Ultrasonic transcutaneous energy transfer for powering implanted

Ultrasonics 50 (2010) 556–566
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Ultrasonics
journal homepage: www.elsevier.com/locate/ultras
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 2009
Received in revised form 12 November 2009
Accepted 13 November 2009
Available online 26 November 2009
Keywords:
Transcutaneous energy transfer
Powering implanted devices
Ultrasonic energy
Acoustic impedance matching
a b s t r a c t
This paper investigates ultrasonic transcutaneous energy transfer (UTET) as a method for energizing
implanted devices at power level up to a few 100 mW. We propose a continuous wave 673 kHz single
frequency operation to power devices implanted up to 40 mm deep subcutaneously. The proposed UTET
demonstrated an overall peak power transfer efficiency of 27% at 70 mW output power (rectified DC
power at the load).
The transducers consisted of PZT plane discs of 15 mm diameter and 1.3 mm thick acoustic matching
layer 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 issues
such as recommended operating frequency range, acoustic link matching, receiver’s rectifying electronics, and tissue bio-safety concerns. Global optimization and design considerations for maximum power
transfer are presented and verified by means of finite element simulations and experimental results.
Ó 2009 Elsevier B.V. All rights reserved.
1. Introduction
The last few decades have seen an enormous increase in the
number and variety of medical devices. According to the American
National Institute of Health (NIH), new devices are being added to
the market every year. Some of these devices are designed to be
implanted into the body for monitoring purposes such as biosensors, glucose indicators. Other implanted devices serve therapeutic
purposes, e.g. pacemakers, defibrillators, heart-assist devices, or
miniature mechanisms for the controlled delivery of medications,
such as implanted insulin pumps. All of these devices require electrical energy for their operation. While the majority of modern implanted devices consume low power (in the range of hundreds of
mW), up to 10 W of power may be required in some cases.
In recent years some very interesting methods for intra-body
electrical energy sources have been reported. Refs. [1,2] discuss
the generation of electrical energy by biological fuel cells that consume inter-cellular glucose. The fuel cells generate a voltage of
about 0.5 V with a power density of 50 lW/cm2. Another work
[3] suggests an intra-body ‘‘energy harvesting” system. According
to this approach, the bio-mechanical energy created by the movement of internal organs, such as the peristaltic movement of the
intestines or the contraction of the heart and lungs, is converted
into electrical energy using piezoelectric elements. Evidently, with
these techniques, an average power below 1 mW can be generated.
Devices that consume average power of a few (<10 W) Watts are
powered wirelessly by a mechanism based on the penetration of
* Corresponding author.
E-mail address: shmilo@eng.tau.ac.il (D. Shmilovitz).
0041-624X/$ - see front matter Ó 2009 Elsevier B.V. All rights reserved.
doi:10.1016/j.ultras.2009.11.004
energy through the tissue, a technique called transcutaneous energy transfer (TET).
Presently, existing TET devices rely on electromagnetic energy
transmission and customarily employ a pair of flat spiral coils
[4–6], one external energy transmitting coil and another intrabody receiving coil, with the two coils facing each other. The internal coil is implanted under the fat layer, typically at a depth of 10–
20 mm.
Until now, published papers have dealt with electromagnetic
TET mechanisms that were capable of transferring power through
the tissue at power levels of up to 10 W. This range of power is
suitable for devices such as heart pumps or cardiac assist devices
that increase the volume of blood pumped around the body [7].
These devices are well-suited to electromagnetic TET-based mechanisms. As mentioned, the electromagnetic transmission is based
on two coils facing each other. The separation between the coils
varies from 10 to 20 mm depending on the skin/tissue thickness.
Furthermore, the distance and relative orientation of the coils
may vary as the tissue moves or deforms. Since there is no ferromagnetic core to concentrate the magnetic flux and close the magnetic circuit, the magnetic coupling coefficient k between the two
coils is typically low (0.1). The low coupling requires a primary
current of several Amps which reduces the efficiency and necessitates batteries with a 1–5 A current drain capability.
Several papers have discussed the need to increase the efficiency of the coupling [4,8]. These authors discussed the difficulties
encountered in this effort; among them, the need to increase the
excitation current, which causes excess heating in the primary coil
and in turn leads to an increase in tissue temperature. In reference
[4], the authors suggested a coil geometry that gave a coupling of
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
k = 0.4. An electromagnetic TET device intended for a miniature implanted Bio-MEMS sensor was presented in [8] which detailed the
design of the RF link at 330 MHz. The diameter of the external coil
was 75 mm, but the implanted coil was miniaturized (1 mm2) and
built into the sensor device itself. The received power reached
about 250 mW.
Clearly, Electromagnetic TET is adequate for transferring energy
in the Watt range, but suffers from some drawbacks such as low
coupling (0.1) which might cause interference with the operation
of devices working in close proximity, such as pacemakers. Conversely, the operation of the TET itself may be perturbed, as its coil
may pick up external electromagnetic radiation created by sources
such as MRI devices in hospitals, or as a result of the nearby presence of massive ferromagnetic objects, such as steel doors.
