U.S. patent application number 12/204470 was filed with the patent office on 2009-01-01 for endoluminal implant with therapeutic and diagnostic capability.
This patent application is currently assigned to CARDIOMETRIX, INC.. Invention is credited to George W. Keilman.
Application Number | 20090005859 12/204470 |
Document ID | / |
Family ID | 21841871 |
Filed Date | 2009-01-01 |
United States Patent
Application |
20090005859 |
Kind Code |
A1 |
Keilman; George W. |
January 1, 2009 |
ENDOLUMINAL IMPLANT WITH THERAPEUTIC AND DIAGNOSTIC CAPABILITY
Abstract
An apparatus includes an endoluminal implant, a RF coupling coil
coupled to the endoluminal implant and a therapeutic transducer
electrically coupled to the RF coupling coil and physically coupled
to the endoluminal implant. The RF coupling coil supplies
electrical power to the therapeutic transducer. The therapeutic
transducer has a capability for delivering therapeutic energy to a
lumen disposed within the endoluminal implant in response to
signals coupled via the RF coupling coil.
Inventors: |
Keilman; George W.;
(Woodinville, WA) |
Correspondence
Address: |
DAVIS WRIGHT TREMAINE, LLP/Seattle
1201 Third Avenue, Suite 2200
SEATTLE
WA
98101-3045
US
|
Assignee: |
CARDIOMETRIX, INC.
Bothell
WA
|
Family ID: |
21841871 |
Appl. No.: |
12/204470 |
Filed: |
September 4, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11551673 |
Oct 20, 2006 |
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12204470 |
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09695748 |
Oct 24, 2000 |
7427265 |
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11551673 |
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09028154 |
Feb 23, 1998 |
6231516 |
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09695748 |
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08978038 |
Nov 25, 1997 |
5967986 |
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09028154 |
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08949413 |
Oct 14, 1997 |
5807258 |
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08978038 |
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Current U.S.
Class: |
623/1.42 ;
128/898 |
Current CPC
Class: |
A61N 5/062 20130101;
A61B 5/0031 20130101; A61F 2/07 20130101; A61B 8/56 20130101; A61B
8/06 20130101; A61F 2/82 20130101; A61F 2250/0002 20130101; A61B
5/14532 20130101; A61M 37/0092 20130101; A61F 2002/072 20130101;
A61M 1/3656 20140204; A61N 1/306 20130101; A61B 8/12 20130101; A61B
8/4483 20130101; A61B 5/6876 20130101; A61N 5/0601 20130101; A61F
2250/0001 20130101; A61N 1/325 20130101; A61B 5/6862 20130101; A61F
2002/075 20130101; A61B 2560/0219 20130101; A61B 8/04 20130101 |
Class at
Publication: |
623/1.42 ;
128/898 |
International
Class: |
A61F 5/00 20060101
A61F005/00; A61B 19/00 20060101 A61B019/00 |
Claims
1. A method comprising: coupling, via a magnetic signal, a signal
to a therapeutic transducer contained in an endoluminal implant;
and activating said therapeutic transducer in response to said
magnetic signal.
2. The method of claim 1 wherein activating said therapeutic
transducer includes activating a light source.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of pending U.S. patent
application Ser. No. 11/551,673 filed Oct. 20, 2006 which is a
divisional of pending U.S. patent application Ser. No. 09/695,748
filed Oct. 24, 2000 which is a Divisional of application Ser. No.
09/028,154 filed Feb. 23, 1998, (Issued U.S. Pat. No. 6,231,516)
which is a Continuation-in-Part of Ser. No. 08/978,038 filed Nov.
25, 1997 (Issued U.S. Pat. No. 5,967,986) which is a
Continuation-in-Part of Ser. No. 08/949,413 filed Oct. 14, 1997
(Issued U.S. Pat. No. 5,807,258).
TECHNICAL FIELD
[0002] This invention relates generally to implantable devices,
and, more particularly, to implantable medical devices having
therapeutic or diagnostic functions within a lumen of an
endoluminal implant such as a stent or other type of endovascular
conduit, and methods related to such implantable medical
devices.
BACKGROUND OF THE INVENTION
[0003] In the 1970s, the technique of percutaneous transluminal
coronary angioplasty (PTCA) was developed for the treatment of
atherosclerosis. Atherosclerosis is the build-up of fatty deposits
or plaque on the inner walls of a patient's arteries; these lesions
decrease the effective size of the artery lumen and limit blood
flow through the artery, prospectively causing a myocardial
infarction or heart attack if the lesions occur in coronary
arteries that supply oxygenated blood to the heart muscles. In the
angioplasty procedure, a guide wire is inserted into the femoral
artery and is passed through the aorta into the diseased coronary
artery. A catheter having a balloon attached to its distal end is
advanced along the guide wire to a point where the sclerotic
lesions limit blood flow through the coronary artery. The balloon
is then inflated, compressing the lesions radially outward against
the wall of the artery and substantially increasing the size of its
internal lumen, to improve blood circulation through the
artery.
[0004] Increasingly, stents are being used in place of or in
addition to PTCA for treatment of atherosclerosis, with the intent
of minimizing the need to repeatedly open an atherosclerotic
artery. Although a number of different designs for stents exist in
the prior art, all are generally configured as elongate cylindrical
structures that are provided in a first state and can assume a
second, different state, with the second state having a
substantially greater diameter than the first state. A stent is
implanted in a patient using an appropriate delivery system for the
type of stent being implaced within the patient's arterial system.
There are two basic types of stents--those that are expanded
radially outward due to the force from an inflated angioplasty type
balloon, such as the Palmaz-Schatz stent, the Gianturco-Roubin
stent and the Strecker stent, and those that are self expanding,
such as the Maass double helix spiral stent, the Nitinol stent
(made of nickel titanium memory alloy), the Gianturco stent and the
Wallstent. Problems with the Maass double helix spiral stent and
the Nitinol stent have limited their use.
[0005] Stents are sometimes used following a PTCA procedure if the
artery is totally occluded or if the lesions have occluded a
previously placed surgical graft. Typically, a stent constrained
within an introducer sheath is advanced to a site within the
patient's artery through a guide catheter. For the balloon expanded
type, after the introducer sheath is retracted, a balloon disposed
inside the stent is inflated to a pressure ranging from about six
to ten atmospheres. The force produced by the inflated balloon
expands the stent radially outward beyond its elastic limit,
stretching the vessel and compressing the lesion to the inner wall
of the vessel. A self expanding stent expands due to spring force
following its implacement in the artery, after a restraining sheath
is retracted from the compressed stent, or in the case of the
Nitinol version, the stent assumes its expanded memory state after
being warmed above the transition temperature of the Nitinol alloy
(e.g., above 30.degree. C.). Following the expansion process, when
the balloon catheter is used, the balloon is removed from inside
the stent and the catheter and other delivery apparatus is
withdrawn. The lumen through the vessel is then substantially
increased, improving blood flow.
[0006] After a stent or other endoluminal device is implanted, a
clinical examination and either an angiography or an ultrasonic
morphological procedure is performed to evaluate the success of the
stent emplacement procedure in opening the diseased artery or
vessel. These tests are typically repeated periodically, e.g., at
six-month intervals, since restenosis of the artery may occur.
[0007] Due to the nature of the tests, the results of the procedure
can only be determined qualitatively, but not quantitatively, with
any degree of accuracy or precision. It would clearly be preferable
to monitor the flow of blood through the stent after its
implacement in a vessel, both immediately following the treatment
for the stenosis and thereafter, either periodically or on a
continuous basis. Measurements of volumetric rate and/or flow
velocity of the blood through the stent would enable a medical
practitioner to much more accurately assess the condition of the
stent and of the artery in which the stent is implanted. Currently,
no prior art mechanism is available that is implantable inside a
blood vessel for monitoring blood flow conditions through a
stent.
[0008] Following stent implantation, it is difficult to monitor the
condition of the affected area. Stents often fail after a period of
time and for a variety of reasons. Several of the causal mechanisms
are amenable to drug treatment. It is highly desirable in at least
some of these cases to localize the drug treatment to the site of
the graft or surgery. For example, when thrombus forms in a given
area, thrombolytic drugs are capable of providing significant
assistance in resolving the thrombosis, but may present problems
such as hemorrhaging, if they also act in other portions of the
patient's body.
SUMMARY OF THE INVENTION
[0009] The present invention provides a capability for including a
therapeutic transducer together with an endoluminal implant such as
a stent or stent graft. Therapeutic transducers may include
ultrasonic, magnetic, iontophoretic, heating or optical devices,
which may permit localized drug delivery or localized drug
activation. Provision is made for delivering energy to the
implanted transducers and for coupling signals to or from the
implanted transducers. The present invention also permits inclusion
of diagnostic transducers together with the endoluminal implant and
allows signals to be transmitted from the diagnostic transducers to
an area outside of the patient's body.
[0010] The present invention can allow steps that may be taken to
restore full fluid flow through, e.g., a stent that is becoming
restricted. In these cases, it is desirable to initiate treatment
before the problem proceeds too far to be corrected without stent
replacement or further PTCA treatment. Clearly, it would be
preferable to be able to monitor the condition of a stent without
resorting to invasive surgical procedures and without prescribing
medication that may not be necessary, so that the useful life of
the stent may be extended, problems associated stent failure
avoided and so that medications are only prescribed when required
by the known condition of the stent and associated vasculature.
[0011] Other advantages that may be realized via embodiments of the
present invention including monitoring of other parameters
measurable within a stent or other type of endoluminal implant
using one or more appropriate sensors or transducers according to
embodiments of the present invention. For example, monitoring
pressure at the distal and proximal ends of the lumen in the
implant and determining the differential pressure can provide an
indication of fluid velocity through the lumen. Temperature can
also be used to monitor fluid flow by applying heat to the fluid
within the lumen and monitoring the rate at which the temperature
of the fluid decreases as the fluid flows through the lumen of the
implant. Integrated circuit (IC) transducers are currently known
and available for sensing the levels of many different types of
biochemical substances, such as glucose, potassium, sodium,
chloride ions and insulin. Any of these IC sensors could be
provided in an endoluminal implant to monitor these parameters.
[0012] Since it is impractical to pass a conductor through the wall
of an artery or vessel for long periods of time, use of a
conventional sensor that produces signals indicative of flow
through a stent, which must be conveyed through a conductor that
extends through the wall of the vessel and outside the patient's
body, is not a practical solution to this problem. Also, any active
flow indicative sensor must be energized with electrical power.
Again, it is not practical to supply power to such a sensor through
any conductor that perforates the vessel wall or that passes
outside the patient's body.
[0013] In addition to stents, the generic term endoluminal implant
encompasses stent grafts, which are also sometimes referred to as
"spring grafts." A stent graft is a combination of a stent and a
synthetic graft is endoluminally implanted at a desired point in a
vessel. Helically coiled wires comprising the stent are attached to
the ends of the synthetic graft and are used to hold the graft in
position. Sometimes, hooks are provided on the stent to ensure that
the graft remains in the desired position within the vessel.
Clearly, it is advantageous to monitor the status of flow and other
parameters through a stent graft, just as noted above in regard to
a stent.
[0014] Endoluminal implants are used in other body passages in
addition to blood vessels. For example, they are sometimes used to
maintain an open lumen through the urethra, or through the cervix.
A stent placed adjacent to an enlarged prostate gland can prevent
the prostate from blocking the flow of urine through the urethra.
Tracheal and esophageal implants are further examples of
endoluminal implants. In these and other uses of endoluminal
implants, provision for monitoring parameters related to the status
of flow and other conditions in the patient's body is desirable.
Information provided by monitoring such parameters, and localized
drug delivery or drug activation, can enable more effective medical
treatment of a patient through use of embodiments of the present
invention.
[0015] Another advantage that may be realized through practice of
embodiments of the present invention is to be able to activate a
therapeutic device on the stent or stent graft that would allow the
physician to activate drugs known to be effective in preventing
further tissue growth within the stent or stent graft in situations
where it is determined that tissue ingrowth is threatening the
viability of a stent or stent graft. Again, the therapeutic device
should be able to be supplied with electrical power from time to
time from a location outside the patient's body.
[0016] Yet another advantage that may be realized through practice
of the present invention is the treatment of tumors or organs that
are downstream of the blood vessel that includes a stent that is
coupled to a transducer. The transducer may be remotely activated
to facilitate localized drug delivery or to provide other
therapeutic benefits.
BRIEF DESCRIPTION OF THE DRAWINGS
[0017] The foregoing aspects and many of the attendant advantages
of this invention will become more readily appreciated as the same
becomes better understood by reference to the following detailed
description, when taken in conjunction with the accompanying
drawings, wherein:
[0018] FIG. 1 is a block diagram according to the invention showing
a first embodiment of an implantable electronic circuit for
coupling electrical signals to or from a selected transducer of a
plurality of transducers.
[0019] FIG. 2 is a block diagram of a second embodiment of an
implantable electronic circuit for coupling electrical signals to
or from a transducer using separate multiplexers for transmit and
receive functions.
[0020] FIG. 3 is a block diagram of a third embodiment of an
implantable electronic circuit for coupling electrical signals to
or from a transducer using separate multiplexers and amplifiers for
transmit and receive functions.
[0021] FIG. 4 is a block diagram of a fourth embodiment of an
implantable electronic circuit for coupling electrical signals to
or from a transducer that employs a local transmitter to excite a
selected transducer, and a modulator/transmitter for transmitting
signals from the transducers.
[0022] FIG. 5 is a block diagram of a fifth embodiment of an
implantable electronic circuit for coupling electrical signals to
or from a transducer, where one transducer is selected for
transmitting and receiving, and a modulator/transmitter is used for
transmitting the signal produced by the receiving transducer.
[0023] FIG. 6 is a block diagram of a sixth embodiment of an
implantable electronic circuit for monitoring the status of a stent
or stent graft, wherein one of a plurality of transducers is
selectively coupled to a modulator/transmitter or a receiver.
[0024] FIG. 7 is a cross-sectional view of a radio frequency (RF)
coupling coil in a stent that is implanted in a blood vessel, and
an external coil that is electromagnetically coupled to the RF
coupling coil.
[0025] FIG. 8 is a cross-sectional view of a RF coupling coil in a
stent implanted in a blood vessel, and includes a block that
represents an implanted coil, which is electromagnetically coupled
to the RF coupling coil.
[0026] FIG. 9 is a side elevational view of a woven mesh RF
coupling coil that comprises a wall of a stent.
[0027] FIG. 10 is a cut away side elevational view of a further
embodiment of an external coil and a side elevational view of a
blood vessel in which a stent is implanted that includes a
saddle-shaped RF coupling coil integrated within the wall of the
stent.
[0028] FIG. 11A is a side elevational view (showing only the
foreground) of a portion of a metal tube-type stent with
nonconductive weld joints, illustrating a RF coupling coil wrapped
around the stent in a pre-expansion configuration.
[0029] FIG. 11B is a side elevational view (showing only the
foreground) of a portion of a zigzag wire stent with non-conductive
joints, illustrating a RF coupling coil wrapped around the stent in
a pre-expansion configuration.
[0030] FIG. 12 is a cut-away view of a portion of a limb showing a
stent implanted at a substantial depth within a blood vessel, and
an external coupling coil that encompasses the stent.
[0031] FIG. 13 is a side elevational schematic view of a dual beam
conformal array transducer on an expandable carrier band for use in
a stent.
[0032] FIG. 14 is an end elevational view of the conformal array
transducer of FIG. 13, within a stent.
[0033] FIG. 15 is a plan view of the conformal array transducer
shown in FIGS. 13 and 14, cut along a cut line to display the dual
conformal arrays in a flat disposition.
[0034] FIG. 16A is a cross-sectional side view of a portion of a
stent in which are disposed transversely oriented transducers for
monitoring flow using correlation measurements.
[0035] FIG. 16B is a transverse cross-sectional view of the stent
and transversely oriented transducers shown in FIG. 16A.
[0036] FIG. 17 is an enlarged partial transverse cross-sectional
view of the layers comprising the conformal array transducer
disposed on a stent within a blood vessel.
