U.S. patent application number 14/475295 was filed with the patent office on 2015-03-19 for high power ultrasound wireless transcutaneous energy transfer (us-tet) source.
The applicant listed for this patent is PIEZO ENERGY TECHNOLOGIES, LLC. Invention is credited to Leon J. Radziemski, Inder Raj Singh Makin.
Application Number | 20150080639 14/475295 |
Document ID | / |
Family ID | 52597761 |
Filed Date | 2015-03-19 |
United States Patent
Application |
20150080639 |
Kind Code |
A1 |
Radziemski; Leon J. ; et
al. |
March 19, 2015 |
HIGH POWER ULTRASOUND WIRELESS TRANSCUTANEOUS ENERGY TRANSFER
(US-TET) SOURCE
Abstract
A bio-implantable energy capture and storage assembly is
provided. The assembly includes an acoustic energy transmitter and
an acoustic energy receiver. The acoustic energy receiver also
functions as an energy converter for converting acoustic energy to
electrical energy. An electrical energy storage device is connected
to the energy converter, and is contained within a bio-compatible
implant for implantation into tissue. The acoustic energy
transmitter is separate from the implant, and comprises a
substantially 2-dimensional array of transmitters. The acoustic
energy converter may also provide conditioned power directly to a
load, connected to said energy converter.
Inventors: |
Radziemski; Leon J.;
(Tucson, AZ) ; Singh Makin; Inder Raj; (Tucson,
AZ) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
PIEZO ENERGY TECHNOLOGIES, LLC |
Tucson |
AZ |
US |
|
|
Family ID: |
52597761 |
Appl. No.: |
14/475295 |
Filed: |
September 2, 2014 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13734817 |
Jan 4, 2013 |
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14475295 |
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61585101 |
Jan 10, 2012 |
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Current U.S.
Class: |
600/16 |
Current CPC
Class: |
A61M 2205/8237 20130101;
A61N 1/3787 20130101; A61N 1/37217 20130101; A61M 1/127 20130101;
H02J 50/80 20160201; H02J 7/00034 20200101; A61M 1/12 20130101;
A61M 1/1086 20130101; H02J 7/025 20130101; H02J 50/90 20160201;
H02J 50/15 20160201 |
Class at
Publication: |
600/16 |
International
Class: |
A61M 1/12 20060101
A61M001/12; A61M 1/10 20060101 A61M001/10 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] This invention was made with Government support under grants
number 1R43EB007421-01A1 and number R44EB007421 awarded by the
National Institutes of health. The government has certain rights in
the invention.
Claims
1. A bio-implantable energy capture and storage assembly, for
implantation into tissue comprising: i. an acoustic energy
transmitter and an acoustic energy receiver, said acoustic energy
receiver also functioning as an energy converter for converting
acoustic energy to electrical energy at a rate of at least 5 Watts;
ii. an electrical energy storage device electrically connected to
said energy converter, wherein said acoustic energy
receiver-converter is contained within the device for implantation
in said tissue; iii. an acoustic energy transmitter external to the
body and separate from said implant, and wherein the transmitter
comprises at least one transducer comprising a 2-dimensional array
of elements arranged in a substantially regular 2-dimensional
geometric shape on a support; and iv. a cooling system including a
circulating coolant and heat exchanger for removing heat from the
energy transmitter, tissue, and implant at a rate of at least 3
Watts; wherein said transmitter operates at a frequency in the
range of 0.75 to 1.5 MHz.
2. The bio-implantable energy capture and storage assembly of claim
1, wherein said substantially regular 2-dimensional geometric shape
is selected from the group consisting of a circle, a rectangle, a
square, a pentagon, a hexagon and an octagon
3. The bio-implantable energy capture and storage assembly of claim
1, further including a wireless feedback loop between said implant
and transmitter used to optimize or stabilize the output power,
with an algorithm using successively smaller scanning steps for
monitoring one or more parameters related to an output power of the
receiver.
4. The bio-implantable energy capture and storage assembly of claim
1, wherein the transmitter operates at a frequency of 0.9 to 1.1
MHz.
5. The bio-implantable energy capture and storage assembly of claim
1, further including a heat-pipe device for cooling the energy
transmitter, the tissue and the implant.
6. The bio-implantable energy capture and storage assembly of claim
3, further including sensor transmitters and receivers on the
acoustic energy transmitter, connected in said feedback loop.
7. The bio-implantable energy capture and storage assembly of claim
6, wherein said sensor transmitters and receivers comprise
ultrasonic elements.
8. The bio-implantable energy capture and storage assembly of claim
3, wherein the 2-dimensional array performs lateral alignment by
electronically determining the minimum number of elements to be
powered resulting in maximum power delivery to the energy converter
within the receiver.
9. The bio-implantable energy capture and storage assembly of claim
3, wherein the 2-dimensional array performs angular alignment by
electronically scanning the ultrasound beam over the face of the
receiver determining the beam angle at which the power delivery is
maximized to the energy converter in the receiver.
10. The bio-implantable energy capture and storage assembly of
claim 3, wherein the 2-dimensional array relaxes the criteria on
angular alignment by substituting the width of an array element for
the width of the entire array, thus relaxing the criteria for
angular alignment.
11. A bio-implantable energy capture and storage assembly for
implantation into tissue comprising: i. an acoustic energy
transmitter and an acoustic energy receiver, said acoustic energy
receiver also functioning as an energy converter for converting
acoustic energy to electrical energy at a rate of at least 5 Watts;
ii. an electrical energy storage device electrically connected to
said energy converter, wherein said acoustic energy
receiver-converter is contained within the device for implantation
in said tissue; iii. an acoustic energy transmitter external to the
body and separate from said implant, and wherein the transmitter
comprises at least one transducer comprising a 2-dimensional array
of elements arranged in a substantially regular 2-dimensional
geometric shape on a support; iv. a cooling system including a
circulating coolant and heat exchanger for removing heat from the
energy transmitter, tissue, and implant at a rate of at least 3
Watts; wherein said transmitter operates at a frequency in the
range of 0.75 to 1.5 MHz, and v. a device for providing conditioned
power directly to a load, connected to said energy converter,
wherein said acoustic energy receiver-converter is contained within
a biocompatible implant for implantation in said tissue, wherein
said acoustic energy transmitter is separate from said implant.
12. The bio-implantable energy capture and storage assembly of
claim 11, wherein said substantially regular 2-dimensional
geometric shape is selected from the group consisting of a circle,
a rectangle, a square, a pentagon, a hexagon and an octagon
13. The bio-implantable energy capture and storage assembly of
claim 11, further including a wireless feedback loop between said
implant and transmitter used to optimize or stabilize the output
power, with an algorithm using successively smaller scanning steps
for monitoring one or more parameters related to an output power of
the receiver.
14. The bio-implantable energy capture and storage assembly of
claim 11, wherein the transmitter operates at a frequency of 0.9 to
1.1 MHz.
15. The bio-implantable energy capture and storage assembly of
claim 11, further including a heat-pipe device for cooling the
energy transmitter, the tissue and the implant.
16. The bio-implantable energy capture and storage assembly of
claim 13, further including sensor transmitters and receivers on
the acoustic energy transmitter, connected in said feedback
loop.
17. The bio-implantable energy capture and storage assembly of
claim 16, wherein said sensor transmitters and receivers comprise
ultrasonic elements.
18. The bio-implantable energy capture and storage assembly of
claim 13, wherein the 2-dimensional array performs lateral
alignment by electronically determining the minimum number of
elements to be powered resulting in maximum power delivery to the
energy converter within the receiver.
19. The bio-implantable energy capture and storage assembly of
claim 13, wherein the 2-dimensional array performs angular
alignment by electronically scanning the ultrasound beam over the
face of the receiver determining the beam angle at which the power
delivery is maximized to the energy converter in the receiver.