For devices consuming large amounts of power, the electromagnetic TET mechanism is probably still the most suitable in spite of
the shortcomings mentioned above. However, for the power range
of tens of mW (to power smart devices such as those of the BioMEMS devices), Ultrasonic TET (UTET) may be a preferable technology due to its power transfer efficiency, compactness, and electromagnetic immunity.
This paper proposes a UTET device in which the energy is transmitted via an acoustic wave as previously suggested to stimulate
bone healing [5]. The UTET operates at a Constant Wave single frequency 673 kHz, and transfer 70 mW to a DC load. Similar to electromagnetic-based TET, there is an implanted intra-body receiver
and a transmitter external to the body. However, in contrast to
electromagnetic TET, the usual coils are replaced with ultrasonic
piezoelectric transducers. In this system, the external electrical
power is converted to a pressure wave, which is then transmitted
transcutaneously. The acoustic energy is collected by an implanted
transducer that reconverts the received acoustic energy back into
electrical energy for use in the medical device.
In fact, the idea of using acoustic waves to transmit energy was
proposed long ago in 1958 by Rosen et al. [9]. The author described
a piezoelectric transformer consisting of a primary piezoelectric
section acoustically coupled to a secondary piezoelectric section
to reconvert the elastic waves back to electrical energy. With a
controlled interface medium, piezoelectric transformers can deliver throughput powers of tens of Watts with a conversion efficiency as high as 98%. UTET can also serve in non-medical
applications such as for powering embedded sensors within metal
constructions [11] (as in use in bridges), for energizing pressure
leak sensors in satellites, or for powering sensors embedded inside
fuel tanks [12].
The principal design considerations of UTET are discussed, such
as operating frequency, piezoelectric material selection, acoustic
impedance matching, power conditioning and safety issues.
2. Acoustic power transfer concept
Energy transfer through the skin requires an external transducer (transmitter) attached to the skin surface facing an implanted transducer (receiver), as illustrated in Fig. 1. An electrical
power source energizes the transmitter that converts the electrical
energy into acoustic pressure waves. The acoustic waves carry the
energy through the tissue toward an implanted receiver positioned
within the radiation lobe of the transmitter. The piezoelectric receiver converts the acoustic energy back into electric energy and
a low loss (efficiency >80%) rectifier network rectifies and filters
the output voltage of the receiving transducer, while reflecting at
its input terminals an impedance conjugate to that of the piezoelectric element. In order to minimize the inconvenience caused
to the patient as well as to ensure a close fit to the body (which
is required for good acoustic coupling), the device should be
lightweight and thin so that its center of gravity is as close as possible to the surface of the body. A UTET system is depicted in Fig. 1.
The power throughput in the present application is limited to the
mW range (Pload < 100 mW) in order to avoid tissue damage. Such
a UTET is powered from a small coin like, low capacity (<500 mA h)
battery that is typically characterized by low drain current capability (<50 mA).
In order to preserve the battery capacity, the current consumed
by the switching amplifier (which feeds the transmitter element)
should be as smooth as possible (i.e., with a crest factor close to
1) [13]. A Maximum Power Extracting (MPE) circuit was designed
for both functions: extracting most of the power generated by the
implanted transducer and increasing the receiver-generated voltage to a level suitable for powering typical loads, such as logic
and analog circuitry (Vdc > 3 V).
The UTET link’s power transfer efficiency is affected by the
switching amplifier losses, transducer losses, tissue absorption,
acoustic impedance matching layer losses, rectifier losses, and
the amount of power captured by the receiver. PZT (Lead, Zirconate, Titanate) materials are preferred choice for the implementation of the transducers compared to piezoelectric polymer
Polyvinylidene Fluoride PVDF such as the Piezoflex (manufactured
by Airmar technology corporation, Milford New Hampshire, USA)
due to properties such as high electromechanical coupling and
mechanical Q factor (k33 = 0.76, Qm = 1200 compared to k = 0.3,
Qm < 25 of PVDF). Using PZT results in high electromechanical energy conversion efficiency (some applications such as piezoelectric
transformers achieve up to 95%), and low excitation voltage
(<15Vpk for 100 mW power transfer). Since the acoustic impedance
of soft tissue is much lower than that of a PZT transducer (1.5 MRayls in comparison to about 30 MRayls for PZT), acoustic impedance
matching is required in order to minimize pressure wave reflections and consequent generation of standing waves within the
tissue.
3. Pressure field shape
According to the Huygens principle, each point on the face of
the radiating transducer may be treated as an independent point
source of radiation. The acoustic field pattern in front of the transducer’s face is the vector sum of contributions from all point
sources [10]. In the general case, the pressure field at an observation point L(x, y, z) is given by the Rayleigh integral [10]:
Pðx; y; z; tÞ ¼ q0
Z u_ p x0 ; y0 : t r
c0
S
2p R
dS
ð1Þ
where (x0 , y0 ) are the coordinates of a point source on the transducer; (x, y, z) are the coordinates of the observation point in front
of the transducer; u_ p is the vibration velocity function over the
transducer’s radiating cross section; c0 is the average speed of
sound (wave’s phase velocity) in the medium; q0 is the average
medium density; and S is the transducer area.