[0037] FIG. 18 is an enlarged partial cross-sectional side view of
a tilted-element transducer array disposed within a stent.
[0038] FIG. 19A is an isometric view of an integrated circuit (IC)
transducer mounted on a tubular stent.
[0039] FIG. 19B is an enlarged partial cross-sectional side view of
the implantable IC transducer mounted on the tubular stent.
[0040] FIG. 19C is an isometric view of the implantable IC
transducer mounted on a woven mesh stent.
[0041] FIG. 19D is an enlarged partial cross-sectional side view of
the implantable IC transducer mounted on the woven mesh stent.
[0042] FIG. 20 is a side elevational schematic view showing an IC
strain sensor and sensing filaments disposed on a stent.
[0043] FIG. 21A is side elevational schematic view of a stent
outline showing a deposit and ingrowth IC sensor and sensing
filament.
[0044] FIG. 21B is a cross-sectional view of a lumen of the stent
in FIG. 21A, illustrating fatty tissue ingrowth.
[0045] FIG. 22 is side elevational view of a portion of a branching
artery in which a stent graft that is used for providing
therapeutic functions is implanted.
[0046] FIG. 23 illustrates an ultrasonic transducer configuration
integrated with a stent or stent graft.
[0047] FIG. 24 illustrates an embodiment of a dual frequency
ultrasonic transducer.
[0048] FIG. 25 illustrates one embodiment of a coil integrated into
a stent.
[0049] FIG. 26 illustrates another embodiment of a coil integrated
into a stent.
[0050] FIG. 27 illustrates an embodiment of an iontophoretic system
for local drug delivery.
[0051] FIG. 28 illustrates an embodiment wherein light emitting
transducers are coupled to a stent.
DETAILED DESCRIPTION OF THE INVENTION
[0052] The present invention is employed for providing therapeutic
functions proximate to an endoluminal implant. As used herein and
in the claims that follow, the term endoluminal implant broadly
encompasses stents, stent grafts (sometimes referred to as "spring
grafts") and other types of devices that are inserted into a lumen
or body passage and moved to a desired site to provide a structural
benefit to the lumen. To simplify the disclosure of the present
invention, most of the following discussion is directed to
embodiments comprising a stent.
[0053] In one embodiment, parameters are monitored via implanted
diagnostic transducers, where the monitored parameters are directed
to determining the status of the fluid flow through the endoluminal
implant, and therapeutic transducers may be activated in response
to the data collected from the implanted diagnostic transducers.
For example, the rate or velocity of fluid flow through a body
passage in which the stent has been positioned can be monitored to
determine the extent of tissue growth or fatty deposits in a blood
vessel in which the stent has been implanted to treat
atherosclerosis. By monitoring these parameters, which are
indicative of blood flow through the lumen of the stent and the
blood vessel in which it is implanted, a medical practitioner can
evaluate the need for further treatment or determine whether
restenosis has occurred, and can locally activate drugs to control
restenosis when it is determined to have occurred. This may be
possible without additional surgery and without some of the
complications associated with systemic administration of drugs.
Moreover, other physical and biological parameters can be monitored
using one or more appropriate sensors attached to a stent.
[0054] When implanted therapeutic transducers are to be activated
for an extended period of time or following an extended delay, the
stent will likely need to receive electrical power from an external
source to energize the implantable electronic circuitry used to
activate the implanted therapeutic transducers. Similarly, when the
status of fluid flow through a stent that has been implanted in a
patient's vascular system (or some other parameter that is sensed
proximate the stent) is to be monitored for an extended period or
following an extended delay, the implanted circuitry associated
with the stent will likely need to receive electrical power from an
external source. This power may also be needed to convey data
indicating the status of fluid flow (or other parameter) from the
implanted stent to a monitoring device that is disposed outside the
patient's body. In many cases, it may be desirable to monitor one
or more parameters at multiple stents or at multiple locations on a
single stent, or to provide therapeutic functions at more than one
stent or to multiple locations within or associated with one stent.
Thus, the specific transducer employed to provide a therapeutic
function or transducer or sensor employed to monitor a desired
parameter must be selectable so that the data signal indicating the
parameter can be transmitted outside the patient's body. However,
in some cases, only a single transducer (which may be operable
without any implanted control electronics) may be required to
provide a therapeutic function or to monitor a parameter such as
fluid volumetric flow or velocity, which is indicative of the
internal condition of the stent and of the blood vessel in which it
is implanted.
[0055] FIG. 1 illustrates a first embodiment of an implantable
electronic circuit for providing one or more therapeutic functions
or for monitoring one or more parameters, applicable to the
situation in which n transducers 44-46 are included on one or more
stents implanted in the patient's body. Variations of the
implantable electronic circuit shown in FIG. 1 are discussed below
to accommodate specific conditions. In addition, other embodiments
of implantable electronic circuits are illustrated in FIGS. 2
through 6. These embodiments, like that of FIG. 1, are useful for
providing power to transducers that provide therapeutic functions,
e.g., that activate drugs or that assist in localized drug
delivery, or that monitor fluid flow or velocity through a stent
and also for transmitting data signals from the transducers to
locations outside a patient's body, e.g., to an external remote
monitoring console Some of these implantable electronic circuits
are better suited for certain types of therapy or measurements than
others, and again, variations in the implantable electronic
circuits are discussed below, where appropriate, to explain these
distinctions. Examples of implantable telemetry systems are
discussed in A Telemetry-Instrumentation System For Monitoring
Multiple Subcutaneously Implanted Glucose Sensors by M.C. Shults et
al., IEEE Trans. Biomed. Eng., Vol. 41, No. 10, October 1994, pp.
937-942 and in Integrated Circuit Implantable Telemetry Systems by
J.W. Knutti et al., Eng. in Med. and Bio). Magazine, March 1983,
pp. 47-50.
[0056] Each of the implantable electronic circuits shown in FIGS. 1
through 6 are intended to be implanted within the patient's body
and left in place at least during the period in which a therapeutic
function may be needed or the flow conditions through one or more
stents or other parameters are monitored. Although separate
functional blocks are illustrated for different components of the
implantable electronic circuits in these Figures, any of the
implantable electronic circuits can be implemented in one or more
application specific integrated circuits (ASICs) to minimize size,
which is particularly important when the implantable electronic
circuits are integral with a stent. The implantable electronic
circuits can be either included within the wall of a stent, or may
be simply implanted adjacent to blood vessel(s) in which the
stent(s) is/are disposed. However, if not integral with the stent,
the implantable electronic circuits must be electromagnetically
coupled to the transducers, since it is impractical to extend any
conductor through a wall of the blood vessel in which a stent is
implanted, to couple to circuitry disposed outside the blood
vessel. Therefore, in some embodiments, the implantable electronic
circuits are integral with the stent so that they are implanted,
together with the stent, inside the blood vessel.
[0057] Each of the implantable electronic circuits shown in FIGS. 1
through 6 includes a RF coupling coil 30, which is coupled via
lines 34 and 36 to a RF-to-DC power supply 32. In one embodiment,
the RF coupling coil 30 is part of the expandable structure of the
stent body or may instead be added to a stent, for example, by
threading an insulated wire through the expandable wall of a stent.
In some embodiments, the RF coupling coil 30 comprises a helical
coil or saddle-shaped coil, as explained in greater detail below.
The RF-to-DC power supply 32 rectifies and filters a RF excitation
signal supplied from an external source to the RF coupling coil 30,
providing an appropriate voltage DC power signal for the other
components of the implantable electronic circuits illustrated in
these Figures. In the simplest case, the RF-to-DC power supply 32
would only require rectifiers and filters as appropriate to provide
any needed positive and negative supply voltages, +V.sub.s and
-V.sub.s.
[0058] However, it is also contemplated that the RF-to-DC power
supply 32 may provide for a DC-to-DC conversion capability in the
event that the electromagnetic signal coupled into the RF coupling
coil 30 is too weak to provide the required level of DC voltage for
any component. This conversion capability would increase the lower
voltage produced by the direct coupling of the external RF
excitation signal received by the RF coupling coil 30, to a higher
DC voltage. Details of the RF-to-DC power supply 32 are not shown,
since such devices are conventional. It is also contemplated that
it may be necessary to limit the maximum amplitude of the RF input
signal to the RF-to-DC power supply 32 to protect it or so that
excessive DC supply voltages are not provided to the other
components.
[0059] Alternatively, each component that must be provided with a
limited DC voltage supply may include a voltage limiting component,
such as a zener diode or voltage regulator (neither shown). In
another embodiment, the RF coupling coil 30 and the RF-to-DC power
supply 32 of FIGS. 1 through 6 may be replaced by a hard-wired
connection to supply DC or AC power in applications where the
implant is needed for a relatively short duration, where the
inconvenience of the cables supplying the power is tolerable and
the risk of infection is manageable. An example of a hard-wired
transcutaneous connection for chronic implants is described in
Silicon Ribbon Cables For Chronically Implantable Microelectrode
Arrays by J.F. Hetke et al., IEEE Trans. Biomed. Eng., Vol. 41, No.
4, April 1994, pp. 314-321.
[0060] The RF-to-DC power supply 32 may include a battery or a
capacitor for storing energy so that it need not be energized when
providing a therapeutic function or monitoring the flow status, or
at least, should include sufficient storage capability for at least
one cycle of receiving energy and transmitting data relating to the
parameter being monitored. Neither a battery nor power storage
capacitor are illustrated in the Figures, since they are
conventional also.
[0061] Implantable electronic systems using battery power may only
require the ability to receive data and control signals, and may
include the ability to transmit signals. As a result, they do not
necessarily require access to the skin, which access facilitates
efficient coupling of power signals. A battery-powered system may
result in a very compact implantable system. Alternatively, a
battery-powered system that also is capable of recharging the
battery via power signals coupled through an implanted coil can
permit continuing treatment without requiring that a physician be
present throughout the treatment or requiring the patient to be in
the medical facility.
[0062] An element that is common to each of the implantable
electronic circuits shown in FIGS. 1 through 3 is a RF decode
section 40, which is used for generating control signals that are
responsive to information encoded in the external RF excitation
signal received by the RF coupling coil 30. This information can be
superimposed on the RF excitation signal, e.g., by amplitude or
frequency modulating the signal.
[0063] In regard to the implantable electronic circuits shown in
FIGS. 1 through 3, when used for monitoring fluid velocity or flow,
the RF excitation frequency is the same as the frequency used to
provide energy for therapeutic functions (e.g., localized drug
activation) or to excite a selected ultrasonic transducer to
produce an ultrasonic wave that propagates through a lumen of the
stent being monitored and for conveying data from the transducer
44-46 that receives the ultrasonic waves. This approach generally
simplifies the implantable electronic circuitry but may not provide
optimal performance.
[0064] Therefore, FIGS. 4 and 5 disclose implantable electronic
circuitry in which the RF excitation frequency used to provide
power to the RF-to-DC power supply 32 and to provide control
signals to the RF decode section 40 is decoupled from the frequency
that is used for exciting the transducers 44-46 and modulating any
data that they provide for transmission to a point outside the
patient's body. Although other types of transducers 44-46 may be
employed that are energized with a RF excitation frequency, such as
surface acoustic wave transducers that are used for sensing
chemical substances, many transducers 44-46 only require a DC
voltage to sense a desired parameter such as pressure or
temperature or to provide a static magnetic field, heat or light
for therapeutic purposes.
[0065] Implantable Electronic Circuits
[0066] Referring now to FIG. 1, a line 36 from the RF coupling coil
30 is coupled to a multiplexer (MUX) 38 to convey signals from a
selected one of a plurality of n transducers 44-46 (which are
disposed at different points on a stent) that are coupled to the
MUX 38. To select the transducer 44-46 that will provide a
therapeutic function or a data signal related to a parameter at a
specific location on a stent, the RF decode section 40 provides a
control signal to the MUX 38 through MUX control lines 42. The
control signal causes the MUX 38 to select a specific transducer
44-46 that is to be excited by the RF signal received by the RF
coupling coil 30 and further, causes the MUX 38 to select the
transducer 44-46 that will provide the data signal for transmission
outside the patient's body (or at least outside the blood vessel in
which the stent is disposed) via the RF coupling coil 30.
[0067] In addition to ultrasonic transducers 44-46, the implantable
electronic circuit shown in FIG. 1 can also be used in connection
with pressure transducers 44-46, For ultrasonic transducers 44-46,
the circuit is perhaps more applicable to the Doppler type for use
in monitoring fluid velocity through a stent. If a single-vessel
pulse Doppler transducer 44-46 is used, the same transducer 44-46
can be used for both transmission and reception of the ultrasonic
wave, thereby eliminating the need for the MUX 38. In the event
that the transducers 44-46 shown in FIG. 1 are used for transit
time flow measurements, it will normally be necessary to use the
MUX 38 to switch between the transducer 44-46 used for transmitting
the ultrasonic wave and that used to receive the ultrasonic
wave.
[0068] For a single-vessel transit time measurement, a pair of
opposed transducers 44-46 that are disposed on opposite sides of
the stent are typically used. In order to acquire bi-directional
fluid flow data, the direction of the ultrasound wave propagation
must be known, i.e., the direction in which the ultrasound wave
propagates relative to the direction of fluid flow through the
vessel. In this case, the MUX 38 is required. However, for
single-vessel applications in which the fluid flow is in a single
known direction, the transducers 44-46 that are disposed on
opposite sides of the stent can be electrically coupled in parallel
or in series, eliminating any requirement for the MUX 38. The
RF-to-DC power supply 32 and the RF decode section 40 could also
then be eliminated, since the retarded and advanced transit time
signals are superimposed on the same RF waveform transmitted by the
RF coupling coil 30 to a location outside the patient's body (or
outside the blood vessel in which the stent is disposed, if an
internal coil is implanted adjacent the blood vessel near where the
stent is implanted). Although this modification to the implantable
electronic circuit shown in FIG. 1 would not permit the direction
of fluid flow through a stent to be determined, the retarded and
advanced transit time signals interfere over time, and their
interference can be used to estimate the magnitude of fluid flow
through the stent.
[0069] In some applications, a single transducer 44-46 or group of
transducers 44-46 may be employed, in which case the implantable
electronic circuit of FIG. 1 may be simplified by coupling the
transducer(s) 44-46 directly to the RF coupling coil 30 and
eliminating the MUX 38. In this embodiment, the RF decode section
40 and the RF-to-DC power supply 32 are optional; when the
transducer, for example, requires DC excitation or other excitation
different than that which may be provided directly via the RF
coupling coil 30, the RF-to-DC power supply 32 may be desirable.
Similarly, some sensors may have more than one function and then
the RF decode section 40 may also be desirable. Similarly, the
implantable electronic circuits of FIGS. 2 through 6 may be
modified to provide the desired or required functionality.
[0070] In FIG. 2, an implantable electronic circuit is shown that
uses a transmit multiplexer (TX MUX) 50 and a receive multiplexer
(RX MUX) 54. In addition, a transmit (TX) switch 48 and a receive
(RX) switch 52 couple line 36 to the TX MUX 50 and the RX MUX 54,
respectively. The RF decode section 40 responds to instructions on
the signal received from outside the patient's body by producing a
corresponding MUX control signal that is conveyed to the TX MUX 50
and the RX MUX 54 over MUX control lines 56 to select the desired
transducers 44-46.
[0071] When ultrasonic signals are being transmitted by one of the
selected transducers 44-46, the TX switch 48 couples the RF
excitation signal received by the RF coupling coil 30 to the
transducer 44-46 that is transmitting the ultrasonic signal, which
is selected by the TX MUX 50. The TX switch 48 is set up to pass
excitation signals to the selected transducer 44-46 only if the
signals are above a predetermined voltage level, for example, 0.7
volts. Signals below that predetermined voltage level are blocked
by the TX switch 48. Similarly, the RX switch 52 couples the
transducer 44-46 selected by the RX MUX 54 to the RF coupling coil
30 and passes only signals that are below the predetermined voltage
level, blocking signals above that level. Accordingly, the RF
signal used to excite a first transducer 44-46 selected by the TX
MUX 50 passes through the TX switch 48 and the lower amplitude
signal produced by a second transducer 44-46 selected by the RX MUX
54 in response to the ultrasonic signal transmitted through the
stent is conveyed through the RX MUX 54 and the RX switch 52 and
transmitted outside the patient's body through the RF coupling coil
30.
