phosphate ions in body fluids and local alkalization lead to the deposition of calcium phosphate on the metal oxide layer during the degradation process, enabling cells to adhere to the surface and form tissues.³⁵

Figure 11.

Mechanisms of metal biodegradation. (A) Environment-specific biodegradation processes of metals based on their reaction with body fluid. Reproduced with permission from ref (317). Copyright 2019 Elsevier. (B) Specific biodegradation processes of metals based on their reaction with body fluids. Reproduced with permission from ref (283) Copyright 2018 Elsevier under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/). (C) Macroscopic observation of Mo degradation in a Kokubo SBF solution at 37 °C. Reproduced with permission from ref (285). Copyright 2020 Elsevier. (D) Optical micrographs of degraded pure Fe and Fe-based alloys after implantation into a growing rat skeleton. Reproduced with permission from ref (318). Copyright 2014 Elsevier. (E) Cell viability test according to the biodegradation of Mg and alloys to examine cytotoxicity and biocompatibility. Reproduced with permission from ref (282). Copyright 2016 PLOS under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/).

The degradation process of Mg involves biomineralization, with the participation of various ions and oxides.²⁸³ An overview of current knowledge on the degradation products of Mg under physiological conditions is shown in Figure 11B. Thus, the selection of an appropriate solution for in vitro investigation depends on the specific objectives of the study.³²¹ To screen for materials and compare degradation rates, simulated body fluid (SBF) and Earle’s balanced salt solution (EBSS) in combination with CO₂/HCO₃^– buffering are recommended.³²² In contrast, cell culture media containing fetal bovine serum (FBS), such as modified eagle medium (MEM), Dulbecco’s MEM, and α-MEM, are recommended to study the behavior and mechanism of degradation.^323,324 The latter culture media offer a degradation behavior closer to that of in vivo conditions but are more complex, require more technical and experimental efforts, and are not typically found in material science-oriented laboratories.³²¹ As the degradation rate is sensitive to changes in temperature and pH, in vivo investigations on the lifetime and degradation of metallic materials should be conducted for practical applications.

Zn,³²⁵ Mo,^227,285 and Fe^318,326 also have high biocompatibility and considerable biodegradation rates showing similar biodegradation mechanisms, which has led to the intensive commercialization of various implants including bioresorbable stents (Table 6, Figure 11C and D). The biodegradation of metallic components inevitably produces salts as byproducts, changing the pH of the surrounding tissue. Thus, biocompatibility and biosafety assessments of bioresorbable materials should be validated by investigating their cytotoxicity, pH changes, genotoxicity, and other factors. Mg has gained significant attention as a biomaterial because of its excellent biocompatibility and good resorbability.³²⁷ Extensive studies have been conducted on its biocompatibility under in vivo and in vitro conditions, revealing that its alloys are biocompatible regarding cell metabolism (Figure 11E).²⁸²

4. On-Demand Transient B-TENGs

The previous sections presented a general overview of bioresorbable materials and B-TENGs based on passive operation, which involves continuous degradation of materials upon implantation. The lifetime of bioresorbable systems relies solely on predetermined characteristics, including material properties, specimen dimensions, surface morphology, and environmental factors, and most current bioresorbable implants employ passive operation.^30,40,44 Therefore, substantial efforts have been dedicated to understanding the kinetics of physicochemical degradation in order to predict and control the lifespan of bioresorbable devices. This is achieved by choosing or modifying the constituent materials, ultimately paving the way for their practical applications in the biomedical field.^{172,174,175,177,228,230} Despite the noticeable accomplishments, bioresorbable implants with passive operation have inherent limitations regarding stable performance, noninvasive lifetime adjustment, and biosafe device removal due to gradually degrading characteristics. Recent advancements in the understanding of biodegradation processes, together with an increasing demand for novel material platforms, have inspired additional research on active operation to overcome the limitations of conventional bioresorbable implants.

Active operation and on-demand transience involve the rapid degradation of bioresorbable materials at an intended time by combining a noninvasive triggering event and stimulus-responsive materials.^46,55,70 This process allows the period between the end of the functional period of the device and its disappearance to be shortened, enabling more precise control over the device lifespan and minimizing the potential risks of prolonged exposure to residues. In contrast to passive operation, active operation is characterized by two distinct phases during the lifetime of the device.^55,56,71,328 During the first phase, the implanted materials maintain their functionality and lose mass at a slower rate. In the second phase, a physical or chemical stimulus triggers their rapid disintegration and dissolution. Recent studies have demonstrated triggered transience using two components: a stimulus-responsive encapsulation layer that controls biofluid infiltration and a rapidly degrading inner device that supports the overall device function.¹¹⁷ When the stimulus is applied, the encapsulation layer breaks down and the inner device becomes saturated with biofluid and ceases to function, thus leading to quick and simultaneous disappearance.

The ability to determine the end point of an active operation while possessing the benefits of passive systems is highly advantageous.^{58,117,123,329} First, an active system reduces patient burden by eliminating residual devices in the body. The passive system still has discomfort because it requires follow-up for monitoring degradation and, potentially, surgical device removal if the device lifetime exceeds the time frame of a relevant process, thus hindering its commercial viability.^10,13,40 In contrast, implants based on active operations can be immediately dissolved if no longer needed, simplifying the planning and execution of the medical process. In an ideal scenario, the only relevant information necessary for managing bioresorbable IMDs with active systems is their maximum functioning time under physiological conditions not imposing triggering events. As such, these devices provide improved biosafety and more precise control over their lifespan.^56,58,61,328 They ensure dependable maintenance, enable on-demand device removal, and diminish worries about unanticipated rapid dissolution or potential negative health effects from lingering residues in tissue. Incorporating active operation into B-TENGs has recently gained huge interest owing to its potential for temporally powering IMDs or self-powered biomedical applications. This section highlights recent research trends in on-demand bioresorbable systems. It covers not only B-TENGs but also other bioresorbable electronics and materials to unearth unexplored opportunities. We outline the use of stimulus-responsive materials and device designs, classifying them according to their respective triggering mechanisms.

4.1. Ultrasound-Triggered Transience

Ultrasounds are inaudible sounds with frequencies above 20 kHz. They are widely used in medical applications, such as diagnostic ultrasound or noninvasive power transfer, delivering high and low air pressure inside the body based on high tissue permeability. There have been reports of ultrasound-responsive drug delivery systems in which a remedial ultrasound mechanically disintegrates a polymer matrix, breaks the weak chemical bonds of polymers, or induces the controlled release of embedded degradation factors. Ultrasound-mediated technology exhibits considerable potential for an effective and safe active operation because of its high energy permeability and biosafety.

Lee et al. developed a fully bioresorbable ultrasound TENG consisting of a poly(3-hydroxybutyrate-co-3-hydroxyvalerate)-PEG (PHBV/PEG) polymeric composite triboelectric membrane, a PHBV encapsulation layer, and a Mg electrode, all of which could be noninvasively eliminated by applying ultrasounds (Figure 12A).⁵⁸ This device had the structure of a single-electrode-mode TENG, incorporating the characteristics of an ultrasound-driven implantable TENG reported in 2019.¹³⁸ The device, with an active area of 4 cm², generated a stable high-output voltage (4.51 V) and output current (27.86 μA) at an ultrasound intensity of 0.5 W·cm^–2. This performance was achieved through the ultrasound-mediated vibration of the PHBV membrane submerged in water. The output power of the B-TENG at 0.5 W·cm^–2 reached 17.24 μW·cm^–2 at an optimized electrical impedance of 20 kΩ. High-intensity ultrasound (HIU) can trigger mechanical disintegration of the device by intensifying acoustic pressure at the micropores in a PHBV film, facilitating on-demand transience (Figure 12B). The output power of the B-TENG drastically decreased when an HIU (3 W·cm^–2) was applied, and no electrical signal was observed after 40 min. The potential of this B-TENG as an on-demand transient triboelectric implant was demonstrated by its output generation and successful fracture under ex vivo conditions (in porcine tissue). An output voltage of 544 mV was achieved under 0.5 mm of porcine skin and was almost maintained for 21 d. The device stopped generating power by triggering events for 120 min, indicating sufficient mechanical disintegration. The recent advances in ultrasound technology are anticipated to enable more sophisticated on-demand transience while ensuring biosafety, thereby achieving commercial viability. Given the variety of available waveforms, frequencies, and focusing options, the use of ultrasounds could potentially enhance both the rate of active degradation and biological safety in the future.^330−332

Figure 12.

Ultrasound-triggered on-demand transience in B-TENGs and polymers. (A) Mechanism of triggered biodegradation by intensified acoustic pressure in the micropores of a PHBV encapsulation layer. (B) Photographs of degrading PHBV and PHBV/PEG films over time. (A and B) Reproduced with permission from ref (58). Copyright 2022 American Association for the Advancement of Science. (C) Structural design of an on-demand B-TENG for bacterial inactivation and (D) improved biodegradation rate of its constituent membranes by applying high-intensity ultrasounds (HIU). (C and D) Reproduced with permission from ref (117). Copyright 2023 Wiley-VCH under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/). (E) Self-clearance mechanism of an alginate hydrogel-TiO₂ NPs composite. Reproduced with permission from ref (329). Copyright 2022 American Chemical Society. (F) Degradation mechanisms of mechanically gated degradable polymers and decrease in molecular weight when subjected to sonication. Reproduced with permission from ref (339). Copyright 2020 American Chemical Society.

Further work by Imani et al. explored the application of B-TENGs as antibacterial agents, implementing active operation.¹¹⁷ This study introduced a novel approach for eradicating microorganisms deep within tissues to prevent surgical site infections (SSIs) using an ultrasound-driven fully bioresorbable B-TENG (Figure 12C). Once the device reaches the end of its programmed lifetime, it can be dissolved on-demand by controlling the intensity of the ultrasound, eliminating the need for surgical removal. Both Mg and PHBV, the primary component materials, degraded gradually over 8 weeks, maintaining their structures (Figure 12D). However, when 5 W·cm^–2 of ultrasound irradiation was applied, the PHBV fully disintegrated within 120 min and the PVA triboelectric layer inside rapidly dissolved within 20 min. Thus, the B-TENG lost all its weight and disappeared shortly after PHBV disintegration.

Wang et al. reported a self-clearance hydrogel composite that can be dissolved by ultrasonic triggering events (Figure 12E).³²⁹ The researchers prepared a hydrogel composite comprised of a sodium alginate hydrogel and titanium dioxide nanoparticles (TiO₂ NPs). The sodium alginate hydrogel, cross-linked with a reactive oxygen species (ROS)-cleavable thioketal, acts as a ROS-responsive material, and the TiO₂ NPs generate ROS in response to ultrasonic waves. The on-demand degradation is accomplished through a combination of two elements: a ROS generator and a ROS-responsive matrix.^333,334 More specifically, the acoustic pressure from the ultrasound induces a piezoelectric potential in the TiO₂ NPs, triggering an electrochemical reaction with water to generate ROS.^335,336 These ROS, in turn, break the thioketal bonds within the hydrogel matrix through radical reactions.^337,338 The degradation process continues as free radicals multiply through a series of radical-forming reactions. As a result, the hydrogel composite transforms from a colloid to a semifluid and dissolves completely within 60 days. The researchers successfully demonstrated the efficacy of this mechanism by using the self-clearing hydrogel composite to reversibly block the sperm duct in a mouse model, creating an efficient contraception device. The relatively low and biosafe intensity of ultrasound required to generate ROS, combined with the existence of many ROS-generating species (such as barium titanate or BTO) and ROS-cleavable chemical bonds, opens a wide range of potential applications for these active operation systems.

Mechanically gated degradable polymers were reported by Lin et al. (Figure 12F).³³⁹ In this study, a cyclobutane (CB) mechanophore was used as a mechanical gate in the creation of acid-degradable polymers by regulating an acid-sensitive ketal integrated into the polymer structure. In the absence of acid, the polymer remained intact and could only be degraded by high mechanical forces, and ultrasonication was used to reduce its molecular weight to 28 kDa. The addition of acid after ultrasonication led to a further decrease in molecular weight to 2.5 kDa. The mechanical strength of the ungated ketal was comparable to that of traditional polymer backbones, and a force of 2 nN was required to activate the CB gate in 100 ms according to single-molecule force spectroscopy.

4.2. Thermally Triggered Transience

Thermoresponsiveness is the ability of a polymer to drastically change its physicochemical properties with the temperature.⁴³ In bioresorbable devices, physicochemical properties of thermoresponsive materials such as solubility in biofluids, water content, and wettability can be changed with temperature variations, initiating decay and decomposition. Thermally triggerable transient bioresorbable devices are most studied in this field, with many different reported thermoresponsive mechanisms (Figure 13A). An increase in temperature is commonly considered a direct triggering event that can cleave primary or secondary bonds of materials.^203,340 In other instances, the heating or cooling of materials can also prompt phase changes in polymers, such as those exhibited by lower and upper critical solution temperature (LCST and UCST, respectively) polymers.^341,342 Changes in temperature also can stimulate a coil-to-globule transition, and vice versa, enabling polymers to retain significant quantities of water and thus degrade more rapidly. When these materials are incorporated into a drug delivery device, heat can trigger degradation by facilitating the release of degradation factors stored within an independent reservoir.^73,343Figure 13B summarizes thermoresponsive transient materials according to the applied temperature and duration of application.^{60,61,68,73,203,340,342,343} The green region indicates the time–temperature thresholds for thermal injury of the human skin.³⁴⁴ While many of these methods could potentially harm living organisms due to high triggering temperatures and prolonged triggering times, there are some notable instances where biosafety has been achieved. These include the triggered transience of methyl cellulose,³⁴² wax,⁶¹ and gelatin/chitosan composites.²⁰³ Therefore, with rigorous research efforts aimed at reducing the triggering temperature and time, it is anticipated that the implementation of biosafe and effective active operation devices based on thermoresponsive polymers will be achievable in the near future.⁵⁵

Figure 13.

Thermally triggered transience. (A) Diagram of the strategies employed to implement thermally responsive degradation and (B) review of triggering temperature and time of thermoresponsive polymers compared to the threshold curve for human skin thermal injury. (C) Photothermally tunable degradation of a PLA-based B-TENG using laser treatment. Reproduced with permission from ref (137). Copyright 2016 PLOS under the terms of the Creative Commons Attribution License. (D) Photothermally tunable degradation of a chitosan-based B-TENG using laser treatment. Reproduced with permission from ref (122). Copyright 2018 Wiley-VCH.

Li et al. focused on the development of fully bioresorbable TENGs as sustainable power sources for healthcare devices with the capability of photothermally tunable biodegradation (Figure 13C).¹³⁷ The authors fabricated B-TENGs with a hemisphere-array-based structure that included a POC triboelectric layer, electrodes, and gold nanorods (Au NRs). These TENGs showed promising results, with peak output voltages of 28 (in vitro) and 2 V (in vivo) when external pressure was applied. The degradation process was effectively controlled through the incorporation of Au NRs that responded to near-infrared (NIR) light. Au NRs convert NIR to heat energy, increasing the local temperature of the device. The small radius of Au NRs increased the conversion efficiency and generated sufficient heat to trigger the biodegradation of the POC. Treatment with a NIR laser for 5 min (808 nm wavelength, 5.2 W·cm^–2) resulted in a temperature change of 13.2 °C. The B-TENG based on PLGA stopped working after 3 (in vitro) and 28 d (in vivo) of applying NIR light. These results showed that a NIR light can trigger the erosion of SBPs, such as PLGA, PLA, and PCL in a rat model, demonstrating the effectiveness of photothermally tunable TENG biodegradation. Want et al. fabricated TENGs with tunable biodegradation rates (Figure 13D),¹²² and a Q-switched neodymium-doped yttrium aluminum garnet (Nd:YAG) laser (532 nm) was used to trigger the biodegradation process. The partially bioresorbable TENGs were constructed from chitosan-based film containing 10% acetic acid, lignin, starch, or glycerol. Laser treatment with a larger number of pulses significantly changed the surface morphology to be rougher and more wrinkled.

Despite promising prospects and ongoing explorations in the field of bioresorbable electronics and chemistry, the use of thermoresponsive materials for triggered transience has been seldom applied to TENGs, with only a few examples in existence. However, the current state of research suggests that there are abundant opportunities for B-TENGs to explore the use of materials that have been previously studied in other fields but remain unexplored. This can further expand the potential of TENG technology and facilitate advancements in triggerable transient materials and devices given the wide variety of material options available. In this context, the following sections provide a comprehensive introduction to triggerable biodegradation mechanisms, including relevant examples from various fields.

