IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 60, NO. 8, AUGUST 2013 2107

Far-Field RF Powering of Implantable Devices:Safety Considerations

Rebecca A. Bercich∗, Daniel R. Duffy, and Pedro P. Irazoqui, Member, IEEE

Abstract—Far-field RF powering is an attractive solution to thechallenge of remotely powering devices implanted in living tissue.The purpose of this study is to characterize the peak obtainablepower levels in a wireless myoelectric sensor implanted in a pa-tient while maintaining safe local temperature and RF poweringconditions. This can serve as a guide for the design of onboardelectronics in related medical implants and provide motivation formore efficient power management strategies for implantable inte-grated circuits. Safe powering conditions and peak received powerlevels are established using a simplified theoretical analysis andFederal Communications Commission-established limits for radi-ating antennas. These conditions are subsequently affirmed andimproved upon using the finite-element method and temperaturemodeling in bovine muscle.

Index Terms—Far field, implantable, intramuscular, myoelec-tric, RF powering, sensors, specific absorption rate (SAR),wireless.

I. INTRODUCTION

IN the world of chronically implanted sensors and stimula-tors, remotely powered devices offer a number of advantages

over their battery-powered and wired counterparts. Wirelesspowering negates the need for recurrent surgeries to replace bat-teries and sidesteps the obvious issues with lead wires that riskinfection by breaching the skin barrier and can break or dislodgeas a result of fatigue and micromotion. Wireless myoelectricsensors have been previously developed that utilize near-fieldinductive coupling to receive power and communicate with anexternal link [1], [2]. Far-field powering offers some practicaladvantages over inductive powering [3], [4], which requires ac-curate alignment of the implant’s primary and secondary coilsin order to operate efficiently [5]. Understanding the receivepower capabilities of implantable devices that operate in the farfield will help decide the value of far-field powering relative toalternative wireless powering schemes.

One of the limitations of far-field powering and telemetry infree space is that power density decreases as 1/r2 (energy spread-

Manuscript received July 18, 2012; revised October 31, 2012 and January 23,2013; accepted January 27, 2013. Date of publication February 12, 2013; dateof current version July 13, 2013. This work was sponsored by the Defense Ad-vanced Research Projects Research Agency (DARPA) MPO under the auspicesof Dr. Jack Judy through the Space and Naval Warfare Systems Center, Pa-cific Grant/Contract No. N66001-11-1-4029. Asterisk indicates correspondingauthor.

∗R. A. Bercich is with the Biomedical Engineering Department, Purdue Uni-versity, West Lafayette, IN 457906 USA (e-mail: rbercich@purdue.edu).

D. R. Duffy and P. P. Irazoqui are with the Biomedical EngineeringDepartment, Purdue University, West Lafayette, IN 47906 USA (e-mail:dduffy@purdue.edu; pip@purdue.edu).

Color versions of one or more of the figures in this paper are available onlineat http://ieeexplore.ieee.org.

Digital Object Identifier 10.1109/TBME.2013.2246787

ing). Far-field powering of implantable devices introduces theobstacle of power density attenuation as the electromagnetic(EM) waves pass through tissue and tissue boundaries. Absorp-tion and reflection of radiated energy in addition to the freespace path loss—all diminish the power return at the receivingantenna embedded in tissue.

Additionally, there are exposure limits set by the FederalCommunications Commission (FCC) that restrict how muchpower can be incident on the surface or absorbed in a givenvolume of tissue. It will also be important to demonstrate thatlocal temperature spikes resulting from the heating of implantedmicroelectronics will not cause biological tissue damage.

A previously developed application-specific integrated circuit(ASIC) designed to retrieve and wirelessly transmit intramus-cular myoelectric signals will be used throughout the analysisas a case study to explore the potential for far-field RF poweringof implantable sensors. This ASIC has an average power con-sumption of 670 μW and will be wirelessly powered at 2.4 GHz.Additionally, average receivable power for implants using alter-native industrial, scientific, and medical (ISM) band frequencies(900 MHz and 5.8 GHz) will be explored and reported.