[0072] The implantable electronic circuit shown in FIG. 3 is
similar to that of FIG. 2, but it includes a transmit amplifier (TX
AMP) 58 interposed between the TX switch 48 and the TX MUX 50, and
a receive amplifier (RX AMP) 60 interposed between the RX MUX 54
and the RX switch 52. The TX AMP 58 amplifies the excitation signal
applied to the transducer 44-46 selected by the TX MUX 50 for
producing the ultrasonic wave that is propagated through the
interior lumen of a stent. Similarly, the RX AMP 60 amplifies the
signal produced by the transducer 44-46 selected by the RX MUX 54
before providing the signal to the RX switch 52 for transmission
outside the patient's body (or at least, outside the blood vessel
in which the stent is implanted). Again, the implantable electronic
circuit shown in FIG. 3 is most applicable to transit time flow
measurements and employs the same frequency for both the RF
excitation signal that supplies power to the RF-to-DC power supply
32 and the signal applied to a selected one of the transducers
44-46 to generate the ultrasonic wave propagating through the
stent.
[0073] In contrast to the implantable electronic circuits shown in
FIGS. 1 through 3, the implantable electronic circuit shown in
FIGS. 4 through 6 enables the RF excitation frequency applied to
the RF-to-DC power supply 32 to be decoupled from the frequency of
the signal applied to excite any selected one of the transducers
44-46. Similarly, the signal produced by the transducer 44-46
receiving ultrasonic waves propagating through the stent is at a
different frequency than the RF excitation frequency applied to the
RF-to-DC power supply 32. In FIG. 4, a transmitter (XMTR) 62 and a
receive modulator/transmitter (RX MOD/XMTR) 64 are coupled to and
controlled by a RF decode/control section 66. The RF decode/control
section 66 determines when the excitation frequency is generated
for application to a selected transmit transducer 44-46 and when
the signal produced by the transducer selected to receive the
ultrasonic wave is used for modulating the RF signal applied to the
RF coupling coil 30. An advantage of this approach is that the RF
power delivered to the RF coupling coil 30 is at an optimal
frequency for penetration through the patient's body, thereby
improving the efficacy with which the RF energy couples to a
specific depth and location within the body. Another reason for
using this approach is to enable selection of a particular
frequency as necessary to comply with radio frequency allocation
bands for medical equipment. Similarly, the frequency applied to
any selected transducers 44-46 to stimulate their production of
ultrasonic waves can be optimal for that purpose. Assuming that the
two frequency bands, i.e., the RF excitation frequency band for the
signal applied to the RF-to-DC power supply 32, and the frequency
band of the signals applied to excite the transducers 44-46, are
sufficiently separated, the RF power delivery can occur
simultaneously with the excitation of a selected transducer 44-46
and the reception of the ultrasonic waves by another selected
transducer 44-46. Accordingly, more RF power can be coupled into
the system from the external source than in the implantable
electronic circuits shown in FIGS. 1 through 3. In some
embodiments, including those where a battery is used, the RF
decode/control section 66 may also include a RF oscillator for
providing the RF signals to the transducers 44-46 or for coupling
signals from the transducers 44-46 to external electronic
apparatus.
[0074] The control signals that are supplied to the RF
decode/control section 66 via the RF coupling coil 30 can be
conveyed using nearly any kind of modulation scheme, e.g., by
modulating the RF excitation that powers the device, or by sending
a control signal on a separate and distinct RF frequency. Also, the
signals that are received from the transducer 44-46 in response to
the ultrasonic wave that is propagated through the stent can be
transmitted through the RF coupling coil 30 at a different
frequency than the incoming excitation frequency, thereby reducing
the likelihood of interference between the power supply and data
signal transmission functions.
[0075] The implantable electronic circuit shown in FIG. 4 is
applicable to transit time flow measurements in which pairs of
transducers 44-46 are selected for transmitting and receiving the
ultrasonic wave that propagates through the one or more stents on
which the transducers 44-46 are installed. The RF decode/control
section 66 can be employed to control the TX MUX 50 and the RX MUX
68 to interchange the transducers 44-46 used for transmission and
reception of the ultrasonic wave on successive pulses. Using this
technique, the direction of the ultrasonic wave propagation through
the stent is changed on alternating pulses of ultrasonic waves,
enabling transit time difference information to be gathered without
requiring further multiplexer programming information to be
transmitted between successive ultrasonic wave pulses. This
approach greatly improves the data gathering efficiency of the
implantable electronic circuit shown in FIG. 4 compared to the
previously described implantable electronic circuits of FIGS. 1
through 3.
[0076] To further improve the implantable electronic circuit shown
in FIG. 4 for use in sensing fluid velocity through a stent using a
Doppler technique, the modification shown in FIG. 5 is made. In
FIG. 5, a TX/RX switch 72 is added so that the implantable
electronic circuit transmits and receives through the same
transducer 44-46. As a result, the separate transmit 50 and receive
54 multiplexers of FIG. 4 are not required. Instead, the MUX 38 is
used to select the specific transducer 44-46 for receiving the RF
excitation signal produced by the XMTR 62 so that the transducer
44-46 produces an ultrasonic wave and then receives the echo from
fluid flowing through the stent to produce a received data signal
that is output through the RX MOD/XMTR 64. The TX/RX switch 72
prevents the signal applied by the TX AMP 58 from overdriving the
input to the RX AMP 60, effectively isolating the RX AMP 60 during
the time that the RF signal is applied to the transducer 44-46 to
excite it so that it produces the ultrasonic wave. However, the
echo signal received by the transducer 44-46 is allowed to reach
the RX AMP 60 when the TX/RX switch 72 changes state (from transmit
to receive). Generally, the implantable electronic circuit shown in
FIG. 5 has the same benefits as described above in connection with
the implantable electronic circuit shown in FIG. 4. The RF
decode/control section 66 responds to the information received from
outside the patient's body that determines which one of the
transducers 44-46 is selected at any given time by producing an
appropriate MUX control signal that is supplied to the MUX 38 over
the MUX control lines 56.
[0077] It is also contemplated that the RF decode/control section
66 may cause the MUX 38 to select a different transducer 44-46 for
producing/receiving the ultrasonic waves after a predefined number
of transmit/receive cycles have elapsed. For example, a different
transducer 44-46 may be selected after eight cycles have been
implemented to transmit an ultrasonic wave into the stent and to
receive back the echoes from the fluid flowing through the stent.
By collecting data related to the status of flow through a stent in
this manner, it becomes unnecessary to send programming information
to the RF decode/control section 66 after each cycle of a
transmission of the ultrasonic wave into the fluid in the stent and
reception of the echo. Also, by carrying out a predefined number of
transmit/receive cycles for the given transducer 44-46 that has
been selected by the MUX 38 and averaging the results, a more
accurate estimate of fluid velocity through the stent can be
obtained than by using only a single transmission and reception of
an ultrasonic wave. Since the signal required to instruct the RF
decode/control section 66 to change to the next transducer 44-46 is
only required after the predefined number of cycles has been
completed, the data gathering efficiency of the implantable
electronic circuit is improved.
[0078] As noted above, the transducers 44-46 shown in FIGS. 1
through 5 need not be ultrasonic transducers; FIG. 6 illustrates an
electronic circuit that is particularly applicable for use with
transducers 44-46 comprising pressure sensors. Silicon pressure
sensors designed to be installed on the radial artery are available
from the Advanced Technologies Division of SRI of Palo Alto, Calif.
Such pressure sensors could be disposed within the wall of a stent
to sense the pressure of fluid flowing through the stent at one or
more points. The MUX 38 is used for selecting a specific pressure
transducer to provide a data signal that is transmitted to the
outside environment via the RF coupling coil 30. In the implantable
electronic circuit shown in FIG. 6, a modulator/transmitter
(MOD/XMTR) 70 receives the signal from the transducer 44-46
selected by the MUX 38 in response to the MUX selection signal
provided over the MUX control lines 56 from the RF decode/control
section 66 and, using the signal, modulates a RF signal that is
supplied to the RF coupling coil 30. The RF signal transmitted by
the RF coupling coil 30 thus conveys the data signal indicating
pressure sensed by the selected transducer 44-46. In many cases, it
will be preferable to monitor the pressure at the upstream and
downstream ends of a stent in order to enable the differential
pressure between these ends to be determined. This differential
pressure is indicative of the extent to which any blockage in the
lumen of the stent is impeding fluid flowing through the lumen. In
most cases, parameters such as fluid flow or velocity are better
indicators of the status of flow through the stent.
[0079] RF Coupling Coil and External Coil Embodiments
[0080] FIGS. 7 through 12 illustrate details of several different
embodiments for the RF coupling coil 30 that is part of the stent
implanted within a patient's body. The RF coupling coil 30 is for
receiving RF energy to provide power for the implantable electronic
circuits of FIGS. 1 through 6 and for transmitting data relating to
the condition of flow and/or other parameter(s) sensed by
transducers coupled to one or more stents that have been installed
within the patient's vascular system. Optimization of RF coupling
between the RF coupling coil 30 on the stent and an external coil
is partially dependent upon the propagation characteristics of the
human body. Since body tissue is largely water, the relative
dielectric constant of mammalian soft tissues is approximately
equal to that of water, i.e., about 80. Also, the permeability of
body tissue is approximately equal to one, i.e., about that of free
space. The velocity of propagation of a RF signal through the body
is proportional to the inverse square root of the dielectric
constant and is therefore about 11% of the velocity of the signal
in free space. This lower velocity reduces the wavelength of the RF
signal by an equivalent factor. Accordingly, the wavelength of the
RF signal transferred between the implanted RF coupling coil on a
stent and the external coil would be a design consideration if the
separation distance between the two is approximately equal to or
greater than one-quarter wavelength. However, at the frequencies
that are of greatest interest in the present invention, one-quarter
wavelength of the RF coupling signal should be substantially
greater than the separation distance between the RF coupling coil
30 on the stent and the external coil.
[0081] One method for optimizing coupling between an implanted coil
and a coil that is external to the body is described in
High-Efficiency Coupling-Insensitive Transcutaneous Power And Data
Transmission Via An Inductive Link by CM. Zierhofer and E. S.
Hochmair, IEEE Trans. Biomed. Eng., Vol. 37, No. 7, July 1990, pp.
716-722. This approach allows the frequency of the signal linking
the implanted and external coils to vary in response to the degree
of coupling between the two coils. Other methods are suitable for
coupling signals between the two coils as well.
[0082] When the implantable electronic circuit includes the RF
coupling coil 30 and a transducer 44-46, but does not include
active electronic circuitry, the external system (e.g., external
power supply and patient monitoring console 100, FIG. 8, below)
senses a parameter related to the electrical input impedance of the
external coil. When the external and internal coils are aligned,
the inductance and the resistance of the external coil are
maximized. The frequency of the signal that is used for adjusting
the alignment may be different than the frequency that is used to
provide electrical signals to the transducer.
[0083] The implantable electronic circuit may include an additional
component to facilitate sensing of alignment between the two coils.
For example, a metal disc in the implant may be detected and
localized by inducing an eddy current in the disc. The external
power supply and patient monitoring console may then detect the
magnetic field generated by the eddy current in the disc, much as a
metal detector operates. Using different frequencies for the
location and therapeutic functions may avoid energy losses caused
by the eddy currents.
[0084] When the implantable electronic circuitry does include
active electronic circuitry, a circuit may be included with the
therapeutic transducer and RF coupling coil that measures the
amplitude of the signal from the external power supply and patient
monitoring console that is induced in the RF coupling coil. A
signal is transmitted from the implantable electronic circuitry to
the external power supply and patient monitoring console, where a
display provides an indication of the coupling. The operator may
adjust the position of the external coil to optimize coupling
between the two coils.
[0085] The penetration of RF fields in the human body has been
studied extensively in conjunction with magnetic resonance imaging
(MRI) systems. RF attenuation increases with frequency, but
frequencies as high as 63 MHz are routinely used for whole-body
imaging, although some attenuation is observed at the center of the
torso at this upper frequency limit. In addition, MRI safety
studies have also provided a basis for determining safe operating
limits for the RF excitation that define the amplitude of
excitation safely applied without harm to the patient.
[0086] It is contemplated that for stent implants placed deep
within the torso of a patient, RF excitations and frequencies used
for communicating data related to the fluid flow through a stent
and/or other parameters sensed proximate the stent can be up to
about 40 MHz, although higher frequencies up to as much as 100 MHz
may be feasible. At 40 MHz, the wavelength of the RF excitation
signal in tissue is about 82 cm, which is just that point where
wavelength considerations become an important consideration. For
shallow implants, RF excitation at a much higher frequency may be
feasible. For example, to provide energy to stents that are
disposed within a blood vessel only a few millimeters below the
epidermis and to receive data from transducers associated with such
stents, excitation frequencies in the range of a few hundred MHz
may be useful. The dielectric properties of tissue have been
studied to at least 10 GHz by R. Pethig, Dielectric and Electronic
Properties of Biological Materials, Wiley Press, Chichester, 1979
(Chapter 7). Based on this study, no penetration problems are
anticipated in the frequency range of interest. The relative
dielectric constant of tissue decreases to about 60 at a frequency
of 100 MHz and is about 50 at 1 GHz, but this parameter has little
effect on power/data signal coupling.
[0087] An external coil 90 and a RF coupling coil 30A shown in FIG.
7 represent one embodiment of each of these components that can be
used for coupling electrical energy and conveying data signals
across a skin interface 102 for applications in which the RF
coupling coil 30A is implanted relatively close to the surface 102
of the skin. For example, the RF coupling coil 30A and the external
coil 90 may provide, via magnetic flux lines 112, the coupling
required for a system used to monitor a stent implanted in an
artery near the skin surface 102. A winding 92 is wrapped around a
core 94 forming the external coil 90 and each end of the winding 92
is coupled to a power source through a cable 98.
[0088] Although the external coil 90 and the RF coupling coil 30A
need not be identical in size, it is generally true that coupling
will be optimal if the two devices are of approximately the same
dimensions and if the longitudinal axis of the external coil 90 is
generally adjacent and parallel to that of the RF coupling coil
30A. By observing the strength of the signal transmitted from the
RF coupling coil 30A, it should be possible to position the
external coil 90 in proper alignment with the RF coupling coil 30A
so that the efficiency of the magnetic coupling between the two is
optimized.
[0089] To function as the core 94 for the external coil 90, the
material used should have a relatively high magnetic permeability,
at least greater than one. Although ferrite is commonly used for
core materials, sintered powdered iron and other alloys can also be
used. Since the magnetic characteristics of such materials are
generally conventional, further details of the external coil 90 and
the core 94 are not provided.
[0090] A housing 96 on the external coil 90 provides RF shielding
against electromagnetic interference (EMI). In one embodiment, the
housing 96 for the external coil 90 is conductive, grounded and
surrounds the external coil 90 except where the surfaces of the
generally "C-shaped" core 94 are opposite the RF coupling coil 30A.
The RF shield comprising the housing 96 is attached to an internal
braided shield 99 of the cable 98. Inside the power supply and
patient monitoring console (not shown in FIG. 7) to which the cable
98 is coupled, the shield 99 is connected to ground. The RF shield
on the external coil 90, along with shields provided around
transducers 44-46 on the stent 106, minimizes external EMI
radiation due to the use of the present invention within a
patient's body.
[0091] For the embodiment shown in FIG. 7, the external coil 90 is
magnetically coupled to a spiral winding 108 in the stent 106 that
is implanted in a blood vessel 107. The spiral winding 108
comprises the RF coupling coil 30A for the stent 106 and the
opposite ends of the spiral winding 108 are coupled to an
electronic circuit 110, which may comprise any of the implantable
electronic circuits, described above in connection with FIGS. 1
through 6. Not shown in FIG. 7 are the one or more transducers
44-46 that are included within the stent 106 to monitor one or more
parameters.
