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In vivo cardiac power generation enabled by an integrated helicalpiezoelectric pacemaker leadLin Donga, Andrew B. Clossona, Meagan Oglesbyb, Danny Escobedob, Xiaomin Hana, Yuan Niea,Shicheng Huanga, Marc D. Feldmanb,∗, Zi Chena,∗∗, John X.J. Zhanga,∗∗∗

a Thayer School of Engineering, Dartmouth College, Hanover, NH, 03755, USAbDivision of Cardiology, Department of Medicine, The University of Texas Health Science Center at San Antonio, San Antonio, TX, 78229, USA

A R T I C L E I N F O

Keywords:Cardiac energy harvestingHelicalBiomedical devicesPiezoelectricMesoporous thin film

A B S T R A C T

Long-term energy supply for electronic systems is challenging for implantable biomedical devices, like cardiacpacemakers. Energy harvesting can significantly extend the lifetime of these devices, however, no clinical trans-lational technologies can efficiently convert the mechanical energy of the heart into electrical power without athoracotomy and interfering with the cardiovascular functions. Almost all reported implantable cardiac energyharvesting designs sutured devices directly onto the epicardium or pericardium with potential risks to the patients.Here, we report a cardiac energy harvesting strategy, which is integrated into part of the existing pacemaker leadand otherwise with no direct contact of heart, by utilizing porous piezoelectric thin films in a bioinspired self-wrapping helical configuration for flexible integration with existing implantable medical devices. We demonstratethat this compact design can be seamlessly coupled with current leads without introducing additional implantationsurgeries. In vivo studies under various conditions (anchoring, pacing, and calcium chloride infusion) are presentedthat demonstrate clinical translation in a porcine model. Both theoretical studies and in vitro experiments are alsopresented to validate the energy harvesting process. The scalability of the design is discussed, and the reportedresults demonstrate a 10×10 array of helical EH devices wrapping all through the lead (a mixed pattern of inseries and parallel connections) would extend the lifetime of the pacemaker battery by 1.5 years. This innovativecardiac energy harvesting strategy represents a significant step forward for clinical translation without a thor-acotomy for patients, suggesting a paradigm for biomedical energy harvesting in vivo.

1. Introduction

Implantable medical devices (IMDs) have become increasinglyprevalent as treatment and diagnostic tools [1]; they include cardiacpacemakers, neurostimulators, left ventricular-assist devices, drug de-livery systems, and bone growth stimulators. In particular, cardiacpacemakers and implantable cardioverter defibrillators (ICDs) are ef-fective means for treatments of ventricular dysrhythmias and con-gestive heart failure [2,3]. However, pacemakers and ICDs are typicallypowered by lithium-ion batteries, leading to limited energy density andlarge device sizes [4]. A typical cardiac pacemaker lasts 5–8 years be-fore its battery needs replacement [5], which requires surgery and canresult in infection and increased healthcare costs. Although technolo-gies are currently being developed that allow for wireless recharging,auto-recharging is a preferable approach for the patient [6–8]. The

present work is an effort to develop a sustainable energy solution forIMDs, such as cardiac pacemakers and ICDs.

With the advances in low power consumption for IMDs [9–13], thereis a growing interest to allow these devices to become fully autonomousand self-sustainable with their own renewable power supply [14]. Onepotential solution is energy harvesting (EH), where the ambient energy isconverted to useful electrical energy at levels sufficient to power IMDs.Harvesting energy in the body has become increasingly attractive as ameans to convert various forms of available energy into long-termelectrical power, e.g., human motion (legs [15] and breathing [16,17])and periodic expansions and contractions of organs such as the heart[18–23] and lungs [24]. For short-term in vivo biomechanical energyconversion, biodegradable EH devices have been designed to be de-graded and resorbed after completing work cycles [25]. Although anabundant amount of research has been published on various EH devices,

https://doi.org/10.1016/j.nanoen.2019.104085Received 3 August 2019; Received in revised form 21 August 2019; Accepted 30 August 2019

∗ Corresponding author.∗∗ Corresponding author.∗∗∗ Corresponding author.E-mail addresses: feldmanm@uthscsa.edu (M.D. Feldman), zi.chen@dartmouth.edu (Z. Chen), john.zhang@dartmouth.edu (J.X.J. Zhang).