Fukada et al. reported a thermally degradable gelatin/chitosan hydrogel as the substrate of bioresorbable inductors (Figure 14A).²⁰³ They developed a cross-linked gelatin/chitosan hydrogel and tailored its thermal response by optimizing the concentration of the cross-linker, genipin, to prevent dissolution. The gel resisted transitioning to a liquid state upon heating, and with genipin concentrations exceeding 1 mM it maintained its shape in warm water (37 °C) containing a gelatin/chitosan mix (2.45 g/0.05 g). Consequently, the degradation rate could be managed, as the hydrogel dissolves in response to a sufficiently high, preprogrammed temperature. This study demonstrates that stimuli-responsive, on-demand bioresorbable systems can be fashioned from a passive system by accurately adjusting the degradation kinetics.

Figure 14.

Thermoresponsive transient electronics. (A) Thermally degradable inductors based on gelatin–chitosan hydrogel films. Reproduced with permission from ref (203). Copyright 2022 American Chemical Society. (B) Temperature-dependent transience originating from the LCST behavior of a Ag NW/methylcellulose composite. Reproduced with permission from ref (342). Copyright 2017 American Chemical Society. (C) Schematic diagram of the phase transition via LCST behavior. (D) Thermally triggered transience using a wax-encapsulated acid. Reproduced with permission from ref (61) Copyright 2015 Wiley-VCH. (E) Wireless transient microfluidic system with a heat-expandable polymer for controlled release. Reproduced with permission from ref (343). Copyright 2015 Wiley-VCH.

Phase transitions instigated by heating or cooling can offer an attractive mechanism for triggering rapid degradation. This approach enables precise control over the critical temperature by adjusting the composition, effectively providing effective transience while preventing any undesired triggering event.^345−349 LCST and UCST polymers are two types of thermoresponsive polymers that undergo phase transitions in response to temperature changes. LCST polymers, possessing a lower critical solution temperature, transition from a soluble state to an insoluble state as the temperature rises above the LCST. Conversely, UCST polymers, characterized by an upper critical solution temperature, shift from an insoluble state to a soluble state as the temperature increases beyond the UCST. Those transitions do not occur gradually even at temperatures around the threshold, thus they are advantageous in ensuring selectivity.³⁵⁰

Zhang et al. developed thermoresponsive conducting composites, combining methyl cellulose with LCST behavior and conductive Ag NWs (Figure 14B).³⁴² The LCST is insoluble above a threshold temperature but soluble below that temperature, as the polymeric binder becomes hydrophilic and dissolves in water (Figure 14C).^348,349 The electrical conductance of the composite was maintained for over 24 h at body temperature and drastically decreased after 5 s at a lower temperature of approximately 20 °C. The methylcellulose was cooled by immersion in water. Although a specific method to cool implanted devices in vivo was not suggested, the highly selective transience of LCST polymers is noticeable. In contrast to LCST polymers, UCST polymers undergo hydrophobic-to-hydrophilic transitions upon heating. Thus, they can also be utilized in active operations. In some respects, UCST polymers may be more feasible than LCST polymers in a clinical setting, given that it is often simpler to increase the temperature of implants than to decrease it.^341,351 However, UCST behavior is only observed in specific polymer systems, such as poly(acrylonitrile) and poly(N,N-diethylacrylamide).^341,351 Therefore, there are limited options for UCST polymers that meet other requirements, such as mechanical toughness and output performance, resulting in fewer reported studies on their use in active operation.

Figure 14D and E present active operation systems integrated with drug delivery.⁶¹ Over the past few decades, drug delivery technology has seen significant advancements, enhancing the efficacy and safety of pharmaceutical treatments. The integration of drug delivery systems with bioresorbable devices offers immense potential for executing on-demand transience, coupled with the controlled release of chemicals that initiates device removal.^352−354 In this context, Park et al. employed thermally responsive transient electronic devices for on-demand drug delivery applications (Figure 14D).⁶¹ This study highlights the use of a wax composite that melts upon exposure to sufficient heat and releases the encapsulated acid, which facilitates the rapid destruction of the device through acidic degradation of the electronic components. The sequestration and encapsulation of methanesulfonic acid (MSA) were accomplished by melt casting an acid/wax emulsion, and a wireless RF heater^345−347 was designed as a serpentine-shaped resistor and spiral inductor coils. In addition to the degradation of Mg traces, the Si-based devices underneath were also destroyed by the MSA via acid-triggered depolymerization of the cyclic poly(phthalaldehyde) (cPPA) substrate. A 40% MSA/wax was triggered at 55 °C, releasing acid that degraded the Mg electrodes and the cPPA substrate within 20 s. The device fully disappeared within 1 min of triggering. This temperature-sensitive approach precisely controls the drug release kinetics, enabling on-demand transience for a targeted and efficient therapeutic intervention.

Lee et al. reported a wireless microfluidic system with three reservoirs of metal or ceramic material etchants able to release stored etchants through serial communication with an infrared light-emitting diode (Figure 14E).³⁴³ An infrared (IR) transmitter and receiver were employed to regulate the Joule heating of heater elements positioned beneath the reservoirs and the heat-expandable polymers. These three separate heater elements were crafted from gold serpentine traces (300 nm thick) on an FR4 substrate. The FR4 substrate was selected due to its conventional use in microelectronics and its low thermal conductivity (0.4 W m^–1·K^–1), enabling it to reach high temperatures with minimal power input. A battery supplied approximately 286 mW of power to each heater, which triggered Joule heating and resulted in peak temperatures around 100 °C within about 20 s of wireless activation using an IR remote control unit. The subsequent expansion of the heat-expandable polymers created mechanical pressure on the reservoirs, pushing the enclosed chemicals through the outlets. This mechanism facilitated the ejection of approximately 2.64 μL of water through the microfluidic channels, across the contained chemical powders, and ultimately onto the targeted electronics. The produced etchants were then capable of partially or fully dissolving layers of silicon (Si), silicon dioxide (SiO₂), aluminum (Al), and other functional materials.

Although the development of a bioresorbable reservoir remains uninvestigated, incorporating a reservoir that can be gated by external stimuli into a bioresorbable device provides several advantages for implementation in an active operation, enabling fast and immediate dissolution by the drugs or etchants in the miniaturized reservoir. Additionally, device removal can be initiated by minimized stimulation; the gate is considerably smaller than the entire device and requires a much lower triggering energy than that required to degrade all device materials, minimizing any adverse health effects. However, ensuring the high mechanical durability of the gate to prevent accidental elimination of the device remains challenging. Furthermore, its biosafety must be evaluated, as the sudden release of the stored drug may damage tissues or lead to acute inflammation.

4.3. Light-Triggered Transience

In addition to the aforementioned triggers, many different mechanisms lead to on-demand transience. Active operational mechanisms can also be triggered by light. Hernandez et al. fabricated phototriggerable transient electronics on cPPA.⁵⁶Figure 15A shows the transience of silicon-based active devices on 2-(4-methoxystyryl)-4,6-bis(trichloromethyl)-1,3,5-triazine (MBTT)/cPPA films exposed to ultraviolet (UV) light (379 nm) for 230 min. MBTT films are photoacid generators (PAGs) that release chlorine radicals upon exposure to UV light. These radicals abstract hydrogens from the environment to generate hydrochloric acid (HCl). HCl then reacts with the acetal backbone of cPPA, initiating a degradation process that ultimately causes the electronics to disintegrate and be destroyed. Integrating PAGs with acid-degradable polymers can be a promising method for implementing light-triggered active operation. There are various PAGs available with different responsiveness and acid-generating wavelengths, such as N-hydroxynaphthalimide triflate (254 nm)³⁵⁵ and Rhodorsil Faba (365–405 nm).³⁵⁶ Hence, several material options are available for imparting UV light.

Figure 15.

Light-triggered transient electronics. (A) Photographs of photoresponsive transient electronics based on MBTT/cPPA films and schematic diagrams of working mechanisms. Reproduced with permission from ref (56). Copyright 2014 Wiley-VCH. (B) Photographs of a hydrogel that transitions from gel to sol by UV light, optical microscopy images of a Mg electrode that consequently undergoes hydrolysis, and schematics of the working mechanisms. Reproduced with permission from ref (70). Copyright 2018 American Chemical Society.

Figure 15B illustrates the UV light-triggered transience and the corresponding mechanism of light-responsive hydrogel and oxide.⁷⁰ Upon exposure to UV light (365 nm and 300 mW·cm^–2), a hydrogel based on 4-arms-polyethylene glycol-NH₂ undergoes a gel-to-sol phase transition. This shift is the result of photoinduced cleavage of azo bonds causing the degradation of the three-dimensional network and thereby increasing water permeation through the oxide layer, which in turn accelerates degradation. This process offers numerous advantages over other forms of triggered transience. A key advantage lies in the ability to achieve precise spatiotemporal control over the degradation process. By modulating the wavelength, intensity, and duration of light exposure, the material’s degradation can be selectively targeted to a specific area and controlled meticulously over time. While light-triggered transience might be limited to superficial implants due to low light transmittance in deeper tissues, and potential risk for heat injury, the capability to precisely manage light incidence and focus presents a valuable option.

5. Material Design for Large Output Power

Device miniaturization is a crucial requirement that can reduce physical harm and negative consequences to the surrounding tissues (e.g., cytotoxicity, genotoxicity, and immune responses), thus achieving seamless implantation.^29,44 Small formfactors enable the utilization of IMDs in confined areas in vivo, thereby broadening their possible clinical applications.^27,33 However, a diminished active area of triboelectric layers results in a low output, which significantly impairs the viability of IMDs in commercial and medical fields. Therefore, achieving a high output power per device size is a pressing demand in this field.¹⁴⁰ The high surface charge density of triboelectric layers is highly contributed to the large output power; thus, intensive research efforts have been directed toward developing triboelectric materials that have large amounts of charges via a triboelectric event.^{146,357−359} Despite significant advancements in the output performance of nondegradable conventional TENGs in the past decade and ongoing efforts in the B-TENGs field, there is still a need to considerably enhance the output level of B-TENGs, and developing high-performance, bioresorbable triboelectric materials remains a challenge due to the short history of B-TENGs and the relatively limited choices of triboelectric materials available. Diverse material development strategies have been tried to achieve large-power-output bioresorbable tribo-materials, which include unearthing high-charge polymers/polymer blends, functional group modification, embedding NPs, ion doping, surface nanostructuring, and changing surface chemical structures through surface functionalization.^{199,209,225,360,361} In this section, we present various research efforts and achievements, each categorized under a specific material design strategy, to provide a roadmap for the future development of high-performance bioresorbable triboelectric materials.

Figure 16.

Representative strategies to develop large power output bioresorbable triboelectric materials.

5.1. High-Charge Polymers

NBPs have been widely used as triboelectric membranes for B-TENGs due to their inherent biocompatibility, however, their triboelectric properties still need to be enhanced.^118,201,362 The triboelectric polarity of NBPs can be improved by decorating them with highly electron-donating/attracting functional groups or manufacturing novel polymer composites. Figure 17A shows a facile method to modify the functional groups in molecular structures.³⁶⁰ Chitosan, usually obtained by the alkaline deacetylation of chitin, was employed due to its exceptional biocompatibility and abundance of hydroxyl and amino functional groups, enabling the modification of its chemical structure. First, the chitosan precursor was dissolved in citric acid and thus hydrolyzed. The acidic conditions facilitated the hydrolysis of both main -chain glycosidic linkages and side=chain N-acetyl linkages.^363,364 Following fragmentation, chitosan, and citric acid established a 3D network with cross-linking junctions between the fragments.³⁶⁵ The creation of hydrogen bonds also contributed to the construction of the 3D network.²⁰⁵ The composition of chitosan and citric acid was optimized to mass ratios from 1:2 to 1:6 by investigating the Tyndall effect.^204,366 The composite films were designated according to their mass ratios (1:2, 1:3, 1:4, 1:5, and 1:6) as CC-1, CC-2, CC-3, CC-4, and CC-5, respectively.

Figure 17.

Development of high-charge polymers to achieve large power output B-TENGs. (A) Schematics of chemical structures and 3D networks of a chitosan–citric acid (CC) polymeric composite. (B) Photograph of a transparent and flexible CC-TENG. (C) Comparison of the output current density of CC-TENGs based on CC-1 and CC-4 composites. (D) Changes in the output current density of CC-TENGs at different chitosan/citric acid ratios. (A–D) Reproduced with permission from ref (360) Copyright 2019 Wiley-VCH under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/). (E) Optical microscopy and SEM images of chitosan-based composite membranes with different nature-derived additives, demonstrating their transparency and changes in surface morphology. (F) Electrical output current of chitosan-based TENGs along with the type of additives incorporated. (E and F) Reproduced with permission from ref (122). Copyright 2018 Wiley-VCH. (G) 3D network of nature-driven ϰ-carrageenan–agar composites. (H) Kelvin probe force microscopy (KPFM) measurement showing an increase in surface potentials of composite films. (G and H) Reproduced with permission from ref (118). Copyright 2022 Elsevier. (I) Schematics and SEM images of cellulose-loaded PVA film (CPF) containing microcrystalline cellulose (MCC) particles and a PVA matrix polymer. (J) Output currents of B-TENGs based on bare PVA, CPF, and microarchitectured CPF (MACPF) that demonstrate highly improved performance. (I and J) Reproduced with permission from ref (371). Copyright 2019 Elsevier. (K) Enhanced output voltages of partially bioresorbable TENGs based on chemically functionalized cellulose nanofibrils (CNF) with nitro and methyl functional groups. Reproduced with permission from ref (202). Copyright 2017 Wiley-VCH.

Figure 17B shows a photograph of a partially bioresorbable chitosan–-citric acid TENG (CC-TENG) with excellent transparency and flexibility. The electrode layers of the CC-TENG were prepared by mixing the as-prepared chitosan-citric solution with AgNWs and then drying it. The CC-TENG, featuring a double-electrode structure, generated an electrical output during periodic contact-separation motions with a polyimide film serving as a counter triboelectric layer. The authors measured the output current density of CC-TENGs with different composite mass ratios. As a result, a higher citric acid content increases the output power of the CC-TENG, as shown in Figure 17C. The CC-TENG containing the CC-4 composite film (mass ratio of 1:5) generated a current density of 125 μA·m^–2, while the device containing the CC-1 composite film only generated 17 μA·m^–2. The output current densities at different citric acid concentrations are demonstrated in Figure 17D, reaching the maximum value at CC-4 and decreasing at a higher content level (CC-5). The authors stated that the degradation of output performance at CC-5 could be attributed to the uneven surface morphology of the composite.

Another strategy to enhance the triboelectric properties of chitosan involves forming polymeric composites with other nature-derived polymers. This method allows the improvement of the surface charge density of chitosan while maintaining low toxicity and cost using an easy synthetic process. Figure 17E illustrates an investigation of high-performing bioresorbable triboelectric polymeric composites that comprise chitosan with nature-derived additives including lignin, starch, glycerol, and acetic acid.¹²² These additives are all environmentally friendly materials: lignin supports the structure of plants,³⁶⁷ starch is obtained from abundant food sources (e.g., potato and rice),³⁶⁸ glycerol is a common component of lipids in plants and animals,³⁶⁹ and acetic acid can be produced through bacterial fermentation.³⁷⁰ These polymeric composites were obtained by dissolving their precursors under weakly acidic conditions using 2% (v/v) acetic acid. The resulting solution was stirred until homogeneous and drop-cast onto a plastic mold at room temperature. The formation of intramolecular and intermolecular hydrogen bonds between hydroxyl (−OH) and amine (−NH) groups contributes to the high strength and low ductility of the as-prepared composite films. Figure 17E presents the optical images of transparent chitosan-based composite films. However, the film containing lignin was not transparent due to its inherent opacity and insolubility in water. The SEM images of the surface topography of the composites also revealed that the composite containing lignin had a segregated area of dense particles when compared to the others.