20. The bio-implantable energy capture and storage assembly of
claim 13, wherein the 2-dimensional array relaxes the criteria on
angular alignment by substituting the width of an array element for
the width of the entire array, thus relaxing the criteria for
angular alignment.
Description
CROSS REFERENCE TO RELATED APPLICATION
[0001] This application is a continuation-in-part of our co-pending
U.S. application Ser. No. 13/734,817, filed Jan. 4, 2013, which
application in turn claims priority from U.S. Provisional
Application Ser. No. 61/585,101, filed Jan. 10, 2012, the contents
of which are incorporated herein in their entireties.
BACKGROUND OF THE INVENTION
[0003] The present invention relates to systems for powering
implanted devices. The invention has particular utility for systems
for powering current implanted devices requiring 24/7 operation and
tens of watts of electrical power for applications such as
heart-assist devices and will be described in connection with such
utility, although other utilities are contemplated.
[0004] The present invention addresses a critical barrier to a
major increase in the availability of heart assist-devices to
patients in need: the present method of providing 24/7 continuous
power to these devices, either as bridges to transplants or as
permanent implants. Many publications cite the shortcomings of the
existing method, which uses percutaneous links to provide
electrical power to Mechanical Circulatory Support Systems (MCSS),
for example, left- or right-Ventricular Assist Devices (VADs).
These include Slaughter and Myers (2010), Si et al (2008), Danilov
et al (2008), and Franco and Vernier (2003). The percutaneously
placed wires provide pathways for infection (Franco and Verrier
place the rate of infections at 40-45%), they periodically break,
potentially create adhesions, and they limit the life style of
patients because of measures they must take to avoid infections. A
2001 news release from NIH about the 1998-2001 REMATCH clinical
study of percutaneously powered LVADs cited the probability of
infection within 3 months of implantation to be 28%. As a result,
at this time, the use of VADs is limited to bridge-to-transplant
patients, those with extreme loss of heart capability.
[0005] Wireless Transcutaneous Energy Transfer (TET) across tissue
is the much-preferred, less-invasive method of providing power to
these devices. The impacts of a TET power system are that it 1)
overcomes a major disadvantage of the present percutaneous method
of providing power, namely high susceptibility to infection,
opening up a lifesaving technology to hundreds of thousands who
suffer from heart failure, and 2) supports the increased use of
presently implanted heart-assist devices, and 3) fosters new
devices targeted to improving human health.
[0006] Powering of MCSS over long periods of time solely by
implanted batteries is not possible with the batteries available
today because of the continuous high power requirement, which in
turn dictates a large storage capacity and heavy battery. A TET
system could deliver power directly to the application, while also
charging an implanted battery which could take over for periods of
1 to 2 or a few hours. Over the past 50 years much effort has been
expended in trying to make an electromagnetic method of TET
(EM-TET) work for MCSS. U.S. Pat. No. 6,579,315, to Weiss discloses
an EM-TET system for an artificial heart. U.S. Pat. No. 5,630,836
to Prem discloses an EM-TET system for both an artificial heart and
a ventricular assist device. Papers (Mehta et al., 2001; Schuder,
2002; Slaughter and Myers, 2010; Danilov, 2010) disclose elements
of an EM-TET system and even some clinical trials. Nevertheless
only a few if any devices based on this principle are commercially
realized. Issues that hold back EM-TET adoption include 1) heating
of tissue due to misalignment of transmitter and receiver coils
which expose metal to magnetic and electric fields that cause
eddy-current heating, 2) heating due to losses in the coils, 3)
loss of transmission efficiency with depth of penetration, due to
decreased coupling of transmitter and receiver, and 4) decoupling
due to perturbation of the inductance of the coils when they
interact with nearby metallic or magnetic materials.
[0007] U.S. Pat. No. 8,082,041 (Radziemski) describes an ultrasound
system suitable for providing low power to devices such as
pacemakers, defibrillators and neurostimulators, primarily to
recharge implanted batteries. It is well known in the art that
batteries for such low power devices are charged for periods of
minutes to hours at a rate of once per day to once per month or
even less frequently. The patent also contains a description of the
prior art with regard to medical-ultrasound power transmission,
which is included by reference. These applications typically
require a few Watts of input power and typically less than a
half-Watt of power at the point of application and do not require
addressing the new issues which must be resolved for high-power
applications. Specifically the aforesaid patent teaches a
bio-implantable energy capture and storage assembly, including an
acoustic energy transmitter for contact with the skin, and an
acoustic energy receiver converter for converting acoustic energy
to electric energy; a battery or capacitor connected to the energy
converter; signals upon which one may base alignment of transmitter
and receiver; and a method of cooling the assembly. The acoustic
energy receiver/converter, which employs ultrasound, is contained
within a biocompatible implant.
[0008] Although methods for providing signals for alignment of
transmitter and receiver are taught in Radziemski, the actual
physical methods of aligning those elements is not taught. Absent
any electronics to perform the alignment, the only option is that
it would be performed manually, by physically adjusting the
orientation of the external transmitter unit. In fact that is the
present state-of-the-art method for the low power EM-TET method
used commercially. In contrast, here is taught a 24/7
high-cooling-capacity element plus a 24/7 non-mechanical alignment
system. Such an alignment system is required because solely manual
alignment of transmitter and receiver over 24 hours of each day is
wholly impractical and unsafe.
[0009] The cooling methods taught in Radziemski only included
thermoelectric, disposable, or reusable coolers on the transmitter
side, and phase change materials in the receiver. These would not
be useful for the 24/7 continuous high-power operation needed for
MCSS applications. This application requires one to two orders of
magnitude more power than the applications discussed in Radziemski,
typically at this time 10 Watts, or 20 Watts or more of electrical
power at the device to be powered. This in turn, because of the
finite efficiencies of all the steps, requires 40, 50, or 60 or
more Watts of acoustic power from the transmitter unit. These
levels of power require new and novel approaches for safely cooling
tissue. Completely passive cooling methods alone, such as a
disposable liquid coolant pack as taught in Radziemski, cannot
dissipate the substantial 24/7 heat load generated in this
high-power application, because such a pack would need to be
changed and reapplied an undue number of times a day--making it
wholly impractical. Phase change materials in the receiver implant
cannot perform continuous cooling in that location because, once
the transition has been made, they need time and lower temperatures
to regain the previous phase. That time is not available in the
24/7 operation of a MCSS. A thermoelectric cooler as taught in
Radziemski is also unusable in the present application because, as
is well known in the art, it generates heat itself in proximity to
itself and the skin which is dangerous to the patient and adds to
the proximate heat load. Likewise cooling systems that operate in
the transmitter, such as taught in Sliva (U.S. Pat. No. 5,560,362)
do not apply here because those were single-ended systems used in
imaging, with low heat loads and low cooling capabilities, and
designed to operate pulsed with low duty cycles, not 24/7. They
could cool the upper layer of the skin, but would not propagate
deeply enough to cool a heat-source in the receiver and the tissue
adjacent to it, as must be performed in the present application.
Hence there exists a compelling need for an ultrasound delivery
system that can deliver 10's of Watts of electrical power
continuously while providing a) reliable non-mechanical alignment
system, and 2) sufficient cooling capacity to dissipate potential
tissue damage.
[0010] Although the present state of the art is to require 10 or
more Watts to the MCSS, with efficiency improvements in the future,
the requirements could be reduced to 5 Watts or less. Also MCSS
placed in infants or young children may require less power as well.
In those cases the demands on US-TET MCSS power delivery and heat
removal will be correspondingly reduced, for example to 5 Watts and
3 Watts respectively.
[0011] The invention described here is a modality for transferring
energy at a high rate (e.g. power) wirelessly and safely across the
skin in quantities sufficient to directly power energy-intensive
implantable medical devices.
[0012] There are few prior references to using ultrasound as a
carrier of energy at the levels needed in heart assist devices.
