In the 1950s, a founder of ASAIO, Dr. Peter Salisbury described the need to treat heart disease with a mechanical device and proposed the idea of a truly artificial heart. His work challenged the scientific community to treat chronic heart disease beyond medicines and the operating table.1 Implantable cardiac devices have revolutionized patient care and outcomes, and are approved as a bridge to transplant, recovery, and also as destination therapy (DT). The number of left ventricular assist device (LVAD) systems implanted worldwide has increased considerably and now exceeds the number of heart transplants. Krabatsch et al. suggest that LVAD implantation, not heart transplantation, be the primary therapy for terminal heart failure.2 Despite the advances in device technology, a sustained, internal power source remains evasive.
Implantable cardiac device technology including pacemakers, automatic implantable cardioverter-defibrillators (AICDs), total artificial hearts (TAHs), and LVADs depend on a safe, reliable, convenient, and continuous power source. Although technology continues to advance, all current devices still require an external power source. Limitations of current devices include, but are not limited to, the need for periodic battery replacement, catastrophic power failure,3 and transcutaneous driveline erosion and infections.4 These major and other minor complications, related to an external power source, decrease overall survival, quality of life, and patient independence.5
A fully implantable device with no external connections and a lifelong energy supply can eliminate a significant source of morbidity and mortality for patients.6 Barriers to the creation of a fully implantable device include device size, biodegradable materials, reliability, and a nontoxic energy source with an appropriate life span. Achieving total implantability would reduce the need for future interventions such as battery replacement and minimize host and device complications.7 We sought to review the limitations of current cardiac device energy sources and discuss the history, current state, and trends of future potential energy sources that are promising in our pursuit toward a fully implantable cardiac device.
Implantable Cardiac Devices
Cardiac Pacemakers and Automatic Implantable Cardioverter-Defibrillators
Initially, pacemakers were powered with a mercury-zinc chemical cell and provided 3–4 years of power.8 In the 1970s, Plutonium 238 (Pu-238) was used to power pacemakers.9 One hundred and thirty-nine patients received Pu-238-powered pacemakers between 1973 and 1987 at Newark Beth Israel Medical Center.9–11 Many patients were lost to follow up, but 11 had a fully functional pacemaker for at least 20 years, with the longest reported device survival of 34 years.11 Based on the institution’s experience, they found the devices to be safe and reliable. Chauvel et al.12 also found that in follow-up of 325 Pu-238-powered pacemakers, the device survival was 97% at 18.5 years. Despite no reported local or systemic consequences to radiation exposure, concerns of extreme toxicity led to the abandonment of nuclear power.9
In the 1970s, power sources for pacemakers and other devices made the transition from mercury–zinc chemical cells to a lithium iodine chemical cell. The lithium iodine cell was smaller in size, more reliable, and provided a longer shelf life than the mercury-zinc cell.8 The lithium iodine battery is able to provide pacemakers with 25 μJ13 for 7–10 years. Because AICDs are required to pace, recognize, and provide high-energy therapy (40 J at 700–800 V within a 10–15 msec period) for tachyarrhythmias, their energy requirement is different.14 The lithium iodide battery was unable to provide low voltage for pacing and deliver high voltage discharges as needed. A lithium–silver vanadium oxide battery is able to accomplish this, and remains as the most commonly used battery for AICDs. Despite these improvements, patients still outlive their device, and require surgical intervention for battery replacement.
Advances in technology (e.g., wireless transmission14) require high currents, which limit the future advancements of lithium batteries that can only supply very low currents.14 These requirements and the need for improved device longevity have led to the pursuit of alternate energy sources. Current cardiac devices and their energy requirements are described in Table 1.
