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Direct Powering a Real Cardiac Pacemaker by Natural Energy of a Heartbeat Ning Li,†,⊥ Zhiran Yi,‡,⊥ Ye Ma,†,⊥ Feng Xie,† Yue Huang,† Yingwei Tian,‡ Xiaoxue Dong,‡ Yang Liu,† Xin Shao,† Yang Li,† Lei Jin,† Jingquan Liu,‡ Zhiyun Xu,† Bin Yang,*,‡ and Hao Zhang*,† †
Institute of Cardiothoracic Surgery at Changhai Hospital, Second Military Medical University, Shanghai 200433, China National Key Laboratory of Science and Technology on Micro/Nano Fabrication, Key Laboratory for Thin Film and Microfabrication of the Ministry of Education, Department of Micro/Nano Electronics, Shanghai Jiao Tong University, Shanghai 200240, China
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‡
S Supporting Information *
ABSTRACT: Implantable medical devices are widely used for monitoring and treatment of severe diseases. In particular, an implantable cardiac pacemaker is the most effective therapeutic device for treating bradyrhythmia, however its surgical replacement is inevitable every 5−12 years due to the limited life of the built-in battery. Although several approaches of energy harvesting have been explored in this decade for powering cardiac pacemakers, the modern, commercial, and full-function pacemaker has never been powered effectively yet. Here, we report an integrated strategy for directly powering a modern and full-function cardiac pacemaker, which can pace the porcine heart in vivo by harvesting the natural energy of a heartbeat, without using any external energy storage element. The generator includes an elastic skeleton and two piezoelectric composites, which could generate a high-output current of 15 μA in vivo over state-of-the-art performance. This study makes an impressive step toward fabricating a selfpowered cardiac pacemaker and resolving the power issue of implantable medical devices by piezoelectric harvesting technology. KEYWORDS: piezoelectric energy generator, implantable harvester, medical device, self-powered, cardiac kinetic energy, PMN-PT, batteryless pacemaker
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the development and clinical application of implantable medical devices. The heart in the human body generates a large hydraulic power,15 indicating that the generators can scavenge the natural energy from a heartbeat inexhaustibly. Previously, different research groups have explored the possibilities to transduce such robust kinetic energy into electric energy.11−13,16−32 The intracorporal energy scavenging technologies include electromagnetic induction,20,33,34 electrostatic,27 piezoelectric generators,11,13,16,21,23,24,35,36 and triboelectric nanogenerators.12,17,25 However, the in vivo outputs of aforementioned devices are still insufficient to be utilized as a power source for implantable cardiac pacemakers or other medical devices. The rigid structure and uneasy miniaturization are the main obstacles for electromagnetic generators. The surface damage13 and additional bias voltage limit the
mplantable medical devices have made a significant contribution to clinical monitoring, diagnosis, and treatment.1−8 An increasing number of intelligent automatic devices are applied in clinical practice for improving the patient’s health condition. Patients with bradyrhythmia need to be treated with implantable cardiac pacemakers to increase heartbeat, while patients who suffer from supraventricular tachycardia, atrial fibrillation/flutter require implantable cardioversion devices to regain sinus rhythm. However, the lifetime of these devices is restricted by the lifespan of their built-in batteries. For example, cardiac pacemakers powered by a lithium iodine battery have a functioning lifespan of 7−10 years, while the lifespan of implantable cardioverter-defibrillators (ICD) is about 4−6 years.9,10 These implantable medical devices have to be replaced periodically once the battery is depleted, and the replacement surgery is inevitably associated with several complications, including infection and bleeding. 11−14 Hence, prolonging the lifespan of the implantable medical devices remains the key challenge for © XXXX American Chemical Society
Received: November 9, 2018 Accepted: January 24, 2019
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DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
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Figure 1. Schematic illustrations, mechanical flexibility, and working principle of the implantable piezoelectric generator (iPEG). (a) The schematic illustration for the iPEG showing the symmetry structure of the iPEG and an expanded view for the piezoelectric composite on the upper side of the elastic skeleton. (b) SEM image of cross-section for the piezoelectric composite on the elastic skeleton. The inset shows the position (A-A′) of the cross-section. (c) Photograph of the piezoelectric composite bent by human fingers. (d and e) Shape changes and the corresponding output of the iPEG at initial, compressing, compressed, and expanding state in one working circle.