[0092] The RF coupling coil 30A used in the stent 106 may be either
an integral part of the stent 106, or it may instead comprise a
separate RF coupling coil 30A that is wound around or through the
structure comprising the wall of the stent 106. To function within
the body of a patient, the stent 106 must be able to bend and flex
with movement of the body, yet must have sufficient surface area
and hoop strength to compress the atheriosclerotic material that is
inside the blood vessel wall radially outward and to support the
vessel wall, maintaining the lumen cross section. Several
manufacturers offer stent designs, each fabricated from wire, bent
back and forth in a periodically repeating "S" shape or zigzag
configuration, forming a generally cylindrical tube. Such stents
are considered ideal for use in practicing the present invention,
since the wire comprising the wall of the stent 106 can be used for
the RF coupling coil 30A. Examples of such stents are the
ANGIOSTENT stent made by AngioDynamics, the stent sold by Cordis
Corporation, the CARDIOCOIL stent produced by Instent and the
WIKTOR stent from Medtronic Corporation.
[0093] FIG. 8 illustrates another embodiment in which an implanted
coil 90A disposed outside the blood vessel 107 adjacent to an
implanted stent 106A is electromagnetically coupled through
magnetic flux lines 112 to the stent 106A using a plurality of
electrically isolated and separate helical windings 109A, 109B,
109C, 109D and 109E forming a woven mesh comprising the wall of the
stent 106A. Not shown are the implantable electronic circuitry and
the transducers 4446 that are coupled to the windings comprising a
RF coupling coil 30B, however, it will be understood that any of
the implantable electronic circuits shown in FIGS. 1 through 6,
discussed above, can be used for this purpose. However, by using
electrically isolated and separate windings 109A-E for the RF
coupling coil 30B, it is possible to avoid multiplexing the signals
from each different transducer 44-46 used in the stent 106A, since
each transducer 44-46 (or sets of transducers 44-46) can transmit
data over its own winding and separately receive an excitation
signal from the implanted coil 90A. This figure shows the implanted
coil 90A coupled to an external power source and monitor through a
cable 98A. The cable 98A can either penetrate the dermal layer of
the patient's body, passing to the outside environment, or
alternatively, may itself be electromagnetically coupled to an
external coil, such as the external coil 90 shown in FIG. 7. When
the cable 98A from the implanted coil 90A penetrates the dermal
layer 102, it is likely that the parameters being sensed by the
stent 106A will only need to be monitored for a relatively short
time, so that the implanted coil 90A can be removed from the
patient's body after the need to monitor the parameters is
satisfied. An advantage of the embodiment shown in FIG. 8 is that
when the stent 106A is implanted deep within the patient's body, it
can be readily energized and the data that it provides can be more
efficiently received outside the body by using the implanted coil
90A as an interface, either directly coupled through the skin 102
or magnetically coupled through an external coil.
[0094] Stents comprising a woven mesh of fine helical wires are
available from certain stent manufacturers. The woven mesh provides
the required hoop strength needed to support the wall of a blood
vessel after the stent is implanted and expanded or allowed to
expand. To maintain the required flexibility for the stent, the
wires comprising the woven mesh of such stents are not joined at
the intersection points. An example is the WALLSTENT stent, which
is sold by Medivent-Schneider. This configuration is also well
suited for practicing the present invention. To be used as the RF
coupling coil 30B, the wires forming the body or wall of the stent
106A must be electrically insulated from the surrounding tissue of
the blood vessel 107 and must be insulated from each other where
they cross except at any node wherein the helical turns are linked
to form one or more sets of coupled turns. The wire used for this
configuration can be either round or flat.
[0095] An embodiment of a RF coupling coil 30C comprising a stent
106B is shown in FIG. 9 to illustrate the configuration discussed
above. The RF coupling coil 30C comprises a woven mesh 132
fabricated from insulated wire so that overlapping segments of the
woven mesh 132 do not electrically connect in the center of the
stent 106B. At each end of the RF coupling coil 30C, the wires
comprising the woven mesh 132 are electrically coupled together at
nodes 134, producing the RF coupling coil 30C. The nodes 134 are
insulated from contact with body fluids or other conductors.
[0096] The couplings at the nodes 134 are preferably not made
randomly or in a haphazard fashion between the various wires
comprising the woven mesh 132. A first wire comprising a helical
coil having, e.g., a first configuration (which may be called a
"right hand spiral" or RHS) has a first end coupled to a first end
of a second wire comprising a helical coil having a second
configuration ("left hand spiral" or LHS; i.e., a mirror image of
the right hand spiral). The voltage induced in the two wires is
equal, but opposite in sign, and the two wires are thus coupled in
series and provide twice the voltage between their second ends than
that produced between the first and second ends of either wire
alone. Accordingly, the second ends of the first two wires cannot
be coupled together at the other end of the woven mesh 132 if these
two wires are to contribute to the total electrical energy derived
from the woven mesh 132. Rather, the wires must be "daisy chained"
in series (i.e., RHS-LHS-RHS-LHS etc.) to provide one embodiment of
the RF coupling coil 30C. Alternatively, a first group of wires all
having the right hand spiral may all be coupled in parallel (i.e.,
have the ends at a first end of the woven mesh 132 coupled
together, and the ends at a second end of the woven mesh 132
coupled together), with wires having the left hand spiral being
similarly treated but in a second group. The groups then may be
combined in series or in parallel, or subsets of the wires may be
grouped and combined.
[0097] When each wire comprising the woven mesh 132 passes around
the central axis of the stent 106B through m degrees, and if there
are a total of n such wires, then the equivalent number of turns in
the RF coupling coil 30C is equal to n.times.m/0.360. Leads 136 and
138 convey signals to and from the nodes 134, coupling the woven
mesh 132 to the implantable electronic circuit 110, which may
comprise any of the implantable electronic circuits of FIGS. 1
through 6.
[0098] The woven mesh structure of the implantable RF coupling coil
30C is often used for stents. However, it should be noted that
currently available woven mesh stents are not woven from insulated
wire, nor are the nodes of the mesh at each end electrically
connected in commercially available stents. In the WALLSTENT stent
by Medivent-Schneider, the ends are instead free floating. It is
also contemplated that an insulated electrical conductor could be
woven into the structure of a commercially available mesh stent.
Alternatively, the RF coupling coil 30C could be fabricated from a
woven mesh or from a plurality of spiral turns of a conductor and
then the mechanical characteristics required of the stent could be
achieved by providing an interwoven wire within the RF coupling
coil 30C. It is also noted that different implantable electronic
circuits can be coupled to separate portions of the woven mesh 132
comprising the RF coupling coil 30C so that the different portions
of the RF coupling coil 30C and the implantable electronic circuits
are electrically isolated from each other, or as a further
alternative, the sections can be coupled in series.
[0099] In FIG. 10, a RF coupling coil 30D in a stem 106C is
illustrated that comprises a plurality of generally saddle-shaped
coils 114 disposed within (or comprising) the wall of the stent
106C. Again, the RF coupling coil 30D is coupled to the implantable
electronic circuit 110. Although only a single layer of
saddle-shaped coils 114 is illustrated, it is contemplated that a
plurality of such interconnected layers could be provided for the
stent 106C.
[0100] For use in electromagnetically coupling with the RF coupling
coil 30D to energize the implantable electronic circuit 110 and to
provide signals to and receive data from the transducers 44-46 (not
separately shown) on the stent 106C, an external coil 90B is
provided that includes a plurality of coils 92B wrapped around a
central portion of a generally E-shaped core 94B. Lines of
electromagnetic flux 112 are thus produced between the central leg
and each of the end legs of the core 94B. It will therefore be
apparent that this embodiment of the RF coupling coil 30D and of
the external coil 90B achieves optimum coupling when the distance
separating the two is minimal. Therefore, the RF coupling coil 30D
and the external coil 90B are best used in applications where the
stent 106C is disposed relatively close to the dermal layer 102 so
that tissue 104 separating the stent 106C from the external coil
90B is only a few centimeters thick. Maximal coupling is achieved
when the central axis of the external coil 90B is aligned with the
central axis of the coil mounted on the stent 106C.
[0101] FIG. 11A illustrates an embodiment of a RF coupling coil 121
that is helically coiled around the circumference of a stent
fabricated by slotting a metal tube 116. The insulated conductor
comprising the RF coupling coil 121 is kinked (or fan-folded) when
wound around the metal tube 116 to accommodate expansion of the
metal tube 116 once implanted in a blood vessel. The insulation on
the RF coupling coil 121 prevents the turns from electrically
shorting by contact with the metal tube 116 or with surrounding
tissue. Although not shown, the RF coupling coil 121 will likely be
adhesively attached to the metal tube 116 at several spaced-apart
locations: The ends of the RF coupling coil 121 are coupled to one
or more transducers or sensors 44-46 (not shown) through an
implantable electronic circuit (also not shown) comprising any of
the implantable electronic circuits shown in FIGS. 1 through 6.
[0102] The metal tube 116 includes a plurality of generally
longitudinally extending slots 117 at spaced-apart locations around
the circumference of the stent. These slots 117 provide the
expansibility and flexibility required of the stent. This design is
similar to the Palmaz-Schatz stent made by Johnson & Johnson
Corporation. To avoid providing a shorted turn with the body of the
metal tube 116, the generally conventional design of the stent is
modified to include a break 118 extending along the entire length
of the metal tube 116. The edges of the metal tube 116 are coupled
at several joints 119 along the break 118 using a non-conductive
material.
[0103] Metal-to-ceramic (or metal-to-glass) welded joints 119 are
commonly employed in medical implants and other electrical devices.
To minimize thermal stress in the joint 119, the metal and the
glass or ceramic must have similar thermal expansion coefficients.
For example, KOVAR alloy, a nickel-iron alloy (29% Ni, 17% Co, 0.3%
Mn and the balance Fe) is one material that can be used to form
glass to metal seals that can be thermally cycled without damage.
This material can be used to form portions of the metal tube 116
disposed along the break 118. Glass or ceramic bonds comprising the
joints 119 then will not experience much thermal stress when the
temperature of the stent changes. This material is commonly used in
lids that are bonded onto ceramic chip carriers in the integrated
circuit industry and thus is readily available.
[0104] An alternative design for a stent formed from a non-woven
wire 145 about which the RF coupling coil 121 is coiled is
illustrated in FIG. 11B. The RF coupling coil 121 is again formed
of an insulated conductor that is helically coiled about the
circumference of the stent. The body of the stent comprises a
plurality of zigzag shapes formed of wire 145 that are joined by
non-conductive (i.e., glass or ceramic) joints 146 at spaced-apart
points that prevent the wires 145 from forming any shorted turns.
This stent configuration is similar to that of the ACS RX
MULTI-LINK stent made by Medtronic and the GFX (AVE) stent produced
by Arterial Vascular Engineering.
[0105] In those cases where stents are implanted relatively deeply
inside the patient's body, at some distance from the surface of the
patient's skin, an alternative external coil 154 can be employed,
generally as shown in FIG. 12. In this example, a stent 144
comprising the RF coupling coil 30C (FIG. 9) is implanted within an
artery 152, which is disposed within a thigh 150 of the patient.
Alternatively, the stent 144 may be implanted, for example, in the
descending aorta, the iliac arteries or to provide therapy to a
tumor that is deeply within the abdomen. To couple with the RF
coupling coil 30C, the external coil 154 includes a plurality of
turns 156 sufficient in diameter to encompass the thigh 150. A RF
shield 160 encloses the outer extent of the external coil 154, so
the external coil 154 is insensitive to capacitively coupled noise.
A lead 158 couples the external coil 154 to a power supply and
monitoring console 101. The external coil 154 can be made
sufficiently large to encompass the portion of the body in which
the implanted stent 144 is disposed such as the torso, a limb of
the patient, or the neck of the patient. Coupling is maximized
between the external coil 154 and the RF coupling coil 30C (or
other RF coupling coil) used on the stent 144 when the central axes
of both the RF coupling coil 30C and the external coil 154 are
coaxially aligned and when the implanted stent 144 is generally
near the center of the external coil 154. Coupling between the RF
coupling coil 30C and the external coil 154 decreases with
increasing separation and begins to degrade significantly when the
implanted stent 144 is more than one external coil 154 radius away
from the center point of the external coil 154. In addition,
coupling is minimized when the central axis of the external coil
154 is perpendicular to the axis of the RF coupling coil 30C.
[0106] Description of the Diagnostic Applications of
Transducers
[0107] An ultrasonic transducer for monitoring flow or fluid
velocity through a stent should be relatively compact and included
in or mounted on the wall of a stent. Typical prior art ultrasonic
transducers include a planar slab of a piezoelectric material
having conductive electrodes disposed on opposite sides thereof.
Since such elements are planar, they do not conform to the circular
cross-sectional shape of a stent. Moreover, prior art transducers
are not compatible for use with a stent that is implanted within a
patient's body and which is intended to be left in place for an
extended period of time. Also, it is apparent that conventional
ultrasonic transducer elements will not readily yield to being
deformed into a compact state for implacement within a blood
vessel, followed by expansion of a stent body to apply radially
outwardly directed force to compress the deposits within a blood
vessel.
[0108] FIGS. 13 through 15 show an embodiment of an extremely low
profile ultrasonic transducer comprising conformal transducer
arrays 174A and 174B, which are disposed on opposite sides of a
stent 168. While the conformal transducer arrays 174A and 174B are
described in conjunction with their application as diagnostic
transducers, the conformal transducer arrays 174A and 174B are also
useful as therapeutic transducers. Since it is contemplated that
this type of ultrasonic transducer assembly might be used on
several different designs of stents, details of the stent 168 are
not illustrated. Instead, only a portion of its outline 170 is
shown. Ideally, each conformal transducer array 174A and 174B
comprises a piezoelectric plastic used as a transduction material
and having sufficient flexibility to allow the transducer elements
to conform to the circular cross section of the wall of the stent
168 when the stent 168 is inserted through the patient's vascular
system and to flex as the stent 168 is expanded within a blood
vessel.
[0109] When used for transit time measurements, as shown in FIGS.
13 and 14, the conformal transducer arrays 174A and 17433 are
disposed generally on opposite sides of the stent 168 and encompass
much of the inner circumference of the stent 168. However, when a
pulsed Doppler measurement is made using the conformal array
transducer 174A, only a single such conformal array transducer 174A
is required, since the conformal array transducer 174A first
produces an ultrasonic wave that is transmitted into the lumen of
the stent 168 and then receives an echo reflected back from the
fluid flowing through the stem 168. If used for continuous wave
(CW) Doppler measurements, the pair of conformal transducer arrays
174A and 174B disposed on opposite sides of the stent 168 are again
needed, one conformal transducer array 174A or 174B serving as a
transmitter and the other as a receiver. In each case, it is
presumed that the fluid has a non-zero flow velocity component
directed along an ultrasonic beam axis of the ultrasonic wave
produced by the conformal transducer array 174A or 174B serving as
a transmitter.
[0110] The conformal transducer arrays 174A and 174B shown in FIGS.
13 through 15 produce ultrasonic beams 178 that are tilted relative
to the transverse direction across the stent 168 in substantially
equal but opposite angles with respect to the longitudinal axis of
the stent 168. Since dual beam transit time measurements are
implemented by the conformal transducer arrays 174A and 174B, the
results are self-compensating for tilt angle errors. This form of
self-compensation is only required where the alignment of the
conformal transducer arrays 174A and 174B relative to the
longitudinal axis of the stent 168 may be imperfect. For transit
time measurements made on stents 168 wherein the alignment of the
conformal transducer arrays 174A and 174B relative to the
longitudinal axis of the stent 168 remains accurately known, an
opposed pair of conformal transducer arrays 174A and 174B disposed
on opposite sides of the stent 168 is sufficient so that the added
complexity of the dual beam transducer geometry is not required for
self compensation.
