Ultrasonics 50 (2010) 556–566

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Ultrasonics

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Ultrasonic transcutaneous energy transfer for powering implanted devices

Shaul Ozeri, Doron Shmilovitz *

School of Electrical Engineering, Tel-Aviv University, Tel-Aviv 69978, Israel

a r t i c l e i n f o

Article history:Received 7 August 2009Received in revised form 12 November 2009Accepted 13 November 2009Available online 26 November 2009

Keywords:Transcutaneous energy transferPowering implanted devicesUltrasonic energyAcoustic impedance matching

0041-624X/$ - see front matter � 2009 Elsevier B.V.doi:10.1016/j.ultras.2009.11.004

* Corresponding author.E-mail address: shmilo@eng.tau.ac.il (D. Shmilovit

a b s t r a c t

This paper investigates ultrasonic transcutaneous energy transfer (UTET) as a method for energizingimplanted devices at power level up to a few 100 mW. We propose a continuous wave 673 kHz singlefrequency operation to power devices implanted up to 40 mm deep subcutaneously. The proposed UTETdemonstrated an overall peak power transfer efficiency of 27% at 70 mW output power (rectified DCpower at the load).

The transducers consisted of PZT plane discs of 15 mm diameter and 1.3 mm thick acoustic matchinglayer made of graphite. The power rectifier on the implant side attained 88.5% power transfer efficiency.

The proposed approach is analyzed in detail, with design considerations provided to address issuessuch as recommended operating frequency range, acoustic link matching, receiver’s rectifying electron-ics, and tissue bio-safety concerns. Global optimization and design considerations for maximum powertransfer are presented and verified by means of finite element simulations and experimental results.

� 2009 Elsevier B.V. All rights reserved.

1. Introduction

The last few decades have seen an enormous increase in thenumber and variety of medical devices. According to the AmericanNational Institute of Health (NIH), new devices are being added tothe market every year. Some of these devices are designed to beimplanted into the body for monitoring purposes such as biosen-sors, glucose indicators. Other implanted devices serve therapeuticpurposes, e.g. pacemakers, defibrillators, heart-assist devices, orminiature mechanisms for the controlled delivery of medications,such as implanted insulin pumps. All of these devices require elec-trical energy for their operation. While the majority of modern im-planted devices consume low power (in the range of hundreds ofmW), up to 10 W of power may be required in some cases.

In recent years some very interesting methods for intra-bodyelectrical energy sources have been reported. Refs. [1,2] discussthe generation of electrical energy by biological fuel cells that con-sume inter-cellular glucose. The fuel cells generate a voltage ofabout 0.5 V with a power density of 50 lW/cm2. Another work[3] suggests an intra-body ‘‘energy harvesting” system. Accordingto this approach, the bio-mechanical energy created by the move-ment of internal organs, such as the peristaltic movement of theintestines or the contraction of the heart and lungs, is convertedinto electrical energy using piezoelectric elements. Evidently, withthese techniques, an average power below 1 mW can be generated.Devices that consume average power of a few (<10 W) Watts arepowered wirelessly by a mechanism based on the penetration of

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z).

energy through the tissue, a technique called transcutaneous en-ergy transfer (TET).

Presently, existing TET devices rely on electromagnetic energytransmission and customarily employ a pair of flat spiral coils[4–6], one external energy transmitting coil and another intra-body receiving coil, with the two coils facing each other. The inter-nal coil is implanted under the fat layer, typically at a depth of 10–20 mm.

Until now, published papers have dealt with electromagneticTET mechanisms that were capable of transferring power throughthe tissue at power levels of up to 10 W. This range of power issuitable for devices such as heart pumps or cardiac assist devicesthat increase the volume of blood pumped around the body [7].These devices are well-suited to electromagnetic TET-based mech-anisms. As mentioned, the electromagnetic transmission is basedon two coils facing each other. The separation between the coilsvaries from 10 to 20 mm depending on the skin/tissue thickness.Furthermore, the distance and relative orientation of the coilsmay vary as the tissue moves or deforms. Since there is no ferro-magnetic core to concentrate the magnetic flux and close the mag-netic circuit, the magnetic coupling coefficient k between the twocoils is typically low (�0.1). The low coupling requires a primarycurrent of several Amps which reduces the efficiency and necessi-tates batteries with a 1–5 A current drain capability.

Several papers have discussed the need to increase the effi-ciency of the coupling [4,8]. These authors discussed the difficultiesencountered in this effort; among them, the need to increase theexcitation current, which causes excess heating in the primary coiland in turn leads to an increase in tissue temperature. In reference[4], the authors suggested a coil geometry that gave a coupling of

S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566 557

k = 0.4. An electromagnetic TET device intended for a miniature im-planted Bio-MEMS sensor was presented in [8] which detailed thedesign of the RF link at 330 MHz. The diameter of the external coilwas 75 mm, but the implanted coil was miniaturized (1 mm2) andbuilt into the sensor device itself. The received power reachedabout 250 mW.