To confirm the improvement in the power generation capability, partially bioresorbable TENGs were fabricated by assembling chitosan-based composite films with a polyimide counter triboelectric layer and aluminum electrodes. A series of triboelectric output measurements were then conducted (Figure 17F). The chitosan–10% acetic acid composite film hit the largest output voltage and current, possibly due to the positive triboelectric polarity of acetic acid.³⁷² Moreover, the excessive content of acetic acid protonates the chitosan, thereby further enhancing its charge-donating ability. Next, the chitosan–lignin composite and pure chitosan exhibited similar triboelectric properties, resulting in comparable triboelectric output performances for the TENGs based on them. Next, the chitosan–starch composite was positioned on the negative side of the triboelectric series, demonstrating electron-attracting behavior as indicated by the reversed output peak direction. This is due to starch being highly tribo-negative, which is attributed to its high hydroxyl group content. Finally, TENGs using chitosan–glycerol composite films generated higher triboelectric outputs than others; however, the electrical output signals were not stable, exhibiting more noise-like waveforms (Figure 17F). This is explained by the stickiness of the chitosan–glycerol surfaces. Consequently, a maximum power density of 17.5 μW·m^–2 was achieved using the 10% acetic acid–chitosan composite.

A nature-derived ϰ-carrageenan–agar (ϰC–agar) composite was used as the triboelectric material to fabricate a fully bioresorbable TENG.¹¹⁸ Both ϰC and agar are edible polysaccharides obtained from red seaweed.^198,373 The 3D aggregated structure of the ϰC–agar composite exhibited an abundance of charge-trapping sites composed of calcium cations (Ca²⁺) and sulfate ester groups (Figure 17G), thus leading to highly electron-donating characteristics. The ϰC–agar composite film was simply prepared by a drop-casting method using an aqueous solution of the precursors and glycerol plasticizer. Figure 17H illustrates changes in the surface potential of the composites via the composition measured using Kelvin probe force microscopy (KPFM). As a result, surface potential notably increased to 1.15 V at 80% ϰC concentration, revealing the composite’s superior charge-donating ability. A fully bioresorbable TENG was fabricated to demonstrate the potential for an implantable temporally powering solution, employing Mg metal plates as electrode layers and pristine PCL as the counter triboelectric layer. As a result, the optimized TENG generated the highest root-mean-square (RMS) output current density of 0.45 mA·m^–2 and a RMS output power density of 0.15 mW·m^–2 at 9.5 MΩ external impedance.

PVA has also been widely explored as a bioresorbable triboelectric material because of its cost-effectiveness, low cytotoxicity, and easy fabrication; however, its weak triboelectric polarity has hampered achieving a large output.²⁰² As shown in Figure 17I and J, loading microcrystalline cellulose (MCC) in a PVA matrix significantly increases the surface charge density.³⁷¹ MCC powder was obtained from refined and crystallized squandered cotton by isolating them through centrifugation and drying. Then, to fabricate a flexible cellulose-loaded polymer film (CPF), the MCC powder was added to a PVA solution and poured into a mold, followed by drying. KPFM measurements were conducted to determine the work functions and evaluate triboelectric polarities of the CPF samples at different MCC contents. The work function reached to a maximum value of approximately 4.8 eV at 2.5 wt % cellulose concentration, indicating the superior triboelectric charge transfer. For further enhancement, microarchitectured CPF (MACPF) films that possess large surface areas were prepared by drop-casting a PVA/MCC solution onto a micropyramidal textured silicon mold. Then, to verify the enhancement in output power, partially bioresorbable TENGs based on bare PVA, CPF, and MACPF were fabricated by assembling each with Al electrode layers and a polyimide counter-triboelectric layer. Figure 17J shows the triboelectric output current of prepared TENGs. Upon applying a mechanical force of 4–5 N at 5 Hz, the CPF-based TENG generated a V_OC of approximately 400 V and an I_SC of approximately 25 μA. The MACPF-based TENG exhibited the highest triboelectric energy generation capability, with a V_OC of approximately 600 V and an I_SC of approximately 40 μA. The MACPF-based TENG had a power density of 84.5 W·m^–2, presenting high mechanical durability and reliability in energy generation.

Figure 17K shows that chemically functionalized cellulose nanofibrils (CNF) exhibited large triboelectric polarity.²⁰² CNF films were treated with aqueous HNO₃ and H₂SO₄ solutions to prepare nitro-functionalized CNF films (nitro-CNF). Methyl-functionalized CNF films (methyl-CNF) were prepared by methylating refined CNF using dimethyl sulfate. The introduction of functional groups through polymer grafting significantly altered the surface potential of the CNF-based films, as assessed by KPFM measurements. The nitro-CNF film exhibited a more negative triboelectric polarity, exhibiting an average surface potential of −1.3 V, which is more negative than the −330 mV surface potential of the untreated CNF films. In contrast, the methyl-CNF film demonstrated tribo-positive properties, with a measured surface potential of 1.1 V. The CNF-based films were then integrated with indium tin oxide (ITO)/PET electrode layers to construct partially bioresorbable CNF-based TENGs, then the output of TENGs was measured, as shown in Figure 17K. The TENG based on pristine CNF film had the smallest V_OC output of approximately 0.8 V. In contrast, the nitro-CNF TENG and the methyl-CNF TENG generated over 4× larger V_OC outputs of 4.9 and 3.7 V, respectively. In precise terms, the nitro-CNF possessed a negative surface charge density of 85.8 μC·m^–2. Conversely, the methyl-CNF carries a positive surface charge density of 62.5 μC·m^–2. When compared to fluorinated ethylene propylene (FEP), one of the most tribo-negative materials, these values represent 71% and 52% increases, respectively. Hence, the polymer grafting technique proved to be an efficient strategy for modulating the triboelectric polarity of CNF films and, potentially, other bioresorbable tribo-materials.

5.2. Nanoparticle Composites

The output performance of a B-TENG can be enhanced by the inclusion of high-permittivity nanoparticles (NPs). This enhancement is due to the high dielectric constant of the nanocomposite, which amplifies the electrical potential difference induced in the working electrodes. Moreover, as presented by some previous studies, these NPs can boost the triboelectric charge-trapping capabilities of the nanocomposite, leading to an increase in the surface charge density. Equation 2 serves as the theoretical basis for understanding a proportional relationship between the relative permittivity, surface charge density of the triboelectric materials and the triboelectric output performances.

2

Here V is the open-circuit voltage, σ is the surface charge density; Q is the triboelectric charges on the dielectric surface; d and ε are the thickness and relative permittivity of the triboelectric layer, respectively; S is the active area; X(t) is the distance between the triboelectric materials; and ε₀ is the vacuum permittivity.^158,258

First, a BTO-embedded cellulose aerogel paper was explored (Figure 18A–D).¹⁹⁹ The SEM image shows a microscopic view of the BTO-embedded cellulose film, in which BTO NPs are evenly distributed (Figure 18B). This is due to the strong hydrogen bonds between BTO NPs and hydroxyl groups in the cellulose chain networks, therefore preventing NP aggregation and sedimentation. The BTO concentration was optimized by electrical characterization of the nanocomposites. Nanocomposites with different mass ratios (0:1, 1:1, 3:1, and 5:1) of BTO NPs and precursor cellulose were designated as pure cellulose, C/BT-1, C/BT-3, and C/BT-5, respectively. To explore the charge-trapping capabilities of the nanocomposites, their dielectric constants were experimentally measured in the 10²–10⁶ Hz range (Figure 18C). The superior relative permittivity of the BTO NPs led to dielectric constants of the nanocomposites that were higher than those of pure cellulose. The C/BT-5 nanocomposite exhibited a dielectric constant of 6.25 at 10³ Hz, which is more than double the value observed for pure cellulose (3.06). Electrical characterization for partially bioresorbable TENGs based on the nanocomposite was conducted to verify the influence of NPs on the power generation capability (Figure 18D). This was conducted using a linear motor to apply an external mechanical force of 12 N at an operating frequency of 2 Hz. As the BTO NP loading in the aerogel paper increases, both the V_OC and I_SC rise gradually, peaking at values of 50 V and 5.1 μA, respectively, for the C/BT-3-based TENG. However, further increments in BTO NPs resulted in a decrease in these values. This is possibly due to the presence of exposed BTO NPs on the surface of the composite, which decreased the effective contact area of cellulose during the triboelectric process.

Figure 18.

NP-embedded composites for large power output. (A) Photograph and (B) SEM images of a cellulose aerogel film containing BTO NPs. (C) Dielectric constants of C/BT-1,3,5 and pure cellulose and (D) improved electrical output voltage of the cellulose aerogel/BTO-based TENG. (A–D) Reproduced with permission from ref (199). Copyright 2020 Wiley-VCH. (E) Schematic illustrations of PCL/GO-based B-TENGs. (F) SEM image of a PCL fibrous membrane with 4% GO. (G) Improved output current of PCL/GO−based TENGs owing to the presence of GO NPs. (E–G) Reproduced with permission from ref (145). Copyright 2019 Elsevier. (H) Schematic illustration of a cellulose acetate/nano-Al₂O₃ (CA/Al₂O₃) nanocomposite-based TENG. (I) Short-circuit charge transfer and output voltage of a CA/Al₂O₃ nanocomposite-based TENG at different Al₂O₃ nanofiller contents. (H and I) Reproduced with permission from ref (374). Copyright 2020 American Chemical Society. (J) Schematic illustration of a TENG based on a cellulose filter paper (CFP)-Ti_0.8O₂ NSs composite. (K) Increase in the output current density of a CFP composite-based TENG through the addition of Ti_0.8O₂ NSs. (J and K) Reproduced with permission from ref (200). Copyright 2020 Wiley-VCH.

Next, graphene oxide (GO) was exploited as a filler to create a high-performing composite (Figure 18E).¹⁴⁵ GO nanosheets (NSs) play a critical role in improving the permittivity of GO–polymer matrix composites.^242,375 Additionally, owing to its abundance of oxygen functional groups, GO possesses excellent charge-trapping capabilities. Thus, GO NSs were embedded in a PCL polymer matrix. GO NSs were dispersed in PCL solutions at various concentrations for further optimization of the synthesis. The GO-dispersed PCL solution was electrospun to obtain nanofibrous structures with a high surface area and flexibility. A fully bioresorbable TENG with a book-shaped structure was fabricated by integrating the electrospun PCL/GO membrane with a Au electrode layer (2 × 4 cm²) for electrical characterization (Figure 18E). The SEM image shown in Figure 18F and microscopic analysis confirmed that the electrospun layer with 4 wt % GO had a reduced fiber diameter (365 ± 9 nm) and a broader size distribution. However, excess doping with 8 wt % GO resulted in an enlarged fiber diameter (626 ± 12 nm) owing to the high viscosity of the PCL/GO precursor solution. As GO NSs content increased, output voltage and current increased and reached maximum values of 22.73 V and 165 nA at 4 wt %, respectively. However, further increment in GO NSs led to a decrease in output owing to aggregation of the GO NSs (1.51 V and 8 nA at 8 wt % GO NSs).

Next, Bai et al. designed a porous nanocomposite fabric (PNF) with a strong charge accumulation capacity that comprised Al₂O₃ nanofillers in cellulose acetate networks (Figure 18H).³⁷⁴ To prepare PNF, Al₂O₃ nanofillers were mixed with cellulose acetate (CA) precursors in acetone/water (85:15, v/v) to produce a CA/Al₂O₃ dispersion, followed by dry-casting. The addition of Al₂O₃ to cellulose acetate networks resulted in the formation of a porous structure in the dry-cast membrane, which endows the material with a large contact area and high mechanical flexibility and durability. Conductive fabric layers and low-temperature vulcanized (LTV) silicone were used to fabricate partially bioresorbable CA/Al₂O₃-based TENGs. The TENG with 10 wt % nanofiller reached the highest V_OC and short-circuit charge transfer (σ_SC) during repeated cycles of contact and separation with the LTV counter-triboelectric layer, generating values of 400 V and 55 nC, respectively (Figure 18I). The Al₂O₃ concentrations above 10 wt % significantly decreased the output due to saturation of the porous structure by nanofillers and mechanical instability, which possibly interrupted charge transfer between the CA/Al₂O₃ composite and LTV layers.

Finally, Ti_0.8O₂ NSs were also studied to enhance the triboelectric charge-trapping effect of cellulose filter paper (CFP) (Figure 18J).²⁰⁰ Ti_0.8O₂ NS-coated CFP membranes were fabricated by dip-coating the CFP films in colloidal Ti_0.8O₂ one, three, and five times. This nanocomposite membrane was encapsulated using PDMS elastomers to prevent the exposure of the Ti_0.8O₂ NSs on the surface. The dielectric constant exhibited a positive correlation with the number of layers of embedded Ti_0.8O₂ NSs, with a notable increase for the constructs with three and five layers. Specifically, at a frequency of 1 × 10⁵ Hz, the dielectric constants of Ti_0.8O₂ NSs/CFP/PDMS composites featuring one, three, and five layers of Ti_0.8O₂ NSs were approximately 2.2×, 4.5×, and 5.2× higher, respectively, than that of the film without any Ti_0.8O₂ NSs. Integrating dielectric Ti_0.8O₂ NSs into the CFP notably enhances charge generation, and the inclusion of Ag NPs establishes an electrically conductive pathway for efficient charge transportation, thus leading to improvement of the output performance. The triboelectric outputs were measured during periodic contact and separation between the composite and an Al electrode layer (Figure 18K). The composite obtained from performing the dip-coating process three times (3-Ti_0.8O₂ NSs/CFP) generated a V_OC of 20 V and a maximum positive output current density reaching 0.5 μA·cm^–2, which were ∼30× and ∼17× higher, respectively, than those of pristine CFP. Meanwhile, composites with higher number of Ti_0.8O₂ NSs layers exhibited slightly lower triboelectric outputs due to the aggregation of NSs that dissipated triboelectric charges on the surface.

5.3. Ion Doping

Ion doping involves the introduction of dopant ions into polymeric triboelectric materials to modify their electrical properties and enhance charge transfer.³⁷⁶ This technique has been applied to various triboelectric materials, increasing the output power of TENGs. However, careful optimization is required to achieve the desired properties, as the choice of dopant ions and doping method can significantly affect TENG performance. Ryu et al. first used a solid polymer electrolyte (SPE) to obtain partially bioresorbable TENGs with large output power.²²⁵ They investigated the effects of adding CaCl₂, HCl, and H₃PO₄ on the triboelectric properties of PVA, which is one of the most positive and bioresorbable polymers (Figure 19A). KPFM measurements were used to analyze the changes in the surface potential of PVA due to ion doping at different concentrations (Figure 19B). The concentration of CaCl₂ in the CaCl₂–PVA-based SPE varied from 0.25 to 0.75 M, the concentration of H₃PO₄ in H₃PO₄–PVA-based SPE varied from 0.5 to 1.5 M, and the concentration of HCl in HCl-PVA-based SPE was 1 M. The cation/anion ratios for HCl, CaCl₂, and H₃PO₄ were 1:1, 1:2, and 1:0.99, respectively. The addition of H₃PO₄ reduced the surface potential of PVA from +247 to −285 mV, whereas CaCl₂ increased the surface potential from +247 to +705 mV. Furthermore, the HCl–PVA-based SPE, with the symmetric pairing of cations and anions, had a near identical surface potential compared to that of pristine PVA. Figure 19C shows the additional electron unoccupied states in the SPE and the electron transfer between the SPE and its base polymer during contact electrification. The additional energy bands originating from anions or cations in PVA can affect ionic doping. The CaCl₂ electrolyte, which has more anions, creates high-energy electron-charged states in the PVA.^377−379 The additional electron-charged states in 0.75 M CaCl₂–PVA-based SPE enhanced their charge transfer in the contact electrification processes. To compare the output performance, 4 × 4 cm² TENGs were fabricated using PTFE and nylon/Al/SPE as the triboelectric layers. The output performance of Al-PTFE TENG was twice that of the nylon-PTFE TENG owing to the presence of free electrons in metals. Remarkably, the CaCl₂–PVA 0.75 M SPE-PTFE had superior performance compared to that of Al-PTFE TENG, with peak outputs of 211 V and 27 μA, respectively.

Figure 19.

Ion-doped polymers for large power output. (A) Schematic illustration of the triboelectric charge transfer between bare PVA and PVA-based solid polymer electrolytes (SPEs). (B) Surface potentials of ion-doped PVA-based SPEs ddepending on salts and concentration, measured by KPFM. (C) Energy band diagrams presenting electron transfer between bare PVA and PVA:CaCl₂ SPE during a triboelectric event. (A–C) Reproduced with permission from ref (225). Copyright 2017 Wiley-VCH. (D) Output voltages and (E) charge densities of PVA−MClx SPEs-based TENGs. (F) KPFM measurement for PVA−LiCl SPEs to identify their CPDs at different LiCl concentrations. (F) KPFM images of PVA-LiCl SPEs with different LiCl concentrations. (D–F) Reproduced with permission from ref (226). Copyright 2019 Elsevier. (G) Photographs of microstructured ion-doped starch films. (H) Chemical network of a starch:CaCl₂ composite polymer. (I) Output current densities of starch:CaCl₂-based TENGs at different concentrations of CaCl₂. (G–I) Reproduced with permission from ref (380). Copyright 2019 Elsevier.