capsular structure. The profile looks like a compressed number “8”. This shaped PET structure is used as the elastic skeleton for the capability of being compressed and quickly rebounding, which could induce a robust deformation of piezoelectric composites. The piezoelectric composite consists of a piezoelectric layer, beryllium-bronze foil, and Cr/Au electrodes. Briefly, single-crystalline (72%)Pb(Mg1/3Nb2/3)O3-(28%)PbTiO3 (PMN-PT) with high piezoelectric coupling coefficient (e31 = 3.91 C/m2, e33 = 18.03 C/m2) and electromechanical coupling factor (k31 = 0.87, k33 = 0.95) is employed as the piezoelectric layer with a thickness of 50 μm. Each side of the piezoelectric layer is sputtered a layer of Cr/Au (50 nm/ 200 nm thick) as top and bottom electrodes. The berylliumbronze foil (50 μm thick) is used to supply uniform stress distribution to the piezoelectric layer. Then two piezoelectric composites are symmetrically bonded on the top and bottom surfaces of the above-mentioned elastic skeleton for achieving switchable output. Copper wires are prepackaged in medicalgrade silicone as electrical leads and bonded to the Cr/Au electrodes utilizing conductive epoxy. A PDMS film (40 μm thick) is deployed by spin coating to obtain the leak-proof characteristic and reduce surface roughness. To further enhance the stability of the device and avoid potential erosion in the complicated in vivo environment, a parylene film is deposited onto the PDMS film to form a compact and holefree coating layer of 10 μm thick. The cross-sectional scanning electron microscopy (SEM) image of the piezoelectric composite and the elastic skeleton is shown in Figure 1b, which clearly demonstrates the morphologies of the different layers. Figure 1c shows the flexibility of the piezoelectric
development of the triboelectric generators and electrostatic generators in practical applications, respectively. In view of piezoelectric generators, generally the output voltage is easy to reach 5 V for most of integrated circuit chips, but the output current always is difficult to meet some practical requirements, especially the implantable medical devices.37 In addition, the operating voltage and current are diverse for different devices, and the energy harvesting system with constant output has low adaptability for this requirement. Here, we report an integration strategy for directly powering a cardiac pacemaker, which could pace the porcine heart in vivo by harvesting the kinetic energy of heartbeat in vivo with a high-performance, implantable piezoelectric energy generator (iPEG). The generator is designed with an elastic skeleton and two piezoelectric composites, which produces a high and switchable electric output after being implanted into the pericardial sac of an adult Yorkshire swine. The generated energy directly powers a modern and full-function cardiac pacemaker in this big animal, which makes an impressive step toward fabricating a self-powered or batteryless cardiac pacemaker.
RESULTS AND DISCUSSION Device Design, Fabrication, and Working Principle. Figure 1a provides an expanded schematic diagram of the iPEG which is composed of elastic skeleton, piezoelectric composites, and biocompatible encapsulated materials (dimensions shown in Figures S1 and S2). A polyethylene terephthalate (PET) sheet (100 μm thick) is constructed into a B
DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
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Figure 2. Experimental studies of the electric behavior of the iPEG. (a) Diagram of the iPEG being applied a periodic loading force by a simulated heart. The iPEG is connected to an oscilloscope through a paralleled resistor or a capacitor. The series-parallel mode of the two piezoelectric composites can be controlled by Switch 1. In vitro output of (b) the peak-to-peak open-circuit output voltage (Voc) and (c) short-circuit current (Isc) in series and parallel modes under different magnitudes of loading forces with the frequency of 2 Hz. (d) Closedcircuit ouput power as a function of external resistance under a loading force of 0.3 N in two modes. (e) Charging curves of the iPEG in two modes for a 47 μF capacitor under the same loading force. (f) Stability measurement, the Voc in parallel mode is recorded for 5 × 106 cycles under the loading force of 0.3 N with the frequency of 2 Hz. (g) Schematic illustration of the test platform for the water bath set up. (h) Effect of temperature and humidity on the stability of the iPEG, which is recorded continuously in a water bath with varied temperature.
composite. No visible damage in the piezoelectric layer can be detected even when it is bent to a curvature of ∼200 m−1. The working principle of the device is illustrated in Figure 1d,e. In short, the deformation process of the iPEG can be divided into four states, that is, initial, compressing, compressed, and expanding states. Accordingly, an alternating current (AC) is generated during this process, and the output of the iPEG reaches its positive maximum in the compressing state and the negative one in the expanding state. Performance of the Device in Vitro. Figure 2a illustrates the dynamic test platform for evaluating the performance of the iPEG in vitro. A heart model is employed to imitate the action pattern of a heartbeat, which is manufactured by PDMS and driven by the sinusoidal loading force of a linear motor. Switchable output is achieved via transforming the connection ways of two piezoelectric composites: series or parallel mode (Switch 1). Figure 2b,c shows the peak-to-peak values of open-
circuit voltage (Voc) and short-circuit current (Isc) of the two modes under different loading forces. The output performances in series and parallel modes and the deformation of the midpoint of the iPEG at different applied loading forces are shown in Figures S3 and S4. With the loading force of 0.3 N and 2 Hz, the Voc reaches its maximum of 36 V, and the Isc is 22 μA in series mode. In contrast, the Voc and Isc reach 22 V and 40 μA, respectively in parallel mode. The short-current waveforms in different connections are shown in Figure S5. This switchable capability of the output voltage/current provides multiple choices for various requirements of implantable medical devices on operating voltage/current. When the Switch 2 is turned on, the closed-circuit output power as a function of the external resistance is described in different modes, as shown in Figure 2d. The maximal output power of ∼33 μW at the optimal external resistance of 700 kΩ in parallel mode and ∼14 μW at 2 MΩ in series mode can be C
DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
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Figure 3. Biocompatibility tests of the iPEG. (a) Fluorescent images of human pericardial fibroblasts which are cultured on the culture dishes as control (upper) and encapsulation layers of the iPEG (lower), showing cytoskeleton (red) and nuclei (blue). (b) Normalized viability of human pericardial fibroblasts after being cultured for continuous 5 days. (c) Fluorescent images of neonatal rat cardiac myocytes which are cultured on the culture dishes (upper) and encapsulation layers of the iPEG (lower), showing cytoskeleton (red) and nuclei (blue). (d) Beating frequencies of neonatal rat cardiac myocytes after being cultured from days 2 to 5. (There are no beating myocytes on day 1, and therefore the data are not shown, *p > 0.05.)
incubated mouse cardiac myocytes in both groups. The densely distributed nucleus in both groups reveals the excellent adhesive rate and the tendency of growing into clusters of cardiac myocytes. In addition, the cardiac myocytes start beating on day 2, and the beating rates show no significant differences in the two groups during the following days, as shown in Figure 3d, demonstrating the safety to the cell function of the device. The classical cell models such as human embryonic kidney cells (HEK293) and mouse fibroblasts (L929) were commonly used to verify biocompatibility of encapsulation layers in previous studies. These cells could not completely mimic the cells of normal tissues in structure and function.12,13 In our study, two different cell types derived from adjacent tissues of the implantation sites are chosen as cellular models, including neonatal rat cardiac myocytes and human pericardial fibroblasts, which may provide more accurate evaluation method for biocompatibility. Additionally, the pericardial sac is unlike other tissue such as subcutaneous, which seldom has cells and vessels inside. Lack of inflammatory cells leads to a low level of inflammatory reactions including tissue hyperplasia and fibrosis. Therefore, the implanted device in the pericardial sac will be engulfed fibrino exudation, which will form a soft and thin layer, rather than severe fibrosis. Hence, the performance of the device would not decrease. The implanted generator will be encapsulated by a soft and thin fibrinous layer. The fibrinous layer will be a natural barrier that can keep the device from causing any tissue damage and necrosis to the heart. Performance of the Device in Vivo. To identify the optimal location for harvesting the kinetic energy of heart in vivo, the iPEG is implanted in four implantation sites of the pericardial sac, that is, apex (AP), anterior wall (AW), posterior wall (PW), and lateral wall (LW), as illustrated in Figure 4a. Figures 4b and S7 exhibit the output Voc and Isc at
gained. The conversion efficiency can be calculated as 2.75%, which can be expressed by the ratio of the effective output electrical power to the input mechanical power.38 When the Switch 3 is turned on, a capacitor of 47 μF could be charged to 4.1 V during 140 s in parallel mode and 6.7 V during 430 s in series mode by the device under a loading force of 0.3 N with the frequency of 2 Hz (Figure 2e). To investigate the stability, the device is acted with the heart model for 500,000 cycles with a loading force of 0.3 N and 2 Hz. No attenuation is detected for its output Voc (Figure 2f). Cross-sectional SEM images of the piezoelectric unit before and after the fatigue test are shown in Figure S6. Additionally, in order to explore the effects of temperature and humidity, the glass container is filled with water until the generator is completely submerged (Figure 2g). The vibrator produces a peak loading force of 0.3 N with the frequency of 2 Hz. The water, with the temperature monitored by a thermometer, is heated from 12 to 40 °C on a hot plate for 1 h. Then, the heated water is cooled to 21 °C naturally. During the whole process of temperature varying, output voltages of the generator are recorded by the dynamic signal analyzer, as shown in Figure 2h, which demonstrates the temperature has little effect on the output performance of the iPEG at 100% relative humidity when the temperature varied from 12 to 41 °C. Biocompatibility Test of the Device. Figure 3a shows the fluorescent images of human pericardial fibroblasts incubated on the culture dish/device on days 1, 3, and 5. The normal cellular morphology, intact cytoarchitecture, and unbroken nuclei can be observed clearly in both groups. Moreover, there are comparable cell densities between the two groups. The similar cell viabilities of the cultured pericardial fibroblasts in Figure 3b indicate the robust proliferation activity in both groups. Figure 3c shows the fluorescent images of the D
DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
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Figure 4. In vivo evaluation for the performance of the iPEG. (a) Schematic view illustrating different implantation sites: apex (AP), anterior wall (AW), posterior wall (PW), and lateral wall (LW). (b) Voc of the implanted iPEG in series mode (upper) and Isc in parallel mode (lower) from different implantation sites. The output Voc and Isc reaches the peak values of 20 V and 15 μA, respectively, when the iPEG is fixed at AP. (c) Photographs of the iPEG in the pericardial sac fixed at AP, showing that the iPEG expands during cardiac systolic phase (left) and then is compressed by the heart during cardiac diastolic phase (right). (d) Magnified and overlapped views of the ECG and the corresponding Voc (upper) and Isc (lower) waveforms of the iPEG at AP.
the four sites. As a result, the maximal output Voc and Isc are obtained when the device is implanted at AP where the motion amplitude of the heart is largest compared to other sites. Respectively, the maximal output Voc of ∼20 V is yielded in series mode, and the maximal output Isc of ∼15 μA is yielded in parallel mode, which are higher than those of all reported approaches in previous studies.11−13 Figure 4c shows the implanted iPEG at AP in the pericardial sac. The device expands when the heart contracts in systolic phase (left panel), while it is compressed by the heart in diastolic phase (right panel). To investigate the relationship between the output of the implanted generator and heart motion, representative waveforms of electrocardiogram (ECG) and the corresponding Voc (upper) and Isc (lower) are magnified and overlapped in Figure 4d. In detail, the output Voc and Isc reach their peak values (PV and PI) closely following the QRS complex, which represents the early systolic phase of the heart. Then the output returns to zero at the end of T wave, which indicates the termination of systolic phase. After that, the output gains its negative peak values (PV′ and PI′) during the diastolic phase of heart. It should be noted that the output PV and PI are much higher than PV′ and PI′, respectively, which could be explained by the fact that the implanted device expands rapidly along with the active contract of the heart during the systole phase. In contrast, the device is slowly compressed when the heart expands during the passive diastole phase. Direct Powering a Modern and Full-Function Cardiac Pacemaker. Figure 5a shows a modern and full-function cardiac pacemaker (Jade 3, Vitaron Corp., U.S.A.) with the
built-in lithium battery being removed. The photograph and circuit diagram of the designed iPEG-powered pacemaker are displayed in Figure 5b,c. According to previous reports,37 the operating energy consumption per pacing stimulus of modern cardiac pacemaker is 0.12 s) followed the pacing signal, which means that the ventricular myocardium is captured by the pacing stimulus and the ventricle starts its contraction. These characteristics of the ECG definitely demonstrate that the heart is successfully paced by the iPEG-powered modern fullfunction pacemaker, which is a great step forward toward achieving a self-powered cardiac pacemaker and resolving the power issue of implantable medical devices. In the experiment, the pacing rate of the commercial pacemaker is set to 80 bpm, which is quicker than the original heart rate (slowed down by administration of verapamil to 60− 70 bpm) and can successfully pace the heart. This is similar to a bradycardia model to reasonably demonstrate the capability of the device, although the heart rates were not low enough to reach the diagnostic standard of bradycardia ( 0.05, then the difference is caused by sampling error. If p ≤ 0.05, then it can be considered that the difference is not caused by the sampling error. The test statistic value t = |X̅ − μ| , where
exempt the patients from surgical replacement, or at least, less frequently.
METHODS Device Fabrication. Fabrication process of the implantable piezoelectric generator mainly consists of three parts: fabricating piezoelectric composite, bonding the piezoelectric composite with elastic skeleton, and encapsulating of the overall generator. The piezoelectric composite includes piezoelectric layer, beryllium-bronze foil, and Au electrodes. First, a bulk (72%)Pb(Mg1/3Nb2/3)O3(28%)PbTiO3 (PMN-PT) (TRSX2A, TRS Ceramics, USA) sheet (300 μm) was thinned to 50 μm by chemical mechanical polish process. Then, Cr/Au (50 nm/200 nm) layer was sputtered at each side of the PMN-PT as electrodes through magnetic sputtering (TRP450, SKY Technology Development Co., Ltd., China). The oriented PMN-PT sheet with electrodes was thermally bonded on the polished beryllium-bronze foil (10 × 25 × 0.05 mm3) by silver epoxy (DAD87, Shanghai Research Institute of Synthetic resins, China) adhesive under 100 °C for 24 h as the piezoelectric composite. Finally, the PMN-PT was polarized by applying an electric field of 3.5 kV/cm at 20 °C for 10 min through a high-voltage pole source (CS9911BI, Nanjing Allwin Instrument Science And Technology Co., Ltd., China). Two piezoelectric composites were bonded to the top and bottom surfaces of polyethylene terephthalate (PET) thin film (100 μm thick) with a shape of compressed number “8” in a side view by epoxy adhesive. Four copper wires were prepackaged in medical grade silicone as electrical leads and bonded to Au electrodes utilizing silver epoxy conductive adhesive. A core/shell packaging technology was adopted for achieving the structural sealability of the device. A polydimethylsiloxane (PDMS) (SYLGARD 184, DOW Corning Inc., USA) layer (40 μm thick), which was cured at 80 °C for 1 h, was coated on the assembled generator to prevent it from contacting body fluid and surrounding tissue. To further enhance its leak-proof characteristic, a parylene layer (Sigma-Aldrich, USA) 10 μm thick was deposited through the SCS Labcoter 2 parylene deposition system (PDS 2010, SCS, USA) as the second protecting layer at room temperature. In Vitro Performance Test. The as-fabricated device was fixed on the simulated pericardium and collided by imitated heart periodically with different magnitude and frequency of the loading forces. The series-parallel mode of the two composite layers can be controlled by Switch 1 through changing the connection ways of the output leads, and the open circuit voltage (Voc) and short-circuit current (Isc) were recorded by using an oscilloscope. When the Switch 2 was turned on, the maximum output power density can be calculated by changing the value of adjustable resistor in the series circuit and recording its corresponding closed-circuit voltage. Then, a 47 μF capacitor was utilized to test the charging characteristic of the iPEG when the Switch 3 was turned on. At last, the fatigue test was carried out by continuous measurement of the Voc in 500,000 cycles. In order to explore the effects of temperature and humidity, the piezoelectric generator was put into a water bath with monitored temperature. The PDMS elastomer was utilized to apply mechanical loading to the piezoelectric generator to imitate the cardiac apex and heartbeat. The elastomer was connected to a dynamic force sensor and was fixed on a vibrator, which can produce a reciprocating movement. The dynamic force sensor was used to monitor the contact force between the PDMS elastomer and the piezoelectric generator. The vibrator and the glass container were fixed on the platform, which was put on the hot plate for heating. A function generator and a power amplifier are used to drive the vibrator. The output of the iPEG and the signals of the sensor were recorded by a dynamic signal analyzer in real time. Cell Isolation and Culture. Primary pericardial fibroblasts and neonatal rat cardiac myocytes were separated from donor pericardial tissue during heart transplant and the heart tissue of 1 day-old neonatal rats, respectively. Then, two kinds of cells were planked with proper cell density and cultured for 5 days in Dulbecco modified Eagle medium (DMEM) supplemented with 10% fetal bovine serum (FBS) at 37 °C under a 5% carbon dioxide atmosphere. Neonatal rat
S/ n
X̅ is the mean of the population, μ is the known mean of the population, S is the standard deviation, and n is the number of the samples. Comparing the t-value with t0.05, if t < t0.05, then p > 0.05; otherwise p ≤ 0.05. In Vivo Performance Test. Adult Yorkshire pigs (male, 50 kg) were managed with conformation to the “Shanghai Administration Rule of Laboratory Animal”, and the Institutional Animal Care and Use Committee (IACUC) approved protocol of the Animal Care Center at the Second Military Medical University. First, the swine was sedated with an injection of midazolam (0.25 mg/kg, intramuscularly) and ketamine (8 mg/kg, intramuscularly), and then the animal was restrained on the animal operating table and endotracheally intubated. Mechanical ventilation, right femoral arterial catheterization, and electrocardiograph (ECG) monitoring were performed for proper perioperative management and data recording. Anesthesia was maintained with propofol (4 mg/kg/h, intravenously). Median sternotomy was employed to provide wide and straight access to heart. In vivo performance test involves implantation, affixion, and testing of the device. The pericardium was cut open by using surgical scissors, to facilitate surgical operation and implantation procedures of the device at different sites, and the pericardium was surgically suspended on the chest wall by sutures. Afterward, the iPEG was fixed on the pericardium at different sites by stitching through four holes at both ends of the device, and the copper wires were connected to the data acquisition system (DAQ) for signal collection.
ASSOCIATED CONTENT S Supporting Information *
The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsnano.8b08567. Additional details on the materials parameters, the comparison of the presented generator to the recent reports, the sizes, in vitro performance, the cross-section SEM images of the piezoelectric unit, the output performance of the presented generator at different implantation sites, an analysis of the impulse signal, an evaluation of the output stability, and the electrocardiogram signals at the different pacing periods (PDF)
AUTHOR INFORMATION Corresponding Authors
*E-mail:
[email protected]. *E-mail:
[email protected]. ORCID
Zhiran Yi: 0000-0002-8679-8302 Jingquan Liu: 0000-0003-4140-1516 Bin Yang: 0000-0001-7948-3823 G
DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
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ACS Nano Author Contributions
(11) Dagdeviren, C.; Yang, B. D.; Su, Y.; Tran, P. L.; Joe, P.; Anderson, E.; Xia, J.; Doraiswamy, V.; Dehdashti, B.; Feng, X.; Lu, B.; Poston, R.; Khalpey, Z.; Ghaffari, R.; Huang, Y.; Slepian, M. J.; Rogers, J. A. Conformal Piezoelectric Energy Harvesting and Storage from Motions of the Heart, Lung, And Diaphragm. Proc. Natl. Acad. Sci. U. S. A. 2014, 111, 1927−1932. (12) Zheng, Q.; Zhang, H.; Shi, B.; Xue, X.; Liu, Z.; Jin, Y.; Ma, Y.; Zou, Y.; Wang, X.; An, Z.; Tang, W.; Zhang, W.; Yang, F.; Liu, Y.; Lang, X.; Xu, Z.; Li, Z.; Wang, Z. L. In Vivo Self-Powered Wireless Cardiac Monitoring via Implantable Triboelectric Nanogenerator. ACS Nano 2016, 10, 6510−6518. (13) Kim, D. H.; Shin, H. J.; Lee, H.; Jeong, C. K.; Park, H.; Hwang, G.-T.; Lee, H.-Y.; Joe, D. J.; Han, J. H.; Lee, S. H.; Kim, J.; Joung, B.; Lee, K. J. In Vivo Self-Powered Wireless Transmission Using Biocompatible Flexible Energy Harvesters. Adv. Funct. Mater. 2017, 27, 1700341. (14) Madhavan, M.; Mulpuru, S. K.; McLeod, C. J.; Cha, Y. M.; Friedman, P. A. Advances and Future Directions in Cardiac Pacemakers: Part 2 of A 2-Part Series. J. Am. Coll. Cardiol. 2017, 69, 211−235. (15) Pfenniger, A.; Jonsson, M.; Zurbuchen, A.; Koch, V. M.; Vogel, R. Energy Harvesting from the Cardiovascular System, or How to Get a Little Help from Yourself. Ann. Biomed. Eng. 2013, 41, 2248−2263. (16) Hwang, G.-T.; Park, H.; Lee, J.-H.; Oh, S.; Park, K.-I.; Byun, M.; Park, H.; Ahn, G.; Jeong, C. K.; No, K.; Kwon, H.; Lee, S.-G.; Joung, B.; Lee, K. J. Self-Powered Cardiac Pacemaker Enabled by Flexible Single Crystalline PMN-PT Piezoelectric Energy Harvester. Adv. Mater. 2014, 26, 4880−4887. (17) Ma, Y.; Zheng, Q.; Liu, Y.; Shi, B.; Xue, X.; Ji, W.; Liu, Z.; Jin, Y.; Zou, Y.; An, Z.; Zhang, W.; Wang, X.; Jiang, W.; Xu, Z.; Wang, Z. L.; Li, Z.; Zhang, H. Self-Powered, One-Stop, and Multifunctional Implantable Triboelectric Active Sensor for Real-Time Biomedical Monitoring. Nano Lett. 2016, 16, 6042−6051. (18) Zhang, H.; Zhang, X.-S.; Cheng, X.; Liu, Y.; Han, M.; Xue, X.; Wang, S.; Yang, F.; Smitha, A. S.; Zhang, H.; Xu, Z. A Flexible and Implantable Piezoelectric Generator Harvesting Energy from The Pulsation of Ascending Aorta: In Vitro and In Vivo Studies. Nano Energy 2015, 12, 296−304. (19) Cheng, X.; Xue, X.; Ma, Y.; Han, M.; Zhang, W.; Xu, Z.; Zhang, H.; Zhang, H. Implantable and Self-Powered Blood Pressure Monitoring Based On a Piezoelectric Thin film: Simulated, In Vitro and In Vivo Studies. Nano Energy 2016, 22, 453−460. (20) Zurbuchen, A.; Pfenniger, A.; Stahel, A.; Stoeck, C. T.; Vandenberghe, S.; Koch, V. M.; Vogel, R. Energy Harvesting from The Beating Heart by A Mass Imbalance Oscillation Generator. Ann. Biomed. Eng. 2013, 41, 131−141. (21) Chen, X.; Xu, S.; Yao, N.; Shi, Y. 1.6 V Nanogenerator for Mechanical Energy Harvesting Using PZT Nanofibers. Nano Lett. 2010, 10, 2133−2137. (22) Li, Z.; Zhu, G.; Yang, R.; Wang, A. C.; Wang, Z. L. MuscleDriven In Vivo Nanogenerator. Adv. Mater. 2010, 22, 2534−2537. (23) Lu, B.; Chen, Y.; Ou, D.; Chen, H.; Diao, L.; Zhang, W.; Zheng, J.; Ma, W.; Sun, L.; Feng, X. Ultra-flexible Piezoelectric Devices Integrated with Heart to Harvest the Biomechanical Energy. Sci. Rep. 2015, 5, 16065. (24) Hwang, G.-T.; Byun, M.; Jeong, C. K.; Lee, K. J. Flexible Piezoelectric Thin-Film Energy Harvesters and Nanosensors for Biomedical Applications. Adv. Healthcare Mater. 2015, 4, 646−658. (25) Zheng, Q.; Shi, B.; Fan, F.; Wang, X.; Yan, L.; Yuan, W.; Wang, S.; Liu, H.; Li, Z.; Wang, Z. L. In Vivo Powering of Pacemaker by Breathing-Driven Implanted Triboelectric Nanogenerator. Adv. Mater. 2014, 26, 5851−5856. (26) Wang, J.; Wu, C.; Dai, Y.; Zhao, Z.; Wang, A.; Zhang, T.; Wang, Z. L. Achieving Ultrahigh Triboelectric Charge Density for Efficient Energy Harvesting. Nat. Commun. 2017, 8, 88. (27) Basset, P.; Galayko, D.; Paracha, A. M.; Marty, F.; Dudka, A.; Bourouina, T. A Batch-Fabricated and Electret-Free Silicon Electrostatic Vibration Energy Harvester. J. Micromech. Microeng. 2009, 19, 115025.
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Z.Y., N.L., and Y.M. contributed equally to this work. Z.Y., H.Z., and B.Y. conceived the idea and designed the experiments. Z.Y., Y.T., X.D., and B.Y. designed, fabricated, and tested the iPEG in vitro. Z.Y. performed theory analysis. Z.Y., N.L., F.X., Y.T., Y.M., Y.H., Y.L., X.S., Y.L., and H.Z. performed the animal experiments, and F.X., N.L., Y.M., and H.Z. performed the biocompatibility study. Z.Y., F.X., Y.T., N.L., Y.M., Z.X., J.L., H.Z., and B.Y. analyzed the data. N.L., Z.Y., H.Z., and B.Y. wrote the paper. All of the authors commented on it.
Notes
The authors declare no competing financial interest.
ACKNOWLEDGMENTS This work is supported in part by the 863 Program (2015AA043503) and National Natural Science Foundation of China (grant nos. 61674104 and 81671783), and Natural Science Foundation of Shanghai (17ZR1446200 and 17441904500). REFERENCES (1) Kolodziejczak, M.; Andreotti, F.; Kowalewski, M.; Buffon, A.; Ciccone, M. M.; Parati, G.; Scicchitano, P.; Uminska, J. M.; De Servi, S.; Bliden, K. P.; Kubica, J.; Bortone, A.; Crea, F.; Gurbel, P.; Navarese, E. P. Implantable Cardioverter-Defibrillators for Primary Prevention in Patients with Ischemic or Nonischemic Cardiomyopathy a Systematic Review and Meta-analysis. Ann. Intern. Med. 2017, 167, 103−111. (2) Abraham, W. T.; Stough, W. G.; Pina, I. L.; Linde, C.; Borer, J. S.; De Ferrari, G. M.; Mehran, R.; Stein, K. M.; Vincent, A.; Yadav, J. S.; Anker, S. D.; Zannad, F. Trials of Implantable Monitoring Devices in Heart Failure: Which Design Is Optimal? Nat. Rev. Cardiol. 2014, 11, 576−585. (3) Ritzema, J.; Melton, I. C.; Richards, A. M.; Crozier, I. G.; Frampton, C.; Doughty, R. N.; Whiting, J.; Kar, S.; Eigler, N.; Krum, H.; Abraham, W. T.; Troughton, R. W. Direct Left Atrial Pressure Monitoring in Ambulatory Heart Failure Patients - Initial Experience with A New Permanent Implantable Device. Circulation 2007, 116, 2952−2959. (4) Giada, F.; Bertaglia, E.; Reimers, B.; Noventa, D.; Raviele, A. Current and Emerging Indications for Implantable Cardiac Monitors. PACE 2012, 35, 1169−1178. (5) Dagdeviren, C.; Li, Z.; Wang, Z. L. Energy Harvesting from the Animal/Human Body for Self-Powered Electronics. Annu. Rev. Biomed. Eng. 2017, 19, 85−108. (6) Halamkova, L.; Halamek, J.; Bocharova, V.; Szczupak, A.; Alfonta, L.; Katz, E. Implanted Biofuel Cell Operating in A Living Snail. J. Am. Chem. Soc. 2012, 134, 5040−5043. (7) Dagdeviren, C.; Shi, Y.; Joe, P.; Ghaffari, R.; Balooch, G.; Usgaonkar, K.; Gur, O.; Tran, P. L.; Crosby, J. R.; Meyer, M.; Su, Y.; Webb, R. C.; Tedesco, A. S.; Slepian, M. J.; Huang, Y.; Rogers, J. A. Conformal Piezoelectric Systems for Clinical and Experimental Characterization of Soft Tissue Biomechanics. Nat. Mater. 2015, 14, 728−736. (8) Shi, Q.; Wang, T.; Lee, C. MEMS Based Broadband Piezoelectric Ultrasonic Energy Harvester (PUEH) for Enabling Self-Powered Implantable Biomedical Devices. Sci. Rep. 2016, 6, 24946. (9) Mallela, V. S.; Ilankumaran, V.; Rao, N. S. Trends in Cardiac Pacemaker Batteries. Indian Pacing Electrophysiol. J. 2004, 4, 201−212. (10) Horlbeck, F. W.; Mellert, F.; Kreuz, J.; Nickenig, G.; Schwab, J. O. Real-World Data on the Lifespan of Implantable CardioverterDefibrillators Depending on Manufacturers and the Amount of Ventricular Pacing. J. Cardiovasc. Electr. 2012, 23, 1336−1342. H
DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
Article
ACS Nano (28) Yang, Z.; Zhou, S.; Zu, J.; Inman, D. High-Performance Piezoelectric Energy Harvesters and Their Applications. Joule 2018, 2, 642−697. (29) Deterre, M.; Lefeuvre, E.; Zhu, Y.; Woytasik, M.; Boutaud, B.; Dal Molin, R. Micro Blood Pressure Energy Harvester for Intracardiac Pacemaker. J. Microelectromech. Syst. 2014, 23, 651−660. (30) Jang, J.; Lee, J.; Woo, S.; Sly, D. J.; Campbell, L. J.; Cho, J.-H.; O’Leary, S. J.; Park, M.-H.; Han, S.; Choi, J.-W.; Jang, J. H.; Choi, H. A Microelectromechanical System Artificial Basilar Membrane Based On a Piezoelectric Cantilever Array and Its Characterization Using an Animal Model. Sci. Rep. 2015, 5, 12447. (31) Karami, M. A.; Inman, D. J. Powering Pacemakers from Heartbeat Vibrations Using Linear and Nonlinear Energy Harvesters. Appl. Phys. Lett. 2012, 100, 042901. (32) Dagdeviren, C.; Joe, P.; Tuzman, O. L.; Park, K.-I.; Lee, K. J.; Shi, Y.; Huang, Y.; Rogers, J. A. Recent Progress in Flexible and Stretchable Piezoelectric Devices for Mechanical Energy Harvesting, Sensing and Actuation. Extreme Mechanics Letters 2016, 9, 269−281. (33) Zurbuchen, A.; Haeberlin, A.; Pfenniger, A.; Bereuter, L.; Schaerer, J.; Jutzi, F.; Huber, C.; Fuhrer, J.; Vogel, R. Towards Batteryless Cardiac Implantable Electronic Devices-The Swiss Way. IEEE Trans. Biomed. Circuits Syst. 2017, 11, 78−86. (34) Zurbuchen, A.; Haeberlin, A.; Bereuter, L.; Wagner, J.; Pfenniger, A.; Omari, S.; Schaerer, J.; Jutzi, F.; Huber, C.; Fuhrer, J.; Vogel, R. The Swiss Approach for A Heartbeat-Driven Lead- And Batteryless Pacemaker. Heart Rhythm 2017, 14, 294−299. (35) Dagdeviren, C. The Future of Bionic Dynamos. Science 2016, 354, 1109. (36) Yu, Y.; Sun, H.; Orbay, H.; Chen, F.; England, C. G.; Cai, W.; Wang, X. Biocompatibility and In Vivo Operation Of Implantable Mesoporous PVDF-Based Nanogenerators. Nano Energy 2016, 27, 275−281. (37) Ohm, O. J.; Danilovic, D. Improvements in Pacemaker Energy Consumption and Functional Capability: Four Decades of Progress. PACE 1997, 20, 2−9. (38) Dagdeviren, C.; Hwang, S.-W.; Su, Y.; Kim, S.; Cheng, H.; Gur, O.; Haney, R.; Omenetto, F. G.; Huang, Y.; Rogers, J. A. Transient, Biocompatible Electronics and Energy Harvesters Based on ZnO. Small 2013, 9, 3398−3404. (39) Horikawa-Tanami, T.; Hirao, K.; Furukawa, T.; Isobe, M. Mechanism of The Conversion of a Pulmonary Vein Tachycardia to Atrial Fibrillation in Normal Canine Hearts: Role of Autonomic Nerve Stimulation. J. Cardiovasc. Electr. 2007, 18, 534−541. (40) Li, Y.; Lu, Y.-M.; Zhou, X.-H.; Zhang, L.; Li, Y.-D.; Zhang, J.H.; Xing, Q.; Tang, B.-P. Increase of Autonomic Nerve Factors in Epicardial Ganglionated Plexi During Rapid Atrial Pacing Induced Acute Atrial Fibrillation. Med. Sci. Monit. 2017, 23, 3657−3665. (41) Bikou, O.; Miyashita, S.; Ishikawa, K. Pig Model of Increased Cardiac Afterload Induced by Ascending Aortic Banding. Methods Mol. Biol. 2018, 1816, 337−342. (42) Watanabe, S.; Bikou, O.; Hajjar, R. J.; Ishikawa, K. Swine Model of Mitral Regurgitation Induced Heart Failure. Methods Mol. Biol. 2018, 1816, 327−335.
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DOI: 10.1021/acsnano.8b08567 ACS Nano XXXX, XXX, XXX−XXX