II. SAFETY LIMITATIONS

Before the maximum power available to an ASIC embeddedin tissue can be determined, we first establish how much powercan be applied to the tissue surface without exceeding safetylimitations. The first safety consideration for radiating antennasoperating in the RF band is the time-averaged surface exposureof tissue. The corresponding FCC standard for uncontrolledexposure to an intentional radiator operating at a frequency of915 MHz is 6 W/m2 maximum permissible exposure (MPE) and10 W/m2 MPE at 2.4 and 5.8 GHz (as averaged over 30 min).For the purposes of this safety analysis, we consider only far-field EM exposure; this is a valid simplification since near-field magnetic fields propagate through tissue with significantlyless attenuation, heating, etc. The power flux density at somedistance x from the radiating source can be found using thefollowing relationship [6]:

S(x) =EIRP4πx2 (1)

where EIRP is the equivalent isotropically radiated power fromthe antenna (equated to the product of antenna input power Pt ,antenna gain Gt , and the efficiency of the antenna’s matchingnetwork emn ). By setting S(x) equal to the surface exposurelimit (10 W/m2), a relationship emerges between antenna–tissueseparation and the transmitting antenna characteristics that must

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2108 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 60, NO. 8, AUGUST 2013

be satisfied in order to guarantee safe exposure levels:

x =

√PtGtemn

40π(2)

As an example of this, if the powering antenna is isotropic(gain of 0 dB), has a matching network efficiency of 100%, andis given 1 W average input power, then the minimum separationbetween that antenna and any part of the body must be 8.92 cm.

The second safety consideration for antennas operating at RFfrequencies is a set of basic restrictions for power absorptionin tissue that cannot be exceeded. This includes a whole-bodyaverage specific absorption rate (SAR) limit (0.08 W/kg [7])as well as a peak spatial SAR limit evaluated over 1 g cube oftissue (1.6 W/kg [7]). With regards to the spatial peak SAR overany 1 g cube of tissue, the power absorbed by any volume oftissue having a cubic form factor can be expressed as:

Pabs = a2 · (S0 − Sa) (3)

where S0 is the power flux density entering the cube, Sa is theremaining power flux density that exits the cube on the oppositeside, and a is the side length of the cube. Using the averagedensity of human tissue ρ = 1.04 g/cm3 [8], the side lengtha corresponding to a tissue mass of 1 g is calculated to be9.87 mm. It is also known that the power flux density throughtissue decreases exponentially in accordance with the followingrelationship:

Sa = S0 · e−2 aλd (4)

λd =√

1π · f · μ0 · σ

(5)

where f is the powering frequency, μ0 the permeability of freespace (1.257e−6 m·kg/(s2 ·A2), and σ is the conductivity of theexposed tissue at a given frequency. The term λd refers to theskin depth or the depth at which the amplitude of the EM wave isreduced by a factor of e−1 [9]. Thus, if the powering frequencyis set to 2.4 GHz (at which point the average conductivity ofhuman tissue is 1.81 S/m [8]), the skin depth will be 7.64 mm.Substituting (4) into (3) yields an expression for the powerabsorbed by a cube of tissue depending on the power flux densityentering the cube:

Pabs = a2 · S0 · (1 − e−2 aλd ) (6)

If (6) is rearranged and the peak spatial SAR limit substitutedfor Pabs , then the absolute maximum power flux density throughthe tissue can be calculated as follows:

S0 =Pabs

a2 · (1 − e−2 aλd )

(7)

This generates the following fundamental safety rule for RFpowering at 2.4 GHz: average flux density (over any 30-minspan) cannot exceed 10 W/m2 because of the MPE limit andcannot instantaneously exceed 17.76 W/m2 because of localSAR limits. Additionally, (6) can be used to estimate the peaklocal SAR under the condition that the MPE limit has beenreached by substituting 10 W/m2 for S0 . This yields a peak local

SAR of 0.90 W/kg, which will be used later for comparison withthe simulated numerical results.