Ryu et al. investigated the effects of CaCl₂, a common inorganic salt dopant, on the triboelectric properties of polymeric triboelectric layers.²²⁵ However, the effects of doping with other inorganic salts on the triboelectric properties of polymeric triboelectric layers, particularly those with symmetric ion pairing, had not been studied. Shi et al. examined the improved triboelectric properties of PVA-MCl_X solid polymer electrolytes (SPEs) (MCl_X = LiCl, ZnCl₂, CaCl₂, FeCl₃, and AlCl₃) to identify the best dopant for high-performance SPE-based TENGs.²²⁶

The output voltages and short-circuit current densities of partially bioresorbable PTFE/PVA-MCl_X TENGs were determined and compared (Figure 19D). The PTFE/PVA-FeCl₃ and PTFE/PVA-ZnCl₂ TENGs had increased output voltages of 803 and 982 V and short-circuit current densities of 90.5 and 114.5 mA·m^–2, respectively. The PTFE/PVA-AlCl₃ and PTFE/PVA-CaCl₂ TENGs also had increased output voltages (1075 and 1177 V, respectively) and short-circuit current densities (167.8 and 203.6 mA·m^–2, respectively). The PTFE/PVA-LiCl TENG had an even more noticeable increase in the triboelectric output, with a more than twofold increase in V_oc (1289 V) and J_sc (210.5 mA·m^–2) compared to the PTFE/PVA TENG. The corresponding charge density of these PTFE/PVA-MCl_X TENGs increased from 191 μC·m^–2 for PVA-FeCl₃ to 197 μC·m^–2 for PVA-LiCl, indicating that the cation used can partially influence the triboelectric properties of the PVA film (Figure 19E). Consequently, LiCl was chosen as the optimized electrolyte for increasing the output power. Figure 19F shows KPFM images of PVA-LiCl SPEs with different LiCl concentrations. The PVA-LiCl SPEs films had a smooth surface, indicating that surface morphology does not significantly influence the contact potential difference (CPD). No obvious change in surface morphology was observed when comparing PVA-LiCl SPEs with the pristine PVA film. However, the CPD of the PVA-LiCl SPEs increased from 273 to 327 mV with an increase in LiCl concentration from 0.25 to 1.0 M. A further increase in LiCl concentration from 2.0 to 3.0 M led to a decrease in the CPD from 302 to 239 mV because the more concentrated LiCl adsorbed more moisture from the environment, shortening the contact potential difference between the probe and the SPE.

An additional study was performed to develop low-cost, environmentally friendly TENGs based on starch SPE films (Figure 19G).³⁸⁰ The performance of these films was enhanced by the addition of salt, and a clean room-free fabrication process was employed for their potential use in biomedical devices.³⁸¹ Starch, a natural polymer that is abundant and easy to isolate and process into a film form, was chosen for this study.³⁸² Starch films have a high content of amorphous regions and a large number of hydroxyl groups, creating a suitable matrix for dissolving cations and ions. The facile processability of starch films enabled the incorporation of salt through stirring. The starch used in this study was obtained from Andean white potatoes by solubilization. It was redissolved at 80 °C and cast over sandpaper surfaces, resulting in uniformly cured starch films. Figure 19G shows the casting process. The electrostatic interactions between – OH and Ca²⁺ and those between – H and Cl^– are shown in Figure 19H. Polytetrafluoroethylene (PTFE) and Al, materials on opposite sides of the triboelectric series, were selected to achieve a high output performance. The highest outputs of the partially bioresorbable TENGs were obtained for starch SPE films with 0.5 wt % CaCl₂. Surprisingly, the electrical output at this salt concentration was 3× that of the pristine starch film. Short-circuit current measurements revealed similar results, with a threefold increase in current output for 0.5 wt % CaCl₂ (Figure 19I).

5.4. Surface Nanostructure

Triboelectric charging is a phenomenon that occurs at surfaces in which the separation of two materials after contact results in the transfer of electric charges between them. The quantity of triboelectric charges generated depends on several factors, including the materials utilized, degree of contact, separation speed, and relative humidity of the environment.^383,384 Among them, the surface area one of the most considerable factors that influence the triboelectric charging process. A larger surface area increases the triboelectric charge generated because the increased number of contact points between the materials improves electron transfer. There are various approaches to enhance the output power of TENGs, including the development of multistacked TENGs with large contact areas via integrated layers.^139,384,385 However, this strategy may not be suitable for powering IMDs in terms of device size, as the required higher number of components can increase the volume and, consequently, the associated costs. Although rubbing together two materials with larger dimensions can generate more electrons, from a microscopic view, the effective area where triboelectric charging occurs is considerably smaller than the macroscopic area of the triboelectric layer because of the rough surface morphology.^386,387 This is because microspikes on the surface contact each other, and the physical mismatch of surface profiles between the contacting materials limits the effective charge transfer.

Several studies have been conducted to improve the surface charge density and output performance of TENGs by modifying their surface morphology and forming nanostructures based on microfabrication or etching techniques. The structure of an arch-shaped TENG using electrospun silk fibroin (SF) as the triboelectric layer was presented, as shown in Figure 20A.²⁰⁹ The regenerated SF film has a form of nanofiber-networked structure facilitating triboelectric power generation owing to its ultrahigh surface-to-volume ratio and significantly rougher surface in comparison to bare films (Figure 20B). Electrospun SF nanofibers with 100–200 nm diameters entangled with each other, whereas the cast silk film had a smooth surface with minor unevenness. PI substrates and aluminum sheets were employed to fabricate the partially bioresorbable TENGs, and the V_OC values of the two B-TENGs were compared under basic conditions (Figure 20C). As a result, the peak voltage of the electrospun silk-based TENG was approximately 1.5× higher than that of the cast silk-based TENG. A maximum instantaneous power output of approximately 4.3 mW·m^–2 was obtained at a resistance of 5 MΩ.

Figure 20.

Bioresorbable polymers with nanostructured surfaces for large power output B-TENGs. (A) Schematic illustration of an arch-shaped silk B-TENG with an electrospun silk fibroin (SF) membrane. (B) FE-SEM image of an electrospun silk membrane. (C) Peak voltages of partially bioresorbable TENGs based on electrospun silk and cast silk membranes. (A–C) Reproduced with permission from ref (209). Copyright 2016 Wiley-VCH. (D) Top-view SEM images of a rough gelatin membrane cast on sandpaper and an electrospun PLA membrane. (E) Schematics of B-TENGs (red, smooth/rough gelatin film; blue, smooth/electrospun PLA membrane) and (F) their short-circuit output current density resulting from contact and separation between different membrane pairs. (D–F) Reproduced with permission from ref (51). Copyright 2018 Elsevier. (G) Photograph, (H) SEM image, and (I) AFM topology of an ICP plasma-etched PLA/PLGA film. (J) Plot of the 100% increase in the output current of a B-TENG based on a PLA/PLGA film upon plasma etching-induced nanostructuring. (G–J) Reproduced with permission from ref (388). Copyright 2020 Wiley-VCH.

Pan et al. developed a fully biodegradable high-performance B-TENG based on nanostructured gelatin film and electrospun PLA nanofiber membranes.⁵¹ These triboelectric materials were prepared using molding and electrospinning. The gelatin solutions were spin-coated onto sandpapers of varying grades, resulting in gelatin films with different roughnesses. These films were then glued to Mg foils using a soluble adhesive tape and peeled off from the sandpaper with the rough surface facing outward. To prepare the electrospun PLA nanomembranes, a PLA solution was electrospun onto a Mg foil placed on a polyethylene terephthalate (PET) support layer. Figure 20D shows SEM images of the nanostructured gelatin film and electrospun PLA nanomembrane. The gelatin film had intended rough surface structures with microhole sizes ranging from 15 to 20 μm, as determined by the grade of sandpaper used. The PLA nanofiber diameter ranged from 900 to 1800 nm, and the nanofiber thickness was optimized at approximately 60 μm.

To investigate the influence of nanostructuring on output performance, the output voltage and current of TENGs using bare and nanostructured materials were compared under regular pushing conditions (Figure 20E and F). The V_OC for the bare smooth gelatin–smooth PLA pair was about 16 V, increasing to approximately 33 V when smooth gelatin was replaced with the rough gelatin film. Further, with a pair of an electrospun PLA film and smooth gelatin, the V_OC increased to approximately 175 V. In addition, the V_OC of the TENG reached approximately 500 V with a rough gelatin film–electrospun PLA membrane pair, which is more than 30× higher than that of the bare gelatin–PLA pair, highlighting the importance of the nanofiber structure. The corresponding short-circuit current densities were approximately 0.3, 0.6, 4.2, and 10 mA·m^–2, respectively, which also demonstrated tremendous increases.

Chen et al. investigated nanostructuring PLA/PLGA polymer blend through inductively coupled plasma (ICP) treatment to enhance the output power (Figure 20G–J).³⁸⁸ The PLA/PLGA blend exhibited a relatively slow degradation, with PLA serving as a polyester matrix of long lifespan and PLGA acting as a degradation promoter. Fully seawater-degradable TENGs were fabricated using seven different SDP films (PLGA, chitin, PLA/PLGA, PLA, copy paper (CP), rice paper (RP), and silk) and copper electrodes. Then, the output voltage and current of the TENGs were measured during regular contact and separation events. PLGA had a maximum V_OC of 23 V, which was higher than those of PLA (16 V) and the PLA/PLGA blend (21 V). The maximum V_OC values of chitin, copy paper (CP), rice paper (RP), and silk were 23, 15, 13, and 1.5 V, respectively. As depicted in Figure 20H and I, the application of ICP etching led to the formation of aligned nanowires (NWs), thus considerably augmenting the surface roughness of the PLA/PLGA film. The AFM analysis revealed dense needle-like structures with a depth of approximately 145 nm. Owing to this enhanced surface roughness, the V_OC, I_SC, and Q_SC of the TENG based on the ICP-etched PLA/PLGA film showed a significant increase of 100% compared to the pristine, smooth PLA/PLGA film (Figure 20I).

5.5. Surface Functionalization

The triboelectric effect is widely recognized as a surface-oriented phenomenon in which the amount of charge transferred is proportional to the surface potential difference based on the electron transfer model.^156,157 Accordingly, the surface dipoles and electronic structures are believed to significantly contribute to the triboelectric charge transfer. Surface conditions, particularly the electron affinities of surface functional groups, often determine the direction and quantity of charge transfer.^389,390 Electron-donating groups, such as NH₂, and electron-accepting groups, such as F, are representative examples of functional groups that can impact the triboelectric charge generation process.^391,392 The modification and functionalization of these surfaces to optimize and improve the performance of triboelectric layers has been reported. Experimental techniques, such as plasma treatment and self-assembled monolayer (SAM) formation, have been employed to introduce desired functional groups or tailor the surface properties of conventional TENGs to maximize their output power.^194,393

Recent studies have shown that surface functionalization can enhance the output power of B-TENGs. Sangkhun et al. developed a simple method to create efficient natural textile-based B-TENGs.³⁶¹ Plain natural textiles, such as cotton and silk, were dip-coated with cyanoalkyl (CN) silane and fluoroalkyl (F) silane, respectively, to convert their surface energy into positive and negative triboelectricity (F-cotton, CN-cotton, F-silk, and CN-silk) (Figure 21A). To confirm the successful grafting of the silanes onto the textiles, their chemical states were determined by X-ray photoelectron spectroscopy (XPS) (Figure 21B). The F-cotton sample exhibited an F 1s peak at 687.7 eV and an O 1s peak at 531.3 eV. The C 1s peak was split into two peaks (286 and 288 eV) owing to the interaction between fluorine and carbon atoms, confirming the successful grafting of fluoroalkyl groups onto cotton. Similarly, the CN-cotton sample exhibited an O 1s peak at 532 eV, a N 1s peak at 400 eV, a C 1s peak at 285 eV, and a Si 2p peak at 103 eV, confirming the successful grafting of cyanoalkyl groups onto cotton. The F-silk sample exhibited a F 1s peak at 687.8 eV and an O 1s peak at 532.8 eV, while the CN-silk sample showed an O 1s peak at 532.5 eV, a N 1s peak at 400.5 eV, a C 1s peak at 285.5 eV, and a Si 2p at 103 eV, confirming the grafting of silane onto silk.

Figure 21.

Surface functionalization for large power output bioresorbable tribo-materials. (A) Schematic illustration of the surface functionalization process to prepare fluoroalkylated siloxane-grafted fabric (F-fabric) and cyanoalkylated siloxane-grafted fabric (CN-fabric). (B) XPS spectra of F-cotton. (C) Output currents of natural textile TENGs (N-TENGs) based on surface-functionalized cotton and silk fabrics. (A–C) Reproduced with permission from ref (361). Copyright 2021 Royal Society of Chemistry. (D) Schematic illustrations of the grafting process of fluorinated group to functionalize fish gelatin (FG) films. (E) Structure of a fully sustainable fish gelatin (FSFG)-TENG. (F) XPS spectra and (G) contact angles of fluorinated FG (F-FG), dopamine-doped FG (D-FG), and FG. (H) V_OC and (I) I_SC of the TENGs based on the F-FG film paired with cotton, Al, cellulose, Cu, and D-FG. (D–I) Reproduced with permission from ref (394). Copyright 2021 Elsevier.

A Cu fabric electrode was assembled by stacking cyanoalkylated and fluoroalkylated siloxane-grafted fabrics and used to fabricate fully biodegradable natural textile-based TENGs (N-TENGs). The average output voltages and output currents of the materials were measured and ranked as follows: F-cot/CN-silk > F-silk/CN-cot > F-silk/CN-silk > F-cot/CN-cot (Figure 21C). The F-cot/CN-silk TENG exhibited average surface charges of 8.19 (0.141 μC·cm^–2) and 4.53 μC (0.078 μC·cm^–2), respectively, at the pressing and releasing states of the three-movement cycles. The F-cot/CN-silk TENG also exhibited the highest electrical output voltage (216.8 V) and output current (50.3 μA, 0.87 μA·cm^–2), which were significantly higher than those of the cot-silk N-TENG (39.2 V and 0.2 μA·cm^–2, respectively). This enhanced electrical performance can be attributed to the surface morphology, hydrophobicity, and chemical composition of the material. Double-stacked TENGs of F-cotton and CN-silk were fabricated to further increase the output power, resulting in a maximum voltage of 219.30 V and a maximum current of 84.87 μA (1.46 μA·cm^–2) under a cycle of mechanical tapping. These findings highlight the potential of surface functionalization in B-TENG performance, improving their scope of practical applications. A single N-TENG generated a maximum output voltage and current of 216.8 V and 50.3 μA (0.87 μA·cm^–2), respectively, without nanopatterning. A double-stacked N-TENG exhibited a higher output current of 84.8 μA (1.46 μA·cm^–2), with a maximum power output of 0.345 mW·cm^–2 at an external resistance of 0.42 MΩ.

Sun et al. developed a flexible, transparent, fully bioresorbable, and fully sustainable fish-gelatin-based TENG (FSFG-TENG) and largely improved the output through surface functionalization of fish gelatin (FG).³⁹⁴ They utilized fish-scale kitchen waste to prepare FG films that could be used as friction layers. Two FG triboelectric layers were modified with dopamine and fluorinated silane to serve as positive and negative layers, respectively, of a triboelectric pair. The two modified FG films were designated as D-FG and F-FG, respectively. An FG solution was drop-cast and dried to form FG films, which were then hydroxylated through oxygen plasma treatment for fluorinated oxetane telomer sulfonate (FOTS) functionalization (F-FG) (Figure 21D). The D-FG film was prepared by mixing a FG solution with dopamine hydrochloride and drop-casting. The copper foil was used as the back electrode for both triboelectric layers, and the B-TENG was fabricated by assembling the layers in an arc shape (Figure 21E). XPS confirmed the chemical surface modification through the appearance of an F 1s peak in the F-FG film (Figure 21F). The hydrophobicity of the F-FG film also improved, as demonstrated by the increase in the contact angle from 103° to 123° (Figure 21G). The tailored surface had significantly higher output power (Figure 21H and I). The FG/FG-paired TENG showed a V_OC and an I_SC of 10 V and 0.2 μA, respectively. However, replacing one of the FG friction layers with the D-FG or F-FG film led to an immediate decrease in the V_OC of the TENGs to ∼100 or ∼200 V, respectively. The TENG with the D-FG/F-FG triboelectric pair exhibited a high V_OC (∼310 V) and I_SC (2.5 μA). This increase in output performance can be attributed to the larger difference in electronegativity between the F-FG and D-FG films.