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and comprehensive reviews of in vivo EH strategies have been describedin the literature [26,27], few have been tested in large animal models.The motion of the heart is particularly compelling. For example, im-plantable triboelectric nanogenerators (iTENG) using ZnO have beenemployed to investigate EH from rat hearts [16]. Kim et al. gathered theenergy from a pig heart to power wireless data transmission [20]. Inanother in vivo experiment, a single crystalline (1-x) Pb (Mg 1/3 Nb 2/3)O 3−xPbTiO 3 (PMN-PT) was implanted into an intact rat heart to showfunctional electrical stimulation of the heart [28]. A lead zirconate tita-nate (PZT) EH system was also developed to harvest energy from theheart and lungs in a bovine model [22]. The mechanical energy createdby the heart has also been gathered from blood pulsation due to thecyclic expansion and recoil of the ascending aorta [29]. However, withmany of these devices, placement requires suturing directly onto theepicardium, requiring open-chest surgery (Supporting Information TableS1) [16,20,22,23,28,30–32]. While no detectable change in cardiaccontraction or epicardial motion was found [22], the procedure of su-turing such EH devices directly onto the heart inevitably introducespotential risks to the patient.

The key challenge is that most EH designs have little to no clinicaltranslation—an EH device that requires a thoracotomy is not an ap-proach that will be accepted clinically. Here, we report an EH strategyhas no contact with the heart but can efficiently covert heart energy intoan electrical energy without interfering with the cardiovascular function.The EH design combines advanced porous thin film energy conversionmaterials with a bioinspired self-wrapping helical structure design toharness and convert the mechanical energy from the lead of a pacemakeror ICD into an electrical power. In particular, although cantilever[33–37] and buckled beam [20,22,38] geometry-based EH structureshave been widely explored in the literature, the current study introducesa unique helical configuration, which shows significant advantages forcapturing the complex motions from a cardiac pacemaker lead and alsointegrates with current leads for compact EH applications. This cardiacEH strategy is compatible with the existing pacemaker lead implantationapproach, and thus does not require additional surgeries. Moreover, thesuccessful in vivo demonstration in large animal models suggests that theproposed strategy could provide a practical and broader power solutionto recharge cardiac pacemakers and other IMDs.

2. Results and discussion

2.1. Design of the helical energy harvester

Fig. 1A shows the schematic of a helical EH device using flexible porouspiezoelectric thin films. This device harvests the motion of a pacemaker leadwithin the heart and requires no additional surgery for implantation. Thebottom five cm of the pacemaker lead is coupled with a sealed, self-wrap-ping helical polyvinylidene fluoride-trifluoroethylene (PVDF-TrFE) EH de-vice. A schematic of the sandwich structure of polydimethylsiloxane(PDMS) /PVDF-TrFE /PDMS is shown in Fig. 1B. The PVDF-TrFE thin filmcaptures mechanical energy due to bending and twisting of the lead, whichis implanted into the right ventricular endocardium. Taking inspiration fromnature's tendrils (Fig. 1C), whose mechanism of coiling is either associatedwith helical arrays of cortical microtubules or differential contraction withina bilayer [39], we developed a method for fabricating a sandwich structureof a self-wrapping helical piezoelectric EH device (Fig. 1D and SupportingInformation Fig. S1) by using the strain effect between layers of PDMS.Finite element analysis (FEA) using ABAQUS was performed to simulate theself-assembly of the helical structure using the strain engineering method(Supporting Information, Video S1). The principle curvatures of a generichelical ribbon were obtained by employing a theoretical model based onlinear elasticity theory with differential geometry and stationary principles(see Supporting Information for details). The self-wrapping helical config-uration aids in the device's attachment to the flexible pacemaker leads, andalso allows it to take advantage of the complex heart motion. Specifically,the helical structure transfers the bending and twisting motion of the heart

during the cardiac cycle into strain within the piezoelectric material, whichin turn produces an electric charge. In addition, this helical EH structureshows significant advantages for the integration with current pacemakerleads for compact design. Fig. 1E illustrates the deformations of the helicalribbon with different pitches, and the in-plane strain distributions wereobtained by FEA (Supporting Information).

Supplementary data related to this article can be found at https://doi.org/10.1016/j.nanoen.2019.104085.