Clearly, Electromagnetic TET is adequate for transferring energyin the Watt range, but suffers from some drawbacks such as lowcoupling (�0.1) which might cause interference with the operationof devices working in close proximity, such as pacemakers. Con-versely, the operation of the TET itself may be perturbed, as its coilmay pick up external electromagnetic radiation created by sourcessuch as MRI devices in hospitals, or as a result of the nearby pres-ence of massive ferromagnetic objects, such as steel doors.

For devices consuming large amounts of power, the electromag-netic TET mechanism is probably still the most suitable in spite ofthe shortcomings mentioned above. However, for the power rangeof tens of mW (to power smart devices such as those of the Bio-MEMS devices), Ultrasonic TET (UTET) may be a preferable tech-nology due to its power transfer efficiency, compactness, and elec-tromagnetic immunity.

This paper proposes a UTET device in which the energy is trans-mitted via an acoustic wave as previously suggested to stimulatebone healing [5]. The UTET operates at a Constant Wave single fre-quency 673 kHz, and transfer 70 mW to a DC load. Similar to elec-tromagnetic-based TET, there is an implanted intra-body receiverand a transmitter external to the body. However, in contrast toelectromagnetic TET, the usual coils are replaced with ultrasonicpiezoelectric transducers. In this system, the external electricalpower is converted to a pressure wave, which is then transmittedtranscutaneously. The acoustic energy is collected by an implantedtransducer that reconverts the received acoustic energy back intoelectrical energy for use in the medical device.

In fact, the idea of using acoustic waves to transmit energy wasproposed long ago in 1958 by Rosen et al. [9]. The author describeda piezoelectric transformer consisting of a primary piezoelectricsection acoustically coupled to a secondary piezoelectric sectionto reconvert the elastic waves back to electrical energy. With acontrolled interface medium, piezoelectric transformers can deli-ver throughput powers of tens of Watts with a conversion effi-ciency as high as 98%. UTET can also serve in non-medicalapplications such as for powering embedded sensors within metalconstructions [11] (as in use in bridges), for energizing pressureleak sensors in satellites, or for powering sensors embedded insidefuel tanks [12].

The principal design considerations of UTET are discussed, suchas operating frequency, piezoelectric material selection, acousticimpedance matching, power conditioning and safety issues.

2. Acoustic power transfer concept

Energy transfer through the skin requires an external trans-ducer (transmitter) attached to the skin surface facing an im-planted transducer (receiver), as illustrated in Fig. 1. An electricalpower source energizes the transmitter that converts the electricalenergy into acoustic pressure waves. The acoustic waves carry theenergy through the tissue toward an implanted receiver positionedwithin the radiation lobe of the transmitter. The piezoelectric re-ceiver converts the acoustic energy back into electric energy anda low loss (efficiency >80%) rectifier network rectifies and filtersthe output voltage of the receiving transducer, while reflecting atits input terminals an impedance conjugate to that of the piezo-electric element. In order to minimize the inconvenience causedto the patient as well as to ensure a close fit to the body (whichis required for good acoustic coupling), the device should be

lightweight and thin so that its center of gravity is as close as pos-sible to the surface of the body. A UTET system is depicted in Fig. 1.The power throughput in the present application is limited to themW range (Pload < 100 mW) in order to avoid tissue damage. Sucha UTET is powered from a small coin like, low capacity (<500 mA h)battery that is typically characterized by low drain current capabil-ity (<50 mA).

In order to preserve the battery capacity, the current consumedby the switching amplifier (which feeds the transmitter element)should be as smooth as possible (i.e., with a crest factor close to1) [13]. A Maximum Power Extracting (MPE) circuit was designedfor both functions: extracting most of the power generated by theimplanted transducer and increasing the receiver-generated volt-age to a level suitable for powering typical loads, such as logicand analog circuitry (Vdc > 3 V).

The UTET link’s power transfer efficiency is affected by theswitching amplifier losses, transducer losses, tissue absorption,acoustic impedance matching layer losses, rectifier losses, andthe amount of power captured by the receiver. PZT (Lead, Zircon-ate, Titanate) materials are preferred choice for the implementa-tion of the transducers compared to piezoelectric polymerPolyvinylidene Fluoride PVDF such as the Piezoflex (manufacturedby Airmar technology corporation, Milford New Hampshire, USA)due to properties such as high electromechanical coupling andmechanical Q factor (k33 = 0.76, Qm = 1200 compared to k = 0.3,Qm < 25 of PVDF). Using PZT results in high electromechanical en-ergy conversion efficiency (some applications such as piezoelectrictransformers achieve up to 95%), and low excitation voltage(<15Vpk for 100 mW power transfer). Since the acoustic impedanceof soft tissue is much lower than that of a PZT transducer (1.5 MRa-yls in comparison to about 30 MRayls for PZT), acoustic impedancematching is required in order to minimize pressure wave reflec-tions and consequent generation of standing waves within thetissue.

3. Pressure field shape

According to the Huygens principle, each point on the face ofthe radiating transducer may be treated as an independent pointsource of radiation. The acoustic field pattern in front of the trans-ducer’s face is the vector sum of contributions from all pointsources [10]. In the general case, the pressure field at an observa-tion point L(x, y, z) is given by the Rayleigh integral [10]:

Pðx; y; z; tÞ ¼ q0

ZS

_up x0; y0 : t � rc0

� �2pR

dS ð1Þ

where (x0, y0) are the coordinates of a point source on the trans-ducer; (x, y, z) are the coordinates of the observation point in frontof the transducer; _up is the vibration velocity function over thetransducer’s radiating cross section; c0 is the average speed ofsound (wave’s phase velocity) in the medium; q0 is the averagemedium density; and S is the transducer area.