R is the distance from the point source to the observation point
L(x, y, z):
R¼
qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
ðx x0 Þ þ ðy y0 Þ2 þ z2
ð2Þ
The Rayleigh integral is difficult to solve in the general case of
an arbitrary transducer shape and non-uniform vibration distribution. However, in the case of UTET which employs a Continuous
Wave (CW) sinusoidal excitation, uniformly distributed over the
face 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:
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
Thin coupling layer
(Castor Oil,
Ultrasonic Jell etc .)
PZT
Transducer
Implanted Unit
5 − 50mm
Ultrasonic Traveling
Waves
t
Vo
H Bridge
P
P
Passive
Network
Battery
MPE
+
Network
Load
PZT
Transducer
Switching
Amplifier
Matching
layers
Low Loss Rectifier
Fig. 1. UTET system.
Pðx; y; z : tÞ ¼
jkq0 c0 u0 jxt
e
2p
Z
S
ejkR
dS
R
ð3Þ
where R is the distance from the infinitesimal point source to the
observation point; u0 is the vibration velocity amplitude; k is the
wavelength of the pressure wave in the medium; c0 is the phase
velocity of the wave; q0 is the density of the medium; x is the
angular frequency; and k ¼ x=c ¼ 2p=k is the wave number.
Actually, the wave number k should be complex for a dissipa2
tive medium such as tissue, k ¼ b2 þ a2 where b is the phase
speed b ¼ x=c and a is the absorption coefficient. Assuming
r a and CW sinusoidal excitation, the solution of (3) is:
Pðr; h : tÞ ¼
jaq0 c0 u0 jðxtkrÞ J 1 ðka sin hÞ
e
sin h
r
ð4Þ
where a is the transducer radius; k is the wave number; J1 is the 1st
order Bessel function; u0 is the vibration velocity; r is the distance
between the center of the radiating disc and the observation point;
and h is the angle formed between the observation point to the
acoustic axis (h = 0).
The pressure directivity of the transducer is defined as the ratio
between the pressure at any angle h relative to that at the acoustic
axis (h = 0) for the same range [10] (see Fig. 2).
DðhÞ ¼
2J 1 ðka sin hÞ
ka sin h
ð5Þ
The field should be directed toward the receiver to let the receiver capture most of the radiated energy. The directivity depends on
ka ¼ 2pk a which is the ratio of the transducer’s perimeter to wavelength (Fig. 3). As an example, for a receiver with a = 7.5 mm located 20 mm from the transmitter,
pressure should
the radiated
¼ 20:5 .
be confined to within tan1 7:5
20
If ka is relatively small, i.e. when the transducer’s perimeter is
not large enough compared to the wavelength, the directivity of
the field is poor, and the field diverges like a spherical wave. This
means that only a fraction of the pressure wave energy can be captured by the receiver.
A 2D finite element simulation (Comsol Multiphysics, Comsol
AB, Stockholm, Sweden) of the pressure intensity profile generated
by a disc transducer is illustrated in Fig. 4, which shows the directivity of the radiated pressure wave. (The wave’s intensity depends
on the pressure squared and is referred to later on.) The pressure
field generated by a uniform CW excited disc transducer can be divided into three zones (Fig. 5). The first zone, the Near Field (NF)
zone, is the one closest to the transducer. Within this zone, the
pressure field envelope oscillates, and has multiple minima and
maxima which make the power transfer unpredictable. After the
NF, the pressure field converges to a natural focus. This interval
is the preferred distance at which to locate the receiver. The Near
Field distance L from the transducer depends on the transducer’s
radius a and on the acoustic wavelength k, in the medium through
which the wave propagates [14], and is given by
L¼
ð2aÞ2 k2 a2
4k
k
For ð2a2 Þ k2
ð6Þ
Beyond the zone of the natural focus starts the Far Field (FF) region, where the pressure field becomes a spherically spreading
wave with little internal structure, whose intensity decays with
distance.
The points on the acoustic axis at which the pressure peaks depend on the wavelength and on the transmitter’s cross section, is
given by Eq. (7), where m stands for the order of the pressure peak.
X max ðmÞ ¼
Fig. 2. Calculation of the pressure field generated by a disc-shaped source.
ð2aÞ2 kð2m þ 1Þ2
4kð2m þ 1Þ
m ¼ 1; 2; 3; . . .
ð7Þ
In cases where the receiver has to be implanted within the NF
range, it is possible to either insert a bag (1–5 mm thick) containing acoustic jell, water, oil, etc., between the transmitter and the
skin, or tune the vibration frequency so as to place the closer radiation pressure maxima of the NF, exactly on the receiver, see Fig. 6.
S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
559
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).
d
Z
Fig. 5. Preferred location of implanted receiver.
