We characterized the wireless powering system in both tissue-mimicking liquid and in multilayered animal tissue. Simulations, field mapping and wireless powering experiments (Fig. 2c–e) demonstrated that the phased surface produces a strongly focused field in homogenous media (saline in experiments, homogeneous muscle tissue in simulations), transferring 0.83 mW to the microdevice over 4 cm of saline at 800 mW output (Fig. 2f and Supplementary Fig. 5). In multilayered porcine tissue, both simulation and experiment showed that 0.45 mW was transferred over 4.2 cm (Fig. 2e,f). For a microdevice in this source–receiver configuration, these transfer efficiencies are about two orders of magnitude greater than those achieved in the near-field by the theoretical optimal source10, and meets requirements for many bioelectronic sensing and modulation systems. For comparison, analogue cochlear processors consume 0.2 mW and pacemakers less than 0.01 mW on average1. Phase control is key to performance, as removal of the reactive loading elements on the phased surface destroys the focal spot, reducing the transferred power fivefold at the 4 cm depth (Fig. 2c,f). In contrast, replacing homogenous media with multilayered tissue reduced the transferred power by less than twofold.
A distinctive feature of the phased surface is that it couples energy into the body with high efficiency and selectivity (>96% of outgoing power), strongly suppressing radiation into free-space. Figure 2f shows that over a 4 cm range from the phased surface, the received power is nearly an order of magnitude less in air than in tissue, even though dissipative losses occur only in tissue. For a microdevice requiring 0.5 mW for function, the device can be powered at a depth of 4.6 cm in a field optimized in muscle tissue, compared with 1.2 cm in air. This counterintuitive result highlights the essential role of propagating fields in transporting energy: the enhanced range is a consequence of evanescent-to-propagating field conversion at the air–tissue interface and subsequent interference with the prescribed phases. The conformal design of the phased surface is essential because a rigid device does not provide effective coupling to the body on curved body interfaces (Supplementary Fig. 6).
Stability of the coupling between the phased surface and the body was tested under deformation and physiological movement. Field mapping measurements show that the field shape is maintained for interfaces with radii of curvature greater than 8 cm, which should accommodate most non-peripheral body surfaces (Supplementary Fig. 6). Measurements of the scattering parameters on three human subjects with the phased surface attached by Tegaderm film on the chest and leg show that the coupled power varies by less than 0.02% during the physiological motions of sitting, standing and walking (Supplementary Fig. 7). In contrast, removal of the phased surface beyond 4 cm distance results in a 7% change in power. No significant changes in the conductivity of the Ag traces were observed across the range of physiologically relevant curvatures (Supplementary Fig. 8). The coupling is also relatively stable across a wide range of polydimethylsiloxane substrate thicknesses greater than 2 mm (Supplementary Fig. 9). These results show that the conformal design enables robust electromagnetic coupling with the body, although matching its mechanical properties to that of skin remains an important direction for future work.
We measured the dependence of performance on source–receiver geometry for the wireless powering configuration in Fig. 3a. As energy transfer is mediated by propagating fields, the dependencies on linear displacement are determined solely by the spatial distribution of the field; the half-power displacements are 11.5 mm in the transverse plane and 6.7 mm in depth (Fig. 3b,c). The orthogonal electric and magnetic components of the generated field interact with both the electric and magnetic dipole moments of the helical receiver. Power can thus be extracted for all azimuthal orientations θ without nulls (Fig. 3d and Supplementary Fig. 10a), while the altitudinal angle ϕ follows a cosine dependence (Fig. 3e and Supplementary Fig. 10b) (half-power displacement: 59° azimuthal and 45° altitudinal). These displacements quantify the sensitivity of performance to changes in source–receiver alignment due to movement of the phased surface or the device, or both. In contrast to near-field systems, operation in the mid-field is immune to frequency-splitting effects, and the overall dependencies are less than or similar to previously reported systems24.
Safety characteristics were also studied computationally and experimentally. Simulations on a computational human torso show that at an output power of 800 mW, the dissipated power is localized to the area under the phased surface with a peak specific absorption rate (SAR) of 8.9 W kg−1, averaged over 10 g of tissue (Fig. 3f). This value is below the 10 W kg−1 threshold, under which no adverse effects have been established25. SAR levels observed for arm and neck configurations are also below this threshold (Supplementary Fig. 11). To assess direct thermal effects, operation of the phased surface on the human body was further studied. Thermal imaging on the human abdomen showed that the temperature increase attributable to the radio-frequency (RF) heating is about 1 °C over 6 min of continuous exposure at 800 mW (Fig. 3g,h), and was comparable on the arm and neck (Supplementary Fig. 12). SAR profiles show that the dissipation is localized near the surface, reaching half the peak value approximately 2 cm from the surface. In applications where the microdevice is powered in a pulsatile rather than continuous manner, such as neural or muscular stimulation, the average exposure will be lowered by the duty cycle. These preliminary studies suggest that the system can safely power microdevices, although further studies will be needed to weigh the risk of RF exposure against potential clinical benefits.