[0111] In the case of pulsed Doppler velocity measurements, a
single conformal transducer array 174A would again likely be
adequate so long as the alignment of the conformal transducer array
174A to the stent 168 is accurately controlled. If the alignment of
the conformal array transducer 174A is not controlled or not well
known, a second such conformal transducer array 174B can be used to
gather velocity data along a second beam axis using pulsed Doppler
velocity measurements. Assuming that the second axis is tilted in
an equal but opposite direction as the first axis, the Doppler
measurements made by the two conformal transducer arrays 174A and
174B should be self-compensating for tilt errors. In this case, the
second conformal transducer array 174B could be mounted on the same
or on an opposite side of the stmt from that where the first
conformal transducer array 174A is mounted to implement the Doppler
measurements.
[0112] For CW or pseudo-CW Doppler velocity measurements (in which
a relatively long duration pulse of ultrasonic waves is produced),
the transit signal is applied for a sufficiently long period so
that a second conformal transducer array 174B is needed to receive
the echo signals. In this case, a single set of diametrically
opposed conformal transducer arrays 174A and 174B can be used.
[0113] As perhaps best illustrated in FIG. 14, the conformal
transducer arrays 174A and 174B need not wrap entirely around the
stent 168. In the illustrated embodiment, the conformal transducer
arrays 174A and 174B each span an arc of approximately 60.degree.
around the longitudinal axis of the stent 168 (i.e., about the
center of the circular stent 168 as shown in FIG. 14). This
geometry produces a measurement zone through which ultrasonic beams
178 propagate that is nominally equal to about 50% of the outer
diameter of the stent 168. If used for Doppler velocity
measurements, it is contemplated that the conformal transducer
array 174A need cover only a central portion of the stent 168. As a
result, the span of the conformal transducer arrays 174A and 174B
can be reduced from about 60.degree. to about 45.degree..
[0114] To produce a wide, uniform ultrasonic beam such as that
needed for transit time measurements of flow, the conformal
transducer arrays 174A and 174B must produce ultrasonic waves
having a wave front characterized by a substantially uniform
amplitude and phase. As shown in FIG. 13, lateral projections
through each of a plurality of transducer elements comprising the
conformal transducer arrays 174A and 174B are indicated by straight
lines 176. These straight lines 176 indicate the centers of the
transducer elements and are perpendicular to the axis of
propagation of waves 178 (represented by bidirectional arrows
directed along the axes of propagation of the ultrasonic waves). In
one embodiment, the spacing between the element centers, i.e.,
between the straight lines 176, is approximately equal to a phase
angle of 90.degree. at the excitation frequency of the conformal
transducer arrays 174A and 174B. Thus, starting at the top of FIG.
13 and working downwardly, transducer elements disposed along each
of the displayed straight lines 176 produce acoustic waves that are
successively delayed by 90.degree., or one-quarter wavelength in
the fluid medium through which the ultrasonic waves propagate. For
tissue, a sound velocity of 1,540 meters/second is normally
assumed, so that the physical spacing of the projected straight
lines would typically be defined by the following:
Projected Spacing in millimeters=1.54/(4*F.sub.0),
[0115] where F0 is equal to the center frequency in MHz. If zero
degrees is assigned to the top-most element of the conformal
transducer array 174A, the next element would operate at
-90.degree. relative to the top element, followed by an element
operating at -180.degree., and then one operating at -270.degree.,
and finally by an element operating at 0.degree. relative to the
top electrode. Thus, the conformal transducer array 174A produces a
succession of ultrasonic waves spaced apart by a 90.degree. phase
shift, thereby achieving a desired phase uniformity across the
conformal transducer array 174A.
[0116] While the discussion herein is in terms of phase shifts of
90.degree., it will be appreciated that other types of transducer
element spacings or relative displacements may require different
phase shifts. For example, three phase transducers are known that
employ a phase shift of 120.degree. between adjacent elements.
Additionally, physical displacements of the transducer elements in
the direction of propagation of the acoustic waves may require
different or additional phase shifts between the electrical signals
coupled to the elements. It is possible to phase shift these
signals to provide a uniform phase front in the propagating
acoustic wave using conventional techniques.
[0117] Amplitude uniformity can be achieved in the ultrasonic wave
front by apodization or "shaving" of the elements of the conformal
transducer arrays 174A and 174B. Although shaving could be achieved
in a variety of ways, one embodiment controls shaving by varying
the area of each element.
[0118] In one embodiment, the conformal transducer arrays 174A and
174B are carried on a band 172 made from the piezoelectric plastic
material used for the element substrate, which is sized to fit
snugly around an outer surface of the stent 168 or inserted into
the lumen of the stent 168 (as shown in FIG. 14). The band 172 is
intended to position the conformal transducer arrays 174A and 174B
in acoustic contact with the wall of stent 168, when the band 172
is wrapped around the stent 168, or to maintain the conformal
transducer arrays 174A and 174B against the inner surface of the
stent 168, when the band 172 is inserted into the lumen of the
stent 168. Contact of the band 172 around the outer surface of the
stent 168 assures that the ultrasonic waves produced by the
elements of the conformal transducer arrays 174A and 174B are
conveyed into the fluid flowing through the interior of the lumen.
In one embodiment, the piezoelectric plastic comprising the band
172 is fabricated from a material such as polyvinylidene fluoride
(PVDF), poly(vinyl cyanide-vinyl acetate) copolymer (P(VCN/VAc), or
poly(vinylidene fluoride-trifluoroethylene) copolymer
(P(VDF-TrFE)), available from AMP Sensors of Valley Forge, Pa. In
one embodiment, P(VDF-TrFE) is used because of its high
piezoelectric coupling and relatively low losses.
[0119] Referring now to FIG. 15, further details of the conformal
transducer arrays 174A and 174B are illustrated. In this
embodiment, adjacent elements of the conformal transducer arrays
174A and 174B produce ultrasonic waves differing by 90.degree.. In
the view shown in FIG. 15, a cut line 175 intersects the lateral
center of the conformal transducer array 174B. In practice, any cut
would more likely extend through the band 172 at a point
approximately midway between the conformal transducer array 174A
and the conformal transducer array 174B. Electrodes comprising each
element of the conformal transducer arrays 174A and 174B can be
photolithographically generated on the piezoelectric plastic
substrate comprising the band 172. Alternatively, the elements can
be formed on a non-piezoelectric material comprising the band 172,
and then the material with the elements formed thereon can be
bonded to a piezoelectric substrate in each area where a conformal
array transducer element is disposed. In this latter embodiment, it
is contemplated that a flexible circuit material such as a
polyimide could be employed for the band 172 and that conventional
photolithographic processing methods might be used to fabricate the
conformal array transducer circuitry on the band 172. Further, the
centers of alternating conformal array elements are coupled
together electrically via conductors 180 (shown as dashed lines) in
FIG. 15. Not shown in FIGS. 13 through 15 are the leads that extend
from an implantable electronic circuit used to drive the conformal
transducer arrays 174A and 174B. Any of the implantable electronic
circuits shown in FIGS. 1 through 6 could be used for the
implantable electronic circuits.
[0120] The pattern of elements comprising each of the conformal
transducer arrays 174A and 174B and the boundary of each conformal
transducer array 174A and 174B (top and bottom as shown in FIG.
15), define sinusoidal segments. The period of the sine wave from
which these sinusoidal segments are derived is approximately equal
to the circumference of the band 172. Further, the amplitude of
that sine wave generally depends on the desired beam angle relative
to the longitudinal axis of the stent. For the sinusoidal segment
employed for each electrode, the amplitude is defined by:
Amplitude=D*tan .THETA.
[0121] Similarly, the amplitude of the sinusoidal segment defining
the boundary of each conformal array 174A and 174B is defined
by:
Amplitude=D/(tan .THETA.),
where O is equal to the angle between the longitudinal axis of the
stent 168 (see FIG. 13) and the ultrasound beam axis 178 and D is
equal to the external diameter of the stent 168. Accordingly, it
should be apparent that one sinusoidal template could be used to
draw all of the transducer elements and a second sinusoidal
template (differing only in amplitude from the first) could be used
to draw the boundary of each conformal array transducer 174A and
174B. The transducer elements are displaced or spaced apart from
one another as required to achieve the phase relationship described
above in connection with FIG. 13. In addition, the actual physical
electrode pattern and placement of the elements on the band 172 can
be determined by finding intersection loci between the band 172 as
wrapped around (or within the inner circumference of) the stent 168
and equally-spaced planes. The spacing between these planes is
defined by the equation noted above for the projected spacing.
[0122] The conductors 180 that couple to adjacent transducer
elements differ in phase by 90.degree.. There are two ways to
achieve the 90.degree. phase variation between the ultrasonic waves
produced by successive electrodes in the conformal transducer
arrays 174A and 174B. In the first approach, a uniformly polarized
piezoelectric plastic substrate is used and every fourth element is
coupled together, producing four groups of elements or electrodes
that produce ultrasonic waves having phase relationships of
0.degree., 90.degree., 180.degree. and 270.degree., respectively.
Alternatively, a zone polarized piezoelectric plastic substrate
could be used and every other element can be coupled together (as
shown in FIG. 15). Each of these two groups is then coupled to
provide an in phase and a quadrature phase transceiving system, so
that ultrasonic waves are produced by adjacent elements in each
group have a relative phase relationship of 0.degree. and
90.degree.. In the first approach, a multi-layer interconnect
pattern is required to couple to all traces for each of the
transducer elements in the four groups. In addition, a more complex
four-phase electronic driving system that includes a phase shifter
is required. Specifically, the signal applied to each of the four
groups must differ by 90.degree. between successive elements to
achieve the 0.degree., 90.degree., 180.degree. and 270.degree.
driving signals. The phase shifter, e.g., may be included in the
modulator that drives the conformal transducer arrays 174A and 174B
(which may be included as a part of the RF decode section 40 of
FIGS. 1 through 3 or the RF decode/control section 66 of FIGS. 4
through 6), and provides the phase shifted excitation signals
applied to each successive element of the conformal transducer
arrays 174A and 174B.
[0123] In the second approach, which may be preferred in some
embodiments because it may simplify the electronic package required
and because it may facilitate use of a simpler, double-sided
electrode pattern, the piezoelectric plastic material must be
locally poled in a specific direction, depending upon the desired
phase of the electrode at that location. A poling direction
reversal provides a 180.degree. phase shift, eliminating the need
for 180.degree. and 270.degree. phase-shifted signals. Thus, the
zones of the substrate designated as 0.degree. and 90.degree. would
be connected to the in-phase and quadrature signal sources with the
elements poled in one direction, while zones for elements
designated to provide a relative phase shift of 180.degree. and
270.degree. would be connected to the in-phase and quadrature
signal sources with the elements poled in the opposite direction.
The elements producing ultrasonic waves with a relative phase
relationship of 0.degree. and 180.degree. would comprise one group
(e.g., in phase) and the elements producing ultrasonic waves with a
relative phase relationship of 90.degree. and 270.degree. would
comprise a second group (e.g., quadrature). Poling the different
groups of elements in local regions in opposite directions is
achieved by heating the material above the Curie temperature,
applying electric fields of the desired polarities to each of those
areas and then cooling the material below the Curie temperature
while maintaining the electric fields. This occurs during
manufacture of the conformal transducer arrays 174A and 174B. The
final element wiring pattern required to actually energize the
conformal transducer arrays 174A and 174B when they are employed
for monitoring flow and/or velocity of fluid through the vessel 170
would preclude applying electric fields in opposite polarity.
Accordingly, the required poling relationship would have to be
performed using either temporary electrodes or by providing
temporary breaks in the actual electrode pattern employed in the
final conformal transducer arrays 174A and 174B.
[0124] In one embodiment, to achieve a desired frequency of
operation, it is contemplated that the electrode mass would be
increased to a point well beyond that required for making
electrical connections. This added mass would act together with the
piezoelectric plastic material to form a physically resonant system
at a desired frequency. In this manner, a relatively thinner and
more flexible piezoelectric plastic material can be used for the
substrate comprising the band 172. Use of mass loading is
conventional in the art of ultrasonic transducer design.
[0125] While the fluids within the vessel 170 may provide an
effective ground plane, in one embodiment, a conductive layer 177
(see FIG. 14) is included. The conductive layer 177 may be disposed
on the inside of the band 172 as illustrated (between the conformal
array transducer 174A and the band 172). In one embodiment, the
conformal array transducer 174A comprises a sandwich of two layers
of piezoelectric plastic, with the driven electrodes disposed
between the two layers of piezoelectric plastic, and ground planes
disposed to either outside surface of the conformal array
transducer 174A. The transducer then comprises a ground plane, a
layer of piezoelectric plastic, a layer of driven electrodes, a
layer of piezoelectric plastic and the other ground plane. This
embodiment has the advantage that the conformal array transducer
174A is well shielded and further is electrically isolated from
body fluids. Other arrangements will also be apparent to those of
skill in the art. When the conformal transducer arrays 174A and
174B are used to transmit ultrasonic waves, the conductive layer
177 may be floating (a "virtual ground") or may be coupled to a
ground or common circuit (e.g., 34, FIGS. 1 through 6). When the
conformal transducer arrays 174A and 174B are used to receive
ultrasonic waves, the conductive layer 177 should be coupled to a
common circuit or ground to reduce noise and EMI.
[0126] In FIGS. 16A and 16B, an alternative approach for monitoring
the velocity of a fluid through an interior 250 of a stent 240 is
illustrated. A pair of ultrasonic transducers 242A and 242B may be
realized as conformal transducer arrays, e.g., as described in
conjunction with FIGS. 13 through 15, or may be realized in other
forms, and may find application as therapeutic transducers in
addition to being useful as diagnostic transducers.
[0127] In the embodiment illustrated in FIGS. 16A and 16B, the pair
of ultrasonic transducers 242A and 242B are mounted in relatively
close proximity within a wall 244 of the stent 240. Alternatively,
the ultrasonic transducers 242A and 242B may be disposed externally
in contact with the outer surface of a stent (not shown). The
ultrasonic transducers 242A and 242B each produce a pulse and
receive an echo back from fluid flowing through the interior 250 of
the stent 240, the echoes being scattered from the fluid flowing
therein. In this embodiment, the signal received from the
ultrasonic transducer 242A in response to the echo is correlated
with a similar signal from the ultrasonic transducer 242B,
resulting in a time delay estimate. The velocity of the fluid is
then computed by dividing a distance between the center of the
ultrasonic transducer 242A and the center of the ultrasonic
transducer 242B by the time delay that was determined from the
correlation analysis. This is explained in more detail as
follows.
[0128] The interaction of the blood with the ultrasound, even when
it is moving at constant velocity, gives rise to a moving acoustic
"speckle" pattern. The term speckle, as used herein, has a similar
meaning in ultrasonics as in optics. It results any time that
narrow-band illumination is used. Optical speckle is visible when a
laser (e.g., a pointer) illuminates a plain white wall. When
illuminated with wideband illumination, the wall appears white and
smooth. When illuminated with laser light, the wall appears to have
bright and dark spots, hence the term speckle. Acoustic speckle is
visible in medical ultrasound images, when the system is used to
image homogeneous soft tissues such as the liver. As in optics, the
acoustic speckle pattern is stationary and constant unless the
tissue or flood is moving with respect to the imaging system. The
same phenomenon is exploited in Doppler systems. When the echo
return from moving blood is constant, there is no observable
Doppler shift in the echo signal.
[0129] The blood consists of thousands of scatterers, and the
ultrasound reflects from ensembles of these scatterers. The
amplitude and phase of the echo, at a given range, depends on the
local distribution of scatterers, which is random. The random
signal of echo amplitude and phase at a given depth repeats as the
blood flows past the second ultrasonic transducer 242B, if the
spacing between the two ultrasonic transducers 242A and 242B is
such that the ensembles of scatterers have not changed
significantly, i.e., if the two ultrasonic transducers 242A and
242B are close enough to each other that turbulence has not
significantly disrupted the ensembles of scatterers. Correlation of
nominally identical random patterns that are displaced in time by
an amount equal to the time required for the blood to move from the
first beam to the second one allows the velocity to be determined
when the separation between the two ultrasonic transducers 242A and
242B is known.
