System overview

The SPM consisted of three parts: the energy harvest unit (iTENG), power management unit (PMU), and pacemaker unit. The energy harvest unit could harvest energy from cardiac motion. At first, switch of the PMU was turned off, the electricity generated by the energy harvest unit was stored in the capacitor of the PMU. Then, the switch was turned on by a magnet which was used as wireless passive trigger, the electrical energy could drive the pacemaker unit to produce pacing electrical pulses and control the rate of cardiac contraction (Fig. 1a). There was a core-shell structure of iTENG that consisted of two triboelectric layers, supporting structure and the shell with two encapsulation layers (Fig. 1b–d). Nanostructured polytetrafluoroethylene (PTFE) thin film was employed as one triboelectric layer (Fig. 1e). A three dimensional (3D) elastic sponge (ethylene-vinyl acetate copolymer, EVA) played a role as a spacer (Fig. 1f), and a memory alloy ribbon (highly resilient titanium) was utilized as the keel. The iTENG was entirely packaged by a flexible Teflon film and a polydimethylsiloxane (PDMS) layer to enhance its structural stability and avoid environmental liquid damage to the device.

Fig. 1 Overview of symbiotic pacemaker system. a Illustration of symbiotic cardiac pacemaker system. b Schematic structure diagram of implantable triboelectric nanogenerator (iTENG). c Photograph of iTENG under bending. d Cross-sectional scanning electron microscope (SEM) image of the iTENG (scale bar: 500 μm). e SEM images of the nanostructure on polytetrafluoroethylene (PTFE) film (scale bar: 1 μm). f SEM image of three dimensional (3D) elastic sponge structure (scale bar: 500 μm). g, h Schematic representation of the mechanism of charge transfer. i The model used to estimate the amount of charge separation that can arise from the transfer of charges Full size image

The operating principle of the iTENG was based on the coupling of contact electrification and electrostatic induction. Contact electrification is mainly caused by the transfer of surface electrons35 (Fig. 1g–i, Supplementary Note 1). Then an electrical potential between two triboelectric layers drove the electrons through external loads due to electrostatic induction.

Characterization of the implantable triboelectric nanogenerator

The effective contact area and surface charge density have significant impacts on the TENG output26,36. The nanostructure of PTFE and spacer/keel supporting structure could effectively increase the contact area to improve output performance of the TENG37 (Supplementary Note 2, Supplementary Table 2, Supplementary Fig. 1). Corona discharge method was used to increase the surface charge density of the PTFE triboelectric layer, which can enhance the output of TENG as well (Fig. 2a). The current flowed from corona needle with high potential into the air, by ionizing and creating a region of plasma around the needle. The ions eventually passed charge to areas of lower potential (PTFE film). A mechanical linear motor was employed to characterize the effect of corona polarization on the electrical output of the TENG. The V OC (open-circuit voltage), Q SC (short-circuit transferred charge), and the corresponding I SC (short-circuit current) of group polarized were up to 187 V, 80.2 nC, and 19.5 μA, respectively. In comparison, the non-polarized ones were 67.5 V, 24.8 nC, and 5.9 μA accordingly. The TENGs with unified specification were utilized in this experiment (Fig. 2b–e).

Fig. 2 Polarized polytetrafluoroethylene film based triboelectric nanogenerator. a Sketch of a corona discharge system. b Schematic diagram of the working principle of iTENG. c–e The output voltage, transferred charge and current of polarized and non-polarized PTFE film based TENG driven by a linear motor. Source data of c–e are provided as a Source Data file. All data in c–e are presented as mean ± s.d. Full size image

Then the polarized PTFE-based TENG was hermetically sealed by flexible encapsulation layers to fabricate an iTENG. The linear motor was used to simulate low-frequency biomechanical excitation for testing the in vitro electrical output performance of the iTENG. The average values of V OC , Q SC and the I SC were 97.5 V, 49.1 nC, and 10.1 μA, respectively (Fig. 3a–c). Further investigations of the effective electric power of the iTENG showed that the instantaneous current decreased and voltage rose with increase of the load resistances (Fig. 3d). Hence, a peak power density of 110 mW m−2 was achieved at a load resistance of 100 MΩ (Fig. 3e).

Fig. 3 In vitro evaluation of the implantable triboelectric nanogenerator. a–c Open-circuit voltage, transferred charge and short-circuit current of the iTENG driven by a linear motor. d Voltage and current at different load resistances. e Peak power density at different load resistances. f Stability tests of iTENG. g Fluorescence images of stained L929 cells that were cultured on encapsulation layers of the TENG; the scale bar is 50 μm. h The normalized viability of L929 cells after being cultured for 3 days. Source data of h are provided as a Source Data file. All data in h are presented as mean ± s.d. Full size image

Superior electrical output, biocompatibility, and stability are critical aspects for an in vivo energy harvester. An accelerate fatigue test was employed to evaluate the long-term output performance of the iTENG. After 100 million mechanical stimuli cycles by vibration table, V OC of the iTENG driven by linear motor was maintained stably at 95 V compared with its initial state, exhibiting outstanding durability and stability (Fig. 3f). In addition, to explore the impact of the ionic liquid environment (mimicking the in vivo environment) on iTENG long-term operation, the test environment was replaced with phosphate buffer saline (PBS 1×). After accelerate fatigue test in PBS for 100 million cycles, the output voltage, transferred charge and current of the iTENG were about 93 V, 47 nC, 9 μA, respectively, when tested in dry environment (relative humidity 40–50%) and 91 V, 45 nC, 8 μA when tested in PBS solution (Supplementary Fig. 2a–f). There is no water penetration and damage on the iTENG (Supplementary Fig. 3a–e). The encapsulation layer can effectively avoid negative effects of wet conditions on output of the iTENG.