4. Frequency selection
Proper determination of the operating frequency is of great
importance since it affects factors such as tissue attenuation (with
a frequency dependence of f1–f1.4 [30]), transducer thickness, distance of the natural focus (Rayleigh distance), and the size of the
reactive elements, e.g. inductors and capacitors [13] in both the receiver and transmitter. In order to maximize the power throughput
of the transducer, it is best to operate the transducer close to its
resonance frequency. Resonance frequencies are determined by
the transducer’s geometry and material constants. Thickness vibration resonance occurs at frequencies in which the transducer’s
Fig. 6. Attaining maximum pressure at the receiver surface by frequency tuning.
thickness equals an odd multiplication of half acoustic wavelengths [15]. The 1st resonance frequency of a disc-shaped transducer vibrating in the thickness vibration mode (Fig. 7) is
determined by its thickness t and its frequency constant Nt
[m Hz], as given by
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
λ
2
Fig. 7. First thickness vibration mode of piezoelectric transducer.
fr ½Hz ¼
Nt
t
ð8Þ
For example, for C-2 material (manufactured by Fuji Ceramics
Corporation, Tokyo, Japan) Nt = 2020 m Hz. Thus, according to Eq.
(8), increasing the frequency results in a thinner device. The average wave intensity I can be shown to obey Eq. (9) (by integration of
instantaneous pressure multiplied by instantaneous particle velocity [14,30]):
I
P2pk
W
¼
m2
2Z
ð9Þ
where Z is the acoustic impedance of the medium (tissue) through
which the wave propagates.
The operating frequency also determines the amount of acoustic energy transformed into heat by loss mechanisms in the tissue.
Although skin and the underlying soft tissue layer have acoustic
impedances and phase velocities that are close to those in water,
the attenuation of the pressure field by tissue is much larger than
water attenuation (soft tissue 0.6–1.5 dB/cm compared to water
0.002 dB/cm at 1 MHz) [29–31], and increases as the frequency
and distance increase. Since the intensity depends on the pressure
squared therefore it decreases at twice the rate compared to the
pressure decreasing rate:
Id ¼ Io e2ad
ð10Þ
where I0 is the intensity at the transducer’s radiating surface, a is
the pressure attenuation coefficient per unit length, and Id is the
intensity at a distance d from the transducer’s radiating surface.
So for a distance d = 3 cm and assuming a = 1 dB/cm at 673 kHz
intensity decreases by 6 dB that means half of the power is lost at
the tissue. This loss adds to the power loss due to the geometrical
spread of the intensity that allows the receiver to capture only part
of transmitted power.
In summary, the selection of operating frequency results from a
trade-off between several conflicting requirements.
The external and internal transducer and matching layer thicknesses decrease with increasing frequency which meets a target of
a thin implanted device (1–3 mm thickness). Also the Rayleigh
distance increases with frequency as do the losses due to tissue
absorption (the Rayleigh distance is the distance from the transmitter radiating surface where transition from near field to far field
occurs. This range behaves as a natural focus and it is the preferred
location for the receiver [14]. For aperture a = 7.5 mm and
f = 673 kHz the Rayleigh distance is 25 mm). The various tradeoffs are illustrated in Fig. 8 which also shows the result of the
summation of the equally weighted criteria representing these
trade-offs. This recommends an operating frequency range of
200 kHz–1.2 MHz. Clearly the choice of operating frequency may
include frequencies higher than 1.2 MHz if a thin receiver
(<1 mm) rather than power transfer efficiency is of higher
importance.
5. Acoustic impedance matching
The progressive pressure wave that propagates from the transmitter through the tissue toward the implanted receiver, encounters acoustic impedance mismatches. Impedance mismatch
causes part of the pressure wave to reflect back at the boundary
layers between the PZT elements and the tissue, since the PZT
has much higher impedance (31 MRayl) than soft tissue
(1.5 MRayl). (The Rayl is a unit of acoustic impedance, defined
as: Rayl ¼ mkg
2 s.) The pressure reflection coefficient, U, for a normally
incident plane wave is given by [14]:
Z tissue Z pzt 0:9
jCj ¼ Z tissue þ Z pzt ð11Þ
Pt ¼ ð1 CÞPi is the transferred pressure where Pi is the incident
wave. This means that the pressure wave captured by the receiver
Pt has only (1 U) = 0.1 of the incident pressure amplitude. From
14
14 14
14
0
0
Combined Effects using
equally weighted criteria
12
10
Recommended Frequency Range:
200 kHz- 1.2 MHz
Rayleigh Distance [cm]
8
2
a
= c
RD
6
f
.7
α =0
4
⋅f
1.2
2
t = Nt ⋅ f −1
0
0
0.5
1
1.5
2
2.5
Transducer’s Thickness t [mm]
3
3.5
4
4.5
0
5
Fig. 8. Influence of operating frequency on design criteria, assuming transducer’s aperture a = 7.5 mm, c = 1500 m/s, attenuation coefficient 0.7 db/cm at 1 MHz, and a
frequency constant Nt = 2020 mHz.