[0130] In other words, the first ultrasonic transducer 242A
receives an echo signal that provides a speckle "image"--where the
distance from the ultrasonic transducer 242A is along the vertical
dimension in FIGS. 16A and 16B, and the successive echo returns are
along the horizontal dimension. The two "images" from the two
ultrasonic transducers 242A and 242B are correlated in the
horizontal dimension, and what results is an instantaneous map of
travel time vs. depth.
[0131] The sampling aperture for this system is much shorter than
the time required for a heartbeat. Accordingly, a series of
measurements, which may be taken during the interval between two
successive heartbeats, may be processed or compared to determine
peak, minimum and average blood velocity when these data are
desired.
[0132] Unlike a Doppler system, the echoes in a correlation type
transducer system like that shown in FIGS. 16A and 16B are not
frequency shifted. Instead, the velocity signal is extracted by
correlating the echo amplitude versus time signals for a pair of
range bins. The velocity versus time is independently determined
for each range bin, resulting in a time dependent velocity profile
across the diameter of the stent 240.
[0133] The conformal transducer arrays 174A and 174B of FIGS. 13
through 15 can be formed on the band 172, but alternatively, can be
included within the structure of a stent, i.e., within its wall.
FIG. 17 illustrates a portion of a cross-sectional view of the
conformal array transducer 174A of FIGS. 13 through 15 fabricated
in a stent wall 190. The entire conformal array transducer 174A is
fitted within the stent wall 190. Details of the stent wall 190 are
not illustrated, since it is contemplated that many different types
of stent configurations are suitable for carrying the conformal
transducer arrays 174A and 1748. The stent wall 190 is shown inside
a blood vessel wall 204. A biocompatible outer coating 192
comprises the next layer, protecting the conformal transducer
arrays 174A and 174B from contact with bodily fluids. In one
embodiment, the outer coating 192 comprises PARYLENE.TM. material,
available from Specialty Coating Systems of Indianapolis, Ind.
Outer coatings 192 comprising PARYLENE.TM. material may be grown to
a desired thickness via vapor coating. In one embodiment, the outer
coating 192 is grown to a thickness of between 0.0001'' to 0.0002''
(2.5 to 5 microns). Below the outer coating 192 is an acoustic
backing 194 comprising a conventional, or a syntactic foam, i.e., a
polymer loaded with hollow microspheres, that serves both for
acoustic isolation and dampening and to minimize capacitive
loading.
[0134] In one embodiment, the acoustic backing 194 comprises one
volume of EPOTEK 377 or 301-2 epoxy glue available from Epoxy
Technology of Billerica, Mass. mixed, e.g., with two or more
volumes of microballoons available from PQ Corp. of Parsippany,
N.J. Microbubbles such as PM6545 acrylic balloons having an average
diameter of 100 microns are employed in one embodiment, with the
acoustic backing being 10 to 20 microballoons thick (one to two
mm). The acoustic backing 194 has a relatively low dielectric
constant (e.g., <10), thereby minimizing capacitive loading
between the electrodes and surrounding tissue. The acoustic backing
194 thus insulates the transducer elements from the surrounding
fluid and tissue in a capacitive sense and also in an acoustic
sense. The next layer comprises a rear electrode 196. A front
electrode 200 is spaced apart from the rear electrode by a
piezoelectric plastic layer 198. In one embodiment, the front
electrode 200 is also the conductive layer 177 of FIG. 14. As noted
above, in the embodiment illustrated in FIGS. 13 through 15, the
piezoelectric plastic layer 198 of FIG. 17 comprises the band 172
of FIGS. 13 through 15. The piezoelectric layer 198 (or the band
172) has a relatively low dielectric constant, e.g., from about six
to eight, compared to tissue (approximately 80).
[0135] In one embodiment, the rear electrode 196 and the front
electrode 200 comprise multi-layer structures (although separate
layers are not shown). For example, the electrodes 196 and 200 will
include a metallic layer that bonds well to the piezoelectric
plastic layer 198, for example, titanium, followed by a highly
conductive layer, for example, copper, followed by an oxidation
resistant layer, for example, gold, and includes other metallic
barrier layers, where appropriate, to prevent reaction between
these layers. Such multi-layer systems are conventional and are
suited for use as the electrodes 196 and 200 in the conformal
transducer arrays 174A and 174B.
[0136] In one embodiment, the front electrode 200 is the "common
electrode" for the transducer elements and serves as a RF shield. A
front coating 202 serves as an acoustic coupling between the
conformal transducer arrays 174A and 174B and the fluid in the
lumen of the stent. In addition, the front coating layer 202 serves
as a biocompatible layer, providing a barrier to fluid ingress into
the conformal array transducers 174A and 174B.
[0137] In both the conformal array transducers 174A and 174B
provided in the band 172 (as shown in FIGS. 13 through 15) and the
conformal array transducer 174A included within the structure of
the stent wall 190, as illustrated in FIG. 17, it is contemplated
that adhesive layers (not shown) may be used between the various
layers. However, certain layers such as the front and rear
electrodes 200 and 196 will likely need not be adhesively coupled
to the piezoelectric layer 198 if photolithographically formed on
the piezoelectric layer 198. Other layers may not require an
adhesive to couple to adjacent layers, e.g., if formed of a
thermoset material that self bonds to an adjacent layer when
set.
[0138] As noted above, one of the advantages of the conformal
transducer arrays 174A and 174B is a relatively low profile. In
some cases, a stent may integrally accommodate a relatively thicker
profile transducer assembly. An embodiment of a tilted element
transducer 210 coupled to a stent 203 that is useful as a
diagnostic transducer or as a therapeutic transducer is illustrated
in FIG. 18. Each element comprising the tilted element transducer
210 includes the rear electrode 196 and the front electrode 200
disposed on opposite sides of the piezoelectric material 198.
Conventional prior art transducers for producing an ultrasonic
waves use a single such element that has a substantially greater
width that is often too great for inclusion within a stent
assembly. In contrast, the tilted element transducer 210 includes a
plurality of elements like those shown in FIG. 18 that minimize the
radial height (or thickness) of the tilted element transducer
210.
[0139] An outer coating 195 again serves the function of providing
a biocompatible layer to protect the transducer components
contained therein from exposure to bodily fluids. When the outer
coating 195 comprises PARYLENE alone, an RF shield 193 extends over
the tilted elements, immediately inside the outer coating 195. When
the outer coating 195 comprises a container (as illustrated), it
includes an outer coating of a material such as PARYLENE. When the
outer coating 195 comprises a conductive material, a separate RF
shield such as the RF shield 193 may not be required. The acoustic
backing 194 is disposed below the RF shield 193 or the outer
coating 195.
[0140] An acoustic filler material 212 is disposed between the
front electrode 200 and the front coating 202, on the interior
surface of the stent 203, and is used to fill in the cavities in
front of the transducer elements. The acoustic filler material 212
is characterized by a relatively low ultrasonic attenuation, so
that it readily conveys the ultrasonic waves produced by the
transducer elements into the lumen of the stent 203. In one
embodiment, in order to minimize reverberations of the ultrasonic
waves in this acoustic filler material 212, its acoustic impedance,
which is related to sound velocity times density, is approximately
equal to that of the fluid in the vessel. The velocity of sound in
the acoustic filler material 212 should also be close to that of
the fluid flowing through the stent 203 so that the sound beam is
not significantly deflected by the acoustic filler material 212. In
another embodiment, the acoustic filler material 212 has a
relatively low sound velocity compared to the fluid. In this
embodiment, the acoustic filler material 212 acts as an acoustic
lens that deflects the sound being produced by the elements of the
tilted element transducer 210. For example, materials such as
silicones or fluorosilicones typically having sound velocities
about 1000 meters per second (compared to a sound velocity of
approximately 1540 meters per second for blood) may be used. Low
velocity lenses are conventional. A benefit of using a low velocity
acoustic filler material 212 is that the elements of the tilted
element transducer 210 can be tilted about 30% less than would be
required otherwise. As a result, the overall height of the tilted
element transducer 210 portion of the stent 203 can be made about
30% thinner than would be possible without the low velocity
acoustic filler material 212. In combination, the plurality of
tilted elements of the tilted element transducer 210 produce an
ultrasonic wave 214 that propagates at an angle relative to the
longitudinal axis of the stent, which is represented by a center
line 216 in FIG. 18.
[0141] FIG. 19A is an isometric view of an implantable integrated
circuit (IC) transducer 220 mounted on a tubular stent 222, which
may comprise a stent similar to that described in conjunction with
FIG. 11A above. Wires form a RF coupling coil 223 coupled to the
implantable IC sensor 220 via wires 225. The wires comprising the
RF coupling coil 223 are formed in a zigzag shape to allow for
expansion of the tubular stent 222 when it is installed. The
implantable IC sensor 220 may include diagnostic or therapeutic
transducers. In one embodiment, the sensing apparatus of the
implantable IC sensor 220 faces the interior of the tubular stent
222, as is described more fully with respect to FIG. 19B below.
[0142] FIG. 19B illustrates the implantable IC sensor 220 mounted
on the tubular stent 222, so that the implantable IC sensor 220
overlies a sensor window opening 224 in the tubular stent 222.
Conductive adhesive or solder 228 couples the implantable IC sensor
220 contacts to the tubular stent 222 (or to conductors that are
coupled to one of the implantable electronic circuits shown in
FIGS. 1 through 6). A biocompatible coating 226 (analogous to the
biocompatible coating 192 of FIG. 17) encloses the implantable IC
sensor 220, except in the area of the sensor window opening 224
through which the implantable IC sensor 220 is in contact with the
fluid flowing through the lumen of the tubular stent 222. The
portion of the tubular stent 222 on which the implantable IC
transducer 220 is mounted may be made rigid, e.g., by thickening
it, to prevent damage to the implantable IC transducer 220 during
the installation of the tubular stent 222. Optionally, a circuit
board (not illustrated in FIG. 19B) may be included between the
implantable IC transducer 220 and the tubular stent 222 to
facilitate making electrical interconnections to the RF coupling
coil 223.
[0143] FIG. 19C illustrates an embodiment wherein the implantable
IC transducer 220 is coupled to a woven mesh stent 222A. The woven
mesh stent 222A comprises wires woven to form a mesh comprising a
RF coupling coil 223A as described in conjunction with FIGS. 8 and
9 above. Wires 225A couple the RF coupling coil 223A to the
implantable IC transducer 220. The implantable IC sensor 220 may
include diagnostic transducers, and the woven mesh stent 222A may
include therapeutic transducers (not illustrated). In one
embodiment, the sensing apparatus of the implantable IC transducer
220 is held in place via an encapsulant 226A to face the interior
of the woven mesh stent 222A, as is described more fully with
respect to FIG. 19D below.
[0144] FIG. 19D is an enlarged partial cross-sectional side view of
the implantable IC transducer 220 mounted on the woven mesh stent
222A of FIG. 19C. The implantable IC transducer 220 is coupled to
the wires comprising the body of the woven mesh stent 222A by the
encapsulant 226A which may also serve as a biocompatible fluid
barrier and as an insulator. This keeps the implantable IC
transducer 220 from contacting body fluids except at the sensing
interface which is mounted within an opening 224A in the wall of
the woven mesh stent 222A. A flexible circuit substrate 227
optionally is employed to provide mechanical attachment and
electrical coupling to the implantable IC transducer 220 via solder
bumps 22$ or other conductive and mechanically robust
interconnection. In the embodiments of FIGS. 19A through 19D, the
implantable IC transducer 220 may comprise the implantable
electronic circuits of any of FIGS. 1 through 6 and the stents 222
and 222A may also include other transducers such as diagnostic or
therapeutic transducers.
[0145] It is contemplated that the implantable IC transducer 220
might be used for measuring parameters such as pressure,
temperature, blood gas concentration and insulin level or the
levels of other metabolite such as glucose or sodium in the blood
stream of a patient in which a stent that includes the IC sensor
220 is implanted. As explained above, the implantable IC sensor 220
is electrically energized with electrical power that is
electromagnetically coupled to the RF coupling coil 223A that
comprises the stent body 222A or which is incorporated as one or
more separate insulated windings within the stent wall structure.
Signals produced by the IC-sensor 220 are converted to data
signals, which are electromagnetically coupled to a monitor outside
the patient's body, also as explained above. In certain
applications of implantable IC sensors 220, it may be advantageous
to perform a differential measurement between two spaced apart
locations on the stent body 222 or 222A. Thus, to monitor fluid
flow through the lumen of a stent 222 or 222A, a differential
pressure measurement made by transducers respectively disposed
adjacent the proximal and distal ends of the stent 222 or 222A
provide an indication of blood flow and of any blockage with the
lumen of the stent 222 or 222A.
[0146] If an external source of heat is applied to heat the blood
or other fluid flowing through the lumen of a stent 222 or 222A,
flow can be determined by monitoring the temperature of the fluid
with IC sensors 220 that are responsive to that parameter. An
external source of RF energy electromagnetically coupled into the
stent 222 or 222A, as disclosed above, can both provide the
electrical power for the components of the stent transducer system
and provide the power for heating the fluid. To avoid tissue
damage, the maximum stent temperature should remain below
42.5.degree. C., which is well established as the temperature above
which hyperthermia and irreversible tissue damage occur. By
analyzing the resultant temperature vs. time "thermal washout"
curve, the flow rate of fluid through the stent 222 or 222A can be
determined. A differential temperature measurement made by
temperature sensors disposed adjacent the opposite ends of the
stent 222 or 222A could also be used to determine flow through the
stent lumen. Using the signals from these sensors, two temperature
vs. time curves can be developed simultaneously. Differences in the
observed thermal washout curves should be primarily a function of
flow through the lumen and thus indicative of that parameter.
[0147] Other methods can be employed to determine flow based on
temperature measurements. For example, by modulating the RF power
used to heat the stent 222 or 222A, the temperature vs. time curves
will exhibit the modulation frequency. The temperature vs. time
curves produced by spaced apart temperature sensors can be filtered
with a relatively narrow bandwidth filter. The phases of the two
filtered signals are compared to extract a flow velocity through
the stent 222 or 222A. The signal processing concept of this
approach is conceptually similar to that used for measuring cardiac
output using a catheter-mounted heater and temperature sensors, as
disclosed in U.S. Pat. No. 5,277,191 entitled Heated Catheter For
Monitoring Cardiac Output.
[0148] Several types of IC sensors 220 that might be incorporated
within a stent in accord with the present invention are disclosed
in previously issued U.S. patents. For example, U.S. Pat. No.
4,020,830 (and re-examination certificate U.S. Pat. No. 4,020,830)
entitled Selective Chemical Sensitive FET Transducers and U.S. Pat.
No. 4,218,298 entitled Selective Chemical Sensitive FET Transducer
describe chemical field effect transistor (FET) transducers that
are sensitive to specific chemical substances or to their
properties. U.S. Pat. No. 4,935,345 entitled Implantable
Microelectronic Biochemical Sensor Incorporating Thin Film
Thermopile discloses an implantable microelectronic biochemical
sensor that incorporates a thin film thermopile for use in
monitoring concentrations of glucose or other chemicals present in
the blood stream. Various types of pressure sensing devices
appropriate for incorporation in the wall of a graft are readily
available from a number of different commercial sources, including
SRI Center for Medical Technology of Palo Alto, Calif.
[0149] Other prior art devices are potential candidates for use as
IC sensors 220 on stents 222 or 222A. In Evaluation of a Novel
Point-of-Care System, the I-Stat Portable Clinical Analyzer,
CLINICAL CHEMISTRY, Vol. 39, No. 2, 1993, K. A. Erickson et al.
describe a blood analyzer based on disposable IC biosensors that
can quantify sodium, potassium, chloride, urea, nitrogen and
glucose levels. A good overview of acoustic wave biosensors is
provided by J. C. Andle et al. in Acoustic Wave Biosensors,
published in the 1995 IEEE Ultrasonics Symposium Proceedings, IEEE
cat. no. 0-7803-29406/95, pp. 451-460. Other types of IC biosensors
are described in the art.
[0150] However, it is sufficient for this disclosure to recognize
that such IC sensors 220 are well known in the art and are
generally available or readily fabricated for use on stents 222 or
222A (or other stent designs) as described above.