R is the distance from the point source to the observation pointL(x, y, z):

R ¼ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiðx� x0Þ þ ðy� y0Þ2 þ z2

qð2Þ

The Rayleigh integral is difficult to solve in the general case ofan arbitrary transducer shape and non-uniform vibration distribu-tion. However, in the case of UTET which employs a ContinuousWave (CW) sinusoidal excitation, uniformly distributed over theface of a disc-shape transmitter as illustrated in Fig. 2, Eq. (1)may be simplified (taking advantage of certain symmetry features)to give Formula (3) for the pressure field at the observation point L:

PZTTransducer

Matchinglayers

MPENetwork

Load

Implanted Unit

PZTTransducer

Low Loss Rectifier

Thin coupling layer (Castor Oil,

Ultrasonic Jell etc .)

P

5 − 50mm

P

Ultrasonic Traveling Waves

Battery

SwitchingAmplifier

H Bridge

Passive Network

t

+Vo

Fig. 1. UTET system.

558 S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566

Pðx; y; z : tÞ ¼ jkq0c0u0

2p ejxtZ

S

e�jkR

RdS ð3Þ

where R is the distance from the infinitesimal point source to theobservation point; u0 is the vibration velocity amplitude; k is thewavelength of the pressure wave in the medium; c0 is the phasevelocity of the wave; q0 is the density of the medium; x is theangular frequency; and k ¼ x=c ¼ 2p=k is the wave number.

Actually, the wave number k should be complex for a dissipa-tive medium such as tissue, k2 ¼ b2 þ a2 where b is the phasespeed b ¼ x=c and a is the absorption coefficient. Assumingr� a and CW sinusoidal excitation, the solution of (3) is:

Pðr; h : tÞ ¼ jaq0c0u0

rejðxt�krÞ J1ðka sin hÞ

sin hð4Þ

where a is the transducer radius; k is the wave number; J1 is the 1storder Bessel function; u0 is the vibration velocity; r is the distancebetween the center of the radiating disc and the observation point;and h is the angle formed between the observation point to theacoustic axis (h = 0).

The pressure directivity of the transducer is defined as the ratiobetween the pressure at any angle h relative to that at the acousticaxis (h = 0) for the same range [10] (see Fig. 2).

DðhÞ ¼ 2J1ðka sin hÞka sin h

ð5Þ

Fig. 2. Calculation of the pressure field generated by a disc-shaped source.

The field should be directed toward the receiver to let the recei-ver capture most of the radiated energy. The directivity depends onka ¼ 2pa

k which is the ratio of the transducer’s perimeter to wave-length (Fig. 3). As an example, for a receiver with a = 7.5 mm lo-cated 20 mm from the transmitter, the radiated pressure shouldbe confined to within � tan�1 7:5

20

� �¼ �20:5�.

If ka is relatively small, i.e. when the transducer’s perimeter isnot large enough compared to the wavelength, the directivity ofthe field is poor, and the field diverges like a spherical wave. Thismeans that only a fraction of the pressure wave energy can be cap-tured by the receiver.

A 2D finite element simulation (Comsol Multiphysics, ComsolAB, Stockholm, Sweden) of the pressure intensity profile generatedby a disc transducer is illustrated in Fig. 4, which shows the direc-tivity of the radiated pressure wave. (The wave’s intensity dependson the pressure squared and is referred to later on.) The pressurefield generated by a uniform CW excited disc transducer can be di-vided into three zones (Fig. 5). The first zone, the Near Field (NF)zone, is the one closest to the transducer. Within this zone, thepressure field envelope oscillates, and has multiple minima andmaxima which make the power transfer unpredictable. After theNF, the pressure field converges to a natural focus. This intervalis the preferred distance at which to locate the receiver. The NearField distance L from the transducer depends on the transducer’sradius a and on the acoustic wavelength k, in the medium throughwhich the wave propagates [14], and is given by

L ¼ ð2aÞ2 � k2

4k� a2

kFor ð2a2Þ � k2 ð6Þ

Beyond the zone of the natural focus starts the Far Field (FF) re-gion, where the pressure field becomes a spherically spreadingwave with little internal structure, whose intensity decays withdistance.

The points on the acoustic axis at which the pressure peaks de-pend on the wavelength and on the transmitter’s cross section, isgiven by Eq. (7), where m stands for the order of the pressure peak.

XmaxðmÞ ¼ð2aÞ2 � kð2mþ 1Þ2

4kð2mþ 1Þ m ¼ 1;2;3; . . . ð7Þ

In cases where the receiver has to be implanted within the NFrange, it is possible to either insert a bag (1–5 mm thick) contain-ing acoustic jell, water, oil, etc., between the transmitter and theskin, or tune the vibration frequency so as to place the closer radi-ation pressure maxima of the NF, exactly on the receiver, see Fig. 6.

Fig. 3. Pressure amplitude directivity of a disc-shaped transducer for various radii a, at a constant frequency of 500 kHz.