Excellent biocompatibility is essential for IMEDs to avoid an adverse influence on the surrounding tissue. As an overview for the biocompatibility and cytotoxicity of iTENG, we observed growth, and viability of Mouse fibroblast (L929s) on the above encapsulation layer material and the cell culture dish. L929s adhered to both of the materials, with similar spreading, and intact detectable cellular structures, i.e. cell nucleus and actin microfilaments (Fig. 3g). The MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-2H-tetrazolium bromide) value of the experiment group was similar to the value of the control group after 3 days of culture (Fig. 3h). These results demonstrated the good cytocompatibility of the iTENG.

In vivo performance of the implantable triboelectric nanogenerator

To evaluate the performance of the iTENG as the energy harvest unit in vivo, a large animal model (Adult Yorkshire porcine, male, 45 kg) was employed in this work (Fig. 4a–e). The iTENG was placed between the heart and pericardium, and the PTFE side faced the left ventricular wall. Cardiac motion caused periodic contact and separation of the two triboelectric layers. The electrical energy generated by iTENG was stored in a 100 μF capacitor through a rectifier. As a result, the voltage of capacitor could be charged from 0 to 3.55 V within 190 min under conditions of blood pressure of ~100/70 mmHg and heart rate of ~77 bpm (Fig. 4f, Supplementary Fig. 4, Supplementary Movie 1). Here, in vivo V OC was up to ~65.2 V, Q SC was ~13.6 nC, and the corresponding I SC was ~0.5 μA (Fig. 4g, Supplementary Fig. 5a–c).

Fig. 4 In vivo energy harvest and electrical characterization. a, b The iTENG implantation process in animal experiments. c, d The iTENG was driven by the diastole and systole of the heart. e Schematic of in vivo experimental electrical characterizations. f Charging curve of a 100 μF capacitor charged by iTENG. g In vivo open-circuit voltage, transferred charge, short-circuit current of the iTENG and simultaneously recorded electrocardiography (ECG). h In vivo output open-circuit voltage and simultaneously recorded ECG signals. i The relationship between ECG signals and the voltage of a 1 μF capacitor charged by iTENG. j Statistics-analysis of minimum voltage, maximum voltage, and the voltage difference. k Statistics-analysis of minimum transferred charge, maximum transferred charge, and the transferred charge difference. Source data of j, k are provided as a Source Data file. All data in j, k are presented as mean ± s.d. Full size image

In addition, the electrical output of iTENG was completely synchronized with the corresponding electrocardiography (ECG) (Fig. 4h). The rise inflection point of the voltage of the iTENG was consistent with the peak of R wave. When the atrioventricular (A–V) valves closed, the heart turned into phase of systole. The iTENG was compressed and produced a voltage pulse with the width equivalent to phase of systole (Supplementary Fig. 6a, b). Furthermore, the rise inflection point of the current of the iTENG was consistent with the peak of R wave, and the peak of the current signal was synchronized with S wave (Supplementary Fig. 6d, e, Supplementary Note. 3). The voltage of the capacitor was simultaneously stepwise increasing with each cardiac cycle (Fig. 4i, Supplementary Fig. 7). The generated energy during each cycle was E max= 0.495 μJ. Here, the average values of V OC,max , V OC,min , and ΔV OC were 65.2 V, −7.7 V, and 72.9 V, respectively. The average values of Q SC,max , Q SC,min and ΔQ SC were 13.6 nC, −1 nC, and 14.6 nC, respectively (Fig. 4j, k, Supplementary Note 4).

To present an intuitive view of the iTENG powering electronic devices, the energy harvested by the iTENG from cardiac motion was stored in a capacitor (100 μF, 3.55 V) to drive a commercial pacemaker. The commercial pacemaker generated a series of pacing electrical pulses with a voltage of about 4 V and a pulse width of 0.9 ms (Supplementary Fig. 8a–d, Supplementary Note 8). A light-emitting diode (LED) was also directly connected to the implanted iTENG. The LED blinked synchronously with heart beating (Supplementary Fig. 9a–c, Supplementary Movie 2, Supplementary Note 9).

Pacing in large animal model by symbiotic pacemaker in vivo

The iTENG was connected to the PMU via wires and rectifier. The electrical energy generated from the iTENG was stored in the capacitor of the PMU, which started to power the pacemaker unit after the magnet placed outside the body turned on the reed switch (Fig. 5a, b, Supplementary Fig. 10). The generated electrical pulses from the pacemaker unit can induce myocardial contraction and regulate heart rate through pacing electrodes. The output voltage and duration of the electrical pulses were 3 V and 0.5 ms respectively (Fig. 5c). The rate of the electrical pulses was preset to 130 bpm, in consideration of the high heart rate of the pig during the experiment.