[0151] In the embodiments of FIGS. 19A through 19D, the implantable
sensor IC 220 senses the concentration of a particular substance or
another parameter that was determined to require monitoring prior
to implanting the implantable sensor IC 220 or that is selected
from a plurality of sensing capabilities provided on the
implantable sensor IC 220 in response to control signals coupled
via the RF coupling coil 223 or 223A. This diagnostic information
is then used in conjunction with the therapeutic transducers of any
of FIGS. 13 through 15, 17, 18, 20, 21 and 23 through 28.
[0152] A stent may include other types of sensors beside the
ultrasonic transducers and the IC sensor 220 noted above. FIG. 20
illustrates an outline of a stent 232 that includes a strain sensor
comprising strain sensing filaments 230 mounted on the stent 232.
In the disclosed embodiment, strain sensing filaments 230 are wound
around the stent 232 to measure displacement that is converted to a
signal for transmission outside the body by an implantable IC 220A.
The filaments 230 exhibit a change in electrical resistance with
strain and are therefore usable to sense the strain experienced by
the stent 232 when it is expanded inside a blood vessel. It is
contemplated that the strain sensing filaments 230 be used only for
strain sensing, so that their dimension, disposition and metallurgy
can be optimized for that function. Alternatively, the strain
sensing filaments 230 can comprise part of the structural body of
the stent 232 so that they also provide a mechanical function
related to the conventional function of the stent 232. It is also
contemplated that strain gauges (not separately shown) can be used
instead of the strain sensing filaments 230. The strain gauges can
be mounted to the stent 232 at selected spaced-apart locations to
measure displacement. Metallized polyimide substrate strain gauges
are suited to this application, by wrapping the substrates around
the body of the stent 232 and attaching the substrates to the body
of the stent 232 at selected spaced-apart points. By monitoring the
size of stent 232 as it is expanded, strain gauges or other strain
measuring sensors can determine when a desired expansion of the
stent 232 has been achieved. Alternatively, the strain data can be
employed to assess the elasticity of the stent 232 and blood vessel
structure by monitoring the dynamic strain over cardiac cycles,
i.e., with successive systolic and diastolic pressure levels.
[0153] Referring to FIG. 21A, an implantable IC sensor 220B that
detects fatty deposits and tissue growth inside the lumen of the
stent 232 is illustrated as being disposed within the body of the
stent 232, coupled to a pair of dielectric sensing filaments 234.
The implantable IC sensor 220B detects fatty deposits and tissue
ingrowth within the lumen of the stent by measuring the dielectric
and/or resistive properties of any material in contact with the
sensing filaments 234, which are, e.g., helically coiled around the
inner surface of the stent 232, from about one end of the stent 232
to its opposite longitudinal end. Alternatively, the sensing
filaments 234 can be incorporated into the body or wall of the
stent 232 itself. For example, when the stent 232 is fabricated
with a woven mesh, a portion of the mesh can be utilized for making
the dielectric and/or resistive measurement, while the remainder is
used for a RF coupling coil.
[0154] In another embodiment, the sensing filaments 234 may be
spatially more limited to allow assessment of where blockage is
occurring within the stent 232. A plurality of localized sensing
filaments 234 may permit assessment of more than one area within
the stent 232, by taking a series of measurements and communicating
the results of the series of measurements to the attending
physician. This may provide data relevant to determining what form
of treatment is appropriate.
[0155] For measuring the dielectric properties, the implantable IC
sensor 220B is energized with power electromagnetically coupled
from an external source into the RF coupling coil (not illustrated
in FIG. 21A) of the stent 232 and produces signals indicative of
tissue ingrowth that are electromagnetically coupled to the
external monitoring system through the RF coupling coil of the
stent 232. In one embodiment, an RF signal at a frequency of from
10 to 100 MHz is applied to the sensing filaments 234. At such
frequencies, tissue has the properties shown in the following Table
1. FIG. 21B illustrates an exemplary cross section of a lumen 235
within the stent 232, showing the ingrowth of fatty tissue 236,
which is in contact with the sensing filaments 234.
TABLE-US-00001 TABLE 1 Tissue Type Relative Permittivity
Resistivity (Ohm-cm) Fat 6 to 20 2000 to 3000 Blood 80 to 160 80 to
90 Muscle 60 to 130 100 to 150
[0156] The permittivity of tissue is closely related to its water
content. Water has a relative permittivity of about 80. Since fat
and fatty deposits of the type found inside blood vessels contain
much less water than other tissue types, the permittivity of fat is
much lower than that of muscle or blood. The wall of a blood vessel
is muscular and highly perfused and will therefore have a much
higher permittivity than a fatty deposit. Similarly, fatty deposits
have a much higher resistivity than either blood or muscle.
Therefore, a measurement of the dielectric and/or resistive
properties of tissue inside the stent 232 can differentiate fatty
deposits from either blood or muscular tissue ingrowth into the
lumen. The measurement can include a determination of capacitance,
resistance or a combination of the two.
[0157] Further information can be obtained from the frequency
dependence of the capacitance and resistance measured inside a
stent lumen. For example, blood has a relatively flat resistivity
vs. frequency characteristic curve, compared to that of muscle.
[0158] FIG. 22 illustrates a stent graft (or spring graft) 260 that
includes embodiments of the present invention. The stent graft 260
differs from a conventional synthetic graft in the method of
delivery. Conventional grafts are installed surgically, while stent
grafts 260 are installed using an endovascular delivery system. The
entire stent graft 260 must be collapsible onto a delivery catheter
(not shown). At a minimum, the stent graft 260 comprises a
synthetic graft section 264 with an expandable stent 262 or 266
disposed at one or both ends. The stents 262 and 266 retain the
synthetic graft section 264 in position. Some stent grafts 260 have
stents 262 and 266 disposed along the entire length of the
synthetic graft section 264, and some may include metal hooks at
one or both ends to firmly attach the synthetic graft section 264
to the vessel wall.
[0159] The stent graft 260 is of a type that is used to repair
arteries near a bifurcation of the artery into two small branches
268 and 270. However, it should be noted that the present invention
can be used with almost any type of stent graft and is not in any
way limited to the bifurcated type shown in the figure. The TALENT
spring graft system available from World Medical Manufacturing is
similar to the stent graft 260. The term "spring graft" is used
with this type of stent graft 260 because the stent portions 262
and 266 may be self-expanding, comprising Nitinol springs acting as
stents 262 and 266 that are embedded into polyester (DACRON.TM.) or
FIFE synthetic graft section 264. The larger diameter aortic
section typically comprises DACRON and the smaller branch portions
typically comprise FIFE. The material comprising the synthetic
graft section 264 is stitched to the Nitinol stents 262. Although a
Nitinol stent is normally self-expanding, a balloon (not shown) may
be included in the delivery system to perform one or more
functions, including expansion of the stent 262, placement at the
desired location, flow occlusion and straightening blood vessels to
aid advancement of the assembly to the desired location.
Electrically insulating ceramic joints 276 couple sections of each
stent 262 and 266 to break any current loop that could reduce the
efficiency of the RF coupling coil. An insulated wire 272 is wound
around the outside of the graft 264 and, in one embodiment, is
formed of kinked or zigzag wire to enable expansion of the graft
264. The wire 272 is coupled to a sensor/electronic circuit 274.
Stent grafts suitable for use in the embodiment shown in FIG. 22
are made by Sulzer Vascutek and W. L. Gore. The ANEURX stent graft
from Medtronic, and the WALLGRAFT stent graft from
Medivent-Schneider, which includes a woven mesh stent within its
wall, are also suitable for this embodiment.
[0160] Description of Therapeutic Transducers
[0161] A variety of therapeutic transducers may be implanted that
are responsive to and/or powered by the signals coupled into the
implantable electronic circuits of FIGS. 1 through 6. One class of
therapeutic transducers 44-46 provide utility by enabling localized
delivery or activation of specific drugs for specific purposes. Two
distinct applications where therapeutic transducers implanted
within endoluminal implants provide therapeutic advantages are as
adjunctive therapy and as primary therapy.
[0162] In adjunctive therapy, the therapeutic transducer is
intended to realize localized drug activation and delivery in the
vicinity of the stent or stent graft. This could be to maintain
flow capability through the lumen by reducing restenosis due to new
deposits of atherosclerotic material or to inhibit tissue ingrowth.
Alternatively, in at least some cases, the same therapeutic
transducer may aid in reducing thrombosis that is causing lumen
blockage by activating appropriate drugs.
[0163] In primary therapy, the stent with the therapeutic
transducer is implanted specifically to provide local drug
activation and delivery to tissue in the vicinity of and downstream
from the stent. For example, a stent could be implanted in an
artery that feeds blood to a tumor site. Systemically administered
chemotherapeutic agents that are not toxic until activated may be
activated during passage through the stent by energy provided by
the therapeutic transducer. The blood containing the activated drug
then proceeds downstream to the tumor site to locally administer
the activated drug. This approach can provide significantly greater
drug concentrations at the tumor site than are obtained
systemically. Similarly, other drugs used to treat a variety of
diseases may be locally activated at the region of interest. In
some cases, modified genetic material may be locally concentrated
in response to therapeutic transducer activation.
[0164] One advantage to localized activation or delivery of drugs
is that the side effects associated with the drugs may be reduced
by only providing the drug at the site requiring treatment. This is
advantageous in many situations, including chemotherapy, where the
drugs are toxic or may have other potentially detrimental side
effects.
[0165] For example, drug activation phenomena have been reported
using ultrasound to break precursor substances down into drug
molecules and other by-products. In this case, one or more of the
transducers 44-46 of FIGS. 1 through 6 are ultrasonic transducers,
several of which are described with respect to FIGS. 13 through 18.
Sonochemical activation of hematoporphyrin for tumor treatment is
described by S. I. Umemura et al. in Sonodynamic Activation of
Hematoporphyrin: A Potential Modality For Tumor Treatment,
published in the 1989 IEEE Ultrasonics Symposium Proceedings, IEEE
cat. no. 00905607/89/0000-0955, pp. 955-960. Ultrasonic
potentiation of adriamycin using pulsed ultrasound is described by
G. H. Harrison et al. in Effect Of Ultrasonic Exposure Time And
Burst Frequency On The Enhancement Of Chemotherapy By Low-Level
Ultrasound, published in the 1992 IEEE Ultrasonics Symposium
Proceedings, IEEE cat. no. 1051-0117/92/0000-1245, pp. 1245-1248.
Similarly, increased toxicity of dimethlyformamide has been
reported in conjunction with ultrasound by R. J. Jeffers et al. in
Enhanced Cytotoxicity Of Dimethylformamide By Ultrasound In vitro,
published in the 1992 IEEE Ultrasonics Symposium Proceedings, IEEE
cat. no. 1051-0117/92/0000-1241, pp. 1241-1244.
[0166] Sonodynamic activation at one or more specific body sites to
provide local drug delivery is possible when one or more of the
transducers 44-46 of FIGS. 1 through 6 are designed to provide
suitable ultrasonic signals and are implanted at the locations
where drug activation provides therapeutic benefits. Sonodynamic
effects are nonlinear effects associated with the peak compression
and expansion portions of the wave cycle; at lower frequencies, the
time that the peak portions of the wave have to act is greater. For
this reason, lower frequencies are preferred in some embodiments
Other embodiments increase peak forces by combining two or more
ultrasonic waves. Several such transducers are described in
connection with FIGS. 23 and 24 below.
[0167] FIG. 23 illustrates an ultrasonic transducer configuration
278 integrated with a stent 279. The ultrasonic transducer
configuration 278 is specifically designed to provide sonodynamic
therapy via standing waves providing sonochemical activation of
blood-borne drug precursors. This occurs in response to control
signals coupled from the implantable electronic circuitry of FIGS.
1 through 6 by lines 281. The ultrasonic transducer configuration
278 is useful where local drug activation is desired in order to
deliver the drug to the vessel having the stent 279 therein. The
ultrasonic transducer configuration 278 is also useful when the
downstream vasculature or an organ or tumor that is supplied blood
via the downstream vasculature is the intended target for the
activated drug.
[0168] The stent 279 includes an implantable ultrasonic transducer
280 on a first surface and a device 282 on a second surface. The
device 282 may be either another ultrasonic transducer similar to
the transducer 280 or an acoustic reflector. The ultrasonic
transducer 280 may be coupled to implantable electronic circuits
using any of the approaches described in connection with FIGS. 1
through 6. In one embodiment, the layer structure described in
connection with FIG. 17 is applicable to the transducer 280. The
standing acoustic wave, represented by the dashed parallel lines in
FIG. 23, that is realized between the transducer 280 and the device
282 results in greater peak acoustic field strength for a given
input energy level, which increases the rate of sonochemical drug
activation and reduces the power levels required for sonochemical
drug activation. Peak acoustic pressure increases of three- to
fivefold are likely in most clinical settings:
[0169] The piezoelectric material forming the transducer 280 may
comprise piezoelectric plastic materials such as PVDF, P(VCN/VAc)
or P(VDFTrFE), available from AMP Sensors of Valley Forge, Pa., or
any of the piezoelectric ceramics, e.g., lead zirconium titanate.
In one embodiment, PZT-4 material available from Morgan-Matroc of
Bedford, Ohio provides high electroacoustic coupling and low
acoustic losses. In another embodiment, the piezoelectric plastic
P(VDF-TrFE) provides high electroacoustic coupling and low acoustic
losses.
[0170] The transducer 280 (and, when the device 282 is a
transducer, the device 282) may be of the type described, for
example, with respect to FIGS. 13 through 18, or may be a slab type
ultrasonic transducer, or may be similar to that shown and
described in connection with FIG. 24, described below. In this
application, the alignment between the transducer 280 and the
device 282 must be maintained in order to preserve parallelism of
the surface of transducer 280 that surfaces the device 282 and the
surface of the device 282 that faces the transducer 280. It is also
important to keep these surfaces opposed to each other, i.e.,
relative lateral motion of the transducer 280 and the device 282
must be inhibited. The result of maintaining this alignment is to
form an acoustic cavity analogous to an optical Fabry-Perot
resonator.
[0171] FIG. 23 also shows biocompatible coatings 284 surrounding
both the transducer 280 and the device 282. The biocompatible
coatings 284 are analogous to the biocompatible outer coating 192
of FIG. 17. The transducer 280 may also include an acoustic backing
analogous to the acoustic backing 194 of FIG. 17, disposed on the
transducer 280 as described in conjunction with FIG. 17. When an
acoustic backing is employed with the transducer 280 (or in the
device 282), it is important that the surface of transducer 280
that faces the device 282 (and the surface of the device 282 that
faces the transducer 280) not be coated with the acoustic backing
material.
[0172] When the device 282 is chosen to be an acoustic reflector,
either a low impedance reflector (i.e., providing an acoustic
reflection coefficient approaching -1) or a high impedance
reflector (Le., providing an acoustic reflection coefficient
approaching +1) may be employed. Low-density foams (e.g., analogous
to the acoustic backing material 194 of FIG. 17) or aerogels
provide low acoustic impedances suitable for use in acoustic
reflectors, while rigid bodies such as metals or ceramics provide
high acoustic impedances suitable for use in acoustic reflectors.
Setting the thickness TR of the acoustic reflector to be an odd
multiple of one quarter of an acoustic wavelength, as measured in
the acoustic reflector material, increases the reflection
coefficient of the acoustic reflector.
[0173] Alternatively, methods for localized delivery of medication
include encapsulation of medications in delivery vehicles such as
microbubbles, microspheres or microballoons, which may be ruptured
to locally release the medications via localized energy provided by
implanted transducers. In some embodiments, the delivery vehicles
may include magnetic material, permitting the delivery vehicles to
be localized via an applied magnetic field, as described in U.S.
Pat. No. 4,331,654 entitled Magnetically-Localizable. Biodegradable
Lipid Microspheres.