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
the power transfer perspective, this is even worst since the power
intensity depends on P2t which is proportional to ð1 CÞ2 . Furthermore, an acoustically unmatched transducer is characterized by a
high unloaded mechanical quality factor (Qm = 500–1800) [15]. This
would result in a very narrow range of operating frequencies close
to the transducer’s series resonance frequency fr, implying a serious
difficulty in tuning the resonance frequencies of the transmitter–receiver pair. Furthermore, reflected pressure waves generate pressure standing waves in the tissue [10], which may generate peak
pressure levels that exceed the tissue safety limit. Thus, it is mandatory to match the transducer impedance to the tissue impedance.
There are several impedance matching techniques based either
on single or multiple matching layers. The matching layers function as a mechanical transformer that reflects a matched acoustic
load towards the transducer device. Multiple matching layer techniques have been proposed in [16]. The simplest matching technique involves the use of a single matching layer. This technique
is based on insertion of a k/4 thick layer between the PZT and
the tissue. A suitable matching material must have acoustic impedance that is close to the calculated one, low losses at the operating
frequency (<1 dB/cm), and must be biocompatible. The acoustic
impedance of the single-layer matching material is chosen according to:
Z matching ¼
qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
Z pzt Z tissue ¼ 30:7 106 1:5 106
¼ 6:8 MRayls
ð12Þ
The main disadvantage of the single-layer quarter-wavelength
matching technique for a PZT-tissue application is the resulted
impedance of about 6.8 MRayl, a value which restricts availability
to only a few materials such as biocompatible pyrolytic carbon
[17–19]. Furthermore, with single-layer based matching, the layer
of adhesive that glues the matching layer to the piezoelectric element is not accounted for and degrades the quality of matching.
A better acoustic matching approach is the multiple matching
layer technique, which allows more freedom in choosing the
matching layer materials [20,21]. With multilayer matching, the
adhesive layer’s thickness and acoustic properties are taken into
account. The technique is based on the multiplication of a chain
of transfer matrices in such a way that each matching layer including adhesive is represented by a 2 2 matrix [21]. Each passive
matching layer is represented by a 2 2 complex matrix Tn used
to represent a linear network having complex two port properties
as described by [21]:
"
Tn ¼
cos hn
j
Zn
sin hn
jZ n sin hn
cos hn
#
ð13Þ
:
where tn is the thickness of the nth layer; kn is the wavelength in the
nth layer; n = 1, 2, 3, 4; and hn ¼ 2p ktnn .
The overall equivalent matrix Cequ is achieved by multiplying
the matrices in a chain as in
T equ ¼ T 1 T 2 T 3 T 4 ¼
C 11
C 12
C 21
C 22
ð14Þ
Looking from the transducer-medium border layer, the equivalent acoustic impedance of the transducer comprising the PZT element and the matching layers is given by (15), [21]. Acoustic
matching is achieved when Zequ = Zm (Zm is the impedance of the
medium).
Z equ ¼
C 11 Z PZT þ C 12
C 21 Z PZT þ C 22
ð15Þ
Two layers acoustic matching consisting of Cyanoacrylate glue
(0.02 mm), and graphite (1.3 mm) was modeled (Fig. 9) using the
Piezocad simulator (Sonic Concepts Inc., Bothell WA, USA) This
technique allows the designer to choose the matching materials
from a broad range of available materials, such as metals, glass,
and plastics. Once the materials have been chosen, the required
thickness of each layer is calculated.
In addition, a backing layer is not required, as in the case of
ultrasound imaging, since the lack of a backing layer results in total
reflection of energy from the transducer’s back surface (air backing), making almost all power available for transmission toward
the receiver.
6. Power conditioning
The Power conditioning stage of a UTET device plays an important role in determining the overall efficiency of the UTET. On the
transmitter side, it has to drive the device at the exact operating
frequency without exciting harmonic modes. On the receiver side,
the circuit should interface with the PZT receiver so as to extract
maximum power.
Fig. 9. Reflection coefficient when using two layer matching: Cyanoacrylate glue (0.02 mm), and graphite (1.3 mm).
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
t
Q1
+3.6V
+
Resonant
Tank
Lr
22uH
+3.6V
+
Q2
DC
Cb Blocking
Matching
Layer(s)
VB1 5.6nF
Cr
2700pF
t
VB1-VB2
Medium
t
t
Q3
Lr
Out of Phase
Q4
Lm
VB2
Depends
on PZT
material
Cb
Cr
PZT
Transducer
C-2, C-204, C601
Fig. 10. Two out-of-phase resonance half-bride legs for the excitation of a piezoelectric transducer.