The fluorination of a surface typically increases the hydrophobicity of the material, requiring only a small amount of treating material.³⁹⁵ Considering many bioresorbable polymers are tribo-positive, amine or other functional groups can be utilized to tailor the surface chemical structure and enhance the electron-donating property.³⁹² Although surface functionalization is an effective tool to improve the output power of bioresorbable materials, the potential impact of altering surface characteristics on the surface’s interaction with the surrounding physiological environment must be considered. Currently, no studies have validated and examined the biocompatibility of functionalized bioresorbable materials through either cytotoxicity tests, genotoxicity tests, or in vivo inflammation. Therefore, a comprehensive investigation of the biosafety of surface-tailored bioresorbable materials is required. Furthermore, considering that previous studies have improved triboelectric charge and output power through charge injection or plasma treatment, the same methods can be explored to increase the output power of B-TENGs. To the best of our knowledge, there are no studies on this topic, and further research is necessary.

6. Biomedical Applications of B-TENGs

B-TENGs provide an exceptional clinical experience by providing a power source and by eliminating the need for subsequent device removal, thus driving IMDs and reducing the physical and psychological risks of the procedure.^58,124 There have been trials to power batteries with B-TENGs, however, the use of B-TENGs as energy sources for IMDs is still limited. As discussed in section 5, bioresorbable materials have been developed with improved triboelectric properties; however, their energy generation capabilities have not yet reached the energy consumption of conventional IMDs in most cases.²⁴⁵ The nontransient nature of conventional IMDs also limits their practical use; these nontransient IMDs still need to be removed at the end of treatment. Bioresorbable circuit components could replace nontransient IMD components but require further research on material and device fabrication. In contrast, the self-powered functionality of B-TENGs offers distinct advantages for clinical applications. Their simple and thin structures minimize physical stress on the surrounding tissues and reduce negative health effects such as inflammation.^46,164 Furthermore, their uncomplicated material configuration results in a lower cost and simplified lifetime control.⁵⁸ B-TENGs also facilitate the collection of high-quality physiological signals and the application of electrical impulses to targeted affected areas with high selectivity because miniaturized B-TENGs can be placed closer to the target location. Consequently, B-TENGs have gained attention as self-powered implants that require less energy to function, including physiological sensors, electroceutical devices, and sterilizers (Figure 22).^30,117,123 In this section, we provide an overview of B-TENGs specifically designed for each of these applications.

Figure 22.

Overview of self-powered bioresorbable IMDs based on B-TENGs.

6.1. Self-Powered Physiological Sensors

Diagnosis is an essential part of the healthcare industry, examining physical symptoms and determining the underlying conditions. Bioresorbable electronics can adhere to organs or tissues to monitor ambient signals. For example, a bioresorbable arterial-pulse sensor has been used to measure arterial blood flow.³³ This sensor consists of an inductor coil and a flexible capacitive pressure sensor. The inductor coil transmits data wirelessly using a RF coupling method. The sensing layer of the pressure sensor wrapped around the artery is composed of a pyramid-structured PGS that exhibits mechanical softness, with a Young’s modulus of 0.12 MPa. This device detects a capacitive change in response to the vessel diameter, which leads to a shift in the resonant frequency of the inductor–capacitor–resistor circuit. The monitoring performance of this sensor was evaluated both in vitro using a customized artery model and in vivo demonstrations, but its practical use is limited by the short transmission distance (<10 mm). Additionally, its performance can be hindered by an unfavorable alignment between transmitting and receiving inductors.

A bioresorbable force sensor was also reported using a piezoelectric mechanism, which converts mechanical deformation into electrical energy.³⁹⁶ The poly-l-lactide polymer acted as a piezoelectric material, forming a sandwich structure with the Mo electrodes. This device identified mechanical pressures in a 0–18 kPa range. The piezoelectric sensor was implanted in the abdominal space of a mouse to monitor its periodic diaphragmatic motion. Compared with a commercial piezoelectric force sensor, the bioresorbable sensor exhibited high sensitivity to mechanical stimuli. Piezoelectric sensors require high piezoelectric responses to mechanical stimuli to perceive ambient physiological events. However, bioresorbable polymers typically have inferior piezoelectric properties, hindering the implementation of bioresorbable piezoelectric sensors.

6.1.1. Cardiovascular Postoperative Care

Soft bioresorbable sensors became prevalent due to their favorable conformability with soft biological tissues and organs.^15,397 The successful implementation of these sensors also requires suitable mechanical properties, flexibility, and biocompatibility.^8,396,398 Various bioresorbable polymers, metals, and oxides with electrical characteristics for bioelectronics have been reported.^{63,218,399−401} The critical parameters of bioresorbable implantable devices have progressed in parallel with advancements in materials science, electronics, and manufacturing.^402−404 However, their implementation for postoperative care applications remains challenging, primarily because of the insufficient device lifetime of most implanted biodegradable sensors (usually 4–10 d).⁴⁰⁵ The most common biodegradable encapsulation materials (e.g., PVA, silk fibroin, and PLGA) have relatively fast transient rates, emphasizing the need for novel bioresorbable materials with long service lives.^124,406 The use of self-powered sensors with triboelectric and piezoelectric mechanisms for wearable and implantable applications has recently gained attention because of the diverse material selection, simple fabrication process, and superior electrical output performance.^407,408 The triboelectric effect of bioresorbable materials can reach maximum output voltages of several volts to hundreds of volts under external mechanical forces.^122,209 Thus, B-TENGs are promising bioresorbable sensors for in vivo disease diagnosis, requiring significant improvements in their sensitivity, linearity, and stability. Furthermore, the triboelectric effect can circumvent the limitations of pressure-sensitive devices.^{146,245,256,257}

The concept of an implantable, fully bioresorbable triboelectric sensor (BTS) has recently been reported (Figure 23A).³⁰ The proposed sensor proved effective on canine subjects, effectively detecting occlusions in the vascular system. The triboelectric layer of the device consisted of a single PLA–4% chitosan (PLA/C) membrane with a nanostructured surface. A Mg layer was deposited on the back, serving as an electrode. The front of the membrane was also coated with Mg as both an electrode and a triboelectric layer. The entire sensor was packaged in PLA/C, using the bioresorbable elastomer POC as the adhesive layer. The two triboelectric layers contact upon the application of an external force, resulting in the transfer of electrons from the Mg surface to the PLA/C surface owing to contact electrification. Variations in the external forces change the distance between the triboelectric layers. Thus, the sensor produces an electrical signal proportional to the mechanical motion.⁴⁰⁹ The nanostructured surface of the device improved the contact area during friction, thereby yielding a more significant electrical signal and improving the sensitivity of the sensor. The performance of the BTS as a mechanical sensor was evaluated by attaching it to a human throat and monitoring vocal cord vibrations. Voltage signals of 4–9 mV were obtained when a volunteer pronounced the letters “B”, “T”, and “S.” The results from two trials with different accents were consistent, demonstrating the reliability of the device. A Fourier transform analysis of the signal was conducted to clarify the distinctions between the pronunciations of other letters. In addition to its mechanical sensor performance, the BTS detected 1 kHz vibration signals from a small loudspeaker and converted them to electrical signals in real time, with a response time of 0.5 ms. The stability of BTS in nonliquid environments was also measured. Its maximum output reached 4.2 V when driven by a linear motor (20 N), remaining stable under a 1–5 Hz mechanical force. The BTS remained stable for over 450 000 operating cycles of mechanical stimuli. A linear motor was employed as a source of mechanical stimuli (20 N, 1.35 Hz) to evaluate the long-term stability of the BTS. Its output voltage was measured at 18.5 h intervals, corresponding to approximately 9 × 10⁴ cycles. The output performance of the BTS was highly significant, with a sensitivity of 11 mV·mmHg^–1 and a linearity close to R² = 99.3%, which was calibrated using a pressure testing system. The observed pressure range (0–170 mmHg) covers most pressures encountered in vivo.

Figure 23.

Physiological sensing using B-TENGs. (A) Expanded structure of an implantable bioresorbable triboelectric sensor (BTS) for cardiovascular postoperative care. (B) Output variation of a BTS implanted in a small animal according to the respiratory event identification. (C) Identification of vascular occlusion events following BTS implantation in a large animal. (A–C) Reproduced with permission from ref (30). Copyright 2021 Wiley-VCH. (D) Working mechanism of TENG-integrated vascular grafts (VG-TENG). (E) Experimental setup images and (F) electrical properties of VG-TENGs under different blood flow conditions. (D–F) Reproduced with permission from ref (127). Copyright 2023 Elsevier. (G) Structure and material design of transient TENG (T²ENGs). (H) Photographs of an implanted T²ENG under the subdermal dorsal region in an in vivo experiment. (I) Electrical output of T²ENG to demonstrate the epilepsy monitoring function. (G–I) Reproduced with permission from ref (128). Copyright 2018 Wiley-VCH.

The BTS was implanted under the skin of the abdominal cavity of a mouse model to identify abnormal respiratory events (Figure 23B). Respiratory events in the mouse drove the contact/separation motions of the BTS. Inhalation generated a voltage output, while exhalation resulted in a decreased voltage output. Therefore, using this respiratory monitoring system, the BTS can easily detect abnormal respiratory events (e.g., dyspnea). Dyspnea causes chest wall expansion and lung inflation, increasing the compression force applied to the BTS and inducing a higher voltage output. For comparison, respiratory events were measured in vitro using commercial mechanical sensors.

In contrast to BTS, the commercial sensors exhibited reduced voltage output at higher compression forces. Next, the ability of the BTS to identify abnormal cardiovascular events (e.g., arrhythmia) was investigated (Figure 23C). The BTS was attached to the vascular wall to measure ambulatory blood pressure signals. A commercial mechanical sensor was used in vitro as a reference. The contact/separation motions of the BTS along the diastole and systole phases were examined. The BTS was compressed in the systole phase, generating a voltage output during its release in the diastole phase. Thus, BTS was demonstrated to suitably monitor abnormal vascular occlusion events. An implantable balloon was used as a vascular occlusion model. The inflated balloon obstructed the blood vessel, resulting in a decrease in blood pressure.

Consequently, the voltage output generated from the BTS was reduced during this period. Furthermore, the voltage output was reduced when the balloon was deflated. Thus, integrating BTS with B-TENGs is a promising arrhythmia monitoring system.

6.1.2. Sensing Hemodynamics of Vascular Grafts

Vascular grafts, common therapeutic solutions to vascular diseases, can be fabricated using diverse methods (e.g., electrospinning and 3D printing).^410,411 However, they may fail to identify unexpected disorders, leading to serious health problems.⁴¹² Thus, using an implantable sensor system is vital to characterize the hemodynamics of vascular grafts in advance.⁴¹³ Although vascular grafts are generally observed using magnetic resonance angiography and computed tomography, these methods are not convenient because of their lack of readiness and real-time communication.⁴¹⁴ Wang et al. investigated a partially bioresorbable TENG-integrated vascular graft (VG-TENG) that wirelessly detected hemodynamic conditions.¹²⁷ PHB was adopted as a tribo-positive material because of its excellent biocompatibility and biodegradability.⁵⁸ An electrospinning process was used to form a micro/nanoscale fiber-structured PHB membrane with improved triboelectric properties. PTFE was used as a contact material because of its strong tribo-negative properties.²⁶² Expanded PTFE (ePTFE), fabricated via the extension of PTFE, is a widely used commercial vascular graft with considerable flexibility and elasticity. A double-electrode -mode TENG was fabricated using electrospun PHB and ePTFE and electrically characterized (Figure 23D). The TENG showed superior triboelectric output performance, generating a voltage output of 440 V, a current density of 19.5 mA·m^–2, and a transferred charge density of 44 μC·m^–2 using an applied force of 30 N. The electrical output performance remained stable for 10 000 operating cycles, proving the long-term durability of the triboelectric materials.

A VG-TENG with a conduit structure was next developed to monitor vascular obstructions (Figure 23E). The PHB tube contacted the ePTFE tube initially, inducing triboelectric charges in each tube. It then separated from the ePTFE tube as the native blood vessel became constricted. In this process, the Cu electrode loses free electrons to maintain the charge equilibrium. When the native blood vessel inflates again, the Cu electrode gains free electrons, whereas the PHB tube moves toward the ePTFE tube. The voltage outputs of the VG-TENG were measured to evaluate its performance under in vitro conditions, assuming the presence of an obstruction. As shown in the right side of Figure 23E, the voltage output of the VG-TENG decreased with an obstruction. Figure 23F plots the voltage outputs at different rotation speeds and obstruction conditions. The voltage output was generally reduced when the tube was obstructed. A higher rotation speed increased the pressure, which increased the voltage output. Thus, the bioresorbable VG-TENG is an effective self-powered monitoring system for hemodynamic sensing.

6.1.3. In Vivo Monitoring of Epileptic Seizures

Epilepsy is one of the most common neurological disorders and is characterized by severe unexpected convulsions.⁴¹⁵ There is no cure for this disease, leading to significant sociological problems. Detecting epilepsy symptoms is desirable for reducing its social impact. A transient TENG (T²ENG) for self-powered epilepsy monitoring was investigated (Figure 23G).¹²⁸ T²ENG could be being fully biodegraded by leveraging silk fibroins (SFs) and Mg as triboelectric layers and electrodes. SFs were modified using microgratings, thereby controlling their roughness and surface morphology. Thus, the silk fibroins exhibited high triboelectric performance and could be optically observed via their unique diffractive patterns using noninvasive laser illumination.⁴¹⁶ Taking advantage of the improved triboelectric properties of the silk fibroins, the T²ENG exhibited enhanced electrical output performances with a maximum voltage output of approximately 60 V and a current output of approximately 1.0 μA. The T²ENG was designed to be optically observed during its degradation process in vivo using its diffraction pattern. The device was immersed in deionized water at room temperature to evaluate its transient performance. The Mg electrode layers dissociated from the device after 6 h of immersion.

The device was implanted beneath the dermis of a mouse model to confirm the viability of the T²ENG-based epilepsy monitoring system in vivo (Figure 23H). The condition of the implanted device was assessed after 3 weeks of implantation. Here, the transient performance of the T²ENG was varied by controlling the crystalline region of the silk fibroins. A T²ENG composed of silk fibroins with low crystalline components was absorbed after 3 weeks because of its high transient rate. In contrast, the T²ENG with highly crystalline silk fibroins exhibited a low transient rate, maintaining its structure. When the T²ENG generates irregular electric signals because of an epileptic symptom, its integrated monitoring system transmits this information. In vivo experiments were conducted to evaluate the performance of this self-powered epilepsy monitoring system. The implanted devices did not generate electrical signals while the mouse rested. A penicillin G (PNG) sodium solution was injected to induce an epilepsy symptom, causing the T²ENG to generate notable electric signals. As shown in Figure 23I, when the electric signals exceed the threshold voltage, the real-time monitoring system sends warning messages saying, “Epilepsy warning! Help!”. After the mouse is treated, the T2ENG stops generating electrical signals, and the monitoring system says, “Epilepsy is alleviated.”

6.2. Therapeutic Use of B-TENGs

Electroceutical devices, also known as electronic drugs, are alternatives to pharmacological methods that use electrical signals for treatment and tissue engineering. Electroceutical devices offer a unique clinical experience with targeted and on-demand therapies. For example, cardiac pacemakers implanted in the chest have been developed to control the heart rhythm and treat arrhythmia.^12,28,417 A high-frequency electric field has been reported as an effective alternative to chemotherapy or radiation therapy for nearly incurable brain tumors.^418,419 Bioresorbable materials have been used to construct electroceutical systems. Electroceutical devices from bioresorbable materials are suitable for short-term electrotherapy, as they are absorbed in human tissue after their service life. However, the absence of a practical energy source with transience and biosafety remains a significant challenge. Wireless bioresorbable power transmission systems have been investigated using RF coupling.^26,41 They are equipped with bioresorbable inductors integrated with cuff electrodes to treat peripheral nerve injuries in the spinal cord and muscular tissues. These devices were designed to be absorbed within 40 d, with a time of operation of several days. Despite providing significant inspiration for bioresorbable electroceutical systems, their use is limited because of the short transmission distance (<10 mm).⁵⁸ B-TENGs are promising energy sources for therapeutics. They can generate electrical impulses in deep tissues by harnessing biomechanical motions (e.g., heartbeat and muscle contraction) or through external sources. In this section, we present previous research on applying B-TENGs to therapeutics.