Suzuki, et al (2003) describe a hybrid magnetic-ultrasonic device
that employs magnetostrictive materials to generate the pressure
waves that carry energy across the skin. That paper mentions
ultrasound, but refers to a different and more complex system that
only demonstrated .about.5 W of output power. High power ultrasound
non-medical applications are well known in that field.
[0013] An important theoretical and practical advantage of US-TET
is the ability to mitigate the effects of lateral and angular
misalignment by non-mechanical electronic means via a two
dimensional array of transmitter transducers, leading to a
completely self-aligning system that does not require patient
intervention. Also, the ultrasound beam, in the near field which is
our case, does not diverge significantly, hence losses due to depth
of the implant are minimal. Both of these advantages accrue to
ultrasound because of its wave nature, and the fact that for power
transfer, the ultrasound wavelength at useful frequencies is much
smaller than the dimensions of the ultrasound transducers. In
EM-TET the converse is true, ruling out the use of non-mechanical
alignment by this principle. Willis (US2008/0294208) teaches the
use of a two-dimensional ultrasound array to search for a receiver
located in or on the heart and provide pacing level voltages to the
heart wirelessly. Willis (U.S. Pat. No. 8,364,276) estimates the
energy per pacing pulse provided as 0.17 microJoules in a 0.5
millisecond pulse. Assuming a pulse rate of 60 per second, this
converts to an average power of 0.17 microWatts. TET-MCSS
applications need on the order of 10-20 Watts continuously (10-20
Joules per second). Hence Willis' array without cooling could not
be used in the present application. Also, in the MCSS application
there is no need for a location function or signal. Willis is
trying to find a small receiver some variable distance or
orientation with respect to the transmitter array. In the MCSS
application the plane transmitter and receiver faces will likely be
10 to 50 mm apart, and very closely parallel to begin with, their
diameters being up to 75 mm or more, hence significantly larger
than the distance separating them. An unfocused beam will suffice
to correct any misalignment by changing the angle of the
transmitted wave front so that it is incident upon the receiver
closely parallel to the plane face of the receiver, thereby
optimizing power delivery.
[0014] It is thus an object of the present invention to provide new
and novel wireless power transfer techniques which alleviate
distress, pain, complications, and operations associated with
infections suffered by patients who would instead have to use the
present method of power delivery to heart assist devices.
SUMMARY OF THE INVENTION
[0015] The present invention provides a new method and apparatus
for powering an implanted device, such as a heart-assist device,
and more particularly to an ultrasound wireless Transcutaneous
Energy Transfer (US-TET) source generating 40, 50 or 60 or more
Watts of acoustic power to operate an implanted device 24/7. In one
aspect, an external transducer is connected to a battery-driven
controller that modulates the power provided to the transmitter. In
another aspect, the power may be supplied by other means, for
example from an electrical power outlet fixed in a home or any
other location. In another aspect, the power may be supplied
intermittently by an internal battery which is kept charged during
the continuous operation of the transmitter. In this latter case
the internal battery takes over while the depleted external
batteries are being replaced by completely charged ones, or for
occasional bathing or other patient conveniences. The external
controller will receive several radio frequency feedback signals
wirelessly from the implant in order to regulate the transmitter
power and frequency, to stabilize the power provided to the MCSS at
an adequate level, and provide peak power as necessary.
[0016] Traditionally semi-solid ultrasound gels are introduced
between the transducer face and the skin surface to attain adequate
acoustic coupling between the transducer and tissue. This approach
is limiting for the MCSS application since application of the
semi-solid gel is untidy, and over the time of 24/7 operation the
gel can dry. In a preferred embodiment, the transducer surface is
coupled to the skin surface using newer direct-coupling means, such
as a transducer face impregnated with silicone oils, vegetable
oils, castor oil, or one of several other natural and synthetic
media. The front face of the transducer is soft and forms a "boot"
made out of an acoustic impedance matched thickness, using
materials such as polyurethane, polyethylene glycol, polyethylene
oxide, or other materials with acoustic impedance between that of
the piezo-electric material and soft tissue.
[0017] At MCSS power levels, which are high for medical devices, a
light-weight, active or active-passive combination cooling
technique is necessary to keep the temperature of the patient's
skin, intervening tissue, and tissue surrounding the implant within
safe bounds. Well-developed off-the-shelf desktop computer CPU
coolers have the capacity and are of size and weight to be useful
as or models for the 24/7 coolers needed for MCSS. They typically
have power requirements of 2-10 Watts, and so can be powered from
the external batteries that are providing power to the ultrasound
transmitter. The most suitable types are circulating liquid coolers
and heat pipe coolers. They have low-profile round or square
cooling pads that have a circulating working fluid and transport
the heat away to an area where an air-cooled heat exchanger rejects
the heat to ambient. The latter may be achieved passively through
heat sinks or actively through radiator-fan assemblies. Liquid
coolers come with flexible hoses and circulate a liquid working
fluid while the heat pipes typically have rigid pipes and the
working fluid undergoes a phase change in a closed-loop cycle for a
highly efficient heat transport. These coolers can cope with
desktop CPUs which generate between 30 W and 120 W, which as will
be seen below, is more than adequate cooling for this
application.
[0018] Temperature sensing devices within the transmitter and
receiver relay temperatures to the external controller, which will
then apply the correct power to the cooling device in order to keep
the temperature of the transmitter, receiver, and intervening
tissue at safe values. The piezoelectric elements which are the
heart of the transmitter and receiver may geometrically be
monolithic single elements, or a one- or two-dimensional array of
small piezoelectric elements. Capacitively Machined Ultrasound
Transducers (CMUTs), composite or polymer piezoelectric materials,
or other mechanisms for inducing ultrasound vibrations are an
alternative to conventional piezoelectric elements. In a preferred
embodiment, a 2-dimensional array can be used to provide
non-mechanical alignment of transmitter and receiver in response to
optimization signals generated within the implant and relayed back
to the transmitter. Details are discussed below in connection with
the relevant figures. The ultrasound receiver is contained within
an implantable case, the external surface of which is biocompatible
material. It may be implanted at a functionally appropriate
distance below the skin surface, e.g. typically about 10 mm to
about 50 mm below the skin surface, or some distance larger,
between, or smaller than those distances. The front, flat face of
the implant is fixed in the tissue approximately parallel to the
front, flat face of the transmitter. Experiments with curved
transducer faces, such as used for focusing, showed poor efficiency
because of the sensitivity of the curved transmitter and receiver
faces to misalignment. Within the implant case are components for
wireless radio frequency communication with the external
controller, electronics for converting the ultrasound to electrical
power, methods for monitoring the output power, sensors for
monitoring the temperature at various points within the implant,
sensors for monitoring and obtaining the optimum conversion
efficiency, and output devices to 1) an implanted battery and 2)
directly to the implanted MCSS.
[0019] There are two geometrical issues affecting alignment of a
transmitter over a receiver in both the electromagnetic and
ultrasound methods. The first is lateral translation over the
implant, and the second is angular misalignment between the
transmitter and receiver. With ultrasound, the use of an array
transmitter enables compensation for both of these misalignments.
The power out of the receiver, or a quantity proportional to the
power such as voltage or current, is a signal fed back to the
external controller which initiates the algorithm directing the
array transmitter to search for the optimum alignment. In another
embodiment, an imaging ultrasound system is added to the
transmitter unit to provide the feedback on the depth and
orientation of the implanted receiver, thereby assisting alignment.
In another embodiment, some elements on the periphery of the array
can be selectively energized to reflect ultrasound signals off the
nearby receiver, either continuously or in pulse-echo mode. This
will provide separation information in at least four quadrants,
which is then used to correct for misalignment. While such signals
have been used before, they have not been a simple by-product of
having an initial 2-D array, hence their employment in this case
does not require additional parts. Details are discussed below in
connection with the relevant figures.