Fig. 4. 2D Comsol Multiphysics simulation of a circular transducer’s intensity profile (radius a = 7.5 mm, k = 2.23 mm).

Z

d

Fig. 5. Preferred location of implanted receiver.

Fig. 6. Attaining maximum pressure at the receiver surface by frequency tuning.

S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566 559

4. Frequency selection

Proper determination of the operating frequency is of greatimportance since it affects factors such as tissue attenuation (witha frequency dependence of f1–f1.4 [30]), transducer thickness, dis-tance of the natural focus (Rayleigh distance), and the size of thereactive elements, e.g. inductors and capacitors [13] in both the re-ceiver and transmitter. In order to maximize the power throughputof the transducer, it is best to operate the transducer close to itsresonance frequency. Resonance frequencies are determined bythe transducer’s geometry and material constants. Thickness vibra-tion resonance occurs at frequencies in which the transducer’s

thickness equals an odd multiplication of half acoustic wave-lengths [15]. The 1st resonance frequency of a disc-shaped trans-ducer vibrating in the thickness vibration mode (Fig. 7) isdetermined by its thickness t and its frequency constant Nt

[m Hz], as given by

2λ

Fig. 7. First thickness vibration mode of piezoelectric transducer.

560 S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566

fr ½Hz� ¼ Nt

tð8Þ

For example, for C-2 material (manufactured by Fuji CeramicsCorporation, Tokyo, Japan) Nt = 2020 m Hz. Thus, according to Eq.(8), increasing the frequency results in a thinner device. The aver-age wave intensity I can be shown to obey Eq. (9) (by integration ofinstantaneous pressure multiplied by instantaneous particle veloc-ity [14,30]):

IWm2

� �¼

P2pk

2Zð9Þ

where Z is the acoustic impedance of the medium (tissue) throughwhich the wave propagates.

The operating frequency also determines the amount of acous-tic energy transformed into heat by loss mechanisms in the tissue.Although skin and the underlying soft tissue layer have acousticimpedances and phase velocities that are close to those in water,the attenuation of the pressure field by tissue is much larger thanwater attenuation (soft tissue 0.6–1.5 dB/cm compared to water0.002 dB/cm at 1 MHz) [29–31], and increases as the frequencyand distance increase. Since the intensity depends on the pressuresquared therefore it decreases at twice the rate compared to thepressure decreasing rate:

Id ¼ Ioe�2ad ð10Þ

where I0 is the intensity at the transducer’s radiating surface, a isthe pressure attenuation coefficient per unit length, and Id is theintensity at a distance d from the transducer’s radiating surface.

0 0.5 1 1.5 2 2.50

2

4

6

8

10

12

14

T

Recommended Frequency Range:200 kHz- 1.2 MHz

Comequ

1t

t N f −= ⋅

Fig. 8. Influence of operating frequency on design criteria, assuming transducer’s apefrequency constant Nt = 2020 mHz.

So for a distance d = 3 cm and assuming a = 1 dB/cm at 673 kHzintensity decreases by 6 dB that means half of the power is lost atthe tissue. This loss adds to the power loss due to the geometricalspread of the intensity that allows the receiver to capture only partof transmitted power.

In summary, the selection of operating frequency results from atrade-off between several conflicting requirements.

The external and internal transducer and matching layer thick-nesses decrease with increasing frequency which meets a target ofa thin implanted device (�1–3 mm thickness). Also the Rayleighdistance increases with frequency as do the losses due to tissueabsorption (the Rayleigh distance is the distance from the trans-mitter radiating surface where transition from near field to far fieldoccurs. This range behaves as a natural focus and it is the preferredlocation for the receiver [14]. For aperture a = 7.5 mm andf = 673 kHz the Rayleigh distance is �25 mm). The various trade-offs are illustrated in Fig. 8 which also shows the result of thesummation of the equally weighted criteria representing thesetrade-offs. This recommends an operating frequency range of200 kHz–1.2 MHz. Clearly the choice of operating frequency mayinclude frequencies higher than 1.2 MHz if a thin receiver(<1 mm) rather than power transfer efficiency is of higherimportance.

5. Acoustic impedance matching

The progressive pressure wave that propagates from the trans-mitter through the tissue toward the implanted receiver, encoun-ters acoustic impedance mismatches. Impedance mismatchcauses part of the pressure wave to reflect back at the boundarylayers between the PZT elements and the tissue, since the PZThas much higher impedance (�31 MRayl) than soft tissue(�1.5 MRayl). (The Rayl is a unit of acoustic impedance, definedas: Rayl ¼ kg

m2 s.) The pressure reflection coefficient, U, for a normallyincident plane wave is given by [14]:

jCj ¼ Ztissue � Zpzt

Ztissue þ Zpzt

� 0:9 ð11Þ

Pt ¼ ð1� CÞPi is the transferred pressure where Pi is the incidentwave. This means that the pressure wave captured by the receiverPt has only (1 � U) = 0.1 of the incident pressure amplitude. From