Using the peak power flux density and the relationship in(1), the corresponding antenna–tissue separation needed to meetlocal SAR limits can be found. If the same powering conditionsas before are applied (isotropic antenna with 100% efficiencyand 1 W input power), this separation is calculated to be 6.69 cm.This is less than the antenna–body separation required to meetthe MPE limits (8.92 cm), so it can be said that these poweringconditions meet all safety standards so long as the antenna–body separation is >8.92 cm. This series of calculations canbe performed on any set of powering conditions in order toestablish the minimum antenna–body separation or input powerthat will satisfy the FCC RF powering safety limits.

As a practical example of this, imagine that the desiredantenna–body separation is 1 cm; the average input power tothe isotropic antenna would need to be reduced by −19 dB inorder to satisfy the MPE limit and the free-space loss would,likewise, decrease by −19 dB. For an isotropic antenna, thiswould be a good power-saving strategy; however, for anten-nas with directive gain moving, the antenna closer to the bodymay significantly restrict the volume of tissue in which a de-vice can receive adequate power. For this reason, the radiationpattern, antenna–tissue separation, input power, and exposurelimits must all be taken into consideration when designing apractical RF powering system for implantable devices.

The final safety consideration is in regards to focalized tem-perature increases in tissue near the implant as a result of heatdissipation from onboard microelectronics. It is important thatthe temperature of the tissue surrounding the ASIC does notincrease more than 1–2 ◦C. Various studies have shown that atemperature increase of this magnitude has adverse effects onliving cells [10]–[16].

III. THEORETICAL LINK BUDGET

It will be useful to apply the previously established safe pow-ering conditions to the task of RF powering an IC within thebody to gain a better understanding of the limits to receivablepower. For the purposes of this theoretical analysis, we assumethat all power incidenting on the tissue surface is absorbed. Byusing the known average power flux-density limit at tissue sur-face S0 (10 W/m2) and (4), the average power flux density at thedepth of the implant S2 can be found. The amount of useablepower received by the circuit (after rectification) Pr will dependon this power flux density as well as the efficiency of the con-version circuit erec , the efficiency of the matching network emn ,and the receiving antenna gain Gr , according to the followingrelationship [6]:

Pr = S2erecemnGrλ

2

4π(8)

Clearly, the level of power received by the on-board elec-tronics of an implanted device can be greatly influenced by theRF-to-DC conversion circuit efficiency as well as the gain char-acteristics of the receive antenna. It is worth noting that conver-sion circuit efficiency tends to increase as the amplitude of the

BERCICH et al.: FAR-FIELD RF POWERING OF IMPLANTABLE DEVICES: SAFETY CONSIDERATIONS 2109

input RF power increases. This behavior has been observed inpreviously developed rectenna (rectification circuit plus receiv-ing RF antenna) assemblies designed to operate at 2.45 GHz,which have demonstrated RF-to-DC conversion efficiencies thatrange from 48% at −5 dBm [17] to 15.7% at −20 dBm [18]. Todemonstrate the practical use of (8), a previously developed mul-tistage CMOS rectification circuit designed for an implantablemyoelectric sensor having a demonstrated conversion efficiencyof 20% at −10 dBm will be used to calculate the theoretical re-ceive power levels.

Receive antenna gain can also vary significantly based on thegeometry and material of the antenna [17]. Due to the size con-straints of small implantable devices, it may not be physicallypossible to use the most efficient antenna design (particularlyfor lower frequencies that require larger dimensions to achieveresonance). For the purposes of estimating receive power levels,a worst case estimate of −20 dB gain will be used.

Assuming an ideal matching network efficiency of 100%, aconversion circuit efficiency of 20%, and a receive antenna gainof −20 dB, the average received power of a device implanted at1 cm into tissue and operating at 2.4 GHz is 1.88 μW.

These calculations can be reversed to determine the averagepower flux density incident on the tissue surface necessary tocontinuously power a device having a known power consump-tion. The average power consumption of this ASIC is 670 μW.The corresponding average power flux density that would beneeded at the surface of the tissue to continuously power thisdevice is 723 W/m2 . This corresponds to a peak spatial SAR of65 W/kg. This device clearly consumes too much power to besafely and continuously powered at any significant depth withintissue but serves to illustrate the means by which an implantableIC with known power consumption can be assessed for safe use.