On both sides, the power conditioning circuits must have efficiency >80% as they affect the overall efficiency of the energy transfer link.
losses, a capacitor which accounts for the zero strain dielectric
capacitance C s0 and a frequency-dependent reactance:
2
X¼
6.1. Transmitter resonance driver
The UTET device is powered by a low voltage, low power source
such as 3.6 V/800 mA h Lithium battery (EL1CR2 manufactured by
Energizer Holdings Inc. St. Louis Missouri, USA). In order to prolong
the battery’s service life, the power amplifier should be as efficient
as possible, so a soft switching topology was chosen for two reasons: it has low switching losses (<10% depending on operating
frequency) and generates a sinusoidal voltage with low Total Harmonic Distortion (THD) (<3%) [23] that avoids excitation of undesired vibration modes in the transmitter. Various resonance
topologies are available, such as Zero Current Switching (ZCS)
and Zero Voltage Switching (ZVS) [22]. A resonance topology based
on two half-bridge legs operating out-of-phase is illustrated in
Fig. 10.
This inverter topology exhibits a typical conversion efficiency of
97%. Furthermore, it generates a low THD sinusoidal voltage of 2%
[23] due to the high quality factor Qe (10) of the resonance tank
circuit which serves as a low pass filter [22]. The use of two outof-phase half-bridge legs further attenuates the harmonics and
doubles the voltage across the transducer [23].
6.2. Receiver power processing
A simplified electrical equivalent circuit of the receiver based on
the KLM model [15,27] includes a voltage source Vg in series with a
complex impedance network, as illustrated in Fig. 11. The source’s
series impedance is composed of a resistor rg which accounts for
frequency dependent
Reactance
s
rg
Xg
Co
h33 sinðbtÞ
x2 Z 0 A
here h33 is the piezoelectric stress constant; b the wave number
b ¼ 2kp; k the wavelength in the piezoelectric material; t the thickness of the piezoelectric element; Z0 the piezoelectric acoustic
impedance; A the radiation surface area of the transducer; and x
is the angular frequency.
The frequency-dependent reactance vanishes when the receiver
operates exactly at the resonance frequency, since sin (bt) = 0
when t = k/2. If the receiver does not vibrate exactly at its resonance frequency, the frequency-dependent reactance can be either
positive or negative, so the overall equivalent source impedance
may be either inductive or capacitive.
The voltage generated by the receiver must also be rectified and
filtered so a low loss (>80%) rectifier network should be employed
[24]. Although rectification is a non-linear process, in order to
maximize the harvested electrical energy, the rectifier network
should present input impedance Zin that is conjugate to the source
impedance. Switched mode active power factor correction topologies [22] are not applicable in the present application due to the
relatively high frequency of the rectified voltage (673 kHz of the
UTET receiver’s generated voltage compared to 100 Hz (120 Hz)
rectified line voltage). This would require a switching frequency
for a boost converter or other active correction topologies of 5–
10 MHz which would impair the efficiency due to power switching
loss on the MOSFET switches [22].
A passive rectification network has been developed as illustrated in Fig. 12. The rectifier consists of a series resonance LC tank
circuit containing L1, C1 to filter out the current harmonics generated by the Schottky diode rectifiers, and an LC output filter consisting of L2, L3, and Cout. Measurement results showed that the
rectification circuit operated with 88.5% efficiency (data presented
in the experimental paragraph).
7. Experimental set-up and results
7.1. Test tank circuit
Vg
Zero Strain Dielectric
Capacitance
Fig. 11. Simplified electrical model of the receiver’s transducer based on the KLM
model.
To verify the performance of the UTET link, measurements of
ultrasound radiation and energy transfer were conducted inside a
homemade water tank, using water at room temperature (25 °C).
Distilled water was used as the medium for acoustic wave
propagation, since its acoustic impedance is close to that of soft
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
frequency dependent
Reactance
s
rg
Xg
Co
Vg
Zero Strain Dielectric
Capacitance
Series Resonant Tank
I rectifier
Vpzt
220uH
1.2nF
L1
C1
.
500pF
470 uH
Simplified Receiver Model
.
.
Vout
.
.
D1
470 uH
CP
MBR340
L2
Cout +
47nF
L3
R load
.
D2
MBR340
Zin
Fig. 12. Schematic of the receiver’s electrical equivalent circuit and the passive rectifier.
biological tissue (1.45–1.55 MRayls) and allows the hydrophone to
be easily moved in order to map the radiation pattern. This serves
as a first-order proof-of-concept. Since the water attenuation is
much lower than that of actual tissue (water 0.002 dB/cm compared to 0.6–1.5 dB/cm of soft tissue at 1 MHz [29–31]), measurements were also carried out with pig tissue of various thicknesses
[25]. The experiments were conducted inside a test water tank fabricated out of 6 mm thick Perspex plates and dimensions of
40 20 20 cm3, see Fig. 13.