[0020] The present invention in one aspect provides a
bio-implantable energy capture and storage assembly, for
implantation into tissue comprising: [0021] i. an acoustic energy
transmitter and an acoustic energy receiver, the acoustic energy
receiver also functioning as an energy converter for converting
acoustic energy to electrical energy at a rate of at least 5 Watts;
[0022] ii. an electrical energy storage device electrically
connected to the energy converter, wherein the acoustic energy
receiver-converter is contained within the device for implantation
in tissue; [0023] iii. an acoustic energy transmitter external to
the body and separate from the implant, and wherein the transmitter
comprises at least one transducer comprising a 2-dimensional array
of elements arranged in a substantially regular 2-dimensional
geometric shape on a support; and [0024] iv. a cooling system
including a circulating coolant and heat exchanger for removing
heat from the energy transmitter, tissue, and implant at a rate of
at least 3 Watts; [0025] wherein the transmitter operates at a
frequency in the range of 0.75 to 1.5 MHz.
[0026] In one embodiment the transmitter is comprised of a
2-dimensional array of elements arranged on a support, preferably
selected from the group consisting of a circle, a rectangle, a
square, a pentagon, a hexagon and an octagon.
[0027] In another embodiment, the bio-implantable energy capture
and storage assembly includes a wireless feedback loop between the
implant and transmitter used to optimize or stabilize the output
power, with an algorithm using successively smaller scanning steps
for monitoring one or more parameters related to an output power of
the receiver.
[0028] In one embodiment the transmitter operates at a frequency of
0.9 to 1.1 MHz
[0029] In yet another embodiment, the bio-implantable energy
capture and storage includes a heat-pipe device for cooling the
energy transmitter, the tissue and the implant, and may also
further include sensor transmitters and receivers on the acoustic
energy transmitter, connected in said feedback loop. In such case,
the sensor transmitters and receivers preferably may comprise
ultrasonic elements.
[0030] In still yet another embodiment, the transmitter is
comprised of a 2-dimensional array of elements arranged on a
support, wherein the 2-dimensional array performs lateral alignment
by electronically determining the minimum number of elements to be
powered resulting in maximum power delivery to the energy converter
within the receiver, or the 2-dimensional array performs angular
alignment by electronically scanning the ultrasound beam over the
face of the receiver determining the beam angle at which the power
delivery is maximized to the energy converter in the receiver.
[0031] In another embodiment of the invention the transmitter is
comprised of a 2-dimensional array of elements arranged on a
support, wherein the 2-dimensional array relaxes the criteria on
angular alignment by substituting the width of an array element for
the width of the entire array, thus relaxing the criteria for
angular alignment.
[0032] The present invention also provides a bio-implantable energy
capture and storage assembly for implantation into tissue of a body
of a living animal, comprising: [0033] i. an acoustic energy
transmitter and an acoustic energy receiver, the acoustic energy
receiver also functioning as an energy converter for converting
acoustic energy to electrical energy at a rate of at least 5 Watts;
[0034] ii. an electrical energy storage device electrically
connected to the energy converter, wherein the acoustic energy
receiver-converter is contained within the device for implantation
in said tissue; [0035] iii. an acoustic energy transmitter external
to the body and separate from the implant, and wherein the
transmitter comprises at least one transducer comprising a
2-dimensional array of elements arranged in a substantially regular
2-dimensional geometric shape on a support; [0036] iv. a cooling
system including a circulating coolant and heat exchanger for
removing heat from the energy transmitter, tissue, and implant at a
rate of at least 3 Watts; [0037] wherein said transmitter operates
at a frequency in the range of 0.75 to 1.5 MHz, and [0038] v. a
device for providing conditioned power directly to a load,
connected to said energy converter, wherein the acoustic energy
receiver-converter is contained within a biocompatible implant for
implantation in tissue, wherein said acoustic energy transmitter is
separate from said implant.
[0039] In one embodiment the transmitter is comprised of a
2-dimensional array of elements arranged on a support, preferably
selected from the group consisting of a circle, a rectangle, a
square, a pentagon, a hexagon and an octagon
[0040] In another embodiment, the bio-implantable energy capture
and storage further includes a wireless feedback loop between the
implant and transmitter used to optimize or stabilize the output
power, with an algorithm using successively smaller scanning steps
for monitoring one or more parameters related to an output power of
the receiver.
[0041] In one embodiment, the transmitter operates at a frequency
of 0.9 to 1.1 MHz.
[0042] In another embodiment the bio-implantable energy capture and
storage assembly further includes a heat-pipe device for cooling
the energy transmitter, the tissue and the implant and may also
further include sensor transmitters and receivers on the acoustic
energy transmitter, connected in said feedback loop. In such case,
the sensor transmitters and receivers preferably may comprise
ultrasonic elements.
[0043] In one embodiment of the invention, the transmitter is
comprised of a 2-dimensional array of elements arranged on a
support, wherein the 2-dimensional array performs lateral alignment
by electronically determining the minimum number of elements to be
powered resulting in maximum power delivery to the energy converter
within the receiver, or the 2-dimensional array performs angular
alignment by electronically scanning the ultrasound beam over the
face of the receiver determining the beam angle at which the power
delivery is maximized to the energy converter in the receiver.
[0044] In yet another embodiment of the invention, the transmitter
is comprised of a 2-dimensional array of elements arranged on a
support, wherein the 2-dimensional array relaxes the criteria on
angular alignment by substituting the width of an array element for
the width of the entire array, thus relaxing the criteria for
angular alignment.
BRIEF DESCRIPTION OF THE DRAWINGS
[0045] Further features and advantages of the present invention
will be seen from the following detailed description, taken in
connection with the following detailed description, wherein like
numerals depict like parts, and wherein:
[0046] FIG. 1 is a schematic of the system of the present
invention;
[0047] FIG. 2A is a block diagram of the components contained
within the external controller part of the invention, including
non-mechanical alignment, coupling, and cooling;
[0048] FIG. 2B is a block diagram of the components contained
within the transmitter assembly part of the invention;
[0049] FIG. 2C is a block diagram of the components contained
within the implant assembly part of the invention;
[0050] FIG. 3A is a schematic of the transmitter, tissue, and
receiver part of the system of the invention;
[0051] FIG. 3B shows the possible arrangement of the controller,
transmitter, receiver and batteries on the body of a patient,
frontal and side views;
[0052] FIG. 4 shows the typical parts of an ultrasound
transducer;
[0053] FIG. 5 illustrates the efficiency of ultrasound power
transmission as a function of frequency in accordance with the
present invention;
[0054] FIG. 6 show lateral alignment of an array transmitter and
receiver in accordance with the present invention;
[0055] FIGS. 6A-6D show other geometric arrangements for lateral
alignment of an array transmitter;
[0056] FIG. 7A illustrates ultrasound beam turning using an array
of transmitters with a phase difference imposed between the
elements of the array in accordance with the present invention;
[0057] FIG. 7B illustrates the desensitization of power delivery to
angular misalignment as the number of elements in a linear array
increases, for transducers of 25 mm diameter and a frequency of 1
MHz in accordance with the present invention. The narrowest trace
corresponds to one element, the broadest to ten elements;
[0058] FIG. 7C illustrates the desensitization of power delivery to
angular misalignment as the number of elements in a linear array
increases, for transducers of 75 mm diameter and a frequency of 1
MHz in accordance with the present invention. The narrowest trace
corresponds to one element, the broadest to thirty elements;
[0059] FIG. 8 illustrates the effect of cooling on the temperature
at the face of the implant in accordance with the present
invention;
[0060] FIG. 9 is a schematic block diagram of the components of the
wireless communication link in accordance with the present
invention.
DETAILED DESCRIPTION
Overall Assembly
[0061] FIG. 1 is an overall block diagram of an US-TET system in
accordance with the present invention. FIGS. 2A, 2B, and 2C are
block diagrams the items within the external controller 100, the
transmitter assembly 200, and the implant assembly 400 and are
discussed in later sections. Referring to FIG. 1, two possible
sources of power can operate the system. They are either a direct
current (DC) power supply 50 such as a battery, typically worn by
the patient, or a conventional room alternating current (AC) source
51. Circuitry within the external controller 100 determines whether
the input power is low frequency AC. If so, it proceeds through a
DC converter and then through circuitry 120 which converts it to
high frequency ultrasound. The external controller 100 controls the
level of input power, frequency of the ultrasound, alignment
algorithm, and cooling level. These can be operated in two modes,
manually and automatically, the latter via a feedback loop 130 and
450 made possible by the wireless communication system 500, which
has external 150 and internal 430 components. The output of the
external controller 100 is connected to the transmitting assembly
210, which is disposed adjacent to the skin of the subject. After
transmission through human tissue 300 the ultrasound is incident on
the receiver 410, which is disposed on or under the face of the
implant 400 adjacent to internal tissue.
[0062] After conversion back to electrical power via circuitry 420
residing within the implant 400, the power is directed to an
implanted controller which modulates the current and other sensors
for the operation of the MCSS, and as necessary, to replenish an
internal DC source such as a battery. The internal battery is used
to power the MCSS for short periods of time such as a few hours,
while the patient removes the external supply to bathe or for other
conveniences. A radio frequency wireless communication system 500
between the external controller and the implant, such as a Zarlink
or other brand over the 405 MHz medical-band system, provides a
means of monitoring functions of the receiver and implant, issuing
performance commands to the elements within it, and maintaining one
or more feedback loops 130 and 450 for optimization of
performance.
[0063] FIG. 3A shows a schematic arrangement of the
transmitter-tissue-implant part of an US-TET system. The
transmitter transducer 210 transmits acoustic energy which is
continuous, via sine waves, square waves, triangular waves or an
arbitrary repetitive shape. Continuous power in this context is as
opposed to pulsed power, and does not exclude occasional periods of
no power delivery for whatever reason, during which the internally
charged batteries take over the operation of the MCSS. The power is
transmitted wirelessly through an external acoustic-coupling medium
between transmitter and tissue 230. Occasional quasi-continuous
operation simultaneous with or separate from continuous operation,
may be necessary for alignment or other reasons. Essentially all
air preferably will be excluded, between the skin of the patient
and the ultrasound transmitter, since air strongly attenuates
ultrasound over frequencies of 100 kHz. A cooling system 240, is
deployed as schematically shown. During in vivo tests external
cooling has been observed to penetrate the dermis, cooling the
intervening tissue and the implant as well. After penetrating the
epidermis, dermis, and possibly fat and muscle layers, the
ultrasound is incident on a biocompatible implanted container 400
which has the receiver 410 on or against the inside of the front
face, and other elements packaged within it. The receiver
transducer 410 converts the acoustic to electrical energy. This
energy proceeds via the schematically shown power outlet 470, which
leads to the internal controller, power conditioning circuitry, and
then to an application such as the MCSS.
[0064] FIG. 3B shows how the transmitter, receiver, batteries and
controller may be positioned on a person to deliver power to an
MCSS. In the side view, the transmitter unit 210, with input from
the controller 100, transmits the ultrasound through the
intervening tissue to the receiver unit 400. Those units may be
placed in any location on the body, anterior or posterior, found to
be advantageous. In a preferred embodiment the transmitter and
receiver would be above and near the heart for MCSS applications,
so that wires inside the body can be kept short. The front view
shows the straps as in the present EM-TET method. The external
batteries 310 are attached to the straps, and can be easily removed
and replaced with fresh batteries as needed. Depicted also is the
connection between the receiver 400 and an MCSS device 320 which
assists the heart in its operation.
Transmitter and Receiver Ultrasound Transducers
[0065] An ultrasound transducer is a device which converts
electrical energy to vibrational energy, and vibrational energy to
electrical energy useful in the present invention. In its simplest
form (FIG. 4) it is comprised of a piezoelectric material which
changes its dimensions when an electric field is placed across it.
These include ceramic, crystalline, composite and polymer
piezoelectrics. Other materials may be used, such as
magnetostrictive materials or CMUTs. In one embodiment, a
piezoelectric disk 211 comprised of a ceramic matrix in which are
embedded crystals of Lead-Zirconium-Titanate (PZT) can be the basis
of a transducer. Other materials such as crystalline
Lead-Magnesium-Niobate in Lead-Titanate (PMN-PT) may also be used.
In general the ultrasound transducer may be a single element, or an
array of individual elements. The piezoelectric surfaces are coated
with a conducting film to which electrodes are attached and which
carry the electromagnetic wave to the material, causing it to
shrink or expand slightly at the frequency of the wave. The disk
normally has a backing to augment the conversion, and is housed in
a case made of plastic or aluminum or titanium or other material.
In the implant, the disk and an impedance matching layer are
preferably bonded directly to the inner face of a titanium implant
case 400 which contains all the components of the implanted device,
and which is hermetically sealed. The element between the disk and
the medium through which the vibrations are passing has a thickness
such as to minimize the reflection of the wave, typically a quarter
or full wave thick, and possibly comprised of multiple layers.
[0066] The transmitter 210 and receiver 410 transducers may have a
high-Q (narrow bandwidth) and be designed and manufactured to have
closely matched resonance frequencies. In a second embodiment, one
of the units may have a high-Q resonant frequency and the other a
lower-Q wider bandwidth resonance, making the combination less
sensitive to temperature-induced changes of frequency in either
unit. In a third embodiment, both units may have a lower-Q and
wider bandwidth. It is well known to those skilled in the art that
maximum electrical or acoustic power is transferred between two
objects when their electrical and acoustical impedances are matched
(Woodcock, 1979). Optimization of the transducer impedances is
assisted by impedance matching software and accomplished with the
addition of inductive and capacitive elements in the transmitter
and/or receiver circuits.
[0067] The operating frequency of the transducers is determined by
a variety of constraints. At too low a frequency, below 500 kHz,
there is the increased probability of cavitation which can lead to
embolisms. At higher frequencies above 1 MHz, the absorption of
tissue increases considerably, and the transducer element becomes
quite thin. A series of experiments whose results are shown in FIG.
5 determined that an optimum frequency is in the range of 0.75 to
1.5 MHz. A narrower band of approximately 200 kHz centered on 1 MHz
is an adequate working range within the wider band, allowing slight
changes in operating frequency with changes in temperature. In
addition to the resonant frequency, the bandwidth is also an
important transducer parameter. Too small a bandwidth, such as in
the kilohertz range, can lead to a lack of overlap of the
transmitter and receiver resonant frequencies due to differential
heating of transmitter and receiver during operation, with a
consequent loss of transmission efficiency. This is an important
tradeoff, central for efficient operation at a constant power
level.
Design of Safe High-Power Transmitter and Receiver Transducers
[0068] A primary consideration in wireless transmission of power
through tissue, whether it be electromagnetic or ultrasound, is the
avoidance of tissue damage. There are well known guidelines to
achieve this for pulsed ultrasound applied to fetal tissue, keeping
the acoustic intensity at the skin at or below a maximum of 0.7
W/cm.sup.2 (AIUM, 1993; Hedrick, 2005; NCRP Report 113, 1992). This
is a very conservative value adopted to avoid significant
temperature rise in critical tissue structures in the fetus during
pulsed obstetrical imaging. Adoption of this metric for our
continuous power delivery, dictates, for a given input electrical
power, the minimum area of a transmitter that applies the power to
a patient.
[0069] An example calculation of a sufficiently large transducer
area follows. Assume a conversion efficiency of electrical to
ultrasound power of 70%. Then 1 W/cm.sup.2 electrical intensity
would produce 0.7 W/cm.sup.2 of acoustic intensity. In passage
through one cm of tissue at 1 MHz about 20% of the acoustic energy
would be absorbed. The efficiency reconversion to electrical energy
at the receiver is assumed be the same, 70%. The total efficiency
then is 40%. In the experimental table shown in the section on
cooling below (Table 2), efficiencies at high powers measured in
proof of principle experiments averaged 30%. Likely sources of
other losses are reflection from interfaces between different
tissue layers and between tissue and the solid surfaces of the
transducers. Assume that 20 Watts of electrical power is necessary
to operate the MCSS. That places a requirement of just under 70
Watts of electrical power at the transmitter, 50 Watts acoustic
power. This requires a transmitter area of 70 cm.sup.2 (diameter of
9.5 cm) to keep the acoustic intensity at 0.7 W/cm.sup.2. An
additional metric for device safety is that tissue temperature
increase due to the TET system application be less than 2.degree.
C. That metric is met by having sufficient cooling capacity.
Another safety concern is mechanical particle motion. Using
conventional expressions for the relationship between ultrasound
intensity and particle motion in water (analogous to soft tissue),
at 0.7 W/cm.sup.2, particle motion is calculated to be a very small
amount.
[0070] The main non-thermal possibility for tissue damage arises
from cavitation, rapid expansion and contraction of air bubbles,
primarily in the lungs. The probability for this effect increases
with ultrasound frequencies below 500 kHZ, and for locations where
ultrasound can interact with lung tissue. Avoiding such locations
and using a frequency around 1 MHz minimizes this possibility.
External Controller
[0071] As shown in FIG. 2A, the external controller 100 contains a
variety of components. When converting from input DC power, it goes
through a DC to DC converter 105 to bring it to a range of useful
current and voltage. It then proceeds to a signal generator 120
such as a variable frequency oscillator or a synthesized signal
generator to condition it to the frequency of interest. When
converting from input alternating current, which may be 120 V, 60
cycle or some other normally used combination, first the electrical
power goes through an AC to DC conversion 105, and then follows the
steps outlined above for a DC power source. In both cases the power
at the appropriate ultrasound frequency then proceeds through an
amplifier 110 to bring it to the level required for the
application. The power level can be set manually by an input
command, or be placed under the control of a feedback loop 130 and
450 which keeps it at the specified value. A useful feedback
parameter, whose value is relayed from the implant to the external
controller, is the output power from the ultrasound receiver.
Typically it would be desirable to keep the output power stable for
optimum operation of the application.
[0072] A second important function of the controller is to monitor
and change the frequency of the ultrasound. Typically the range of
changes are approximately 10% of the resonant frequency, and this
is achieved via a variable frequency oscillator 120 or a
synthesized signal generator 120, methods well known to those
skilled in the art. The frequency can be set manually with an input
command, or can be placed under the control of a frequency feedback
loop 130 and 450.
[0073] Two other important functions are a) monitoring and aligning
the transmitter and receiver faces non-mechanically, b) controlling
the cooling mechanism to regulate the heat removal needed for safe
operation.
[0074] Embedded in the controller is the radio frequency antenna
150 which enables reception of communications from the implant on a
medical communication band. These include receiving values of
temperatures 140 being monitored in various implant locations,
monitoring the efficiency of power conversion 140, and monitoring
transmitter and receiver unit alignment. In one embodiment, a
hybrid National Instruments Signal Express plus C++ code collects
and stores the data automatically and continuously for up to 10
parameters, both for patient information on a user interface 160
and for periodic diagnostic downloading. The latter allows a
variety of charts, comparisons, and figures of merit to be recorded
and analyzed, to monitor the performance of the system.
[0075] Software compares the temperature readings with a preset
regime of safe temperatures and, if necessary, sends a warning to a
user interface 160, similar to a smart phone, which allows the
patient to monitor power efficiency and receive safety warnings.
The user interface communicates with the controller using a
wireless protocol, such as Bluetooth, Wi-Fi, or other advanced
method.
Transmitter Unit and Components
[0076] As shown in FIG. 2B, power from the external controller 100,
at an ultrasound frequency, proceeds to the transmitter assembly
200 and transmitter transducer 210. The transmitter transducer is
preferably a two dimensional array of elements. This activates the
transmitter transducer 210 to convert electrical power to
ultrasound for transmission through human tissue 300. The
transmitter alignment stage 220 contains a method of being fixed to
the patient, a manual adjustment method to approximately align the
transmitter and receiver faces, a non-mechanical adjustment
algorithm and electronics to complete the alignment of the wave
front from the transmitter parallel to the receiver face, a space
for an element 230 which excludes air between the ultrasound
transmitter and the skin of the patient, and a cooling method 240.
The alignment stage may be fixed to the skin by means of a double
sticky tape on the bottom or over the top of the alignment stage
(Mehta et al., 2001, FIG. 3). Another embodiment has a strap or
holster in addition to or in place of the sticky tape to secure the
transmitter unit to the skin. Another embodiment attaches the stage
via a slight suction generated by a boot and clamp method, as used
for affixing items to the inside of an automobile windshield. The
manual adjustment method, in one embodiment, is comprised of a
platform with three screws of fine pitch set in a triangle, which
aligns the platform angularly over the implant. Initial lateral
alignment is performed over the slight protrusion of the implant
which rises from a few millimeters to one centimeter or more over
the adjacent tissue. A lightweight cone on the bottom of the
alignment platform may fit over the protrusion, ensuring secure
lateral alignment.
Implant Unit and Components
[0077] FIG. 2C is a block diagram of the components of the implant
assembly. FIG. 3 illustrates the placement of the implant 400
connected to the tissue 300. The piezoelectric element 410 which is
the key element of the receiver transducer, is placed on the front
face of the implant 400, or underneath it and permanently affixed
to it. Preferably it is a single element transducer, although an
array may be used in another embodiment. Adjacent to that element
is found circuitry 420 which converts the ultrasound to electrical
power, AC or DC, as required by the application which is receiving
the power. The converted power is monitored 440 and the analog data
stored. Embedded in various locations in the implant will be
thermal sensors 460 which enable the temperatures in those
locations to be monitored. Circuitry for analog to digital
conversion of those data 420 are also embedded in the implant, as
are internal radio frequency wireless communication components 430,
including an antenna. The data so transmitted are the input for the
feedback loop 130 and 450. The external controller 100 then resets
parameters such as power, frequency, and alignment in order to
stabilize the power provided to the internal application.
Non-Mechanical Alignment of Transmitter and Receiver
[0078] Alignment of the transmitter and receiver is an important
issue both in EM-TET and US-TET. Even though the transmitter unit
may be affixed securely to the skin over the implant, it is
possible that the implant could move slightly within the somewhat
flexible tissue in which it is placed. Motion of the patient will
affect the alignment as well. Hence a method of both lateral
translations and angular alignment in the post-implanting phase, is
desirable and necessary. Furthermore, it is desirable that the
methods of alignment not depend on the patient's intervention,
because the system will be required to operate even when the
patient is asleep. Ultrasound provides a method for non-mechanical
alignment not available to EM-TET.
[0079] One dimensional arrays of ultrasound transmitter elements
are well known to those skilled in the art. Their principal
applications are for scanning an ultrasound beam in space to image
structures in the body, and for non-destructive testing of
materials and weld integrities. Two dimensional arrays have been
made as well, and the technology is advancing to make inexpensive
2-D arrays (Ranganathan, et al., 2004; Fuller et al., 2009). Willis
(US2008/0294208) has used a two dimensional array to locate a
deeply embedded receiver and to focus very weak ultrasound energy
on it to provide pacing signals to the heart.
[0080] FIG. 6 shows an arrangement for lateral alignment of a
larger circular 2-D array 215 over a smaller circular receiver 410.
Preferably the 2-D array of the transmitters is arranged on a
circular disk, e.g. as shown in FIG. 6, although other regular 2-D
geometric arrangements, e.g. square, pentagonal, hexagonal,
octagonal, etc., shapes may be used as illustrated in FIGS. 6A-6D.
In the algorithm for lateral alignment, a feedback loop 130 and 450
relays the output power level of the receiver back to the
controller 100 that activates a number of elements in the 2-D array
transmitter 215. The controller 100 activates elements sequentially
along one axis, and then along a second axis perpendicular to the
original direction. In this way the centroid of the active elements
that maximizes or optimizes the output power is obtained. Once the
optimum centroid position is determined, the number of array
elements surrounding that point is increased radially until the
output power plateaus, thus minimizing waster energy. That array of
elements remains activated until a significant departure from the
chosen output power is observed with the feedback loop 130 and 450,
leading to a rescanning. The frequency of rescanning depends on the
rapidity of changes in the lateral position, which is likely to be
slow.
[0081] For angular alignment two effects are considered. The first
of these is the turning of the beam wave front from parallel to the
face of the transmitter array, through an angle that makes the wave
front parallel to the face of the implanted receiver. This
compensates for angular misalignment of the faces of the two
transducers. For two dimensional surfaces this needs to be done
along two axes. It is well known to those skilled in the art that
this is accomplished by embedding a constant time differential,
which results in a phase difference, between each element of the
array. The result is shown schematically in FIG. 7A which
illustrates the beam turning 216 by introducing a constant phase
217 between elements of a one-dimensional array 218.
[0082] The second effect deals with decreasing the sensitivity to
alignment of two plane parallel transducers faces. Maximum power
transfer takes place when the incoming wave is at the same phase at
all points on the receiver. In order to keep the incoming wave from
the transmitter in phase across the face of the receiver, the two
must be aligned to within one-half wavelength. For a frequency of
one MHz in tissue that is approximately 1 mm. This alignment
condition becomes more and more stringent as the diameter of the
transducers increase. For a 10 mm diameter transducer, the
alignment condition is that the two surfaces be parallel to 1 mm
out of 10 mm. For a 70 mm diameter transducer, the condition is 1
mm out of 70 mm. This condition is relaxed for an array because the
width of the array element substitutes for the overall width of the
whole array. An array element width can vary from 0.1 mm to several
millimeters. This relaxation is shown in FIG. 7B in a model-based
calculation result for an ultrasound frequency of 1 MHz. There is
plotted the steered power versus the number of array elements for a
pair of 25 mm diameter transducers, where the transmitter is a
one-dimensional array, and the receiver a monolithic single
element. The narrowest trace is for one element, then follow in
increasing width the traces for 2, 3, and 4 elements. For a single
25 mm diameter transmitter element (the whole transducer), the
power falls to 80% within a degree of misalignment on either side
of the center line. Increasing the number of elements per unit area
to 10 spreads the 80% power cone to .+-.8.degree.. That in turn,
reduces the restriction on the angular alignment to retain 80%
power, to .+-.8.degree.. FIG. 7C shows the result of a calculation
for a 70 mm diameter transmitter array with up to 30 elements, and
a monolithic 70 mm diameter receiver. The narrowest trace is for
one element, then follow in increasing width the traces for 2, 3,
and in sequence up to 30 elements. With 30 elements, the 80% power
level is retained to .+-.10.degree.. By combining the relaxation on
alignment due to the array, with a feedback loop, in one embodiment
based on monitoring the output power of the receiver, a
non-mechanical means of aligning the transmitted wave with the
receiver face has been achieved. This method can be used to
maximize power, or to retain a constant power level which is
slightly below the most efficient operation. Hence alignment
becomes a method to retain a very tight tolerance on the output
power. To be effective in operation, it is necessary to have an
array in two orthogonal directions, able to compensate for angular
displacement along each of two axes. In the algorithm for angular
alignment, a feedback loop 130 and 450 relays the output power
level of the receiver back to the controller 100 that inputs the
phase change from element to element in the 2-D array transmitter
215. The controller 100 inputs a series of phase changes
sequentially along one axis, and then along a second axis
perpendicular to the original direction. In this way the two angles
are determined that maximize or optimize the output power. The
angles thus optimized remain activated until a significant
departure from the chosen output power is observed with the
feedback loop 130 and 450, leading to a rescanning.
The Feedback Loop
[0083] The feedback loop 130 and 450 is illustrated in FIG. 2A and
FIG. 2C, connecting the external controller with the implant. The
basic feedback algorithm used to optimize the position of each axis
of the lateral and angular alignments, and the frequency from the
signal generator, is this. First, the angular or lateral position
for each axis or the frequency is swept across its entire range
with a gross step between each position or frequency, while
measuring the level of the receiver power. Next, the positions and
frequency are again swept but across a smaller range centered
around the best position or frequency from the previous sweeps, and
at a smaller step size. The process is repeated until a very fine
step size thus narrowing in on the optimal frequency or position.
Individual power measurements may vary due to electronic noise
effects. With gross steps, it is easy to measure distinct changes,
but as the step size decreases, the noise floor quickly overcomes
the differences in power created by a change in position or
frequency. To get a finer step size and still be able to discern a
clear change in power, an averaging of ten measurements is useful.
In another embodiment, the averaged measurements were filtered for
each location and frequency. From digital signal processing it is
known that an ideal low pass filter in the frequency domain is a
sine function in the time domain. More formally, given the filter
H(.omega.) defined below for the frequency domain
H ( .omega. ) = { 1 , - .omega. .ltoreq. .omega. c .ltoreq. .omega.
0 , else , ##EQU00001##
the inverse discrete time/space Fourier transform h(n) of the
H(.omega.) is equal to
h ( n ) = sin ( .omega. c n ) .pi. n , ##EQU00002##
where h(n) is the impulse response of the filtering system. This
particular function is known as the sine function. The output is
equal to the convolution of the input with the impulse response.
Since this filter is symmetric, convolution with this filter is
equivalent to cross correlation. Thus, the filtered power at a
particular location or frequency n.sub.0 is
y ( n 0 ) = k = N - n 0 N + n 0 x ( k ) h ( k ) ##EQU00003##
where N+1 is equally to the number of coefficients of the symmetric
filter and x is the signal of measured powers. Such a filter
implementation is clearly not ideal because of the finite filter
length of the filter and the finite precision of the digital
values; however, the power measurements are filtered only to
identify a clear peak in the data. At a low angular cut off
frequency of around 0.5 radians (determined empirically) most of
the AC components of the power measurements are removed. By
implementing this filter as part of the algorithm, an optimal
position for each axis and an optimal frequency are obtained in
which adjustments no longer yield perceivably higher powers.
Cooling
[0084] A considerable amount of the input electrical power to the
transmitter piezo elernent(s) is converted to heat because such
elements are, as known in the art, at best typically 70% efficient
in transduction from electrical to acoustic power. A cooling method
will constrain tissue exposure to high temperature. Cooling was
successfully accomplished in animal studies by circulating water
through conduits around the base of the transmitter assembly as
illustrated in FIG. 3, 240. The method provides cooling even
through the intervening tissue to the bottom of the implant and the
tissue adjacent to it, via conduction. This is illustrated in FIG.
8 which shows the temperatures measured in an in vivo porcine
study, at the top of the implanted receiver, approximately 1 cm
deep into the tissue, without (upper) and with (below) external
water cooling, at .about.120 mA of charging current into the
implanted battery. With the water cooling the temperature of the
tissue exposed was well controlled.
[0085] A calculation and experimental result will show the order of
magnitude of the expected heat load. The Table 1 below shows an
estimate of the power lost to heat in the two conversions and
through 1 cm of tissue, with an input electrical power of 50 Watts.
In this case the result is and efficiency of 40%, and 30 Watts lost
to heat.