Total Artificial Hearts and Left Ventricular Assist Devices
Total artificial hearts and LVADs are long-term mechanical circulatory support systems that serve as either a bridge to therapy (BTT), recovery, or DT. The LVADs are attached and augment the native hearts cardiac output, whereas TAHs replace the native heart providing all of the cardiac output. The energy requirement to provide left ventricular flow at 10 L/min at a pressure of 100 mm Hg is 2.2 W.15 The power requirement for both ventricles to sustain a cardiac output of 10 L/min is 3.09 W.15 Total artificial hearts have a higher energy requirement as it powers both ventricles as opposed to one in LVADs, and energy efficiency varies by device generation (Table 1).6,16–18,22,25 Moreover, the form of flow the device delivers affects the energy requirement of the device. Left ventricular assist devices can be categorized as pulsatile versus continuous-flow (nonpulsatile). Pulsatile displacement pumps that imitate the heartbeat have higher energy requirements than continuous flow pumps.29,30 The pulsatility of clinically used TAHs is an attributable factor for their higher energy requirement.30 Continuous flow TAHs are currently under development.17,18 Thus, unfortunately, LVADs and TAHs require much more energy than the predicted 2.2 and 3.09 W, respectively, to accommodate for device inefficiency and different patterns of flow.
Since the inception of TAHs and LVADs, finding suitable energy sources has proven to be problematic. Current models have high rates of complications and cause discomfort to the patient because of the reliance of an external driveline or device unreliability.31,32 The most widely used energy source for TAHs and LVADs are implanted batteries supplied by a percutaneous driveline connected to an external battery. Depending on the device, the internal battery can provide 20–60 minutes of energy when not charged.33 The driveline provides a connection between the external and the internal environment and is a major source of complications, notably infections and trauma. The driveline is a conduit for bacteria, often resulting in multidrug resistant infections that ultimately lead to rehospitalization, higher health care costs, and greater patient burden.34,35 The REMATCH trial demonstrated that device-related infections were responsible for 41% of deaths in one LVAD population.34 Another study revealed that prophylactic antibiotics did not significantly reduce driveline infections, again demonstrating the urgency for an alternative energy source.35 Additionally, the presence of a driveline places many limitations: for example, patients are restricted in activities that can compress, pull, or kink the line; it needs to be connected to an electrical power source while sleeping; and patients are limited with water contact to avoid getting external components wet.36 Even small amounts of mechanical stress can result in damage to the leads internal wires, ultimately leading to pump failure and a need for urgent hospitalization for repair. These complications lead to unplanned hospitalizations resulting in additional economic burden on patients and hospitals with a median cost of $4,857 (range, $207–$627,203) per hospitalization.37
Current efforts are to relocate or replace the driveline to minimize infections and patient burden. In 1980, Phillips et al.38 described two patients where he stabilized the velour-encased drivelines by exteriorizing them through the iliac crest. Both patients died, one from refractory arrhythmia and the other from multiorgan failure, although autopsy showed no infection and both drivelines were healed solidly into the bone. Another example is the Jarvik 2000 LVAD, whose driveline exits through the temporal bone and connects to a power supply behind the ear.39 Bone is highly vascular and resists infection, and minimizes drive live vibration compared with an exit through the abdominal wall.
Transcutaneous energy transmission system (TETS) powers a transmitter coil outside the body with alternating current to build a magnetic field passing through the human skin. This induces a current on a receiver coil inside the body to supply power for the implantable cardiac device.40 Without the need for a driveline through the skin, TETS has potential to prevent driveline-related infections and improve patient’s quality of life.41 The first mechanical circulatory support systems implanted with TET technology with a TET coil in the subclavicular position and external coil secured on patient skin are the AbioCore TAH (Abiomed, Danvers, MA) and the LionHeart 2000 LVAD (Arrow International, Reading, PA).6 Although TET continues to be investigated as a potential method to power implantable cardiac devices, limitations of TET include thermal injury from the external coil, while range and alignment problems can reduce energy transfer. Additionally, patients are also still required to wear heavy portable batteries, as the efficiency of the transmission coil is approximately 30%.39
Another mode of power transfer for implantable cardiac devices that is being investigated is the free-range resonant electrical energy delivery system (FREE-D). This FREE-D uses high-efficiency resonant coupling technology to provide wireless power to an implantable cardiac device without requiring direct contact between the patient and energy source.6 This FREE-D offers multiple advantages over TET systems (which operate based on nonresonant coupled induction) including improved efficiency and patient tether-free mobility in an unrestricted space.6 Reports of FREE-D in the clinical setting are eagerly anticipated. Overall, none of these options meet the requirement of total implantability because of the associated peripherals, even though they are not directly attached to the organism.
Current Trends and Future Perspectives of Implantable Energy Sources
In the pursuit to create a lifelong implantable cardiac device, the energy source needs to be an externally implanted source dense enough to last a lifetime, or harvested internally and remain self-sustaining. Table 2 describes the advantages and disadvantages of potential energy sources for fully implantable biomedical devices. One area of great potential is the utilization of intracorporeal energy from human mechanical, biological, and chemical sources. This can reduce complications and inadequacies seen with traditional energy sources. We will explore future perspectives on implant technology to provide insight into these barriers, and identify areas requiring further development to achieve a totally implantable device.
Mechanically Harvested Energy
Self-powered bioimplantable systems have been studied as an attractive approach to integrate an energy-harvesting device inside the human body to convert biomechanical movements into electrical power. Mechanical energy harvesting techniques include the following: 1) piezoelectric effect, 2) electromagnetic generation, and 3) electromagnetic induction.
Piezoelectricity is the electric charge that accumulates in certain solid materials such as bone, DNA, and proteins.52 Advantages of a piezoelectric system include its small size, lack of moving parts, and its self-sufficiency whereby the system uses the power it generates to operate. The application of flexible piezoelectric thin films as a self-sustainable energy source for eliminating batteries in biomedical devices and as a nano sensor for in vivo diagnosis or therapy have been described.53 A piezoelectric stack generator can be surgically attached between a bone and a muscle-tendon unit. The motor nerve of the muscle will then be electronically stimulated to produce an isometric muscle contraction, which repetitively exerts a force on the piezoelectric generator to generate power with excellent efficiency in vivo.52 Large muscles produce more power than small muscles, but to minimize loss of function and for easier implantation, smaller and redundant muscles can be used for greater yield in power.54
Lu et al.555 reported on the successful utilization of piezoelectricity to harvest biomechanical energy from heart motion and showed the possibility and feasibility of self-powered artificial pacemakers in an animal model. With each heartbeat, the piezoelectric device was bent and electrical energy was converted from the material’s mechanical deformation.42 The electrical energy generated provided the pacemaker the necessary electrical stimuli (0.223 mA, 8.2 V) to pace the heart. This novel idea could not only charge pacemaker batteries, but also provide a self-sustainable energy source for pacemakers.
The application of piezoelectric material has also been extended to implantation of the material on the humeral insertion of the latissimus dorsi, which Trumble et al.56 reported to have produced enough power to drive an LVAD in a canine model. This produced a steady state power of 478 mW per contraction, enough to provide circulatory support to the canine for 37 days before device failure (addressed as a manufacturing flaw). Despite promising results, the application of piezoelectric material requires the selected muscle to be fatigue-resistant – a limitation that warrants consideration. In those canines that were implanted with piezoelectric driven LVADs, the latissimus dorsi, the muscle targeted for energy harvesting, was electrically conditioned for 6–8 weeks to convert the muscle phenotype to small and fast twitch fibers.56,57 In the model, muscle conditioning was successfully accomplished, but the feasibility of replicating this in a human population is questionable where there is no proven long-term stability. As well, the targeted muscle function would be sacrificed and no longer used by the patient. This is less of an issue in patients with previous dysfunction such as paraplegics or quadriplegics, where the functionality of the paralyzed muscle is not expected to return.52
Piezoelectric soles have been demonstrated to be able to harvest mechanical energy from the human motion of walking. Khaligh et al.58 describes a polyvinylidene fluoride (PVDF) insole in a sneaker. When the plastic was bent at a frequency of 0.9 Hz, it produced 1.3 mW. Another material used was a lead zirconate titanate (PZT) bimorph insole, which at a frequency of 0.9 Hz was able to produce an average power of 8.4 mW.58 Even more power was generated with a midsole piezoelectric device producing 8.2 W per foot at 1 step/sec for a 70 kg person.59 As the amount of power generated was minimally influenced by foot contact pattern, this led to a consistency between patient-to-patient variation in power production.59 Strengths of piezoelectricity include the lack of required surgical or invasive procedures. Limitations include the need for continuous motion to produce energy or the requirement for a secondary energy source when idle, and the low energy yield.
Electromagnetic Generation from Hydraulic Power
Cardiac hydraulic power can be harvested through electromagnetic generation to produce electricity. Pfenniger et al.43 describe the mechanism of using the pressure potential of blood converted into mechanical torque energy to rotate an electromagnetic generator, which produces electrical energy. In vitro, a turbine created by tesla was used as an implantable intravascular turbine for a peripheral artery like a stent. Numerical models of the turbine resulted in creating a power output of 840 μW at a flow rate of 237 ml/min.43 Potential sites for turbine placement include the internal thoracic artery (ITA). As the turbine requires a minimum arterial flow rate of 100 ml/min (physiologic baseline 40 ml/min), an arterio-venous (AV) shunt is needed to increase ITA flow rate from 40 ml/min to 160 ml/min.43 The dependence on an AV shunt makes this a less practical option. It is also concerning that the turbine itself is dependent on the device being powered. Additionally, the insertion of any endovascular device is not without complications. Materials used need to be nonthrombogenic and hemodynamically compatible to prevent thrombi/emboli, cell shearing, and platelet activation. Pfenniger et al.43 demonstrated that in theory, the majority of flow shear stress exposure is below the level for platelet activation, although in vivo studies are lacking.
Utilization of cardiac hydrodynamic power would mean that a device would not require a separate pump, but rather rely on the heart as a pump—although these types of devices cannot be utilized for failing heart that requires mechanical circulatory support. Another area of exploration includes the improved oxygenator technology with lower hemodynamic resistance. This may be feasible in a passive, pumpless application in the setting of isolated pulmonary failure providing that normal cardiac function is present. Such application of cardiac hydrodynamic power has already been utilized for devices such as pump less total artificial lungs (TALs).60 Extracorporeal devices, such as the Novalung, remove CO2 when connected to an AV shunt, and operate on passive blood flow supplied by native cardiac function, thus requiring no additional energy source.61 Thus, utilization of native cardiac function with AV shunts remains as a viable alternative energy source.
Another system to harvest energy is electromagnetic induction that converts arterial expansion and contraction to electrical energy. Pfenniger et al.44 developed a prototype flexible coil to surround an artery and provide a magnetic field in parallel to the artery axis. This mechanism uses pulsatile arterial pressure to create energy via the external coil. During systole, the coil expands due to the deformation of the artery and induces an electric field that was able to generate up to 42 nW of harvestable power in an ex vivo swine artery. Limitations of the study included the use of water as fluid in the place of blood, and a fixed systolic/diastolic pressure of 140/60 mm Hg. As the device is external to the artery and in contact with adventitia rather than intima, it can be hypothesized that fewer endovascular complications are expected. However, as with any foreign device, inflammatory reactions may occur, and in this case can cause hardening of the artery. Limitations of this approach include the following: 1) arterial calcification and artery hardening, which can interfere with device performance; 2) large device size; 3) implantation requires an invasive procedure; 4) the application of this technology in patients with cardiovascular disease may be limited because of its dependence on cardiovascular function, where changes in blood pressure such as arterial hypotension will negatively affect device function.44
Biofuel cells use enzymes from living cells rather than metals to oxidize organic compounds and generate electrical energy. The human body is composed of a large amount of oxygen and glucose, which may be harvested for biochemical energy. Selection of enzymes and biofuel should take into account the risks and benefits of competing reactions and byproducts produced. For example, hydrogen peroxide generated from oxygen-dependent oxidases is toxic. The use of biofuel cells for device support such as artificial hearts and pacemakers has been explored, although operational lifetime and power requirements remained barriers to feasibility.45 Over time, significant progress has been made in extending the lifetime and reducing the size of biofuel cells leading to improvements in biocompatibility. For example, biofuel cells to sense and transmit per-implant space temperature, pressure, and chemical concentration were made with only 2 out of 10 components from previous biofuel cells.45 Fewer components allowed them to be small (size <1 mm3, weigh <100 μg) and affordable (<$1 USD) while operating in a physiologic environment, with nontoxic reactants and products.45 A glucose biofuel cell (2.4 ml volume in size) has been implanted in a rat abdominal cavity and produced 38.7 μW to power an LED and thermometer. In the healing process, the cell was covered with vascularized thick adherent adipose tissue with no signs of rejection or inflammation after 110 days.46 Another report shows the successful application of biofuel cells in a snail.47 Electrodes were placed into the snail hemolymph, which is rich in oxygen and glucose, to generate 7.45 μW of power.47 There was a decline in electric output upon operation of the cell; however, these limitations were due to glucose transport and snail exhaustion rather than a decrease in enzyme activity.47 This problem is less likely to be an issue in mammals with a greater glucose and oxygen concentrations in blood compared with hemolymph, and more efficient circulation to restore biofuel stores.47
The application of biofuel cells has already been studied in pacemakers. The biofuel cell, composed of pyrroloquinoline quinone-dependent glucose dehydrogenase and laccase, was connected to a battery-free pacemaker. The device was then placed in human serum solution with a physiologic glucose concentration of 6.4 mM. Under these conditions, the biofuel cell was able to power the pacemaker for at least 5 hours.48 Although results are promising, further research in humans is needed to determine its feasibility as a self-sustaining energy source.
Radioisotope Thermoelectric Sources
Nuclear batteries use energy from particles emitted by radioisotopes.62 Radiation produced by nuclear isotopes creates heat that can be converted to electrical current by a thermopile. A thermopile is a group of thermocouples, which convert thermal energy directly into electrical energy.9 The long life span and stable power output under environmental stress has made nuclear isotopes an desirable alternative energy source for implantable devices.10 Previous cardiac pacemakers with Pu-238 had lasted more than 30 years without any adverse events. Pu-238 was also studied for use in TAHs and LVADs between 1964 and 1973.49 The first totally implanted TAH was in fact anatomically driven by Pu-238.50 This led to the development of a fuel container capable of generating 50 W of energy; however, the idea was abandoned because of public fear of nuclear energy. With a half-life of 89 years, a higher power density and negligible irradiation when encapsulated, it remains an attractive source of energy.14,63
Promethium is another nuclear energy source that has been used in the Betacel pacemaker produced by Medtronic’s. It could provide 50 μW of power for 7–10 years with no harmful effects secondary to radiation.14
A third nuclear source tritium was found to have a life span of >20 years; however, the poor energy output prevented its use.51
Challenges to application in implantable devices.
Barriers remain in place to obtain nuclear materials for further testing and clinical use, where an implant licensed from national nuclear energy authorities would be required.14 In addition, recovery of nuclear materials from deceased patients may pose a challenge if not recovered or returned as demonstrated in those with Pu-238 pacemakers. To ensure appropriate safety measures/handling of nuclear material and prevent catastrophic crematorium explosions, patients must satisfy certain conditions such as having appropriate identification at all times.64 Nuclear energy has the potential to satisfy the requirements of a life-long, fully implantable device, although high costs, availability, and safety concerns have limited nuclear use in humans. Although concerns with radiation use should not be disregarded, studies demonstrate that previous use of Pu-238-based pacemakers had such low radiation that it was found to be harmless in pregnant women and their newborn infants.9,65
Challenges to Full Implantability
Current alternative power sources, except nuclear energy, are all in their early infancy and require further research in human studies. Although nuclear energy has been shown to be a reliable source, concerns regarding costs and safety (e.g., leakage and theft) of nuclear fuel remain among society.66,67 Modern pacemakers and ACIDs require 15–50 μW and 40–60 μW, respectively, whereas LVADs and TAHs require 3.5–24 W and 13–27 W of power, respectively (Table 1). Piezoelectric, hydraulic, mechanical, and nuclear sources are able to provide sufficient power for pacemakers and AICDs (Table 2). Piezoelectric and nuclear sources would be sufficient for LVADs and TAHs; other sources remain inadequate, providing power on the nano to microwatt scale (Table 2). Reliability, efficiency, cost-effectiveness, availability, biocompatibility, power output, size, portability, and quality of life remain as major areas of focus in the choice of suitable energy sources. Ultimately, alternative power sources will need to demonstrate decreased mortality or patient burden versus the standard of care to obtain widespread acceptance. Only time will show what will be the energy source of choice in the future.
The pursuit toward full implantability has been perennial. Beginning from the early experience with the pacemaker to the development of artificial hearts,68 technologic innovations have made significant progress in the field of implantable biomedical devices to improve patient care and outcomes. Although battery and driveline technology continues to improve, the success and acceptance of energy source depends not only on its reliability but also on the suitability and ergonomics of its daily function. An energy source that can be fully implantable and last the duration of the patient’s lifetime can significantly improve patient autonomy, survival, and quality of life. Given that each energy source comes with its strengths and limitations, selection of the ideal energy source needs to be tailored to the length of biomedical device support required by the patient to justify the cost effectiveness of the technology.
1. Salisbury P: Implantation of physiological machines into the mammalian organism. Identificiation of problems connected with the implantation of artificial hearts and of artificial kidneys. Experimental results to date. Trans Am Soc Artif Intern Organs 1957.3: 37–42.
2. Krabatsch T: Should LVAD implantation be the new gold standard for terminal heart disease ? ISHLT Links 2013.4:6–7.
3. Jafar M, Gregoric ID, Radovancevic R, Cohn WE, McGuire N, Frazier OH: Urgent exchange of a HeartMate II left ventricular assist device after percutaneous lead fracture. ASAIO J 2009.55: 523–524.
4. Topkara VK, Kondareddy S, Malik F, et al: Infectious complications in patients with left ventricular assist device: Etiology and outcomes in the continuous-flow era. Ann Thorac Surg 2010.90: 1270–1277.
5. Pereda D, Conte JV: Left ventricular assist device driveline infections. Cardiol Clin 2011.29: 515–527.
6. Wang JX, Smith JR, Bonde P: Energy transmission and power sources for mechanical circulatory support devices to achieve total implantability. Ann Thorac Surg 2014.97: 1467–1474.
7. Bazaka K, Jacob M: Implantable devices: issues and challenges. Electronics 2013.2: 1–34.
8. Sanders RS: Kusomoto F, Goldschlager N, The Pulse Generator, Cardiac Pacing for the Clinician. 2008, pp. 47–71, Springer US.
10. Amar BA, Kouki AB, Cao H: Power approaches for implantable medical devices. Sensors (Switzerland) 2015.15: 28889–28914.
11. Parsonnet V, Driller J, Cook D, Rizvi SA: Thirty-one years of clinical experience with “nuclear-powered” pacemakers. Pacing Clin Electrophysiol 2006.29: 195–200.
12. Chauvel C, Lavergne T, Cohen A, et al: Radioisotopic pacemaker: Long-term clinical results. Pacing Clin Electrophysiol 1995.18: 286–292.
13. Mallela VS, Ilankumaran V, Rao NS: Trends in cardiac pacemaker batteries. Indian Pacing Electrophysiol J 2004.4: 201–212.
14. Mond HG, Freitag G: The cardiac implantable electronic device power source: Evolution and revolution. Pacing Clin Electrophysiol 2014.37: 1728–1745.
15. Chien S, Fung Y: An Introductory Text to Bioengineering. 2008.Singapore, World Scientific Publishing.
16. Mohacsi P, Leprince P: The CARMAT total artificial heart. Eur J Cardiothorac Surg 2014.46: 933–934.
17. Fumoto H, Horvath DJ, Rao S, et al: In vivo
acute performance of the Cleveland Clinic self-regulating, continuous-flow total artificial heart. J Heart Lung Transplant 2010.29: 21–26.
18. Greatrex NA, Timms DL, Kurita N, Palmer EW, Masuzawa T: Axial magnetic bearing development for the BiVACOR rotary BiVAD/TAH. IEEE Trans Biomed Eng 2010.57: 714–721.
19. Moise JC, Foerster JM, Faeser RJ, Hellwig JW: Development of compact thermal and electrical energy converters left heart assist systems. Trans Am Soc Artif Intern Organs 1978.24: 77–83.
20. Schumer EM, Black MC, Monreal G, Slaughter MS: Left ventricular assist devices: current controversies and future directions. Eur Hear J 2015.
21. Kucukaksu DS, Sener E, Undar A, Noon GP, Tasdemir O: First Turkish experience with the MicroMed DeBakey VAD. Tex Heart Inst J 2003.30: 114–120.
22. Westaby S, Katsumata T, Evans R, Pigott D, Taggart DP, Jarvik RK: The Jarvik 2000 Oxford system: increasing the scope of mechanical circulatory support J Thorac Cardiovasc Surg 1997.114: 467–474.
23. Hoshi H, Shinshi T, Takatani S: Third-generation blood pumps with mechanical noncontact magnetic bearings. Artif Organs 2006.30: 324–338.
24. Kishimoto S, Date K, Arakawa M, et al: Influence of a novel electrocardiogram-synchronized rotational-speed-change system of an implantable continuous-flow left ventricular assist device (EVAHEART) on hemolytic performance. J Artif Organs 2014.17: 373–377.
25. Chorpenning K, Brown MC, Voskoboynikov N, Reyes C, Dierlam AE, Tamez D: HeartWare controller logs a diagnostic tool and clinical management aid for the HVAD pump. ASAIO J 2014.60: 115–118.
26. Tayama E, Olsen DB, Ohashi Y, et al: The DeBakey ventricular assist device: current status in 1997 Artif Organs 1999.23: 1113–1116.
27. Cheung A, Chorpenning K, Tamez D, et al: Design concepts and preclinical results of a miniaturized heartware platform: the MVAD system. Innovations (Phila) 2015.10: 151–156.
28. Farrar DJ, Bourque K, Dague CP, Cotter CJ, Poirier VL: Design features, developmental status, and experimental results with the Heartmate III centrifugal left ventricular assist system with a magnetically levitated rotor. ASAIO J 2007.53: 310–315.
29. Yozu R, Golding LA, Jacobs G, Harasaki H, Nose Y: Experimental results and future prospects for a nonpulsatile cardiac prosthesis. World J Surg 1985.9: 116–127.
30. Kaiser LR, Kron IL, Spray TL. Mastery of Cardiothoracic Surgery: Lippincott Williams & Wilkins, 2007.
31. Angud M: Left ventricular assist device driveline infections: The Achilles’ heel of destination therapy. AACN Adv Crit Care 2015.26: 300–305.
32. Horton SC, Khodaverdian R, Powers A, et al: Left ventricular assist device malfunction: A systematic approach to diagnosis. J Am Coll Cardiol 2004.43: 1574–1583.
33. Mancini D, Burkhoff D: Mechanical device-based methods of managing and treating heart failure. Circulation 2005.112: 438–448.
34. Chinn R, Dembitsky W, Eaton L, et al: Multicenter experience: Prevention and management of left ventricular assist device infections. ASAIO J 2005.51: 461–470.
35. Stulak JM, Maltais S, Cowger J, et al: Prevention of percutaneous driveline infection after left ventricular assist device implantation: Prophylactic antibiotics are not necessary. ASAIO J 2013.59: 570–574.
36. Mussivand T, Hendry PJ, Masters RG, Keon WJ: Development of a ventricular assist device for out-of-hospital use. J Heart Lung Transplant 1999.18: 166–171.
37. Quader MA, Green AJ, Shah KB, Cooke R, Kasirajan V: Hospital readmissions after discharge to home with the Total Artificial Heart Freedom driver: readmission reasons, clinical outcomes, and health care costs. J Heart Lung Transplant 2016.35: 251–252.
38. Phillips SJ, Kongtahworn C, Zeff RH, et al: A new left ventricular assist device: Clinical experience in two patients. Med Instrum 1980.14: 288–293.
39. Gerosa G, Scuri S, Iop L, Torregrossa G: Present and future perspectives on total artificial hearts. Ann Cardiothorac Surg 2014.3: 595–602.
40. Fu Y, Hu L, Ruan X, Fu X: A transcutaneous energy transmission system for artificial heart adapting to changing impedance. Artif Organs 2015.39: 378–387.
41. Fang W, Liu W, Qian J, Tang H, Ye P: Modeling and simulation of a transcutaneous energy transmission system used in artificial organ implants. Artif Organs 2009.33: 1069–1074.
42. Hwang GT, Park H, Lee JH, et al: Self-powered cardiac pacemaker enabled by flexible single crystalline PMN-PT piezoelectric energy harvester. Adv Mater 2014.26: 4880–4887.
43. Pfenniger A, Vogel R, Koch VM, Jonsson M: Performance analysis of a miniature turbine generator for intracorporeal energy harvesting. Artif Organs 2014.38: E68–E81.
44. Pfenniger A, Wickramarathna LN, Vogel R, Koch VM: Design and realization of an energy harvester using pulsating arterial pressure. Med Eng Phys 2013.35: 1256–1265.
45. Heller A: Miniature biofuel cells. Phys Chem Chem Phys 2004.6: 209.
46. Zebda A, Cosnier S, Alcaraz JP, et al: Single glucose biofuel cells implanted in rats power electronic devices. Sci Rep 2013.3: 1516.
47. Halámková L, Halámek 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.
48. Southcott M, MacVittie K, Halámek J, et al: A pacemaker powered by an implantable biofuel cell operating under conditions mimicking the human blood circulatory system–battery not included. Phys Chem Chem Phys 2013.15: 6278–6283.
49. Tchantchaleishvili V, Bush BS, Swartz MF, Day SW, Massey HT: Plutonium-238. ASAIO J 2012.58: 550–553.
50. Cooper DK Novitzky D: The Transplantation and Replacement of Thoracic Organs: The Present Status of Biological and Mechanical Replacement of the Heart and Lungs. 2012. New York, NY, Springer Science & Business Media, 451–467.
52. Lewandowski BE, Kilgore KL, Gustafson KJ: In vivo
demonstration of a self-sustaining, implantable, stimulated-muscle-powered piezoelectric generator prototype. Ann Biomed Eng 2009.37: 2390–2401.
53. Hwang GT, Byun M, Jeong CK, Lee KJ: Flexible piezoelectric thin-film energy harvesters and nanosensors for biomedical applications. Adv Heal Mater 2015.4: 646–658.
54. Lewandowski BE, Kilgore KL, Gustafson KJ: Design considerations for an implantable, muscle powered piezoelectric system for generating electrical power. Ann Biomed Eng 2007.35: 631–641.
55. Lu B, Chen Y, Ou D, et al: Ultra-flexible piezoelectric devices integrated with heart to harvest the biomechanical energy. Sci Rep 2015.5: 16065.
56. Trumble DR, Melvin DB, Dean DA, Magovern JA: In vivo
performance of a muscle-powered drive system for implantable blood pumps. ASAIO J 2008.54: 227–232.
57. Trumble DR, Melvin DB, Magovern JA: Method for anchoring biomechanical implants to muscle tendon and chest wall. ASAIO J 2002.48: 62–70.
58. Khaligh a, Zeng P, Zheng C: Kinetic energy harvesting using piezoelectric and electromagnetic technologies #x2014; state of the art. Ind Electron IEEE Trans 2010.57: 850–860.
59. Antaki JF, Bertocci GE, Green EC, et al: A Gait-Powered Autologous Battery Charging System for Artifical Organs. ASAIO J 1995.41: 588–595.
60. Bronzino JD, Peterson DR: Molecular, Cellular, and Tissue Engineering. 2015.Hoboken, CRC Press.
61. Fischer S, Simon AR, Welte T, et al: Bridge to lung transplantation with the novel pumpless interventional lung assist device NovaLung. J Thorac Cardiovasc Surg 2006.131: 719–723.
62. Wei X, Liu J: Power sources and electrical recharging strategies for implantable medical devices. Front Energy Power Eng China 2008.2: 1–13.
63. Parsonnet VA lifetime pacemaker revisited. N Engl J Med 2007.357: 2638–2639.
64. NRC: Information Notice No. 98-12: Licensees’ Responsibilities Regarding Reporting and Follow-Up Requirements for Nuclear-Powered Pacemakers.
65. Laurens P, Haiat R, Gavelle P, Maurice P, Chiche P: [Isotop cardiac pacemaker during pregnancy. 3 cases]. Nouv Presse Med 1976.5: 2997–3000.
66. Poirier V: Will we see nuclear-powered ventricular assist devices? ASAIO J 2012.58: 546–547.
67. Phillips SJ: Nuclear powered devices: is it time to revisit the use of nuclear energy? Artif Organs 2015.39: 201–202.
68. Schumer EM, Ising MS, Slaughter MS: The current state of left ventricular assist devices: Challenges facing further development. Expert Rev Cardiovasc Ther 2015.13: 1185–1193.