One strategy that can be used to significantly reduce ASICpower consumption is power cycling in combination with astorage source. In this way, the device can be designed with avery low duty cycle (for example, powering on for 1 μs out ofevery 1 ms) while wirelessly coupled power is stored during theother 999 μs.

IV. FEM ANALYSIS

The preceding theoretical analysis establishes a good conser-vative approximation for safe powering conditions but is lim-ited by the assumptions imposed. By utilizing software capableof simulating three-dimensional (3-D) EM fields, the safety ofthese conditions can be affirmed with greater certainty and theactual power dispersion can be modeled and analyzed. High-Frequency Structure Simulator 13 (HFSS 13.0) from ANSYS isused to replicate the theoretical powering conditions and simu-late the resulting E-fields.

The simulation environment is set up using a half-wave(length = 1/2λ) center-fed dipole antenna placed parallel tothe z-axis and a three-layered body model representing the wallof the human chest. The three layers of the model are skin, adi-pose tissue, and muscle layers, with thicknesses and dielectricproperties based on previous data [8], [19], [20]. This model wasplaced one wavelength λ, along the y-axis from the center of

Fig. 1. Simulation environment from the perspective of the positive x-axis ata powering frequency of 2.4 GHz. Dipole–tissue separation is one wavelength λ

(328 mm at 915 MHz, 125 mm at 2.4 GHz, and 51.7 mm at 5.8 GHz). Skin layerthickness = 2 mm, adipose layer thickness = 8.4 mm, muscle layer thickness =21.6 mm. Tissue’s height and width were both twice the dipole–tissue separation

Fig. 2. (a) Radiation pattern when ϕ = 90◦ and θ is swept from 0◦ to 180◦.(b) Radiation pattern when θ = 90◦ and ϕ is swept from 0◦ to 180◦. Due to itsdimensions and separation from the antenna, the block of tissue lies within theregion defined by 45◦ ≤ ϕ ≤ 135◦ and 45◦ ≤ θ ≤ 135◦. Thus, the maximumEIRP in the direction of the tissue will be the local maximum −2.3029 dB.

the antenna so that it resided in the far field of the antenna’s EMfield. A side view of the system in the simulation environmentis given in Fig. 1.

The antenna shown in Fig. 1 radiates isotropically in thexy-plane in free space. However, its radiation pattern changeswhen objects are introduced into its field. The previously definedsafety conditions depend on the radiating body having an EIRPof 1 W in the direction of the tissue; therefore, it was necessary tomodel the gain of the antenna within the simulation environmentand adjust the input power accordingly. The radiation patternof the antenna in the simulation environment at a poweringfrequency of 2.4 GHz is shown in Fig. 2. The peak gain in thedirection of the tissue is Gt = −2.3029 dB and lies at ϕ = 90◦

and θ = 90◦ (in the direction of the positive y-axis). Resultantly,the input power to the system Pt was set to 1.7 W to satisfy thecondition that EIRP = PtGt = 1 W. Oftentimes, the resonanceof an antenna will shift when objects are introduced to or movedwithin the radiated field. For this reason, the S11 parameter ofthe dipole antenna was assessed and found to have a bandwidth

2110 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 60, NO. 8, AUGUST 2013

Fig. 3. Poynting vector field (directional power flux density) in tissue at theair–tissue boundary with 1 W EIRP output from half-wave dipole at a poweringfrequency of 2.4 GHz.

of 2.06–2.48 GHz, which means it is still an effective radiatorfor RF powering at 2.4 GHz in the simulation environment.

The distribution of the power flux density incident on thetissue can be imaged by plotting the real part of the Poyntingvector in air at the air–tissue boundary plane. It is expectedthat the peak incident power flux density will be less than theFCC limit (10 W/m2) since the antenna–tissue separation is>8.92 cm. The actual simulated maximum is 8.11 W/m2 , whichindicates that the safe conditions established by the conservativetheoretical analysis can be replicated and affirmed in the simula-tion environment. One of the benefits of using a high-frequencysimulator is that it can easily and accurately model the reflec-tion of energy at the boundary of two materials having uniquedielectric properties. The power flux density at the air–tissueboundary in the skin (after initial reflection) is shown in Fig. 3.The maximum power flux density that passes through the skinis 6.56 W/m2 . This is well below the MPE limit, verifying thatconditions chosen using conservative estimates will satisfy theaverage surface exposure constraint.