TABLE-US-00001 TABLE 1 lost to heat A calculation of heating. Watts
input electrical (Watts) 50 conversion efficiency (ratio) 0.7
resulting input acoustic power (Watts) 35 15 transmission of power
through 1 cm of tissue at 1 MHz 0.8 (ratio) power remaining to
receiver (Watts) 28 conversion efficiency (ratio) 0.7 electrical
power out of receiver (Watts) 19.6 8.4 Total 30.4 (Watts)
To validate the estimates above, many experiments were performed
with 3'' diameter transducers, at input electrical powers of up to
60 Watts, through a 20 mm thick gel pad, while monitoring
temperatures of transmitter and receiver faces with attached
thermocouples. Data from one of these experiments in the Table 2
below illustrates the rapid increase in temperature without
cooling. In 7 minutes the transmitter increased in temperature by
24 C, and the receiver 18 C. The overall efficiencies measured
about 30%, somewhat lower than the 40% calculated. This was likely
due to other losses such as reflections at interfaces.
TABLE-US-00002 TABLE 2 Experimental results on heating. Electrical
Temperature Temperature Electrical Power in transmitter receiver
time Power out efficiency Watts C C Hr:min Watts % 2 20.2 19.9 1:30
0.66 33 10 21.7 21.6 1:32 2.98 30 20 23.3 23.3 1:33 5.99 30 40 29.1
29 1:35 10.58 26 60 44 38 1:37 14.58 24
[0086] Desktop computer CPU coolers are available that are
well-developed off-the-shelf units with 30 to 120 W cooling
capacities that exceed the needs in MCSS applications demonstrated
above. These systems are compact, efficient and relatively quiet in
operation. The circulating pumps are capable of running
continuously in computers for up to six years. (Kang et al.
(2007)). The overall system in Kang et al. included a pump, cold
plates, a heat exchanger and flexible tubing. Liquid cooling
systems can incorporate single phase liquids, or 2-phase media such
as used in heat-pipes. These thermal dissipation schemes are very
feasible in actively cooling a heat source such as the ultrasound
transmitter, either as a single element or, in a multi-element
configuration.
[0087] In a preferred embodiment, the closed-loop liquid cooling
system is attached to the proximal transducer surface or the
housing. A heat dissipating blower fan and fin-array can be used in
the ultrasound source embodiment, without or with the closed-loop
liquid cooling system. In another embodiment, these systems are
split and attached to one or more heat generating surfaces, such as
ultrasound arrays. The current ultrasound MCSS embodiment with
anticipated waste power specifications as calculated above can
easily be accommodated in the design in order to achieve acceptable
source temperatures of 35.degree. C. or lower, over several
years.
[0088] In another embodiment, the above combinations for thermal
dissipation for closed loop circulation, circulating fans, and
conductive fins are augmented by using a refrigerant based
liquid/gas to achieve yet lower temperatures at the ultrasound
source plane adjacent to the skin. This cold front plane propagates
distally to further cool the exposed tissue as well as the receiver
surface. In Kang et al., they show results from a heat-pipe based
system. The input power generated by a CPU chip was approximately
20 W and the heat pipes maintained the temperature typically at
40.degree. C. Without heat pipe operation, the temperature soared
to 90.degree. C. in less than 2 minutes.
[0089] The novelty of this approach to ultrasound cooling lies in
adapting the CPU cooling methods to the MCSS application.
Wireless Radio Frequency Communication System
[0090] The purpose of an RF-Link is to have a wireless,
bi-directional, non-invasive means of communication between a
device implanted in a living human body, and an external
controller. This provides the capability to remotely read out key
parameters in the implant while permanently installed, and control
parameters inside the implant, such as controlling a variable
discharge dummy load to speed up battery discharging. FIG. 9
illustrates schematically the wireless communications RF-link 500
where an external base station 151 in an external controller 100
can communicate bi-directionally in half-duplex mode with the
internal component 430 in the implant 400. The implant transceiver
430 device is paired with a microcontroller for added
functionality. The base station 151 preferably is fitted with the
microcontroller because sufficient power is always available. An
example of a platform suitable for application to implants in
humans, is Zarlink's medical implant communications service (MICS)
band transceivers ZL70102. MICS is the industry standard for
medical implants. It specifies low-power devices operating in the
400 MHz band without license requirement. Operating in the
industrial, scientific and medical (ISM) band at 2.45 GHz is also
license-free.
[0091] The system consists of a base station module 151, an implant
module 430 and the required software package to control the system
and communicate with the user interface. The hardware uses two
microprocessors for the base station transceiver and two
microprocessors for the implant transceiver. Zarlink provided the
source code starting point, a software package that contains
firmware for the microprocessors and an elaborate graphical user
interface (GUI) that allows control of all features of the entire
system from low-level bit addressing of registers to
impedance-matching of the RF stages. The code is written in Visual
C# and developed on the integrated development environment (IDE)
Microsoft Visual Studio 2008.
[0092] The Zarlink chip uses a 2.45-GHz wake-up subsystem
consisting of the 2.45-GHz receiver and the wake-up controller,
plus an ultra-low-power, 25-kHz strobe oscillator that can be used
for timing purposes. The wake-up controller is a digital subsystem
that identifies when the implant module 430 receives a valid
2.45-GHz wake-up data packet from the base station 151, which is
unique for a particular implant. The wake-up controller then powers
up the media access controller (MAC) 431 and the 400-MHz
transceiver 432, so that the implant can respond on 400 MHz and
establish a two-way MICS-band link with the base station 151. While
the 400-MHz link is operative, the 2.45-GHz wake-up subsystem is
powered down. When the implant reverts to the sleep state, the
2.45-GHz wake-up subsystem is periodically re-enabled to listen for
any possible wake-up transmissions.
[0093] In the base station 151, the MAC 152 provides a modulation
signal for the external 2.45-GHz wake-up transmitter 153. The
ZL70102 154 has features to facilitate and optimize a 400-MHz
wake-up mode. A key feature of the ZL70102 is a fast received
signal strength indicator (RSSI) sniff function that is optimized
for sniffing and that leaves out operations that are required only
for a normal wake-up. The bulk data communication takes place in
the 400 MHz band while the wake-up calls are made in the 2.45 GHz
band. The reason for the lower frequency for bulk communication is
that 2.45 GHz electromagnetic waves experience significant
absorption in body tissue, which is mainly water. With less loss at
400 MHz the transmitter power requirements are significantly less,
an important feature for extending battery life.
[0094] When the implant 430 correctly receives the 2.45-GHz wake-up
transmission from the base station 151, it responds using its
400-MHz transceiver 432. Therefore an on-chip, 2.45-GHz transmitter
152 is not needed. The base station 151 uses an external 2.45-GHz
Wake-Up Transmitter module, which is controlled jointly by the
application processor and the ZL70102 154. The wake-up function
uses 2.45 GHz because the band is internationally designated as an
ISM frequency band and so is more generally available on an
international basis at a higher power level than other frequency
ranges. The use of a higher transmitter power allows a reduction in
the sensitivity of the wake-up receiver. Also, the use of a higher
frequency tends to increase the received power available from the
antenna, although this advantage is partly offset by the increased
loss within the patient's body at 2.45 GHz. Taking all these
factors into consideration, the overall result is a significant
advantage in using 2.45 GHz. Zarlink recommends operation under the
requirements for wideband data transmissions, as opposed to RFID
regulations, since the allowable spectrum mask limits permit a
faster rise time for the 2.45-GHz on/off keying. When operating
under regulations for wideband data transmission, it may be
necessary to provide frequency hopping in the 2.45-GHz transmitter
152. The bandwidth of the 2.45-GHz wake-up receiver in the ZL70102
433 is large enough that a substantial frequency spread can be used
without loss of sensitivity caused by the mistuning of the input
network.
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