A scanning electron microscopy (SEM) image of the multilayers EHdevice's cross-section (Fig. 1F) shows the 20 μm thick PDMS sand-wiched between two layers of 45 μm thick porous PVDF-TrFE films. Anaverage pore diameter of 6 μm was measured within the layer of porousfilm. It was also observed in a single layer PVDF-TrFE thin film EHdevice that the pores in the PVDF-TrFE thin film were partially filledwith the spin-cast layer of PDMS, which enhanced piezoelectric outputdue to a higher incompressibility of the infill materials. In addition, thePVDF-TrFE material was characterized using an X-ray diffractometer(Rigaku Rotating Anode XRD); the diffraction pattern had a char-acteristic peak at 19.3° that was associated with the highly piezoelectriccrystalline β-phase (Fig. 1G).

2.2. In vitro experimental characterization

To begin to assess the right ventricular motion [40–43], a test setupwas designed to closely approximate in vivo conditions (Supporting In-formation, Fig. S2 and Video S2). This shaker test platform was built tosimulate the motion of a pacemaker lead coupled with the EH device invivo, particularly the bending motion of the lead due to the heart beating.Based on an anatomical distance from the insertion point at the sub-clavian vein to the tricuspid valve [40,42], a length of 75mm was usedfor the right ventricle (RV), simulating the distance from tricuspid valveto the apex of the heart during diastole [41–43], while a 50mm lengthwas used to simulate the RV during systole. Note that a limitation of thetest setup is that while the platform was designed according to anato-mically-based parameters, the slack of the lead based upon surgical im-plantation techniques may also affect the performance of the EH devicein practice [44]. In vitro measurements with two major motions (vibra-tion and bending of the pacemaker leads) demonstrated that the bendingof the helical piezoelectric thin film contributes to the overall electricaloutput of the EH device. Fig. 2A shows that the voltage output generatedfrom the bending motion of the EH device is significantly larger than thecontribution from vibration (peak-to-peak open-circuit voltage of 0.6 Vversus 0.05 V, respectively). These results were also confirmed by in vivoevaluation of the helical EH device before and after anchoring the pa-cemaker leads to the RV septum in two different in vivo porcine studies.

Supplementary data related to this article can be found at https://doi.org/10.1016/j.nanoen.2019.104085.

Practical in vivo implementation of the EH device on a pacemakerlead requires a way to encapsulate the device, so that it may be isolatedfrom bodily fluids. The experimental results in Fig. 2B show that theencapsulated EH device was successfully isolated from saline, and onlya slight decrease of voltage output in saline was observed. In addition,the cycling stability of the EH device in saline shows that the corre-sponding voltage output of the EH device is stable with no significantdegradation over 104 cycles (Fig. 2C), suggesting that the EH device hasadequate fatigue properties. For long-term clinical translation, the EHwill have to be incorporated inside the pacemaker lead.

Characterization of the charging ability of the single-layer porous filmbased EH device was performed by using a Schottky bridge rectifier(MB12S, Micro Commercial). The alternating current generated from theEH device converted to a direct current (DC) signal, and a peak voltageoutput of 0.28V after rectifying was obtained. A sawtooth voltage curve andcharging curve were measured by adding a 33 μF capacitor, since the ca-pacitor only has time to discharge briefly before the next DC voltage re-charges it back to the peak value. The inset of Fig. 2D shows a schematic ofan equivalent circuit of the EH for measuring the voltage across the

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capacitor. It takes 36 s for the helical EH device to reach 0.28V from theenergy of the pacemaker lead, and the practical power on the capacitor canbe estimated as 0.036 μW ( =P CU t/22 , where C U, and t are the capa-citance, saturation voltage, and the time to reach such saturation point).Given a pacemaker battery capacity of 1 Ah and a longevity of 10 years[45,46], the average energy consumption of a pacemaker is determined tobe 28.75 μW. This indicates that 10 helical EH devices connected in serieswould yield an electric potential of 2.8V, and a 10×10 array of helical EHdevices wrapping all through the lead (a mixed pattern of in series andparallel connections) would extend the lifetime of the pacemaker battery by1.5 years. Note that the longevities and battery capacities of pacemakersvary, according to information provided by manufacturers based on their

own test protocols, and those numbers might also be different from thedevice longevities in real life.