It will also be of interest to calculate the maximum powerabsorbed by 1 g of tissue to determine how well these condi-tions satisfy the basic restriction of peak spatial SAR. HFSS hasa built-in average SAR tool that is useful for single-dielectricmodels but is not accurate for evaluating absorption in volumesthat span multiple dielectrics and include the boundaries be-tween those dielectrics. Instead, the peak local SAR will beevaluated using the software’s fields calculator. This is doneby first identifying the magnitude and coordinates of the peaklocal SAR. Unsurprisingly, this maximum is found to occurat the origin of the axes on the surface of the skin where thedirectional gain of the antenna is largest. For the purposes ofevaluating the total energy absorbed by a 1 g cube of tissue perFCC specifications, the following equation is used:

SARpeak =Pin − Pout

M(9)

where Pin is the total power entering a cube with mass M andPout is the total power leaving the cube. The net power through

a single face on the cube Pnet,F can be found by taking the dotproduct of the Poynting vector S and the unit vector normal tothe face, and integrating it over the area of the cube face F asfollows:

Pnet,F =∫

F

(S · n̂)dF (10)

A 1 g cube of tissue is distinguished in the simulation en-vironment using six square planes. Each face of the cube isassigned a number as follows: face 1 (F1) lies on the xz-planeand transects the y-axis at the origin, face 2 (F2) transects thex-axis at x = a/2, face 3 (F3) transects the y-axis at y=a, face4 (F4) transects the x-axis at x = −a/2, face 5 (F5) transectsthe z-axis at z = a/2, and face 6 (F6) transects the z-axis at z =−a/2. Resultantly, the total power absorbed in the cube Ptot willbe:

Ptot = Pin − Pout = Pnet,F1 + Pnet,F4 + Pnet,F6

− Pnet,F2 − Pnet,F3 − Pnet,F5 (11)

The field calculator in HFSS can be used to calculate thesequantities. By substituting the result of (11) into (9),the peakabsorbed power in 1 g of tissue was evaluated to be 0.36 W/kg,which is much less than the absolute limit of 1.6 W/kg andabout one-third of the analytic estimate of 0.90 W/kg foundusing (6). This discrepancy between the analytic and simulatedpeak local SAR can be attributed to the observance of the re-flection phenomenon as well as individual tissue properties inthe HFSS environment. In order to reach the peak spatial SARlimit, the flux density in skin at the air–tissue boundary wouldhave to be 29.14 W/m2 (much greater than the 17.76 W/m2 esti-mated by the conservative analysis). For this reason, 29.14 W/m2

can be established as the absolute maximum allowable powerflux when powering at 2.4 GHz. However, due to MPE limitsthe time-averaged exposure over 30 min must remain below10 W/m2 when powering at 2.4 GHz. Table I summarizes themaximum average power flux density at the implant and theresulting maximum average power available to an implant’s ICunder the condition that MPE limits have been reached in thesimulation environment. This is shown for three potential pow-ering frequencies within the ISM bands.

V. LOCAL TEMPERATURE

Evaluation of local temperature spikes due to heat diffused byan active implantable IC is performed by dispersing various lev-els of power in 16.4 cm2 samples of bovine muscle, which hasbeen demonstrated to have similar permittivity characteristicsto that of human muscle at the chosen powering frequency [21].This is achieved by supplying a range of voltages across a 1,003-Ω resistor fitted within a medical epoxy package having dimen-sions of 11.85 mm × 5.50 mm × 3.95 mm (similar to theform factor of previously developed implantable myoelectricsensors [1], [2]) implanted at a depth of 10 mm into the tissuesamples. All the samples are maintained at body temperaturethroughout the experiment using an incubator at 37 ◦C. The de-vice is continuously powered for approximately 15 min until thetemperature adjacent to the device stops increasing and reaches