3 3.5 4 4.5 5

Rayleigh Distance [cm]

ransducer’s Thickness t [mm]

bined Effects using ally weighted criteria

0

14

00

1414

1.2

0.7 fα = ⋅

2a

RDf

c=

rture a = 7.5 mm, c = 1500 m/s, attenuation coefficient 0.7 db/cm at 1 MHz, and a

S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566 561

the power transfer perspective, this is even worst since the powerintensity depends on P2

t which is proportional to ð1� CÞ2. Further-more, an acoustically unmatched transducer is characterized by ahigh unloaded mechanical quality factor (Qm = 500–1800) [15]. Thiswould result in a very narrow range of operating frequencies closeto the transducer’s series resonance frequency fr, implying a seriousdifficulty in tuning the resonance frequencies of the transmitter–re-ceiver pair. Furthermore, reflected pressure waves generate pres-sure standing waves in the tissue [10], which may generate peakpressure levels that exceed the tissue safety limit. Thus, it is manda-tory to match the transducer impedance to the tissue impedance.

There are several impedance matching techniques based eitheron single or multiple matching layers. The matching layers func-tion as a mechanical transformer that reflects a matched acousticload towards the transducer device. Multiple matching layer tech-niques have been proposed in [16]. The simplest matching tech-nique involves the use of a single matching layer. This techniqueis based on insertion of a k/4 thick layer between the PZT andthe tissue. A suitable matching material must have acoustic imped-ance that is close to the calculated one, low losses at the operatingfrequency (<1 dB/cm), and must be biocompatible. The acousticimpedance of the single-layer matching material is chosen accord-ing to:

Zmatching ¼ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiZpzt Ztissue

q¼

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi30:7 106 1:5 106

p¼ 6:8 MRayls ð12Þ

The main disadvantage of the single-layer quarter-wavelengthmatching technique for a PZT-tissue application is the resultedimpedance of about 6.8 MRayl, a value which restricts availabilityto only a few materials such as biocompatible pyrolytic carbon[17–19]. Furthermore, with single-layer based matching, the layerof adhesive that glues the matching layer to the piezoelectric ele-ment is not accounted for and degrades the quality of matching.

A better acoustic matching approach is the multiple matchinglayer technique, which allows more freedom in choosing thematching layer materials [20,21]. With multilayer matching, theadhesive layer’s thickness and acoustic properties are taken intoaccount. The technique is based on the multiplication of a chainof transfer matrices in such a way that each matching layer includ-ing adhesive is represented by a 2 2 matrix [21]. Each passivematching layer is represented by a 2 2 complex matrix Tn used

Fig. 9. Reflection coefficient when using two layer matching

to represent a linear network having complex two port propertiesas described by [21]:

Tn ¼cos hn jZn sin hn

jZn

sin hn cos hn

" #: ð13Þ

where tn is the thickness of the nth layer; kn is the wavelength in thenth layer; n = 1, 2, 3, 4; and hn ¼ 2p tn

kn.

The overall equivalent matrix Cequ is achieved by multiplyingthe matrices in a chain as in

Tequ ¼ T1T2T3T4 ¼C11 C12

C21 C22

� �ð14Þ

Looking from the transducer-medium border layer, the equiva-lent acoustic impedance of the transducer comprising the PZT ele-ment and the matching layers is given by (15), [21]. Acousticmatching is achieved when Zequ = Zm (Zm is the impedance of themedium).

Zequ ¼C11ZPZT þ C12

C21ZPZT þ C22ð15Þ

Two layers acoustic matching consisting of Cyanoacrylate glue(0.02 mm), and graphite (1.3 mm) was modeled (Fig. 9) using thePiezocad simulator (Sonic Concepts Inc., Bothell WA, USA) Thistechnique allows the designer to choose the matching materialsfrom a broad range of available materials, such as metals, glass,and plastics. Once the materials have been chosen, the requiredthickness of each layer is calculated.

In addition, a backing layer is not required, as in the case ofultrasound imaging, since the lack of a backing layer results in totalreflection of energy from the transducer’s back surface (air back-ing), making almost all power available for transmission towardthe receiver.

6. Power conditioning

The Power conditioning stage of a UTET device plays an impor-tant role in determining the overall efficiency of the UTET. On thetransmitter side, it has to drive the device at the exact operatingfrequency without exciting harmonic modes. On the receiver side,the circuit should interface with the PZT receiver so as to extractmaximum power.

: Cyanoacrylate glue (0.02 mm), and graphite (1.3 mm).

t

t

PZTTransducer

t

VB1

VB2

VB1-VB2

Q1

Q2

Q3

Q4

Medium

Lr

Lr

Cr

Cr

Lm

Cb

CbOut of Phase

Matching Layer(s)

DC Blocking

Resonant Tank

t

22uH

2700pF

5.6nF

Depends on PZT material

C-2, C-204, C601

+

+

+3.6V

+3.6V

Fig. 10. Two out-of-phase resonance half-bride legs for the excitation of a piezoelectric transducer.

562 S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566

On both sides, the power conditioning circuits must have effi-ciency >80% as they affect the overall efficiency of the energy trans-fer link.

6.1. Transmitter resonance driver

The UTET device is powered by a low voltage, low power sourcesuch as 3.6 V/800 mA h Lithium battery (EL1CR2 manufactured byEnergizer Holdings Inc. St. Louis Missouri, USA). In order to prolongthe battery’s service life, the power amplifier should be as efficientas possible, so a soft switching topology was chosen for two rea-sons: it has low switching losses (<10% depending on operatingfrequency) and generates a sinusoidal voltage with low Total Har-monic Distortion (THD) (<3%) [23] that avoids excitation of unde-sired vibration modes in the transmitter. Various resonancetopologies are available, such as Zero Current Switching (ZCS)and Zero Voltage Switching (ZVS) [22]. A resonance topology basedon two half-bridge legs operating out-of-phase is illustrated inFig. 10.

This inverter topology exhibits a typical conversion efficiency of97%. Furthermore, it generates a low THD sinusoidal voltage of 2%[23] due to the high quality factor Qe (�10) of the resonance tankcircuit which serves as a low pass filter [22]. The use of two out-of-phase half-bridge legs further attenuates the harmonics anddoubles the voltage across the transducer [23].

6.2. Receiver power processing

A simplified electrical equivalent circuit of the receiver based onthe KLM model [15,27] includes a voltage source Vg in series with acomplex impedance network, as illustrated in Fig. 11. The source’sseries impedance is composed of a resistor rg which accounts for

frequency dependent Reactancerg

Vg

Xg Cos

Zero Strain Dielectric Capacitance

Fig. 11. Simplified electrical model of the receiver’s transducer based on the KLMmodel.

losses, a capacitor which accounts for the zero strain dielectriccapacitance Cs

0 and a frequency-dependent reactance:

X ¼ h233 sinðbtÞx2Z0A

here h33 is the piezoelectric stress constant; b the wave numberb ¼ 2p

k ; k the wavelength in the piezoelectric material; t the thick-ness of the piezoelectric element; Z0 the piezoelectric acousticimpedance; A the radiation surface area of the transducer; and xis the angular frequency.

The frequency-dependent reactance vanishes when the receiveroperates exactly at the resonance frequency, since sin (bt) = 0when t = k/2. If the receiver does not vibrate exactly at its reso-nance frequency, the frequency-dependent reactance can be eitherpositive or negative, so the overall equivalent source impedancemay be either inductive or capacitive.

The voltage generated by the receiver must also be rectified andfiltered so a low loss (>80%) rectifier network should be employed[24]. Although rectification is a non-linear process, in order tomaximize the harvested electrical energy, the rectifier networkshould present input impedance Zin that is conjugate to the sourceimpedance. Switched mode active power factor correction topolo-gies [22] are not applicable in the present application due to therelatively high frequency of the rectified voltage (673 kHz of theUTET receiver’s generated voltage compared to 100 Hz (120 Hz)rectified line voltage). This would require a switching frequencyfor a boost converter or other active correction topologies of 5–10 MHz which would impair the efficiency due to power switchingloss on the MOSFET switches [22].

A passive rectification network has been developed as illus-trated in Fig. 12. The rectifier consists of a series resonance LC tankcircuit containing L1, C1 to filter out the current harmonics gener-ated by the Schottky diode rectifiers, and an LC output filter con-sisting of L2, L3, and Cout. Measurement results showed that therectification circuit operated with 88.5% efficiency (data presentedin the experimental paragraph).

7. Experimental set-up and results

7.1. Test tank circuit

To verify the performance of the UTET link, measurements ofultrasound radiation and energy transfer were conducted inside ahomemade water tank, using water at room temperature (25 �C).

Distilled water was used as the medium for acoustic wavepropagation, since its acoustic impedance is close to that of soft

+

Simplified Receiver Model

470uH

470uH

220uH 1.2nF MBR340

MBR340

47nFRload

500pF

Series Resonant Tankfrequency dependent Reactancerg

.

.

.

.

.

.

L1

L2

Vout

Vpzt

Irectifier

Zin

Vg

Xg

Zero Strain Dielectric Capacitance

L3

C1 D1

D2

CP

Cos

Cout

Fig. 12. Schematic of the receiver’s electrical equivalent circuit and the passive rectifier.

S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566 563

biological tissue (1.45–1.55 MRayls) and allows the hydrophone tobe easily moved in order to map the radiation pattern. This servesas a first-order proof-of-concept. Since the water attenuation ismuch lower than that of actual tissue (water 0.002 dB/cm com-pared to 0.6–1.5 dB/cm of soft tissue at 1 MHz [29–31]), measure-ments were also carried out with pig tissue of various thicknesses[25]. The experiments were conducted inside a test water tank fab-ricated out of 6 mm thick Perspex plates and dimensions of40 20 20 cm3, see Fig. 13.

In order to avoid reflections from the test tank walls, they werecovered with 10 mm wide ultrasonic absorber sheets (Aptflex F28)attached to the internal tank walls with APTBOND B1 adhesive(both manufactured by Precision Acoustics, Dorchester UK).