In order to avoid reflections from the test tank walls, they were
covered 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 PZT
elements 15 mm in diameter and 3 mm thick (Fuji Ceramics
Z3T15D-C2, piezoelectric properties close to PZT-4) and were operated in the first thickness vibration mode, see Fig. 14. The other
transducer parameters are: Acoustic impedance 30.7 MRayls;
Nt = 2020 m Hz, density 7600 kg/m3, and Qm = 1200. The multilayer
matching technique assumed a 20 lm thin layer of Cyanoacrylate
glue (Quantum 149 made by Hernon, Sanford Florida, USA) as a
first matching layer and a layer of 1.3 mm graphite as the second
matching layer. The acoustic matching layer was fabricated from
a carbon rod EK 2200 (made by SGL Carbon Group, Wiesbaden
Germany) with a density of 1.82 g/cm3 and a Young’s modulus of
23 GPa. To compensate for the inaccuracy of the glue layer’s thickness (designed to be 20 lm), the graphite layer thickness was calibrated after the cure of the glue (24 h). In general, the physical
properties of pyrolytic carbon fall between those of graphite and
diamond. It is chemically inert and resistance to wear and mechanical fatigue. Pyrolytic carbon has been evaluated in cardiovascular,
dental, soft tissue and orthopedic implants, has been proven to be
biocompatible and hemocompatible, and is the material of choice
for the construction of artificial heart valves [18,19].
7.3. Pressure and power measurement results
The pressure dependence on the lateral distance from the
acoustic axis was measured with a hydrophone probe located at
two sample distances from the transmitter surface: 15 mm (at near
field) and 30 mm (just beyond the Rayleigh distance). The hydrophone probe TC4038 (manufactured by Reson, Slangerup Denmark) 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 power
operational amplifier (manufactured by Apex Cirrus, Austin Texas,
USA) designed to have voltage gain of 10
, and a 5.1 X output
resistor was inserted in series with the amplifier’s output to avoid
amplifier instability caused by the capacitive nature of the PZT device. The small signal (1–2 V) applied to the amplifier’s input was
generated by a waveform generator 33250 A 80 MHz function
Receiver
Electrical
Connections
Transmitter
Electrical
Connections
XY Manipulator
Perspex
Base
Water Level
Tissue
Transducer
Enclosure
Ultrasonic
Absorber
5mm − 35mm
PZT
Graphite
Water
Tissue
Holder
Fig. 13. Test tank construction.
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S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
1.3mm Graphite
2nd layer
15mm
PZT
P
Water
Cyanoacrylate
20um 1st layers
Fig. 14. Construction of transducers.
d: distance between transducers
1
d=30mm
Normalized Pressure
0.9
0.8
d=15mm
0.7
0.6
0.5
0.4
Acoustical Axis
-6
-4
-2
Tx
Source
Xg
d
Rx
Rectifier
Load
27
d=10mm
d=15mm
Pin
Pout
AC
23
DC
19.2
d=20mm
Constant
Pin=260mW
d=25mm
d=30mm
15.4
11.5
lm
7.69
3.85
Fig. 17. Measured DC load power as a function of lateral non-overlapping and
distance d (power transferred through pig muscle tissue).
center resonance frequency. Fig. 18 shows a maximum reduction
of 28% of the load power due to unmatched excitation frequency.
The high efficiency (88.5%) passive rectifier illustrated in Fig. 12
was built and typical waveforms were measured and recorded. The
rectifier 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 measured
anode voltages of the rectifier diodes, and the voltage, current, and
power at the rectifier’s input. The value of Vg was measured by an
open circuit experiment. The component’s values of the source
impedance were evaluated by measuring the voltage and current
under short circuit conditions (across the receiver output terminals) and also while loaded by a 500 X resistor (higher than the
magnitude of the transducer’s impedance which is about 230 X
in order to obtain two measurement results; below and above
the transducer’s impedance).
8. Discussion
0.3
0.2
-8
d=5mm
inAC
1.1
Fig. 16. Measured DC load power as a function of non-overlapping lm and distance
d (at input power, Pin = 260 mW).
oDC
generator (manufactured by Agilent Technologies, Santa Clara California, USA). Electrical waveforms were recorded using a TDS5054,
500 MHz, four channel oscilloscope along with P5050 voltage
probes and CT-2 current probes with a transducer sensitivity of
1 mV/mA (all of which are manufactured by Tektronix, Oregon,
USA).
The hydrophone probe was moved laterally away from the
acoustic 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 UTET
operated in a Constant Wave CW regime to limit the peak pressure
that the tissue is exposed to below the safe level dictated by the
FDA [26,28].
Regarding the Mechanical Index MI, since the peak pressure
measured was 120 kPa, the calculated MI is 0.14 (at 673 kHz)
which is lower than the FDA limit of 1.9 (abdominal). While keeping the same excitation level and frequency of 673 kHz (input
power of 260 mW), the dependency of power transfer through
the water medium on lateral non-overlapping of the transducers
lm was measured at 3 distances (5 mm, 15 mm, 30 mm) and the
results are presented in Fig. 16.
The power transfer through pig muscle tissue was measured as
well [25]. The tissue was immersed in the test tank and placed between the transmitter and receiver. Fig. 17 shows the measured
output power as a function of tissue thickness (1.8 dB/cm at
1 MHz) and lateral non-overlapping. The power transfer is also
sensitive to the accuracy of the excitation frequency.