BERCICH et al.: FAR-FIELD RF POWERING OF IMPLANTABLE DEVICES: SAFETY CONSIDERATIONS 2111

TABLE IPOWER FLUX DENSITY AT FOUR IMPLANT DEPTHS AND CORRESPONDING

MAXIMUM AVERAGE POWER AVAILABLE TO THE IMPLANT’S ICAT 915 MHz, 2.4 GHz, and 5.8 GHz

Fig. 4. Temperature increase in bovine muscle immediately adjacent todummy sensor consuming 10, 20, 30, and 40 mW. Experiment was repeatedthree times per power consumption value (n = 3).

steady state. Tissue temperature is recorded using a thermocou-ple and LabVIEW data acquisition program. The increase inlocal temperature due to heating of the sensor microelectronicscan be seen in Fig. 4.

The results of the local temperature measurements indicatethat up to 40 mW of power can be expelled by an active sensorin muscle and the local temperature will not increase more than1 ◦C, which prior work has demonstrated to be a threshold forthermal damage in tissue [10]–[16]. Since 40 mW is far morepower than can be safely supplied to an implant at any significantdepth (see Table I) and more than the power requirement of thepreviously developed ASIC (670 μW), unsafe elevation in localtemperature surrounding the implant will not be a concern.

VI. CONCLUSION

A theoretical analysis and FEM modeling of tissue exposedto EM fields have generated a relationship between safe far-fieldRF powering conditions and the usable power received by animplant depending on the depth within the tissue. For an IC im-planted 1 cm into a muscle of the chest and an external transmitpower of 1 W at 2.4 GHz, the maximum average receive powercan be estimated at 3.90 μW. It is possible that microelectronicswith higher power consumptions such as the myoelectric sen-sor ASIC having a continuous power consumption of 670 μWcould still feasibly be powered at this depth if effective powercycling at low duty cycles is implemented on chip. Additionally,focalized thermal damage of tissue should not be a concern solong as IC power consumption does not exceed 40 mW.

Assuming constant receive antenna gain, the average powerreceived by a device can be increased significantly by reducingthe powering frequency. In practice, this would require increas-ing the antenna dimensions and, resultantly, the overall size ofthe device. This illustrates one of the fundamental challengesof RF powering subcutaneous devices and that is the tradeoffbetween powering frequency and antenna dimensions needed toachieve a given antenna efficiency.

A powering frequency of 2.4 GHz was used throughout thisanalysis to demonstrate a compromise between the objectivesof reduced energy absorption in tissue and minimizing devicedimensions. However, for devices that demand smaller formfactors, it may be prove to be more efficient to increase thepowering frequency and accept a greater power loss throughtissue absorption in exchange for better antenna efficiency.

The significant losses that occur in tissue as well as throughthe receive antenna and conversion circuit will demand that rig-orous power management strategies be employed on all activesensor microelectronics that utilize far-field RF powering. How-ever, receive power levels can be improved by optimizing thereceive antenna gain, improving the efficiency of the conver-sion circuit, or powering at lower frequencies. For any givenimplantable device application in which wireless powering isemployed, the benefits and tradeoffs in selecting a certain pow-ering frequency and antenna design will need to be carefullyconsidered. This will be crucial in generating the most efficientsolution that also meets the top-level device requirements.

REFERENCES

[1] G. E. Loeb, R. A. Peck, W. H. Moore, and K. Hood, “BIONTM systemfor distributed neural prosthetic interfaces,” Med. Eng. Phys., vol. 23,pp. 9–18, 2001.

[2] R. F. Weir, P. R. Troyk, G. A. DeMichele, D. A. Kerns, J. F. Schorsch, andH. Maas, “Implantable myoelectric sensors (IMESs) for intramuscularelectromyogram recording,” IEEE Trans. Biomed. Eng., vol. 56, no. 1,pp. 159–171, 2009.

[3] E. Y. Chow, C.-L. Yang, Y. Ouyang, A. L. Chlebowski, P. P. Irazoqui,and W. J. Chappell, “Wireless powering and the study of RF propaga-tion through ocular tissue for development of implantable sensors,” IEEETrans. Antennas Propagation, vol. 59, no. 6, pp. 2379–2387, Jun. 2011.