Wireless powering of implanted microdevices
We characterized the performance of the system by wirelessly powering microdevices implanted in adult pig (male, 70 kg) viscera and neck. The microdevice was inserted into the body by tunneling the optical fibre into the peritoneal cavity and the phased surface was placed over the device on the body surface. Initial placement of the phased surface relied on feedback from the microdevice; no imaging guidance was used. Once the surface placement was determined, wireless powering of the device was repeatable and tolerant to placement (±1 cm). Computed tomography showed the relative positions of the phased surface and the implanted microdevice (Fig. 4a,b) in three configurations: (1) right upper quadrant and (2) right lower quadrant of the abdomen, with the microdevice implanted in the peritoneal cavity, and (3) right neck surface, with the microdevice implanted in the carotid sheath. The depth of implantation ranged from 3.8 to 4.1 cm from the surface (Fig. 4c). The power received in each configuration was about 0.6 mW at 800 mW continuous power (Fig. 4d,e), with the profile, traced along the trajectory of the optical fibre, showing localization of energy to a spot with width varying from 1.8 cm (0.096λ) to 2.8 cm (0.149λ; Fig. 4f). The width of the profile measured in the upper abdomen was slightly narrower than the width predicted in homogenous tissue, probably due to small distortion of the field shape across the rib cage. Note that the thicknesses of the porcine tissues evaluated here are significantly greater that the corresponding human average. Optical power densities above 1 mW mm−2 (Supplementary Fig. 4b) can be generated under safe average exposure levels if the duty cycle is below 50%, which meets requirements for use in applications such as optogenetics26.
We also studied the effect of bone structures on energy transfer. Power was wirelessly transferred across the rib cage in the upper abdomen configuration (Fig. 4a). The received power profile (Fig. 4d) showed performance and symmetry comparable to that of the lower abdomen and neck configurations, which do not contain bone. Computational studies showed that the received power decreases by a maximum of 13% when the microdevice is behind a rib-like (1 cm width) structure in otherwise homogenous tissue (Supplementary Fig. 13a,b,d,e). Reflection increases the received power slightly when bone is present behind the microdevice. The robustness of wireless powering to bone is a result of the relatively low impedance mismatch between bone and soft tissue (λbone/λmuscle ≈ 1.6) and diffraction due to the long wavelength (λmuscle = 2.5 cm) relative to the dimensions of the bone. Layered bone structures at thicknesses comparable to the human skull (<1 cm) cause slight broadening of the field shape, although reduced absorption results in an overall increase in received power (Supplementary Fig. 13c,f). These results demonstrate the potential of the phased surfaces to power devices in regions difficult to access through modalities with high bone contrast, such as ultrasound.
We next illustrate operation of the phased surface in vivo by wirelessly powering a miniaturized stimulator for cardiac pacing. The stimulator, shown in Fig. 5a, consists of the microdevice attached to pacing electrodes, yielding a cylindrical device 1.5 mm in diameter and 5 mm in length. The dimensions of the helical coil are consistent with previous experiments. The stimulator was mounted on the tip of a deflectable transvenous catheter and inserted at endocardial stimulation sites in the left ventricle, right atrium andright ventricle (Fig. 5b–d). The stimulation voltage waveform (10 ms width, 600 ms period) was generated by pulsed RF excitation of the phased surface, which produces minimal distortion because of its relatively large bandwidth (60 MHz, Supplementary Fig. 14). Figure 5e shows representative electrocardiogram recordings during 10 s stimulation/rest intervals for each stimulation site. The heart rate is elevated to the target rate following 2 to 4 s of stimulation, immediately returning to baseline once the stimulation pulses cease (Fig. 5f). Different target rates were also achieved at the same site (Supplementary Fig. 15). The minimum average power levels required to successfully pace the heart were 216 mW (left ventricle), 34 mW (right atrium) and 108 mW (right ventricle), which are well below the previously established safety thresholds. The variability in required power is likely due to the different depths of the stimulation sites as well as the excitability of the target tissue. Continuous operation of the device will require methods to achieve long-term reliability, such as through integration of rechargeable energy storage components27. These results, however, indicate that the performance levels are sufficient for microdevices to wirelessly elicit temporary physiological effects from deep organs on demand. Bioelectronic modulation of physiological function through neural stimulation or drug release could provide a therapeutic effect in this mode.
Clinical applications of the system may benefit from a portable or wearable wireless powering source. In contrast to phased arrays, the RF components required to drive the phased surface are minimal and can readily be integrated. Supplementary Fig. 16 shows an example in which commercial RF integrated circuits (oscillator and power amplifier) and a battery are mounted behind a rigid phased surface. The components do not significantly affect the field shape and the focal spot generated from a flat surface is nearly identical to the tethered source. The integrated device can achieve a maximum output power of 1 W, which is sufficient to reach the continuous SAR-limited output. The substrate choice is currently limited by the rigid components; integration of commercial components with overall flexible mechanics represents an important direction for future work.