[0174] In one embodiment, the device 282 is formed from a magnetic
ceramic or a magnetic metal alloy, and is also capable of acting as
an efficient acoustic reflector. This embodiment allows
localization of magnetic delivery vehicles via the static magnetic
field associated with the device 282, followed by insonification of
the delivery vehicles when appropriate via ultrasound emitted by
the transducer 280 in response to signals from any of the
implantable electronic circuits shown in FIGS. 1 through 6. As used
herein, the term "insonify" means "expose to sound" or "expose to
ultrasound"; insonification is used to mean exposure to sound or
ultrasound. Insonification of delivery vehicles can provide
localized heating, can rupture microbubbles to locally release
drugs or drug precursors contained in the delivery vehicles or can
trigger sonodynamic activation of drug precursors that are
blood-borne or that are released when the delivery vehicles
rupture. Microbubbles containing antistenotic agents are described,
for example, by R. L. Wilensky et al. in Microspheres, Semin.
Intervent. Cardiol., 1: 48-50, 1996. Microbubbles of various
compositions and filled with various drugs are developed and
manufactured by ImaRx Pharmaceutical Corp. of Tucson Ariz. An
advantage that is provided by use of an implanted permanent magnet
for localization of magnetic delivery vehicles in this embodiment
and others is that permanent magnets do not require a rechargeable
energy source in order to function. In some embodiments, this can
provide a way of reducing power needs from the RF-to-DC power
supply 32 of FIGS. 1 through 6.
[0175] The frequency of the ultrasound from the therapeutic
transducer can be varied to enhance or to reduce cavitation
resulting from the ultrasound emitted from the transducer.
Suppression of cavitation via frequency modulation is described in
U.S. Pat. No. 5,694,936 entitled "Ultrasonic Apparatus For
Thermotherapy With Variable Frequency For Suppressing Cavitation."
Methods for suppression or enhancement of cavitation are described
in U.S. Pat. No. 4,689,986 entitled "Variable Frequency
Gas-Bubble-Manipulating Apparatus And Method." Enhancing cavitation
to enhance sonodynamic activation, rupture of microspheres,
microballoons or microbubbles, to locally heat tissue or to destroy
tissue is possible by causing the frequency of the emitted
ultrasound to decrease with time. On the other hand, cavitation may
be decreased by causing the frequency of the emitted ultrasound to
increase with time. This may be used to limit tissue damage while
still supplying sufficient ultrasound to accomplish, e.g., a
diagnostic purpose.
[0176] Sonodynamic activation of drugs or sonically-induced
delivery vehicles rupture may occur at reduced power levels when
properly-phased collinear acoustic signals at two different
frequencies are provided. This effect has been shown to be
particularly advantageous when one signal is at a frequency that is
the second harmonic of the other signal and the two signals have an
appropriate phase relationship. Increased tissue damage for a given
intensity of ultrasound has also been reported by S. I. Umemura in
Effect Of Second-Harmonic Phase On Producing Sonodynamic Tissue
Damage, published in the 1996 IEEE Ultrasonics Symposium
Proceedings, IEEE cat. no. 0-7803-36151/96, pp. 1313-1318.
Sonochemical activation of a gallium-deuteroporphyrin complex
(ATX-70) at reduced total power density by use of properly phased
signals comprising a first signal and a second signal at twice the
frequency of the first signal is described by S. I. Umemura et al.
in Sonodynamic Approach To Tumor Treatment, published in the 1992
IEEE Ultrasonics Symposium Proceedings, IEEE cat. no.
1051-0117/92/0000-1231, pp. 1231-1240. An example of a transducer
that is designed to provide for transduction of two ultrasonic
signals, one of which may be the second harmonic of the other, is
now described with reference to FIG. 24.
[0177] FIG. 24 illustrates an embodiment of a dual frequency
ultrasonic transducer 290. The dual frequency transducer 290 is
designed to provide two different frequencies of collinearly
propagating ultrasound, where one of the frequencies may be the
second harmonic of the fundamental transducer frequency, when
supplied with suitable electrical signals. The phases of the two
signals may be adjusted by the implantable electronic circuit of
FIGS. 4 through 6 and this may be in response to signals from the
power supply and patient monitoring console 101 of FIG. 12. The
dual frequency transducer 290 comprises a disc 292 of piezoelectric
material, poled, for example, as indicated by direction arrow 298.
The disc 292 has a diameter D and a thickness TX. Electrode 294 and
electrode 296 are formed on opposed surfaces of the disc 292 as
described in conjunction with the rear and front electrodes 196 and
200 of FIG. 17 above.
[0178] In one embodiment, the diameter D is chosen to provide the
desired fundamental transducer frequency via radial mode coupling,
while the thickness Tx is chosen to provide the second harmonic of
the fundamental transducer frequency via thickness mode coupling.
In this case, the diameter to thickness ratio D/TX may be
approximately 2:1. Conventional mode charts provide more precise
ratios for a variety of materials. The radial mode comprises radial
particle motion primarily into and out from the center of the disc,
i.e., perpendicular to the direction arrow 298, and symmetric about
a cylindrical axis of the disc 292. The surfaces of the disc 292
exhibit longitudinal motion (i.e., parallel to the direction arrow
298) in response to the radial mode oscillation because of the
Poisson's ratio of the material. The thickness mode comprises
particle motion parallel to the direction arrow 298. As a result,
acoustic energy propagating in the same direction at both
frequencies may be coupled out of the disc 292 via the surfaces on
which the electrodes 294 and 296 are formed. In some embodiments,
the acoustic radiating surface emitting the ultrasound does not
include an electrode 294 or 296. For example, electrodes may be
disposed on the sidewalls, with ultrasound being emitted from the
planar surfaces.
[0179] In another embodiment, the radial mode providing ultrasound
at the fundamental transducer frequency may be chosen to be a
harmonic of the lowest radial mode of the transducer 290. The
transducer 290 may then be designed to have a larger diameter D
than is possible when the lowest radial mode corresponds to the
fundamental transducer frequency. This allows a larger area to be
insonified by both ultrasonic signals than is otherwise
feasible.
[0180] In one embodiment, frequencies of 500 kHz and 1 MHz are
chosen as the two output frequencies for the dual frequency
transducer 290. When the disc 292 comprises lead zirconium titanate
(PZT), the diameter D is about 4 mm and the thickness TX is about 2
mm. The resulting dual frequency transducer 290 is small enough to
be incorporated in an implantable device and yet also large enough
to insonify a significant portion of the lumen of many blood
vessels or stents.
[0181] In an alternative embodiment, a rectangular slab may be
substituted for the disc 292. In one embodiment, a lateral mode may
then be used instead of the radial mode associated with the disc
292 to provide the resonance at the fundamental frequency, with the
thickness mode providing the resonance at the second harmonic.
Conventional mode charts are used to select the ratios of the
relevant dimensions.
[0182] Coating a cylindrical sidewall of the disc 292 and one of
the electrodes 294 and 296 with an acoustic isolator 300 (analogous
to the acoustic backing 194 of FIG. 17) allows the other of the
electrodes 294 and 296 to serve as an acoustic radiator. Choosing
the acoustic isolator 300 to have a low relative dielectric
constant reduces capacitive loading of the dual frequency
transducer 290 by the patient's body, which, as noted above, has a
high relative dielectric constant (approaching 80) and which also
includes conductive solutions. Coating the acoustic isolator 300
with a grounded conductor 302, selecting the electrode 296 to be a
grounded electrode and selecting the electrode 294 to be a driven
electrode reduces unwanted radiation of electromagnetic signals
from the transducer 292. A thin biocompatible coating 304
(analogous to the outer coating 192 of FIG. 17) protects the dual
frequency transducer 290 from exposure to biological matter without
preventing radiation of ultrasound from the surface bearing the
electrode 296.
[0183] Other types of localized therapy include coupling a
thermally-activated medication to carrier molecules that have
affinity to tumor tissue. Localized heating of the tumor tissue
enables selective activation of the medication in the tumor tissue,
as described in U.S. Pat. No. 5,490,840 entitled Targeted Thermal
Release Of Drug-Polymer Conjugates. Localized heating may be
effected through ultrasound via an ultrasonic transducer, e.g.,
transducers 4446 (FIGS. 1 through 6) implanted to allow
insonification of the affected area. Higher acoustic frequencies
provide shorter penetration depths, i.e., provide greater control
over where the ultrasound and therefore the resultant heat is
delivered. Additionally, heating is increased by ultrasonic
cavitation in the presence of microbubbles, microspheres or
microballoons. Other methods for providing localized magnetic
forces or heating include electromagnetic or resistive heating
transducers 44-46 comprising coils.
[0184] FIG. 25 illustrates a coil 312 integrated into a stent 310.
The coil 312 comprises saddle-shaped wires 313 integrated into the
stent 310. The coil 312 may be an electromagnetic transducer used
to magnetically capture delivery vehicles bearing drugs. Leads 314
couple the coil 312 to an implantable control IC 315, which may
comprise the implantable electronic circuits of any of FIGS. 1
through 6. The implantable electronic circuits of FIGS. 4 through 6
may provide advantages in this situation because the frequency of
the signal providing power to the implantable electronic circuits
may be different from the frequency of the signals to the
transducers 44-46, such as the coil 312. This may avoid a situation
where the signals providing power to the implantable electronic
circuits also result in release of drugs in the vicinity of the RF
coupling coil 30 that is receiving the electrical power.
[0185] When a suitable current, either AC or DC, is supplied via
the leads 314, a magnetic field represented by flux lines 316 is
generated. The magnetic field captures magnetic delivery vehicles
that have been introduced into the patient's bloodstream. The
increased concentration of delivery vehicles in the target vicinity
can be used to provide local increases in delivery of drugs
contained in the delivery vehicles.
[0186] Microbubbles including medication may be localized via a
magnetic field and ruptured via an oscillating magnetic field as
described in U.S. Pat. No. 4,652,257 entitled
Magnetically-Localizable. Polymerized Lipid Vesicles And Method Of
Disrupting Same. Suitable magnetic fields may be provided via
application of RF or RF and DC electrical energy to the coil 312.
In these embodiments, one or more of the transducers 44-46 of FIGS.
1 through 6 comprise the coil structure 312. In response to signals
coupled to the implantable electronic circuit, the transducer 44-46
that is selected is activated and is supplied with current to
either trap the magnetic delivery vehicles so that they can be
ruptured via signals provided from another selected transducer
44-46 (e.g., an ultrasonic transducer that ruptures microbubbles,
microspheres or microballoons via cavitation), or an oscillating
magnetic field may be superposed on the magnetic fields generated
by the coil 3 12 used to trap the delivery vehicles.
[0187] Referring again to FIG. 25, in another embodiment, a
permanent magnet 311 may be included on or in the stent 310 to
provide a static magnetic field for localization of magnetic
delivery vehicles. An oscillating magnetic field may then be
provided via signals supplied to the coil 312 to rupture the
delivery vehicles under the control of the implantable electronic
circuit of any of FIGS. 1 through 6, where the coil 312 acts as one
of the transducers 44-46. These embodiments may reduce power
requirements for the implantable control IC 315 while retaining
external control over when the drug or drug precursor is released
via signals from the power supply and patient monitoring console
101 of FIG. 12. Other types of coils, e.g., analogous to the RF
coupling coils 30B, 30C or 30D of FIGS. 7 through 10, or 121 of
FIGS. 11A and 11B, may also be used instead of the RF coupling coil
312.
[0188] FIG. 26 illustrates another embodiment of a coil 312A
integrated into a stent 310A. The coil 312A is analogous to the
coil 312 of FIG. 25, but is shaped as a cylindrical coil rather
than as a saddle-shaped spiral. Leads 314A couple wires 313A
comprising the coil 312A to an implantable control IC 315A, which
is analogous to the implantable control IC 315 of FIG. 25. When a
suitable current, either AC or DC, is supplied via the leads 314A,
a magnetic field represented by flux lines 316A is generated. The
coil 312A may be used to capture magnetic delivery vehicles that
have been introduced into the patient's bloodstream.
[0189] In other embodiments, the coils 312 or 312A may form
resistive heating transducers comprising a resistive material and
may, if desired, be wound with bifilar wire to prevent them from
acting as electromagnets or RF coupling coils. In another
embodiment, the coils 312 or 312A may be heated directly by
magnetic fields inducing current in the coils 312 or 312A, or, the
body of the stent 310A may form a resistive heating transducer that
is heated via magnetically-induced currents.
[0190] Stents are typically fashioned from metals that are
biocompatible, such as titanium alloys (e.g., Nitinol, a nickel
titanium alloy), stainless steel (e.g., 316L), platinum/iridium
alloys or tantalum. All of these materials are suitable for
fashioning a stent that is to be directly heated by RF-induced eddy
currents (such stents would not include slots such as slot 118,
FIG. 11A, or insulating couplings such as 119, FIG. 11A, or 146,
FIG. 11B), however, titanium and Nitinol have the highest
electrical resistivity, while platinum/iridium and tantalum have
the lowest electrical resistivity. When a stent body is to be
directly heated by induced eddy currents, titanium or Nitinol may
present advantages.
[0191] When a RF coupling coil is to be fashioned from these
materials, those applications with higher power requirements may
favor the materials with the lower resistivities.
[0192] When a current is passed through coils analogous to RF
coupling coils 312 or 312A but comprising resistive material, or
through a stent body as eddy currents, a local temperature rise is
produced. This local temperature rise may be employed to rupture
microbubbles having a melting point slightly above normal human
body temperatures. One system using microbubbles having a
controlled melting point to facilitate rupture of the microbubbles
at predetermined localized areas within a patient's body is
described, for example, in U.S. Pat. No. 4,558,690 entitled Method
Of Administration Of Chemotherapy To Tumors. The localized heating
may be provided by a structure similar to the cylindrical RF
coupling coil 30A of FIG. 7, the woven mesh coils 30B and 30C of
FIGS. 8 and 9, the saddle RF coupling coil 30D of FIG. 10, the RF
coupling coil 121 of FIGS. 11A and 11B, the coil 312 of FIG. 25 or
the coil 312A of FIG. 26, with the conductors of the coils
comprising a suitably resistive material such as nichrome wire. The
heating may be supplied directly by RF excitation of the coils 30A
through 30D or 121, or it may be effected via the implantable
electronic circuits of FIGS. 1 through 6. This may be in response
to signals from the power supply and patient monitoring console 101
of FIG. 12. Additionally, delivery vehicles such as microbubbles,
microspheres or microballoons can increase localized heating of
tissue via rupture of the delivery vehicles caused by localized
application of ultrasound, as discussed, for example, in Technical
Report: Drug And Gene Delivery, Jul. 2, 1997, ImaRx Pharmaceutical
Corp.
[0193] Transducers may be employed to facilitate drug penetration
through the wall of a stent or stent graft and into the surrounding
vasculature via sonophoresis, i.e., ultrasound enhancement of drug
penetration into body tissues, or via iontophoresis, i.e.,
electrical field enhancement of drug penetration into body tissues,
when suitable transducers are included in the stent or stent
graft.
[0194] Methods and apparatus for localized drug delivery via
sonophoresis or phonophoresis are described in U.S. Pat. No.
4,484,569 entitled Ultrasonic Diagnostic And Therapeutic Transducer
Assembly And Method For Using, U.S. Pat. No. 5,016,615 entitled
Local Application Of Medication With Ultrasound and U.S. Pat. No.
5,267,985 entitled Drug Delivery By Multiple Frequency
Phonophoresis. These patents generally discuss transdermal delivery
of medication to an affected area and note that use of more than
one frequency of ultrasonic energy is beneficial in some
situations.
[0195] An iontophoretic catheter for drug delivery is described in
Iontophoretic Drug Delivery System, by R. G. Welsh et al., Sernin.
Intervent. Cardiol., No. 1, pp. 40-42 (1996). The system uses a
microporous membrane enclosing a drug solution and a drug delivery
electrode. A reference electrode is coupled to the biological
tissue at a site that is separate from the drug delivery electrode.
The reference and drug delivery electrodes are coupled to a power
supply that provides an electrical potential between the two
electrodes. Cationic drugs move from the anode towards the cathode,
while anionic drugs move from the cathode towards the anode, with
the rate being generally proportional to the current. Control over
localized drug delivery is effected via control of the current and
the duration of the current from the drug delivery electrode. One
application is for delivery of antirestenotic agents.
[0196] