## 1 Introduction
Rechargeable batteries, such as those used in cochlear implants, can be powered or recharged transcutaneously using external signals, e.g., radio frequency (RF), ultrasound, infrared light, low-frequency magnetic field, and so on. More recently, internal charging using the energy produced by the physiological environment or natural body motion has been investigated [21; 22]. Further miniaturization can be attained by means of battery-less implants, where energy harvested from natural or artificial power sources surrounding the patient is used directly to power the device [23; 24; 25; 26]. Inductive (or near field) and electromagnetic (or far field) coupling are frequently used for remote powering of such battery-less devices [27]. In the former case,time-harmonic magnetic field generated by the low-frequency alternating current in the external coil generates an alternating current in the implanted component [28], whereas in the latter, electromagnetic waves propagate from the antenna in the far field region to power the implanted chip [29]. Biomedical actuators that do not rely on the conventional wireless delivery, harvesting, accumulation and storage of power in electrical form have also been investigated for such high-energy actuation applications as drug release and mechanical adjustment in prosthetic devices [30]. Recently, [PERSON] and [PERSON] designed stepwise microactuators where incoming ultrasonic waves initiate vibrations in the mechanical oscillator components of the device. The oscillations are then converted into stepwise motion of a mechanical actuator through oblique impact, without the need to convert the energy from mechanical to electrical form. Figure 1: **(A)** Cross-section of a model of the modified hip implant with a metal head. The temperature telemetry with thermistor, electronic circuit and power/data coil are placed inside the neck of the implant [31]. **(B)** Retinal prosthesis: (**a**) the schematic version of a minimally invasive approach; **(b)** photograph of an implant in the eye of a minipig [32,33]. **(C)** An active, flexible device for cardiac electrophysiological mapping: (**a**) circuit design; **(b)** photographic image of the fabricated device in a slightly bent state; the inset shows a magnified view of a pair of unit cells; **(c)** sequential images during the contraction cycle of the heart, with blue lines emphasizing the degree of bending along the device and the black arrow in the left-most image indicating a conventional pacing electrode [34]. **(D)** The implantable microchip-based human parathyroid hormone drug delivery device (54 mm \(\times\) 31 mm \(\times\) 11 mm, \(l\times w\times h\)) containing two microchips with 10 reservoirs each (13.0 mm \(\times\) 5.4 mm \(\times\) 0.5 mm, \(l\times w\times h\)). Schematic cross section of microchip assembly showing drug releasing from one reservoir [35]. Concomitantly, there is a strong emphasis on increasing the functionality and reliability of these electronic devices to support real-time complex _in vivo_ stimulation, data collection, data compression and fast wireless data transmission to external components of the system. This increasing complexity of signal processing electronics further adds to the power budget of the device, which should remain very low if the device is to remain operational for extended periods of time. For instance, a ultra wide band technology offers high-speed data transfer between the implanted device, e.g., implantable electronic cardiovascular devices, and the medical practitioner, and low interference potential, yet its implementation is limited due to its high power consumption [36,37]. Although the wireless programming is highly application-specific, differing significantly between that for a pacemaker and a wireless pulse oxymeter, subjectivity of a particular wireless technology to interference is an important factor to consider, given that the wireless devices operate within an electromagnetically shared environment.