6.2.1. Wound Healing

Electrical stimulation has been shown to enhance the process of wound healing. The internal electric field created by differences in the transepithelial potential plays a crucial part in the repair and renewal of wound surfaces, known as re-epithelialization. Since the first measurement of wound current by Emil Du-Bois Reymond over a century ago,⁴²⁰ successive research has verified that this internal electric field is present in a range of animals, and electrical therapy might effectively achieve synergistic wound healing by reshaping or intervening in the electric field.^421,422 It is vital for epithelialization, as it directs the movement of key electroactive cells, like epithelial cells, during wound repair. TENGs have been highlighted as innovative energy supply systems with low power and flexibility in the form of wearable or implantable medical devices to address the poor portability of conventional bulky electrotherapy equipment and power sources.⁴²³ B-TENGs, being eco-friendly and implantable, are well-suited to wound care and planetary considerations. Their unique properties enable a continuous electrical stimulation source that facilitates faster wound healing without causing any adverse environmental impacts. Additionally, given their implantable nature, they offer a patient-friendly approach to treatment. They could be used in surgical operations to reconstruct or seal significant damage to organs and tissues.

An ultrasound-driven injectable B-TENG (denoted as I-TENG in Figure 24) was developed to treat organ damage and wounds.¹²³ The I-TENG employed PLA as the triboelectric and encapsulation layers and Mg as the counter electrode to secure full biodegradation capability. It had a thin and long shape to be easily injectable (2 cm × 0.2 cm) and generated output power by ultrasound-mediated vibration of PLA and consequent coupling of the triboelectric effect and the electrostatic induction between PLA and Mg (Figure 24A). The voltage output of the I-TENG increased with the increasing incidence energy in PBS (pH = 7.4), reaching a maximum of approximately 632 mV under 1.5 W·cm^–2 of applied ultrasounds. The I-TENG was injected with a needle at 5 mm under the epidermis and, once on-site, generated electricity remotely using ultrasound stimulation (Figure 24A and B). Once the power was turned on, a voltage output signal with the same frequency as the initial ultrasound was produced, with a voltage and current of approximately 458 mV and 2.3 μA at 1 W·cm^–2, respectively.

Figure 24.

Electroceuticals using B-TENGs. (A) Schematic illustration of an I-TENG under the skin. (B) Photographs of I-TENGs placed inside a needle. (C) Scratch assay to demonstrate enhanced migration by equivalent electrical stimulation. (A–C) Reproduced with permission from ref (123). Copyright 2023 Wiley-VCH. (D) Overall procedure for rapid wound closure and hemostasis using BA-TENG. (E) Scratch wound healing experiments using BA-TENG. (D and E) Reproduced with permission from ref (125). Copyright 2023 Wiley-VCH. (F) Schematic illustration of electrical stimulation using BN-TENG and of the progress observed after BN-TENG implantation. (G) Pause time between two beating cycles of the cardiomyocyte cluster, according to the stimulation effect of the BN-TENG. (H) Beating rates of different cardiomyocyte clusters according to the stimulation effect of the BN-TENG. (F–H) Reproduced with permission from ref (208). Copyright 2018 Wiley-VCH. (I) Schematic illustration of electrical stimulation using BD-TENG. (J) Neuron cells oriented by the electric field (the yellow arrow represents the direction of the electric field, scale bar is 50 μm). (K) Neuron cell alignment analysis for different cell angles. (I–K) Reproduced with permission from ref (124). Copyright 2016 American Association for the Advancement of Science under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/).

Cell migration and proliferation are vital to the wound healing process. To evaluate the influence of the electrical impulse generated by the I-TENG on tissue regeneration, a finite element (FEM) simulation and a series of scratch and proliferation assays were conducted (Figure 24C). The FEM simulation results revealed a high-frequency AC electric field of approximately 0.92 V·mm^–1 between two parallel electrodes at 1 W·cm^–2. A cell monolayer was cultured between two electrodes in a cell confocal dish, and a scratching with a 200 μm width was made to conduct the scratch assay. The electrode was connected to a function generator with the same frequency (20 kHz) and amplitude as the I-TENG output signal at 1 W·cm^–2, and the area of the cell-free zone was measured in microscopic images of cell migration using the ImageJ software. After 6 h of stimulation, the area in the experimental group was 19.3%, considerably smaller than that in the control group (29.1%). After 18 h, the linear scratches of the experimental groups became almost invisible, indicating a faster cell migration rate. Proliferation assays were also performed to evaluate the efficacy of I-TENG electrical stimulation in wound healing; the number of cells per unit in the experimental group after 72 h was higher than that after 24 h. This study revealed that electroceuticals using high-frequency I-TENG can be used to treat organ damage and improve wound healing.

Subsequent research examined an ultrasound-driven bioadhesive TENG (BA-TENG) as a portable bioadhesive patch that sealed skin injuries and generated a potent electric field (approximately 0.86 kV·m^–1) to encourage cellular migration and proliferation (Figure 24D).¹²⁵ The fully bioresorbable BA-TENG was engineered with a flexible top layer and a bioadhesive bottom layer, effectively delivering electrical energy and a robust, sutureless sealing on wet tissues. The upper membrane was composed of PCL-based polyurethane (PCL-r-PU) as a triboelectric and encapsulation layer, poly(3,4-ethylene dioxythiophene) polystyrenesulfonate (PEDOT:PSS) as a triboelectric electrode, a poly(acrylic acid) N-hydroxysuccinimide ester (PAA-NHS)–PVA copolymer (PAA-NHS-PVA) as a bioadhesive, and Mo as an electrode.

Rapid and robust adhesion to blood-covered tissues is crucial for maintaining hemostasis and wound sealing in clinical and biomedical settings. The BA-TENG was applied to a 5 mm diameter defect in a wet isolated porcine colon, forming a conformal attachment to the colon tissue within approximately 5 s that was difficult to remove from the wet surface. Upon contact with the wet tissue, the PAV eliminated interfacial water and instantaneously dried the surface through hydration. The NHS ester and carboxylic acid groups of the PAV facilitated the strong and stable adhesion to the tissue, forming covalent cross-linked amide bonds and physically cross-linked hydrogen bonds, respectively. The BA-TENG device exhibited a formidable interfacial toughness of approximately 150 J·m^–2 and a shear strength of approximately 40 kPa. The hemostatic sealing properties of the BA-TENG were further evaluated in live animals using a Sprague–Dawley rat liver incision model. The BA-TENG was applied to a bleeding liver injury, measuring the hemostatic time and blood loss volume. The BA-TENG-treated group demonstrated rapid wound sealing within approximately 5 s, leading to a dramatic reduction in blood loss of approximately 70 mg. In contrast, the untreated injury exhibited a prolonged hemostatic time (>300 s) and substantial blood loss (388 mg).

The triboelectric properties of the PCL-r-PU were investigated using first-principles simulations. The electrostatic potential map of PCL-r-PU revealed that the urethane group is electron-rich, which significantly affected the electron distribution among the molecular chains. Notably, although nitrogen atoms constitute the smallest fraction in PCL-r-PU, they contribute more to forming the highest occupied molecular orbitals than other atoms. The low electron affinity of N and its electron donor properties result in a positively charged PCL-r-PU surface after friction with PEDOT:PSS. Thus, the PCL-r-PU exhibits tribo-positive properties relative to those of PEDOT:PSS. A BA-TENG with dimensions of 1.5 cm × 1.5 cm can generate a stable output performance, with a voltage output of approximately 1.50 V and a current output of approximately 24.20 μA.

Scratch and proliferation assays were conducted to assess the impact of high-frequency electrical impulses with an electric field of 0.86 kV·m^–1 (Figure 24E). The initial relative wound areas in the scratch assay were approximately 31%. The serrated electrode group exhibited the fastest cell migration rate after 16 h due to tip-enhanced effects, resulting in a wound coverage of approximately 12%, smaller than that of the linear electrode group (19%). In contrast, the control group displayed the largest relative wound area (23%). In the proliferation assay, the cell number per unit area in the serrated group reached approximately 35, higher than that of the control group (19), indicating a higher rate of cell proliferation. These in vitro experiments suggest that the 20 kHz electrical impulses generated by the BA-TENG effectively promote tissue healing.

6.2.2. Stimulating Cardiomyocyte Clusters

B-TENGs have also been used to regulate the beating of cardiomyocyte clusters.²⁰⁸ NBPs were electrically characterized to improve the triboelectric performance of the fully bioresorbable TENG. The bioresorbable nature-materials-based TENG (BN-TENG) was integrated with a rectifier and interdigital electrode to apply electrical stimulations (Figure 24F). The interdigital electrode was covered with a 50 μm thick PDMS layer to prevent undesirable electrochemical reactions. After rectification, the optimized BN-TENG generated a voltage output of 18 V, which enabled the formation of a direct current (DC) electric field in the interdigital electrode. FEM simulation using a 300-μm microgap distance confirmed that the DC electric field had an intensity of approximately 8 V·cm^–1. Initially, the primary cardiomyocytes were seeded on the interdigital electrode. Previous studies have indicated that gap junction proteins (connexins) promote the electrical interconnections of cardiomyocytes, which facilitate intercellular communication.^244,424,425 During the incubation period, the cells formed isolated cardiomyocyte clusters through interconnections within the microgap of the electrodes (Figure 24G). After 48 h, the cardiomyocyte clusters (C1, C2, C3, and C4 in Figure 24H) exhibited a slow beating rate with a prolonged pause before BN-TENG-driven electrical stimulation. Using the BN-TENG, electrical stimulation was applied to cardiomyocytes for 30 min at a frequency of 1 Hz. The electrical stimulation accelerated the beating rates of the cardiomyocyte clusters, reducing the pause time from 1.382 to 0.606 s.

Similarly, the contraction time during the beating cycle was reduced from 0.320 to 0.240 s. Figure 24H shows the beating rates of each cardiomyocyte cluster. Notably, the beating rates of C2 and C3 significantly increased by approximately 8.8× after electrical stimulation.

6.2.3. Manipulating Neuron Cell Orientation

Electrical stimulation has also been explored for neuron cell orientation, enabling the clinical use of B-TENGs.¹²⁴ In this study, an interdigital electrode was connected to fully bioresorbable TENGs (denoted as BD-TENGs) to deliver the electrical stimulation (Figure 24I). The Cu-deposited interdigital electrode was encapsulated with a 100 μm thick PDMS layer. This electrode had an interval space of 100 μm, which was designed to produce an electric field of 100 V·mm^–1 at an open-circuit voltage of 1.0 V. Nonetheless, considering the thickness of the PDMS encapsulation layer, FEM simulations confirmed that the actual electric field intensity experienced by the neuron cells was 0.75 V·mm^–1. Primary neuron cells were obtained from the cerebral cortex of an SD neonatal rat via mechanical trituration, trypsinization, filtration, and centrifugation. They were then cultured and seeded on the surface of the PDMS layer at a density of 30 000 cells·cm^–2. Electrical stimulation was performed 30 min·d^–1 with a frequency of 1.0 Hz. After 5 d of stimulation, the cytoskeleton of the neuron cells was observed by laser scanning confocal microscopy (Figure 24J). The electrically stimulated cells were parallel to the direction of the electric field, whereas the unstimulated cells did not have a defined orientation. The alignment of the neuron cells was recorded by measuring the angle between the long axis of the cell and the direction of the electric field. Figure 24K shows the statistical analysis results, in which −cos(2θ) is the index of cell alignment (ICA). An ICA value close to −1 denotes that the cells are nearly parallel to the electric field. Conversely, an ICA value between 0 and +1 indicates that the cells are randomly aligned. Approximately 88% of the neuron cells had ICA values between −1 and 0, which suggests that the electric field effectively aligned the neuron cells. Meanwhile, the ICA values of the unstimulated neuron cells were randomly distributed. Thus, B-TENGs can be promising for facilitating neuroregeneration.

6.3. Rehabilitation Using B-TENGs

In 2019, over 2.4 billion people worldwide had health conditions that would benefit from rehabilitation, representing a 63% increase from 1990.⁴²⁶ The demand for rehabilitation is emphasized by the changing demography and human health.⁴²⁷ For example, improved life expectancy has increased the prevalence of disability and chronic diseases. However, there is a significant gap between the demand for rehabilitation and availability. Although unexpected circumstances, such as conflicts, disasters, and outbreaks, abruptly increase the need for rehabilitation services, more than 50% of the population in low- and middle-income countries is not supported adequately by rehabilitation services.⁴²⁸ Owing to their cost-effectiveness, minimized device volume, and simple structure, B-TENGs have been investigated as attractive options for rehabilitation medicine.^429,430 There are several studies on implanting B-TENGs to rehabilitate musculoskeletal systems (e.g., bone pain, joint, and muscle).

The aging population has also led to an increase in the number of bone fracture cases, particularly those associated with osteoporosis and osteogenesis imperfecta, resulting in a significant social and economic burden.^431−435 Conventional treatments, such as pharmacological methods and stem cell therapy, have significant side effects and lengthy clinical approval processes.^436,437 Thus, alternative approaches are required to address the high prevalence of bone fractures and associated costs.⁴³⁸ Electrical stimulation has been explored for bone tissue regeneration.^439−441 Previous in vitro studies have demonstrated that electrical stimulation from TENGs can trigger osteoblast cell proliferation and differentiation.⁴⁴⁰ However, their nontransient properties restrict their practical use in fracture healing.^442,443

A recent study reported an implantable, fully bioresorbable, ultraflexible bone fracture electrical stimulation device (FED) with bioresorbable characteristics, allowing its absorption in human tissue after its service life (Figure 25A–C). This FED was composed of two elements: a B-TENG producing electrical impulses and a set of interdigitated electrodes spatially forming an electric field (Figure 25A). Both the TENG and interdigitated electrodes were constructed on a PLGA substrate. The TENG unit contained an island-bridge Mg electrode on which a micropyramid-structured PLGA (P-PLGA) layer was attached facing upward, forming a lower triboelectric layer. An additional island-bridge Mg electrode-coated PLGA layer was placed on top of the P-PLGA layer and thermoplastically sealed, constructing the upper triboelectric layer. The micropyramidal structure of the P-PLGA improved the sensitivity and power density of the TENG by increasing contact at the interface. The island-bridge Mg electrodes were connected using a serpentine geometry. The island bridge configuration and serpentine geometry effectively promote structural robustness, decreasing the overall modulus and minimizing the constraints on FED flexibility.

Figure 25.

Rehabilitation and antibacterial activity of B-TENGs. (A) Schematic illustration of a FED structure. (B) Fracture healing process. (C) Improvement of mineral density and bending performance through FED electrical stimulation. (A–C) Reproduced with permission from ref (54). Copyright 2021 National Academy of Science. (D) Schematic illustration of a TENG-based tissue battery. (E) Cartilage repair system for electrical stimulation using a TENG-based tissue battery. (F) Flow cytometry results that demonstrate accelerated cartilage repair using a TENG-based tissue battery structure. (D–F) Reproduced with permission from ref (130). Copyright 2023 Elsevier. (G) Schematic illustration of the antibacterial mechanism of an RSSP Patch. (H) In vivo antibacterial inhibition of S. aureus by the RSSP Patch. (collected after 7 d, n = 20, ***p ≤ 0.001). (G and H) Reproduced with permission from ref (131). Copyright 2018 Wiley-VCH. (I) Schematic illustration of IBV-TENG under the surgical site to prevent SSI. (J) Images of viable bacteria (E. coli and S. aureus) and the ex vivo antibacterial effect with/without electrical stimulation using IBV-TENG. (I and J) Reproduced with permission from ref (117). Copyright 2023 Wiley-VCH under the terms of the Creative Commons Attribution 4.0 (https://creativecommons.org/licenses/by/4.0/).

Consequently, the FED adhered to irregular surfaces, even under significant stress, such as on a bent finger joint or when subjected to multiple twists. The B-TENG generated electrical outputs by repetitive contact–separation motions between the P-PLGA triboelectric layer and the Mg electrode. The TENG was electrically characterized to optimize the Mg island length (L) at a frequency of 1 Hz. The voltage output peaks increased with an increase in L from 200 to 1000 μm, reaching a maximum value of 4.5 V_PP. Further increases in L to 2000 and 4000 μm decreased the voltage output to 3.2 and 1.2 V, respectively.

The FED was subcutaneously implanted in a rat model to determine its in vivo energy generation, positioning the TENG between the knee joint and the hip (Figure 25B). The voltage output was monitored during regular activity, reaching a value of approximately 4 V. The bone fracture healing was evaluated pathologically, comparing three groups fed and grown under similar conditions (Figure 25C). The mineral density of the right tibia in rats was evaluated and compared after 6 weeks of intervention in groups I, S, and F and a control group. No significant difference in mineral density was found between groups I and N and between groups S and F, with a 27% improvement in mineral density observed in group I compared with the control group. The flexural stress of the tibia bone was also measured, revealing 83% improvement in group I, significantly higher than that of the control group and similar to that of the normal group.

Cartilage is a vital component of several joints, such as the knee, hip, ankle, and shoulder, playing a crucial role in joint lubrication and cushioning.^444,445 However, various factors (e.g., high-intensity physical exercise, rheumatoid arthritis, and age-related degeneration) can cause cartilage defects, resulting in pain and reduced joint mobility.^241,446 Given the limited regenerative capacity of cartilage, the treatment of cartilage defects has become a highly investigated area worldwide.^447,448 Several treatment options, including allogeneic and autologous osteochondral transplantations, have been explored.⁴⁴⁹ Both have advantages and limitations, such as the high preservation cost of allogeneic transplants, the possibility of exogenous diseases and immune solid responses caused by allogenic antigens, and the limited availability of raw materials for autologous transplants.^450,451 The field of microfracture self-healing is still in early development, and further research is required to prove its efficacy. Stem cell therapy delivered through intravascular injection, intra-articular injection, or cell transplantation is also a promising alternative, requiring additional long-term studies to determine its efficacy and safety.^452,453 Cartilage tissue engineering, which involves the use of seed cells, engineered scaffolds, and cell growth factors, is a highly researched area ,as it provides an unlimited source of repair materials.^454,455 However, the hyaline cartilage produced by this approach may not withstand the stress of repeated joint movements as well as normal cartilage.^456,457 Therefore, there is a continuous demand for efficient methods to augment chondrocyte proliferation.

Smart electronic scaffolds for the treatment of cartilage defects (tissue battery) were engineered by combining degradable porous scaffolds and a self-powered sensor based on a multiconvex TENG.¹³⁰ An artificial tissue battery was developed for real-time detection and repair stimulation. The fully bioresorbable tissue battery consisted of two components, a porous scaffold (PCL-PLGA-HA; HA = hydroxyapatite) to fill cartilage defects and a TENG-based pressure sensor composed of CA-CS-HA (CA = collagen aggregate; CS = chitosan) and PCL/PDMS/PCL-FA (FA = fluorapatite) as the positive and negative friction layers, respectively (Figure 25D). The physical connection between the scaffold and the sensor was established through rough surface interlocking, hydrogen bonding, and molecular chain entanglement. The electricity generated from the sensor was collected by electrodes (Ag NWs), serving as the energy source for electrical stimulation and the transmission of electrical signals. The scaffold (PCL-PLGA-HA) allowed chondrocyte proliferation within its pores before material degradation and cartilage defect reparation. The pressure sensor was assembled in a hermetically sealed sandwich configuration to ensure stability and sensitivity in the presence of complex body fluids and under mechanical stress from joint movements. The positive friction layer (CA-CS-HA) featured a double-sided microconvex structure, which improved the effective contact area and the sensor signal. The negative friction layer (PCL/PDMS/PCL-FA) comprised multiple microconvexities acting as pillars to increase the maximum detection range. A mortise and tenon structure formed between the pillar and scaffold (PCL-PLGA-HA), enabling real-time feedback on the cartilage repair status.

The tissue battery had a porous PCL-PLGA-HA scaffold and a TENG-based pressure sensor made of CA-CS-HA and PCL/PDMS/PCL-–FA. The scaffold and sensor were interconnected via physical interlocking and hydrogen bonding, facilitating real-time detection and repair induction. The pressure sensor enabled the real-time monitoring of the repair state, ensuring stable and sensitive performance in the presence of complex body fluids and under mechanical stress from joint movements. The positive friction layer, CA-CS-HA, exhibited a double-sided microconvex structure with improved sensor signal and power density, whereas the negative friction layer, PCL/PDMS/PCL-FA, featured multiple microconvexities that increased the contact area and the maximum detection range of the sensor. PCL-PLGA-HA degraded faster than PCL/PDMS/PCL–FA to ensure an adequate device performance. During implantation, the scaffold was positioned facing the inside of the cartilage to facilitate the migration of growth factors and cells. The TENG-based sensor transformed the mechanical energy from body motion into electricity, promoting the proliferation of chondrocytes (Figure 25E). The electrical signals generated by the tissue battery could also be wirelessly transmitted to mobile devices, computers, or servers, enabling patients and healthcare providers to monitor and assess cartilage repair status and take prompt medical action as needed.

The performance of TENG-based pressure sensors in tissue batteries, including their sensitivity, response range, and detection limit, depends on the polarity and design of the positive and negative friction layers. The addition of HA and FA to the sensors improved their response range by increasing the compression modulus of the material, resulting in a greater external force required for the positive and negative friction layers to come into contact. Incorporating CS in CA-CS-HA counterbalanced the electron-donating character of −NH₂ in CA, thereby improving the sensor sensitivity. The presence of electron-withdrawing groups (such as −CH₃ and −F) in PCL/PDMS/PCL-FA further increased the energy density of the negative friction layer. The microconvex and multilayer sandwich structures on the surface of the positive and negative friction layers increased the contact area when subjected to an external force, leading to a higher energy density. The response range of the sensor could also be increased by incorporating micropillars with 30 μm of diameter in the fabrication process of PCL/PDMS/PCL-FA.

The effectiveness of ES in inactive cell activation was demonstrated in rabbit knee cartilage, as evidenced by a voltage increase of 100 mV·mm^–1 (Figure 25F). However, using biomaterials for cartilage repair poses significant concerns regarding the toxicity of the materials and the potential for infection. Tissue batteries were fabricated using FDA-certified biocompatible polymers to address this issue (CA, PLGA, PCL, PDMS, FA, and HA).⁴⁵⁸ A methyl-thiazolyl-tetrazolium cytotoxicity assay, performed to evaluate the biocompatibility of the tissue battery, revealed that the tissue battery had higher cell viability than the PCL-PLGA-HA composite, indicating that the ES enhanced chondrocyte proliferation.

6.4. Antibacterial Activity Using B-TENGs

Infections can occur during the postoperative wound healing, causing physical pain and creating an economic burden on the patient.^459,460 These infections can manifest in various body incisions, such as the organs, muscles, and skin, resulting in localized pain, abscess formation, organ dysfunction, secondary surgeries, or even death.^461,462 Postoperative infections can increase the mortality rate up to 11×, with an estimated half a million cases reported annually in the United States that cost 3–10 billion dollars.^463,464 The inhibition of microorganism proliferation is crucial for preventing surgical infections.⁴⁶⁵ Although various antibiotics have been developed since the introduction of penicillin in 1928, the emergence of antimicrobial resistance has compromised their effectiveness. Thus, infectious diseases are a leading health problem of the 21st century.^466,467 The COVID-19 outbreak has highlighted the limitations of conventional pharmaceutical treatments, as emerging microorganisms may not be treatable by drugs, leading to pandemic events. This perspective has increased the demand for alternative solutions to antibiotics.⁴⁶⁸ Photodynamic and photothermal therapy were explored for the control of microorganism proliferation.⁴⁶⁹ However, these antimicrobial therapies have significant limitations, such as limited selectivity of treatment sites, low efficiency, and tissue damage due to high temperatures and lighting effects.⁴⁷⁰ Sonodynamic therapy uses ultrasounds to trigger its effects, ensuring deep tissue penetration, noninvasiveness, and nonionization. However, its therapeutic efficiency is limited by the hypoxic microenvironment of deep-tissue infections and tumors. Sonothermal therapy is a form of ultrasonic interfacial engineering with toxicity and biosafety concerns.^330,471

Electrical stimulation (ES) can be used as a nonpharmaceutical method for controlling the growth of microorganisms with known antimicrobial properties.⁴⁷² Electroporation, a physical process that utilizes a strong electric field to damage microbial structures, such as bacterial membranes and viral capsids, is a promising approach for microbial disinfection.^{385,473−475} To achieve highly localized electric fields, Ag, CuO, or ZnO NWs are positioned vertically on a flat electrode surface, and the electric field is significantly amplified at the wire tip (>107 V·m^–1) through low drive voltages (several volts).^476−478 The feasibility of NW-assisted electroporation disinfection technology for the bacterial disinfection of aquatic environments has recently been established.^430,479,480 However, research in the fabrication of bioresorbable NWs is ongoing, limiting the practical implementation of bioresorbable electroporation systems. Recent studies have explored ES-assisted microbial disinfection systems that utilize triboelectricity generated from bioresorbable TENGs. This section provides an overview of the structure and operating mechanism of these microbial disinfection devices.

6.4.1. Low-Frequency Triboelectricity-Driven Antibacterial Activity

A genetically engineered biofunctional TENG was developed using recombinant spider silk proteins (RSSP) for in vivo energy harvesting (Figure 25G).¹³¹ The RSSP was genetically engineered to optimize its charge affinity, resulting in improved triboelectric performance and mechanical strength in TENG devices. The uniform chain lengths in RSSP, when compared with naturally harvested proteins such as spider silks, silk fibroins, and collagens, improved its reliability and repeatability for large-scale manufacturing. Furthermore, the transparency of the RSSP enabled parallel optical readout during energy harvesting. An innovative rapid prototyping technique based on water lithography (WL) was introduced to improve triboelectrification further. This technique creates rough patterns on the surface of RSSP using inkjet printing for biocompatible, economically viable, and scalable fabrication. The hydroink used in this method provided a natural vehicle for molecules and particles, enabling both surface and bulk chemical and physical modifications, as well as bulk doping of the RSSP. This technology can be used for various biofunctional TENG applications, including the fabrication of implantable antibacterial patches. RSSP is produced through recombinant expression in Escherichia coli. The resulting purified RSSP solution can be modified with a range of chemical and biological agents, including graphene, CNT, and drugs, through a straightforward mixing procedure. The pure or functionalized spider silk solutions can then be cast or spin-coated onto PET/indium tin oxide (ITO) substrates with a thickness of approximately 2 μm to meet large-scale demand. A WL technique was developed to create rough patterns on the RSSP surfaces and incorporate functional molecules and particles onto the TENG, improving triboelectrification.^215,481,482 The RSSP/PET/ITO substrate was developed for 12 h in an ethanol/water mixture to produce RSSP-WL-TENG. The duration of this process can be reduced by increasing the annealing temperature.²⁰⁷

The triboelectric mechanism relies on the transfer and re-equilibration of electrons between two friction surfaces. The ability of the material to effectively release and retain electrons without dissipation is vital to achieving good triboelectric performance.^483,484 The RSSP protein was optimized to maximize the number of electrons it could carry. The triboelectric performance of RSSP improved with an increasing number of repeating units and a higher molecular weight, reducing the number of chain ends to create a more uniform and mechanically robust protein. In addition, the uniform chain length of RSSP caused RSSP/PET/ITO films to be transparent even after WL. This transparency enables their use in optical applications.⁴⁸⁵ The output performance of the RSSP-WL-TENG was significantly improved through WL and programmable RSSP crystallinity. Various strategies, including surface structuring and ethanol/water annealing, have been employed to enhance the electrical output and robustness of TENGs, particularly under high-humidity conditions. Uniform holes can be formed by carefully controlling the printing parameters (height of approximately 200 nm; width of about 30 μm), whereas rough patterns are obtained by randomly printing multiple layers on top of each other. As demonstrated by FEM simulations, structured surface morphologies increase the electrical potential when the surface carries the same amount of charge, leading to improved triboelectric output.⁴⁸⁶

The antibacterial mechanism of a fully bioresorbable RSSP patch is illustrated in Figure 24G. The application of triboelectric charging creates a potential difference between the bacteria and the positively charged surface of the RSSP patch. This causes extracellular electron transmission between the bacteria and the patch, leading to a morphological impairment of the bacteria, increased production of ROS,⁴⁸⁷ and ultimately bacterial death. The RSSP patch can also serve as a matrix similar to other silk materials.⁴⁸⁸ The antibacterial properties of the RSSP patch can be further improved by introducing GR and Ag NPs.⁴⁸⁹ The high discharging capacity of the RSSP-based nanocomposites enables a significant charge to be maintained by the environment, providing postcharging antibacterial properties. The RSSP-doped patches exhibited antibacterial properties against Gram-negative (Escherichia coli) and Gram-positive (Staphylococcus aureus) bacteria, respectively (Figure 25H). However, RSSP samples without triboelectric charging had negligible effects on bacterial growth, thus proving that the RSSP does not have intrinsic antibacterial properties. The antibacterial efficiency varied with the doping samples. The RSSP/GR/Ag patch was the most effective, killing 93% of Escherichia coli and 58% of Staphylococcus aureus. The difference in efficacy is attributable to the thicker peptidoglycan layer in the cell wall of Gram-positive bacteria, which provides greater resistance to electric fields.

6.4.2. High-Frequency Triboelectricity-Driven Antibacterial Activity

A technology for inhibiting microorganisms in soft tissues has been proposed using an implantable, fully bioresorbable, and high-frequency vibrating TENG (IBV-TENG) powered by ultrasounds.¹⁷ The implanted IBV-TENG, designed to prevent surgical site infections through electric stimulation, is illustrated in Figure 25I. The IBV-TENG was fabricated using PHBV as the encapsulation and vibrating layer and PVA as the counter friction layer; both PVA and PHBV are bioresorbable and biocompatible materials.⁴⁹⁰ A 50 μm thick Mg foil coated with a PHBV solution (1% w/v) and covered with a PVA solution was selected as the electrode because of its rapid hydrolysis rate and high biocompatibility.³²⁷ The device was designed as a single electrode with a flexible PHBV layer to promote friction between the triboelectric layers and prevent ultrasound reflection. The active area and thickness of the device were approximately 1 × 2 cm² and 170 μm, respectively. The electrical output of the IBV-TENG was measured by placing it in deionized water at different distances (1, 3, 5, and 7 mm) from the ultrasound probe (frequency of 20 kHz; probe intensity of >3 W·cm^–2).¹³⁸ A voltage of approximately 4 V and a current of approximately 22 μA were obtained at 40 MΩ and 1 Ω, respectively. The electrical output was also investigated at varying ultrasound probe power intensities and distances. An increase in the power intensity increased the electrical output, whereas an increase in the distance reduced the electrical output.¹³⁸

An in vitro study was performed to determine the antibacterial activity of the IBV-TENG. The IBV-TENG was securely fixed to a Petri dish and placed under an ultrasound probe. The electrode was kept in a separate Petri dish to eliminate any heating effect from ultrasonic irradiation. The Mg electrode was covered with PHBV to prevent direct contact with the bacterial solution, comprising Gram-negative (Escherichia coli) and Gram-positive (Staphylococcus aureus) bacteria at concentrations of approximately 105 colony-forming units per milliliter.⁴⁸⁷ The bacteria were effectively eliminated through electrical stimulation of 4 V for 1 h using IBV-TENG, resulting in 99% and 100% elimination of Escherichia coli and Staphylococcus aureus, respectively (Figure 25K). The IBV-TENG presented higher antibacterial activity against Staphylococcus aureus than Escherichia coli, possibly because of fundamental structural differences between Gram-positive and Gram-negative bacteria, including an additional protective outer membrane in Gram-negative microorganisms.^491,492 The electrical stimulation produced by the IBV-TENG altered the polarity of the bacterial membrane, disrupting electron transfer in bacterial activity to cause structural destruction and bacterial death.^59,385,493

The survival rate was evaluated under variable conditions to determine the factors responsible for bacterial survival after electrical stimulation (ES). Ultrasounds were applied to a bacterial sample at 1–2 W·cm^–2 for 1 h without IBV-TENG. No significant difference was observed compared with the initial survival rate.⁴⁹⁴ The bacteria also proliferated when cultured on a PHBV-covered Mg electrode without ES and ultrasounds for 2 h, maintaining the original survival rate. The pH was also monitored during the ES (1–2 W cm^–2) for 60 min to determine its role in bacterial survival, revealing no significant changes. Thus, the antibacterial activity of IBV-TENG is pH-independent.⁴⁸⁷

Porcine tissue with fat and muscle layers was used as the test substrate to study the survival rate of bacteria ex vivo (Figure 25J). A bacterial solution (30 μL, approximately 105 CFU·mL^–1) was spread on 0.5 × 2.5 cm² of porcine tissue. The IBV-TENG was placed on the contaminated portion of the tissue and covered with a pork skin layer with a thickness of approximately 5 mm. The ultrasound probe was fixed above the skin layer where the IBV-TENG was implanted, and ES was performed on the contaminated tissue. The viability of the bacterial colonies was determined via imaging. The ultrasound treatment was ineffective at killing the bacteria in the initial stage. However, the survival rate significantly reduced after ES, with over 92% and 86% inactivation of Escherichia coli and Staphylococcus aureus, respectively.