[4] E. Y. Chow, C. Yang, and P. P. Irazoqui, “Wireless powering and propa-gation of radio frequencies through tissue,” in Wireless Power Transfer.Aalborg, Denmark: River Publishers, 2012, ch. 9.

[5] M. Catrysse, B. Hermans, and R. Puers, “An inductive power system withintegrated bi-directional data-transmission,” Sens. Actuators A, Phys.,vol. 115, pp. 221–229, 2004.

2112 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 60, NO. 8, AUGUST 2013

[6] D. M. Pozar, Microwave and RF Design of Wireless Systems. Hoboken,NJ, USA: Wiley, 2001.

[7] F. C. Commission, Part 15—Radio frequency devices (47 CFR 15), Title47 of the code of federal regulation, Apr. 9, 2012.

[8] C. Gabriel, “Compilation of the dielectric properties of body tissues atRF and microwave frequencies,” Occup. Environ. Health Directorate, TX,USA, Rep. AL/OE-TR-1996-0037, 1996.

[9] J. S. Seybold, Introduction to RF Propagation. New York, NY, USA:Wiley, 2005.

[10] W. C. Dewey, L. E. Hopwood, S. A. Sapareto, and L. E. Gerweck, “Cellu-lar responses to combinations of hyperthermia and radiation,” Radiology,vol. 123, pp. 463–474, 1977.

[11] T. M. Seese, H. Harasaki, G. M. Saidel, and C. R. Davies, “Characteri-zation of tissue morphology, angiogenesis, and temperature in the adap-tive response of muscle tissue in chronicheating,” Lab. Investig., vol. 78,pp. 1553–1562, 1998.

[12] M. Ueda, J. Bures, and J. Fischer, “Spreading depression elicited bythermal effects of ultrasonic irradiation of cerebral cortex in rats,” J.Neurobiol., vol. 8, pp. 381–393, 1977.

[13] T. Fujii and Y. Ibata, “Effects of heating on electrical activities of guineapig olfactory cortical slices,” Eur. J. Physiol., vol. 392, no. 3, pp. 257–260,1982.

[14] P. S. Ruggera, D. M. Witters, G. von Maltzahn, and H. I. Bassen, “In vitroassessment of tissue heating near metallic medical implants by exposureto pulsed radio frequency diathermy,” Phys. Med., vol. 48, pp. 2919–2928,2003.

[15] J. G. Nutt, V. C. Anderson, J. H. Peacock, J. P. Hammerstad, andK. J. Burchiel, “DBS and diathermy interaction induces severe CNSdamage,” Neurology, vol. 56, pp. 2919–2928, 2001.

[16] J. C. LaManna, K. A. McCraeken, and M. Patii, “Stimulus-activatedchanges in brain tissue temperature in the anesthetized rat,” Metab. BrainDis., vol. 4, pp. 225–237, 1989.

[17] J. Heikkinen and M. Kivikoski, “Performance and efficiency of planarrectennas for short-range wireless power transfer at 2.45 GHz,” Microw.Opt. Technol. Lett., vol. 31, pp. 86–90, 2001.

[18] G. Andia Vera, A. Georgiadis, A. Collado, and S. Via, “Design of a2.45 GHz rectenna for electromagnetic (EM) energy scavenging,” in Proc.IEEE Radio Wireless Symp., pp. 61–64.

[19] R. McLean, M. Richards, C. Crandall, and J. Marinaro, “Ultrasound de-termination of chest wall thickness: implications of needle thoracostomy,”Amer. J. Emerg. Med., vol. 29, pp. 1173–1177, 2011.

[20] T. Sumerling and S. Quant, “Measurement of the human anterior chest wallby ultrasound and estimates of chest wall thickness for use in determinationof transuranic nuclides in the lung,” Radiat. Protect. Dosimetry, vol. 3,pp. 203–210, 1982.

[21] M. M. Brandy, S. A. Symons, and S. S. Stuchly, “Dielectric behavior ofselected animal tissues in vitro at frequencies from 2 to 4 GHz,” IEEETrans. Biomed. Eng., vol. BME-28, no. 3, pp. 305–307, Mar. 1981.

Authors’ photographs and biographies not available at the time of publication.