Fig. 5 symbiotic cardiac pacing in vivo. a Illustration of symbiotic cardiac pacemaker system turned on by wireless passive trigger. b A block diagram of the components in symbiotic cardiac pacemaker system. c Stimulation pulse with different frequencies generated by pacemaker unit. d, e Symbiotic cardiac pacemaker system turned on by wireless passive trigger in animal experiments. f ECG, Femoral Artery Pressure (FAP) and heart rate (HR) and systolic blood pressure (sBP), Stimulus-R wave (S-R) interval during symbiotic cardiac pacemaker system work. g ECG of the intrinsic heart rate, with a normal sBP. h ECG with a pacing stimulus in the refractory period, with a normal sBP. i ECG of successful pacing, with a significantly decreased sBP. j ECG with failed pacing by attenuated stimuli, with a restored sBP Full size image

To achieve cardiac pacing in a large animal model, we implanted the entire SPM system into the chest of a pig. After continuously harvesting energy from cardiac motion in ~200 min, the SPM was switched on by the wireless passive trigger (NdFeB magnet). The electrical pulses (0.5 ms, 130 bpm; setting parameter) were generated and transmitted to the myocardium (Fig. 5d, e). Some typical exhibitions of ECG had appeared during the pacing process by SPM (Fig. 5f). Here, cardiac physiology states of experimental animals can be divided into four periods, including intrinsic rhythm, invalid pacing, valid pacing and invalid pacing period (Supplementary Movie 3).

The typical P wave, QRS complex and T wave could be identified in the ECG of intrinsic rhythm (Fig. 5g). When the SPM was switched on and releasing the pacing stimuli in the refractory period, the heart had no response to the stimuli and showed an ECG of invalid pacing (Fig. 5h). Once the pacing stimuli released in the non-refractory period, a contraction of heart and a stimulated QRS complex could be observed in the ECG. Here, the stimulated QRS complex appeared immediately following the pacing stimulus, indicating the heart was successfully paced by SPM (Fig. 5i). With the consumption of the energy by fast pacing, the voltage of the capacitor of the PMU and the amplitude of the pacing electrical pulse declined concomitantly, leading to failure in pacing the heart. The pacing stimulus could not induce a stimulated QRS complex in ECG and an invalid pacing ECG could be detected (Fig. 5j).

The variations of heart rate and blood pressure of the experimental animal further illustrated the above-described process. The heart rate of experimental animal was increased from 90 bpm in the intrinsic phase to ~130 bpm in the valid pacing phase and then returned to 90 bpm during the invalid pacing phase. The systolic blood pressure declined significantly from ~100 mmHg of the intrinsic phase to ~60 mmHg of the valid pacing phase due to the fast pacing by the SPM. It returned to ~90 mmHg when the heart rate slowed down in the invalid phase. The above results confirmed that the SPM system successfully achieved cardiac pacing in large animal scale.

Correcting arrhythmia in large animal model

To demonstrate the ability of SPM to correct arrhythmia, we performed pacing therapy on an animal model (Adult Yorkshire porcine, male, 35 kg) with sinus arrhythmia induced by sinus node hypothermia38,39. The arrhythmia induced by sinus node hypothermia may deteriorate to sinus arrest and even ventricular fibrillation, which would cause death if not treated promptly (Fig. 6a, Supplementary Fig. 11a–c).

Fig. 6 Correcting arrhythmia on large animal model. a Illustration of symbiotic cardiac pacemaker system correcting arrhythmia on large animal model. b Symbiotic cardiac pacemaker system in animal experiments. c electrocardiography (ECG), Femoral Artery Pressure (FAP) of the animal model and voltage of capacitor during correcting arrhythmia experiment Full size image

The electrical energy generated by the iTENG was stored in a 200 μF capacitor for powering the pacemaker. As a result, the voltage of the capacitor can be charged from 0 to ~4 V within 63 min under the blood pressure of ~110/60 mmHg and heart rate of ~82 bpm (Fig. 6b, c). Here, in vivo V OC was up to ~39 V, Q SC was ~21 nC, and the corresponding I SC was ~0.8 μA (Supplementary Fig. 12a–c). In addition, the output voltage and power have been evaluated under resting (∼50 bpm), active (∼90 bpm), and stressing (∼130 bpm) states (Supplementary Fig. 13, Supplementary Note 6).

Ice cube was used to create sinus node hypothermia, then a typical arrhythmia ECG was observed. The pacing therapy was performed promptly. Sinus arrhythmia was converted to pacing rhythm when the iTENG-based SPM was turned on. Heart rate remained at ~68 bpm, and blood pressure began to recover to the previous level. After about one minute, the power supply voltage dropped to 1.4 V. The SPM stopped working, and the pacing rhythm turned into normal heart rhythm. This result confirmed that SPM successfully corrected sinus arrhythmia and prevented further deteriorating condition (Fig. 6c).