7. Challenges and Future Perspectives

B-TENGs have recently emerged as promising implantable energy solutions, exhibiting rapid advancements owing to an in-depth understanding of their degradation mechanisms and material designs.^58,124,196 This enables a high output and modulation of the device’s lifetime according to the application. B-TENGs with active operation are feasible alternatives with enhanced functionalities, enabling the precise control and adjustment of the lifetime of the triboelectric implant to minimize possible adverse health effects.^117,122,137 The large output power, simple structure, and extensive material options of B-TENGs led to their broad application to power implantable electronics and to act as self-powered bioresorbable electronics. B-TENGs have been used to collect physiological signals for health monitoring, disease diagnosis, tissue and nerve regeneration, rehabilitation, and antibacterial activity to prevent surgical infections.^{30,117,123−125,127,128,131,208,495}

Despite recent advances and future opportunities, B-TENGs face significant challenges that must be addressed to ensure their successful implementation and adoption in healthcare and biomedical applications. We may achieve technological innovation through fundamental materials studies that tackle the central requirements of B-TENGs (e.g., high output power and controlled biodegradation). Despite the considerable advancements made in B-TENGs over recent years to enhance power output and impose an adjustable lifespan in physiological conditions, particular challenges still exist. The limited choices of triboelectric pairs and the need for highly triboelectrically negative bioresorbable materials continue to be major obstacles to achieving breakthroughs in output power. Further, while some studies have reported controlled biodegradation by employing encapsulation polymers and laser-induced thermal transitions, there has yet to be substantial progress in ensuring the safe and effective clearance of B-TENGs. The research in this area remains nascent compared to mechanisms explored in transient electronics. These limitations contribute to challenges in ensuring technological requirements such as functional stability, biocompatibility, and comprehensive functionalities. Preclinical and clinical evaluations of degradation behavior and biological impacts from various angles are pivotal to manifesting the potential of B-TENGs as a novel category of IMDs and power sources. However, most studies have only confirmed their potential in feasibility tests. Therefore, the evolution of next-generation B-TENGs should prioritize advanced bioresorbable materials research to tackle these challenges, necessitating a multidisciplinary approach. Figure 26 provides an overview of the current technological challenges, and numerous research groups are concentrated on addressing these limitations. In this section, we provide ongoing research trends for each of the current challenges of B-TENGs and suggest feasible strategies to overcome them.

Figure 26.

Challenges and required properties of future TENGs for transient electronics and IMDs.

7.1. Effective And Predictable Transience

Despite substantial progress, the control of B-TENG lifetime remains challenging. The varying degradation mechanisms and weight loss rates of different component materials lead to discrepancies between the functioning and disappearance times of the B-TENG, creating a risk for adverse health effects from residual material. These constraints limit the available clinical options for implant adjustment, increasing the need for frequent hospital visits and the reliance on repetitive imaging diagnostics to monitor the transient behavior. This increases the burden on the patient and healthcare provider. Addressing these challenges is crucial to improving the adaptability and practicality of B-TENGs in clinical settings, leading to better patient outcomes and reducing the strain on healthcare resources. Although current research addresses this challenge using stimuli-responsive materials, most studies on integrating active operation to B-TENGs still need to develop practical methods to trigger their transience successfully in a noninvasive manner. This limitation has hindered the broad application of B-TENGs, as stimuli can inadvertently harm the patient. Consequently, further research is required to overcome these obstacles and fully harness the capabilities of B-TENGs for healthcare and biomedical applications.

Using ROS generators and drug delivery systems is promising to address this challenge. ROS generators produce oxygen radicals in response to various triggers such as heat, light, and ultrasound. The combination of ROS generators and ROS-reactive polymers has significant potential for TENG biosafety and effectiveness.^47,496 Given that the chain reaction propagates continuously throughout the entire component materials, only a few instances of low-intensity stimuli are required to destroy the device. This approach can effectively accelerate the clearance of bioresorbable materials using limited stimulation energy, thereby addressing concerns regarding the controlled degradation and lifetime of B-TENGs. Drug delivery technologies have a significant potential for their use in bioresorbable devices. Drug-loaded polymer matrices or composites containing nanocarriers, such as metal–organic frameworks and polymersomes, have been developed to store and release drugs on-demand for clinical applications.^{45,56,328,497−499} These composites can store and release enzymes, etchants, or oxidants to promote degradation, creating a highly effective and reliable active operation system. Integrating wirelessly controlled drug reservoirs into bioresorbable devices, and making small reservoir valves to control drug release, minimizes the trigger energy, improving biosafety. Micro-light-emitting diodes (micro-LEDs) can also trigger transience by combining light-responsive materials or systems.^500,501 These strategies may address the current challenges of B-TENGs, controlling their lifetime and degradation and paving the way for developing more reliable, effective, and safe bioresorbable devices for healthcare and biomedical applications.

7.2. Stability

Selectivity and mechanical durability are essential factors in the development of stable B-TENGs. The triggering mechanism must be selective, only triggering with the desired response. In many cases, selectivity and biosafe triggering cannot be obtained simultaneously. Though stimuli-responsive materials should be sensitive to initiate erosion in response to low-energy stimuli, this may cause unintended transience to other triggers or impacts. Implanting AND gate logic in the stimulus that activates transience, requiring two or more distinct types of stimuli for activation, can help improve selectivity.^502−504 The use of dual frequency bands for power generation and triggering events can enhance the selectivity of active operations,⁵⁰⁵ preventing interference between the two functions to ensure that the TENGs perform optimally.

In addition, it is essential to ensure that TENGs maintain their mechanical durability throughout the degradation process, withstanding environmental stress and pressure without compromising the triggering mechanism. In reservoir-integrated devices, the release of degradation factors from the destroyed drug reservoir can cause the uncontrolled disappearance of the device, affecting its performance and harming the surrounding tissues. Thus, the reservoir and its contents must be carefully designed to minimize the sudden release of degradation factors by mechanical impact. B-TENGs must gradually erode in the physiological medium to maintain a stable performance before activating the transient behavior. This requires an understanding of the mechanisms of molecular degradation and the dominant macroscopic degradation mode of the encapsulation layer, as it can considerably impact the functioning time. According to reported theoretical and experimental degradation models, the dominant degradation mode is determined by several factors, including the diffusion coefficient, specimen thickness, and erosion rate. Thus, the TENG design can be optimized by predicting the dominant degradation mode, ensuring stable performance. This can be achieved by using barrier layers or suitable material processing to improve the chemical durability of the encapsulation layer and minimize gradual erosion, as well as careful monitoring of the performance of the device to ensure that it meets the desired specifications.

7.3. Biocompatibility and Biosafety

The biocompatibility and biosafety of B-TENGs are vital for their application. Most existing studies on B-TENGs focused on cytotoxicity and immune response studies of their components, including the bioresorbable triboelectric polymer, the encapsulation material, and the metal.^45,46,58 However, examining the biocompatibility of the constituent elements is insufficient. It is also necessary to evaluate the cytotoxicity and immune response to the surrounding tissues that may occur due to the transient process. Furthermore, it is required to identify how the device components are absorbed and eliminated in the body. One representative approach to follow their elimination pathway utilizes ICP-MS to detect residual amounts of metallic elements within biological tissues during device degradation.³¹⁷ Finally, verifying that the electrical energy generated does not adversely affect biological systems is necessary.

In addition to evaluating the biocompatibility and safety of B-TENG components, it is vital to examine the biosafety of triggered transience and various triggers, including light, heat, and ultrasounds.^{45,56,59−61,340,356} The use of triggers to initiate the degradation of B-TENGs raises concerns about their potential impact on surrounding tissues and cells. Researchers must consider the intensity, duration, and frequency of trigger application when evaluating the biosafety of triggered transience, as each trigger may have unique interactions with biological tissues, potentially causing thermal effects, mechanical stress, and other unintended consequences. When applied to initiate B-TENG degradation, these triggers must not induce tissue damage, inflammatory responses, or other adverse effects. Furthermore, trigger biosafety should be evaluated at various molecular- and organ-level scales. Here, in vitro and in vivo studies can reveal the potential risks and benefits of using these triggers. In vitro studies can help determine cytotoxicity, cellular uptake, and cellular responses, while in vivo studies can assess broader biological effects and interactions within living organisms.^60,328,340 The biocompatibility and safety of the byproducts generated during the triggered degradation process must be evaluated. The release of degradation products should not cause any harmful effects on the surrounding tissues or the overall biological system.

7.4. Output Power

B-TENGs are promising energy sources for various applications. However, their output performance must be increased for practical application in energy harvesting, sensing, and medicine. To increase the output performance of B-TENGs, researchers are exploring new materials and technologies that can improve energy conversion efficiency and increase output voltage and current. It may require high-performance dielectric materials, advanced fabrication techniques, and novel designs employing nanomaterials, high-k composites, ferroelectrics, high-surface-charge materials, and multistack structures.^{49,58,64,139,159,206,506} Most implantable TENGs driven by the physical motion of organs or low-frequency vibrations of the human body (e.g., heartbeat, gastronomic movements, walking, and arm shaking) exhibit low output performance, limiting the power of implantable electronics. The miniaturization of implantable TENGs, an essential requirement for medical implants, is also challenging owing to their low output power.

Hinchet et al. first reported ultrasound-driven implantable TENGs with approximately 1000× higher output compared with previous studies.¹³⁸ Since then, the potential of ultrasound-driven TENGs has been widely recognized as a promising approach to increasing the output performance of on-demand B-TENGs.^58,138 However, most bioresorbable polymers still have a low surface charge density due to tribo-neutral moieties, which leads to low performance. Bioresorbable high-k composites, ferroelectric materials, high-surface-charge nanomaterials, and multistack configurations can be used to enhance the performance of TENGs.

7.5. Self-Powered Functionalities

B-TENGs generate electrical energy from ambient mechanical stimuli, removing the need for external power sources. This substantially reduces their size, weight, and complexity and minimizes the need for batteries, wires, or other circuit components, rendering them more suitable for implantation.^{52,64,123−125} Moreover, B-TENGs reduce the risk of infection and inflammation associated with the physical stress caused by large-volume implants.^28,139,417 B-TENGs can be implanted close to the affected or monitored area because of their streamlined configuration and miniaturized dimensions.⁵⁸ This eliminates most electrical noise, improving accuracy when delivering electrical impulses and collecting targeted physiological signals. These characteristics ultimately translate to valuable and effective devices for biomedical applications.

However, the self-powered nature of these devices is limited by the availability of mechanical energy sources. Mechanical movement occurs in multiple locations throughout the human body, such as the heart, digestive organs, and muscles.^118,124 B-TENGs generate electrical currents by exploiting this mechanical motion through deformation. However, if a location does not provide sufficient mechanical energy, the B-TENG must be positioned in a suitable location and connected to the working components via lead wires. The intermittent electrical output, lack of control integrated circuits (ICs), and absence of energy storage also restrict their application. Given that B-TENGs only output power in the presence of biomechanical energy, their proper functioning cannot be confirmed without the implantation of additional monitoring devices, limiting their use to the few applications that do not require strict control of the amplitude, frequency, and waveform of electrical impulses.

Several approaches can be implemented to address these limitations and exploit the advantages of B-TENGs. First, the function of the device can be aligned with the biomechanical energy pattern. A previous study had implanted a TENG device onto the outer surface of the stomach, stimulating the vagus nerve when the stomach moved in response to appetite to reduce food intake.¹⁶⁴ B-TENGs can also be powered using externally controlled mechanical energy sources. For example, the output power of ultrasound-mediated TENGs can be intentionally controlled by modulating the intensity and incident pattern of the ultrasounds.¹³⁸ Thus, a database can be built using clinical tests and simulation or machine learning models to determine the output power of US-TENGs in response to a given ultrasound intensity, enabling better control and monitoring of their functions.

7.6. Multiple Combined Functionalities

Integrating bioresorbable power management-integrated circuits and batteries with B-TENGs is crucial for effectively implementing complex and accurate bioresorbable implants in clinical settings. This integration can enable a wide range of applications, including exact and tunable electrical stimulation, monitoring of multiple physiological signals, and wireless data transmission. Compared with existing implants that require high-capacity batteries and occupy considerable space, B-TENGs enable the miniaturization of IMDs using low-capacity batteries and B-TENG recharging. However, several challenges still need to be addressed. First, bioresorbable batteries with high energy density must be developed for miniaturization, as most current bioresorbable batteries have low energy densities and limited durability.^36,507 Second, the output power generated by B-TENGs must be increased to support high-power IMDs and reduced charging times. Finally, bioresorbable power management circuits with high conversion efficiency must be created. However, their development remains challenging because numerous passive and active elements are needed to manage the high-impedance output of TENGs.^357,508 The development of bioresorbable batteries for integration with these devices is also of concern, as their biosafety has yet to be fully proven.

8. Conclusion

The development of B-TENGs is a remarkable breakthrough in biomedical engineering. These devices present a safe and sustainable alternative to conventional power sources in IMDs, obviating the need for extraction surgery and follow-up assessments to monitor residual materials. Incorporating these bioresorbable devices can enhance medical treatment delivery, ultimately leading to improved patient outcomes. Significant challenges are associated with B-TENGs, such as ensuring adequate and predictable transience, improving the stability, biosafety, and output power, and designing innovative functionalities and bioresorbable electronics for integration. Nevertheless, their future appears promising, with vast potential for advancing biomedical technology. Continued research and development efforts in this field are vital for the widespread adoption of B-TENGs, paving the way for a more sustainable healthcare system. By harnessing the potential of B-TENGs, we can transform biomedical engineering and contribute to advancing healthcare worldwide.

Biographies

Minki Kang is currently a postdoctoral researcher in the School of Advanced Materials Science and Engineering at Sungkyunkwan University. His research interests include energy harvesting and biomedical applications based on biocompatible, biodegradable materials.

Dong-Min Lee is currently a Ph.D. candidate in School of Advanced Materials Science and Engineering at Sungkyunkwan University. His research interests focus on advanced skin/tissue-interfacing electronics and their diagnostic and therapeutic use.

Inah Hyun is currently a Ph.D. student in the Department of Materials Science and Engineering at Yonsei University. Her research interests include implantable and bioresorbable energy harvesting devices and their biomedical applications.

Najaf Rubab is currently a research professor in the Department of Materials Science and Engineering, Gachon University. Her research interests include biomedical energy harvesting based on triboelectric technology.

So-Hee Kim is currently a Ph.D. student in the Department of Materials Science and Engineering at Yonsei University. Her research interests include triboelectric nanogenerators for energy harvesting and their applications in medical devices.

Sang-Woo Kim is a YONSEI World-Class Fellow Professor at Yonsei University. His recent research interests are focused on triboelectric/piezoelectric nanogenerators, self-powered sensors, body-implantable devices, and 2D materials. He has published over 350 research papers (with an H-index of 88). He served as the Chairman of the fourth NGPT conference in Korea in 2018. Currently, he holds the positions of Director of the Center for Human-oriented Triboelectric Energy Harvesting (Research Leader Program) and Director of the National Core Materials Research Center, which is supported by the National Research Foundation of Korea.

Author Contributions

^∥ M.K., D.-M.L., and I.H. contributed equally to this work. CRediT: Minki Kang conceptualization, writing-original draft, writing-review & editing; Dong-Min Lee conceptualization, writing-original draft, writing-review & editing; Inah Hyun conceptualization, writing-original draft, writing-review & editing; Najaf Rubab visualization; So-Hee Kim writing-review & editing; Sang-Woo Kim conceptualization, funding acquisition, project administration, supervision, writing-review & editing.

The authors declare no competing financial interest.

Special Issue

Published as part of the Chemical Reviewsvirtual special issue “Wearable Devices”.