The scalability of the helical EH design was also demonstrated. Twohelical harvesting units (a double-helix device) were connected in seriesto increase the voltage. Due to the helical structure of the EH design, asecond harvesting unit could be added between the intervals of the firstunit for compact EH. The results in Fig. 2E confirmed that the strategyof using two harvesting mechanisms contributed to a higher (doubledthe peak-to-peak voltage) output of 1.25 V. Therefore, from a designperspective, a mixed pattern (both in series and parallel connections)could be further pursued to contribute a significant higher electricaloutput by using multiple EH units. Additional in vitro characterization

Fig. 1. Cardiac energy harvesting strategy, helical configuration, and material characterization. (A) Schematic of a helical piezoelectric energy harvester usingflexible porous piezoelectric thin films. The EH device is self-wrapped around a pacemaker lead and ultimately charges the battery of the pacemaker. (B) Schematic ofsandwiched PDMS/PVDF-TrFE/PDMS thin film with gold electrodes (single layer porous PVDF-TrFE EH device). (C) An example of a tendril in nature. (D) Finiteelement analysis for fabricating a sandwich structure of self-wrapping helical piezoelectric EH device by using the strain effect within the layers. (E) The in-plainstrain distribution of the helical ribbon with different pitches by using FEA. (F) SEM image of the cross section of sandwiched two layers of porous PVDF-TrFE thinfilm with the PDMS layer in between. (G) XRD spectra showing the peak associated with the crystalline β-phase at 2θ=19.3°.

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illustrated that a double-layer PVDF-TrFE porous thin film based EHdevice generated about 1.8 times higher voltage output compared withthat from the single-layer PVDF-TrFE porous thin film (Fig. 2F).

2.3. In vivo experimental characterization

In vivo evaluation of the EH device on an existing pacemaker lead(Fig. 3A) requires an effective method to encapsulate the EH device to beisolated from bodily fluids. The excellent biocompatibility of PVDF hasbeen verified through previous in vivo studies [47]. In this study, theporous PVDF-TrFE thin film was encapsulated within PDMS layers, andwas co-cultured with L929s (mouse fibroblast cell line) to study its cy-totoxicity and biocompatibility. The viabilities of L929 cells on the en-capsulated EH devices showed no significant differences compared to thecontrol group after 3 days of cell culture, demonstrating good bio-compatibility of the encapsulated EH device (Supporting InformationFig. S3). Using the standard implantation techniques for pacemakerleads, the sealed helical EH device was inserted through a vascularsheath and advanced into the right ventricle of a porcine heart underfluoroscopic guidance (Fig. 3B and Supporting Information Video S3). Afluoroscopy image in Fig. 3C shows a porcine heart with an implantedpacemaker lead coupled with an EH device. Evaluation of various an-choring conditions of the pacemaker leads were performed to comparethe generated electrical output from the EH device. In vivo testing of thesame helical EH device was measured before and after screwing in the tipof the pacemaker leads in two different porcine right ventricles (Fig. 3Dand Supporting Information Fig. S4). In both animals, the anchored leadsyield much higher electrical output generated by the single helix EHdevice, where the voltage output (peak-to-peak) is 0.5 V in Animal A(Fig. 3D, top) and 0.65 V in Animal B (Fig. 3D, bottom). This behaviorwas also confirmed in the in vitro results in Fig. 2A. The electrical outputgenerated from the complex motion (bending and twisting) of the pa-cemaker lead (anchored in vivo) was, in general, significantly larger than

the contribution from the vibration of the leads (unanchored in vivo).Supplementary data related to this article can be found at https://

doi.org/10.1016/j.nanoen.2019.104085.The right ventricle is much more complex for cardiac EH applications

where heart rate can vary with time. Therefore, in vivo evaluation of theeffect of heart rate on the electrical output of the EH device was per-formed. Electrical output data were collected at a baseline heart rate of94 beats per minute (bpm) (Fig. 4A). An external pulse generator wasconnected to the proximal end of the lead to pace the right ventricle(Fig. 4B). The voltage generated by the implanted EH device was col-lected at 94 bpm (baseline rate), 120 bpm, 140 bpm, and 160 bpm. Theresults in Fig. 4A show that the voltages decrease with increasing heartrate. This behavior can be explained by the fact that the bending andtwisting motion of the pacemaker lead decreases at higher heart rates,since the lead has shorter distance to travel due to the time to recoverbefore the next cycle of deformation, which in turn decreases the elec-trical output of the EH device. In addition, arterial pulse (AP) pressurewas measured by a micro-manometer pressure sensor placed in thedescending thoracic aorta as an indirect measure of cardiac contractility,and electrocardiogram (EKG) was simultaneously measured with surfaceelectrodes. There was a decrease in the arterial pulse pressure with RVpacing consistent with the decreased EH performance (Fig. 4A).Fig. 4C–D presents the in vivo stable voltage results at a heart rate of 140bpm during pacing, and simultaneous EKG and arterial pulse pressure.

We additionally characterized the influence of cardiac contraction onthe EH by injecting Calcium Chloride (5–7mg/kg IV) to increase cardiaccontractility. During the peak increase in aortic pulse pressure, the voltageoutput from the EH device increased by a factor of 2, from 0.6 V up to1.2 V (peak-to-peak) as shown in Fig. 5A. Fig. 5B shows the in vivo voltagefluctuation from the EH device at the transition state from normal to in-creased contractility. Similarly, the trend of arterial pulse pressure wasconsistent with the EH performance, while there were no significantchanges observed in EKG (Fig. 5A). The in vivo results from the EH device

Fig. 2. In vitro experimental characterization. (A) Comparison of voltage results generated by the EH device from the motions of vibration and the bending motionof the pacemaker lead. (B) Comparison of voltage output in air and in saline. (C) Cyclic stability of the EH device in saline over 104 cycles. (D) Charging curve for theEH by connecting with a bridge rectifier and a capacitor, and inset is a schematic of equivalent circuit of the EH for measuring the voltage across a 33 μF capacitor. (E)Comparison of open-circuit voltage output generated by single- and double-helix EH devices. (F) Comparison of voltage output generated by single- and double-layerPVDF-TrFE porous thin films EH devices.

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under different anchoring conditions, varying heart rates, and drugtreatment conditions are consistent with in vitro experimental results.

3. Conclusions

We have developed a helical-designed compact EH device integratedwith an existing pacemaker lead to convert cardiac motion to electricalenergy. Our approach is ideal for clinical translation, since it is compatiblewith the existing lead implantation techniques and does not require anadditional surgery for placement. Advances in material fabrication and abioinspired helical configuration establish the design strategy of the EHdevice. Here, we have developed the porous structure material of PVDF-TrFE, whose piezoelectric output has been enhanced by three-fold throughthe optimization of pore sizes and electromechanical coupling efficiencycompared to solid PVDF-TrFE thin films [48,49]. Additionally, the helicalconfiguration of the EH device shows significant advantages for capturingboth complex vibrational motion and the bending/twisting of the lead dueto the beating heart. Theoretical models (see Supporting Information) weredeveloped to establish design rules for guiding the fabrication approach viamultilayers of the device. The performance of the EH was evaluated in vitroby simulating the right ventricular motion; measurements considered thevibration and bending motion of the lead and included extensive tests forheart rate cycling, EH behavior in saline, and charging capability. A matrixof these helical energy harvesting elements on the pacemaker lead implies a

great potential for charging the pacemaker batteries. In addition, in vivostudies demonstrated clinical translation in a large animal model. Both invitro and in vivo results demonstrate the EH's capability to provide sig-nificant electrical energy directly from the motion of a pacemaker lead.Future directions with this work consist of multiple optimization steps.Multilayer-based EH devices in a mixed connection pattern (in series andparallel) using multiple EH units will be conducted in future work to en-hance the overall energy outputs. In addition, the capabilities along theentire length of the lead will also be investigated for increasing the func-tional area of the EH lead. The proposed cardiac EH strategy represents asignificant step forward for clinical translation without a thoracotomy forpatients, and thus eliminates the need for surgeries of battery replacement,and provides a sustainable power source for pacemakers or any other im-plantable and wearable biomedical devices.

4. Methods

4.1. Fabrication of porous PVDF-TrFE thin film

The PVDF-TrFE solution was prepared from polyvinylidenedi-fluoride-trifluoroethylene powder (molar ratio: 75/25), which was thendissolved at a 15wt % in N, N-Dimethylformide solvent for 8 h undermagnetic stirring conditions. The first layer was solid PVDF-TrFE thinfilm, which was made through consecutive spin coating (1000 rpm,

Fig. 3. In vivoevaluation of anchoringeffect in porcine heart. (A) A schematic ofencapsulated helical piezoelectric EH deviceon a pacemaker lead for in vivo experi-mental characterizations. (B) Photograph ofimplantation of a pacemaker lead with anEH device into a porcine heart. (C) Afluoroscopy image of a porcine heart withthe pacemaker lead implanted. (D) In vivomeasurements of the open-circuit voltagesgenerated by the same helical EH devicebefore and after anchoring the tip of thepacemaker leads in two different porcinehearts. The peak-to-peak voltage output ofthe EH device for the anchored scenario inAnimal A is 0.5 V (top) and the voltage re-sult in Animal B is 0.65 V (bottom).