The transducers were aligned accurately and placed in the test
tank at a distance of 20 mm. The test tank was filled with distilled
water, and a sample of pig muscle 20 mm thick was immersed in
the test tank between the transducers. The sensitivity of the extracted load power to the accuracy of the excitation frequency
was measured by sweeping the excitation frequency around the
0
2
4
6
8
Lateral Distance From Acoustical Axis [mm]
Fig. 15. Measured pressure as a function of distance from acoustic axis at distances
d of 15 mm and 30 mm from transducer radiating surface.
This paper proposes a transcutaneous energy transfer method
using propagating ultrasonic waves. The objectives of the work
were to show that such power transfer is feasible to a range of
up to 40 mm, and measure its sensitivity to the distance, non-overlapping of the transducers, and inaccuracy of the frequency of the
S. Ozeri, D. Shmilovitz / Ultrasonics 50 (2010) 556–566
1.1
1
Normalized Load Power
0.9
0.8
0.7
0.6
0.5
0.4
0.3
0.2
668
669
670
671
672
673
674
675
676
677
678
Excitation Frequency [kHz]
Fig. 18. Measured load power sensitivity to excitation frequency accuracy.
565
to the distance. Fig. 16 shows the measured power through water
at various distances and lateral non-overlapping. Although water
has practically negligible absorption at 673 kHz, peak power transfer efficiency achieved was 0.38 which is the result of working at
matched load condition at the receiver, limiting the maximum possible efficiency to 50%. Power transfer is also sensitive to lateral
non-overlapping caused by moving transducers as illustrated in
Fig. 17. As might be expected, the effectiveness of the device is
greatly adversely affected in the combined situation of distance
and lateral non-overlapping. Although a medium excitation frequency of 673 kHz was chosen (Fig. 8 recommends a frequency
range of 200–1200 kHz lower than used by ultrasonic imaging),
the power transfer efficiency (including rectifier loss) through pig
muscle dropped to 27% and even lower (Fig. 17) due to the fact that
pig abdominal muscle tissue has more than two orders of magnitude higher attenuation compared to water. It explains the added
loss between the results of Fig. 16 (using water) to those depicted
at Fig. 17 (using pig muscle). In addition, the design of the transducer matching layer assumed a medium acoustic impedance of
1.5 MRayl, but pig abdominal muscle used has a mean value of
1.63 MRayl. The slight impedance mismatch probably caused a
small part of the energy to be reflected back at the receiver–tissue
interface contributing to about 15% less power being captured by
the receiver. Fig. 18 illustrates the sensitivity of UTET power transfer to the exact value of the excitation frequency. The frequency
mismatch might be the result of a finite practical accuracy (±1%)
of the manufacturing thickness of the PZT.
Electrical measurements of the rectifier showed a rectification
efficiency of 88.5% as presented in Fig. 19. Rectifier losses consist
of diode losses and inductor losses (mostly copper losses due to
the 10 X series resistance of the small 0805 size RF coils manufactured by Taiyo-Yuden, Tokyo, Japan).
The maximum power throughput imposed by bio-safety limits
was also referred to in the context of the UTET mechanism. The
MI (0.14) is well below the safety threshold (1.9). Furthermore,
the FDA’s requirement indicated by the Ispta (intensity, temporal
average), limits the intensity to 94 mW/cm2 for abdominal and
other tissue (but not ophthalmic, which is not relevant for the
UTET). Meeting these limits, a continuous power of 150 mW can
be transferred using transducers having an area of 1.6 cm2.
9. Conclusions
Fig. 19. Measured input voltage, current and power, and anode voltages of the
rectifier.
exciting source. The paper recommends an operating frequency
range taking into account the tissue absorption and size of the
transducers. As can be seen from Fig. 4 the power transfer is sensitive to operating conditions. The pressure field diffracts at distances longer than the Rayleigh distance, but the fact that the
pressure field is not focused, helps to spread the energy over the
entire cross sectional area of the tissue. (Focused pressure field
although beneficial for imaging and treatment purposes, is not recommended for the UTET since it might increase the intensity above
the safety threshold of 94 mW/cm2). The preferred location of the
receiver is at the natural focus of the pressure field (about 25–
30 mm) as illustrated by Figs. 4 and 15. Load power is sensitive
The ultrasonic transcutaneous energy transfer (UTET) approach
is proposed for the purpose of energizing implanted devices. Measurement results at operating frequency of 673 kHz indicate that
the UTET is an effective method for transferring power at a density
less than 94 mW/cm2 while complying with tissue safety limits. A
power transfer efficiency of 27% is achieved through lossy pig muscle tissue. However, the power transfer efficiency is quite sensitive
to the distance and non-overlapping of the transmitter and receiver. Finally, receiver power conditioning circuitry which is essential for high power transfer efficiency is proposed which gave
88.5% rectification efficiency.
Acknowledgement
This research was partially supported by Israel Ministry of Science and Technology under Grant No. 3-4772.
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