7.2. Construction and characterization of transducers

The UTET transducers were fabricated from disc-shaped PZTelements 15 mm in diameter and 3 mm thick (Fuji CeramicsZ3T15D-C2, piezoelectric properties close to PZT-4) and were oper-ated in the first thickness vibration mode, see Fig. 14. The othertransducer parameters are: Acoustic impedance 30.7 MRayls;Nt = 2020 m Hz, density 7600 kg/m3, and Qm = 1200. The multilayermatching technique assumed a 20 lm thin layer of Cyanoacrylateglue (Quantum 149 made by Hernon, Sanford Florida, USA) as afirst matching layer and a layer of 1.3 mm graphite as the secondmatching layer. The acoustic matching layer was fabricated froma carbon rod EK 2200 (made by SGL Carbon Group, Wiesbaden

Ultrasonic Absorber

Tissue Holder

Perspex Base

ReceiverElectrical

Connections

Tissue

5mm − 35

Fig. 13. Test tank

Germany) with a density of 1.82 g/cm3 and a Young’s modulus of23 GPa. To compensate for the inaccuracy of the glue layer’s thick-ness (designed to be 20 lm), the graphite layer thickness was cal-ibrated after the cure of the glue (24 h). In general, the physicalproperties of pyrolytic carbon fall between those of graphite anddiamond. It is chemically inert and resistance to wear and mechan-ical fatigue. Pyrolytic carbon has been evaluated in cardiovascular,dental, soft tissue and orthopedic implants, has been proven to bebiocompatible and hemocompatible, and is the material of choicefor the construction of artificial heart valves [18,19].

7.3. Pressure and power measurement results

The pressure dependence on the lateral distance from theacoustic axis was measured with a hydrophone probe located attwo sample distances from the transmitter surface: 15 mm (at nearfield) and 30 mm (just beyond the Rayleigh distance). The hydro-phone probe TC4038 (manufactured by Reson, Slangerup Den-mark) has a sensitivity of �224.5 dB ± 2 dB at 673 kHz re 1 V/lPa(�228 dB ± 2 dB at 100 kHz). To drive the transmitter transducer,a laboratory power amplifier has been built using a PA09 poweroperational amplifier (manufactured by Apex Cirrus, Austin Texas,USA) designed to have voltage gain of 10, and a 5.1 X outputresistor was inserted in series with the amplifier’s output to avoidamplifier instability caused by the capacitive nature of the PZT de-vice. The small signal (1–2 V) applied to the amplifier’s input wasgenerated by a waveform generator 33250 A 80 MHz function

PZTGraphite

Transducer Enclosure

Water Level

XY Manipulator

Water

TransmitterElectrical

Connections

mm

construction.

PZT Cyanoacrylate20um 1st layers

Water15mm P

1.3mm Graphite 2nd layer

Fig. 14. Construction of transducers.

Fig. 16. Measured DC load power as a function of non-overlapping lm and distanced (at input power, Pin = 260 mW).

d=5mm

d=10mm

d=15mm

d=20mm

d=25mm

d=30mm

Constant Pin=260mW

oDC

inAC

27

23

19.2

15.4

11.5

7.69

3.85

lm

Load

d RectifierTx RxSource

Xg

Pin PoutDC

AC

Fig. 17. Measured DC load power as a function of lateral non-overlapping anddistance d (power transferred through pig muscle tissue).

564 S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566

generator (manufactured by Agilent Technologies, Santa Clara Cal-ifornia, USA). Electrical waveforms were recorded using a TDS5054,500 MHz, four channel oscilloscope along with P5050 voltageprobes and CT-2 current probes with a transducer sensitivity of1 mV/mA (all of which are manufactured by Tektronix, Oregon,USA).

The hydrophone probe was moved laterally away from theacoustic axis while recording the pressure every half millimeter,see Fig. 15. The maximum lateral distance measured was ±6 mm.Peak pressure measured by the hydrophone was 120 kPa. The UTEToperated in a Constant Wave CW regime to limit the peak pressurethat the tissue is exposed to below the safe level dictated by theFDA [26,28].

Regarding the Mechanical Index MI, since the peak pressuremeasured was 120 kPa, the calculated MI is 0.14 (at 673 kHz)which is lower than the FDA limit of 1.9 (abdominal). While keep-ing the same excitation level and frequency of 673 kHz (inputpower of 260 mW), the dependency of power transfer throughthe water medium on lateral non-overlapping of the transducerslm was measured at 3 distances (5 mm, 15 mm, 30 mm) and theresults are presented in Fig. 16.

The power transfer through pig muscle tissue was measured aswell [25]. The tissue was immersed in the test tank and placed be-tween the transmitter and receiver. Fig. 17 shows the measuredoutput power as a function of tissue thickness (�1.8 dB/cm at1 MHz) and lateral non-overlapping. The power transfer is alsosensitive to the accuracy of the excitation frequency.

The transducers were aligned accurately and placed in the testtank at a distance of 20 mm. The test tank was filled with distilledwater, and a sample of pig muscle 20 mm thick was immersed inthe test tank between the transducers. The sensitivity of the ex-tracted load power to the accuracy of the excitation frequencywas measured by sweeping the excitation frequency around the

-8 -6 -4 -2 0 2 4 6 80.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

1.1

Lateral Distance From Acoustical Axis [mm]

Nor

mal

ized

Pre

ssur

e

d=15mm

d=30mm

d: distance between transducers

Acoustical Axis

Fig. 15. Measured pressure as a function of distance from acoustic axis at distancesd of 15 mm and 30 mm from transducer radiating surface.

center resonance frequency. Fig. 18 shows a maximum reductionof 28% of the load power due to unmatched excitation frequency.

The high efficiency (88.5%) passive rectifier illustrated in Fig. 12was built and typical waveforms were measured and recorded. Therectifier output was loaded with a resistive load and tuned to give+3.9 V DC at a load power of 70 mW. Fig. 19 presents the measuredanode voltages of the rectifier diodes, and the voltage, current, andpower at the rectifier’s input. The value of Vg was measured by anopen circuit experiment. The component’s values of the sourceimpedance were evaluated by measuring the voltage and currentunder short circuit conditions (across the receiver output termi-nals) and also while loaded by a 500 X resistor (higher than themagnitude of the transducer’s impedance which is about 230 Xin order to obtain two measurement results; below and abovethe transducer’s impedance).

8. Discussion

This paper proposes a transcutaneous energy transfer methodusing propagating ultrasonic waves. The objectives of the workwere to show that such power transfer is feasible to a range ofup to 40 mm, and measure its sensitivity to the distance, non-over-lapping of the transducers, and inaccuracy of the frequency of the

668 669 670 671 672 673 674 675 676 677 6780.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

1.1

Excitation Frequency [kHz]

Nor

mal

ized

Loa

d P

ower

Fig. 18. Measured load power sensitivity to excitation frequency accuracy.

Fig. 19. Measured input voltage, current and power, and anode voltages of therectifier.

S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566 565

exciting source. The paper recommends an operating frequencyrange taking into account the tissue absorption and size of thetransducers. As can be seen from Fig. 4 the power transfer is sensi-tive to operating conditions. The pressure field diffracts at dis-tances longer than the Rayleigh distance, but the fact that thepressure field is not focused, helps to spread the energy over theentire cross sectional area of the tissue. (Focused pressure fieldalthough beneficial for imaging and treatment purposes, is not rec-ommended for the UTET since it might increase the intensity abovethe safety threshold of 94 mW/cm2). The preferred location of thereceiver is at the natural focus of the pressure field (about 25–30 mm) as illustrated by Figs. 4 and 15. Load power is sensitive

to the distance. Fig. 16 shows the measured power through waterat various distances and lateral non-overlapping. Although waterhas practically negligible absorption at 673 kHz, peak power trans-fer efficiency achieved was 0.38 which is the result of working atmatched load condition at the receiver, limiting the maximum pos-sible efficiency to 50%. Power transfer is also sensitive to lateralnon-overlapping caused by moving transducers as illustrated inFig. 17. As might be expected, the effectiveness of the device isgreatly adversely affected in the combined situation of distanceand lateral non-overlapping. Although a medium excitation fre-quency of 673 kHz was chosen (Fig. 8 recommends a frequencyrange of 200–1200 kHz lower than used by ultrasonic imaging),the power transfer efficiency (including rectifier loss) through pigmuscle dropped to 27% and even lower (Fig. 17) due to the fact thatpig abdominal muscle tissue has more than two orders of magni-tude higher attenuation compared to water. It explains the addedloss between the results of Fig. 16 (using water) to those depictedat Fig. 17 (using pig muscle). In addition, the design of the trans-ducer matching layer assumed a medium acoustic impedance of1.5 MRayl, but pig abdominal muscle used has a mean value of1.63 MRayl. The slight impedance mismatch probably caused asmall part of the energy to be reflected back at the receiver–tissueinterface contributing to about 15% less power being captured bythe receiver. Fig. 18 illustrates the sensitivity of UTET power trans-fer to the exact value of the excitation frequency. The frequencymismatch might be the result of a finite practical accuracy (±1%)of the manufacturing thickness of the PZT.

Electrical measurements of the rectifier showed a rectificationefficiency of 88.5% as presented in Fig. 19. Rectifier losses consistof diode losses and inductor losses (mostly copper losses due tothe 10 X series resistance of the small 0805 size RF coils manufac-tured by Taiyo-Yuden, Tokyo, Japan).

The maximum power throughput imposed by bio-safety limitswas also referred to in the context of the UTET mechanism. TheMI (0.14) is well below the safety threshold (1.9). Furthermore,the FDA’s requirement indicated by the Ispta (intensity, temporalaverage), limits the intensity to 94 mW/cm2 for abdominal andother tissue (but not ophthalmic, which is not relevant for theUTET). Meeting these limits, a continuous power of 150 mW canbe transferred using transducers having an area of 1.6 cm2.

9. Conclusions

The ultrasonic transcutaneous energy transfer (UTET) approachis proposed for the purpose of energizing implanted devices. Mea-surement results at operating frequency of 673 kHz indicate thatthe UTET is an effective method for transferring power at a densityless than 94 mW/cm2 while complying with tissue safety limits. Apower transfer efficiency of 27% is achieved through lossy pig mus-cle tissue. However, the power transfer efficiency is quite sensitiveto the distance and non-overlapping of the transmitter and recei-ver. Finally, receiver power conditioning circuitry which is essen-tial for high power transfer efficiency is proposed which gave88.5% rectification efficiency.

Acknowledgement

This research was partially supported by Israel Ministry of Sci-ence and Technology under Grant No. 3-4772.

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