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30s) on a clean glass slide. The cast solid film was then evaporated in anoven for 10min at a temperature of 50 °C. The spin-coating process wasrepeated to fabricate porous PVDF-TrFE thin film directly on the firstlayer of solid PVDF-TrFE film. This porous piezoelectric thin film wasformed through a water vapor phase-separation method. By controlling90% relative humidity at a temperature of 37 °C for 6 h in a chamber, asecond layer of porous PVDF-TrFE film was produced. Through thisfabrication method, the PVDF-TrFE film completed phase separationbetween solvent and non-solvent, and the pore diameter and porositywere finely controlled [48]. An average pore diameter of 6 μm wasfound within the porous film. After evaporation, the thin film waspeeled from the glass slide and 20 nm of gold was sputter coated on

either side of the thin film, forming the top and bottom electrodes of thestandard sandwich piezoelectric device. To increase the material crys-tallinity, the annealing process was performed at a temperature of135 °C for 2 h. Then the sample was electrically poled at 100 °C byapplying an electrical field of 80 V/μm for an hour.

4.2. Fabrication of helical EH device

The single layer porous PVDF-TrFE film based helical EH devices werefabricated by stretching the bottom layer PDMS of the sandwich structure,while leaving the rest of the layers at a neutrally strained state. This causeda folding of the sandwiched thin films when the tension was released. As it

Fig. 4. In vivoevaluation of pacing in a porcine model.(A) In vivo measurements of the open-circuit voltages generated by the helical EH as a function of time bychanging the heart rate via a temporary pacemaker from 94 bpm to 160 bpm, and simultaneously-monitored EKG and arterial pulse pressure of the porcine during thepacing process. (B) Schematic of cardiac pacing. (C) In vivo results of voltage output of EH device at a heart rate of 140 bpm. (D) Simultaneously-monitored EKG andarterial pulse pressure of the porcine at a heart rate of 140 bpm.

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is shown in Supporting Information Fig. S1, a spin-coated thin layer ofPDMS was first given tension through mechanical stretching at a stretchratio of 0.3 (30% larger than initial length). Then the piezoelectric thin filmwas placed on the stretched PDMS at a given angle of 30° to the direction ofthe stretching. This piezoelectric thin film was made up of a layer of solidPVDF-TrFE and a layer of porous PVDF-TrFE, sandwiched between twolayers of gold electrodes. A second layer of PDMS was then spin-coated ontop of the piezoelectric thin films. Finally, the sandwiched PDMS/PVDF-TrFE/PDMS film was cut away and released from the stretching, where thepre-stretched bottom layer of PDMS caused it to fold into a helical structure.To integrate the EH sandwich device with the pacemaker lead, the self-wrapping helical structure is placed along the bottom 5 cm of the pace-maker lead, in the approximate area of maximum bending when implantedinto the right ventricular myocardium heart wall. The device is held in placeby utilizing the self-wrapping helical structure of the EH device. The device'swiring runs up along the entire length of the lead and is held in place bysmall pieces of polyolefin heat-shrink tubing. An extra length of the wiringextends beyond the lead to give easy access for analysis. The first 7 cm of thelead, including the EH device and its wiring, are sealed using a two-partscurable liquid silicon rubber, in order to maintain the lead's flexibility. Therest of the pacemaker lead length is then covered in polyolefin heat-shrinktubing to reduce wire movement during implantation.

4.3. In vitro experimental method

A shaker-based test setup was designed to mimic the in vivo condi-tions (see Supporting Information Fig. S2). The testing setup consistedof a shaker (2025E from The Modal Shop), a pacemaker lead (Med-tronic CapSureFix Novus MRI SureScan Model 5076), and a fixtureframe. The tip of the pacemaker lead was connected to the shaker tomimic the heart motion during systole and diastole at a target fre-quency (1–2 Hz) and displacement (<1 cm). This input excitation was

provided by a function generator (Keysight 33500 B) with a poweramplifier (2100E21-400 from The Modal Shop). Real-time electricaloutput generated by the helical EH was collected using a data acqui-sition (DAQ) card (National Instruments, model NI USB-6008) andLabView software. An oscilloscope (Tektronix TDS 2014B) was used toassist monitoring the voltage output by the EH device simultaneously.

4.4. In vivo experimental method

All animal procedures were approved by the Institutional Animal Careand Use Committee at the University of Texas Health Science Center at SanAntonio. Studies were performed on one male pig and one female pig, agedsix months old and an average weight of 258.4±6.2 lbs. Pigs were sedatedvia Telazol (4–8mg/kg IM), Xylazine (1–2.2mg/kg IM), and 3–4%Isoflurane, followed by endotracheal intubation and maintenance on0.5–3% Isoflurane. The animal was placed on the surgical table in dorsalrecumbency. Lidocaine (2–4mg/kg IV) and Amiodarone (10–12mg/kg IV)were maintained throughout the procedure to prevent ventricular fibrilla-tion. Following anesthesia, a vascular cutdown was performed over theright femoral artery, and a micromanometer pressure sensor was advancedto the descending thoracic aorta under fluoroscopic guidance to measureinvasive arterial pulse pressure as an indirect measure of cardiac con-tractility. A vascular cutdown was performed over the right external jugularvein and a 20 French sheath was inserted in the vein via Seldinger tech-nique. A hollow needle was passed into the vein, and a guidewire wasplaced through the needle into the vein. After removing the needle, a ca-theter introducer and sheath were passed over the guidewire into the vein.Then the guidewire and introducer were removed, leaving the sheath,which is then secured in place with suture. A wire was inserted into thepacemaker lead (Medtronic CapSureFix Novus MRI SureScan Model 5076)to increase stiffness and make it easier to manipulate the lead. Prototype EHleads were inserted through the sheath and advanced into the right ventricle

Fig. 5. In vivoevaluation by varying cardiac contractility in a porcine model. (A) In vivo measurements of the open-circuit voltages generated by the helical EHby changing porcine cardiac contractility via Calcium Chloride injection (IV). Real-time monitoring of EKG and arterial pulse pressure, demonstrating the peak ofpressure was observed following drug injection. (B) In vivo results of voltage output from EH device with simultaneously-monitored EKG and arterial pulse pressure atthe beginning of drug injection.

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under fluoroscopic guidance and fixed into the right ventricular septum.The electrical energy generated by the EH devices was measured before andafter screwing in the tip of the pacemaker lead, as shown by fluoroscopy inSupporting Information, Fig. S4A and Fig. S4B, respectively, to compare theenergy generation in these two conditions. A temporary external pulsegenerator (Medtronic, Model 5346DDD) was connected to the proximal endof the lead to pace the right ventricle. A pacing protocol was performed inwhich data was collected at baseline rate, 120 bpm, 140 bpm, and 160 bpm.To collect data during increased contractility, Calcium Chloride (5–7mg/kgIV) was slowly injected. Following data collection, the lead was removedfrom the right ventricle. Real-time EKG and arterial pulse pressure weremeasured throughout these studies.

4.5. Cell culture and cytotoxicity test of the EH device

Biocompatibility of EH devices was investigated using L929 (amouse fibroblast cell line) cells. Cell viability of the experimental groupcontaining the testing samples (PDMS encapsulated porous PVDF-TrFEEH devices) was compared to that of the control group with only cells.The blank group with only cell culture medium was designed to avoidbackground interference. Each group had four replicates. The EH de-vices were prepared and sterilized in 70% ethanol overnight. Cell sus-pension of L929s was prepared and cell number was calculated by a cellauto counter. 100 μL of cell suspension (5× 104 cells/mL) was seededinto each well (96-well plates) of the control and experimental groups.The same amount of cell culture medium was added into wells of theblank group. Then the plates were placed into an incubator (37 °C, 5%CO2) for 12, 24, 48, and 72 h. At the end of each culturing period, 10 μLof CCK-8 solution (from Sigma-Aldrich) was added into each well of thethree groups, and the plates were put back into the incubator for 2–4 h.After that, absorbance of the groups at 450 nm was measured using aUV–Vis plate reader (TECAN, Spark 10M). The survival rate of cells wascalculated using the formula =survival rate A A( )/sample blank

×A A( ) 100%control blank , where Asample, Acontrol, and Ablank is the ab-sorbance of sample, control, and blank group, respectively. Cell viabi-lity was obtained by normalizing the survival rates of the experimentalgroups against that of the control group from the first day.

Acknowledgement

The authors acknowledge financial support from the National Institutesof Health (NIH) Director's Transformative Research Award (R01HL137157,principal investigator (PI): X.J.Z.), National Science Foundation award(ECCS1509369, PI: X.J.Z.) and the startup fund from the Thayer School ofEngineering at Dartmouth. Z.C. also acknowledges the support from theBranco Weiss - Society in Science fellowship, administered byEidgenössische Technische Hochschule (ETH) Zürich and Neukom InstituteCompX Faculty award. L.D. thanks Patricio Sarzosa from Thayer SchoolEngineering at Dartmouth for the assistance with graph design.

Appendix A. Supplementary data

Supplementary data to this article can be found online at https://doi.org/10.1016/j.nanoen.2019.104085.

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Lin Dong is a Research Associate at Thayer School ofEngineering at Dartmouth College. She obtained her PhDfrom Department of Mechanical Engineering at StevensInstitute of Technology, where she was awarded Innovationand Entrepreneurship Doctoral Fellowship. Her researchinterests include advanced materials for implantable andwearable energy devices as well as actuating and sensingdevices using soft materials.

Andrew Closson is a Ph.D. student under the su-pervision of Professor John XJ Zhang at ThayerSchool of Engineering at Dartmouth College. Hereceived his B•Sc. degree in Bioengineering fromthe University of Maine in 2016. His research fo-cuses on the development of materials and devicesfor energy harvesting applications in biomedicaldevices.

Meagan Oglesby received her Bachelor's Degree at TrinityUniversity in 2015. Formerly working under the super-vision of Dr. Marc Feldman at UT Health San Antonio, herresearch focused on novel cardiovascular technology andtherapies. She is currently pursuing a Master's degree inPhysician Assistant studies.

Daniel Escobedo is a research associate in the Departmentof Medicine/Cardiology at University of Texas HealthScience Center at San Antonio. He has over 35 years ofexperience in animal surgery. He has also been a consultantfor Millar Instruments and Science Inc.

Xiaomin Han received his PhD degree at Thayer School ofengineering, Dartmouth College in 2019. He was a PhDstudent under the supervision of Assistant Prof. Zi Chen. Hismain research focuses on the multistability of thin struc-tures.

Yuan Nie received her BS degree (2011) and MS degree(2014) in Mechanical Engineering in P. R. China. She iscurrently a Ph.D. candidate at Thayer School ofEngineering, Dartmouth College, NH, USA. Her researchinterests are the development of microfluidics reactors,synthesis of nanomaterials, on-chip biosensing and cellanalysis.

Shicheng Huang received his BS degree in EngineeringMechanics in P. R. China in 2015. He is currently a Ph.D.candidate of Thayer School of Engineering at DartmouthCollege. His research interests include post-buckling con-figuration of a rod buckling in elastic media, chiralitytransformation of helical ribbons, and the origami struc-tures and properties.

Marc D. Feldman attended College at Duke University andMedical School at University of Pennsylvania. He com-pleted his Internship and Residency at the University ofChicago, and Cardiology training at Harvard - Beth IsraelHospital. He is currently the CCU Director at the Universityof Texas Health Center in San Antonio and the BriscoeProfessor of Medicine. He has published 140 peer reviewedfull length manuscripts, and co-authored 33 issue andpending US Patents.

Zi Chen is an Assistant Professor at Thayer School ofEngineering at Dartmouth College. He received his PhDfrom Department of Mechanical and Aerospace Engineeringat Princeton University and was a Research Scientist atWashington University in St. Louis. His research interestsinclude mechanical instabilities, multistable structures,energy harvesting, origami/kirigami structures, robotics,and tissue/cell biomechanics.

John X.J. Zhang is a Professor at Dartmouth College. Hereceived a Ph.D. from Stanford University, and was aResearch Scientist at MIT. He is a Fellow of AmericanInstitute for Medical and Biological Engineering, and a re-cipient of NIH Director’s Transformative Research Award,NSF CAREER Award, DARPA Young Faculty Award,Facebook SARA award and Sony Faculty Innovation Award.His research focuses on exploring nanomaterials, bio-physics, and nanofabrication technology for biosensing,energy harvesting and wearables.

L. Dong, et al. Nano Energy 66 (2019) 104085

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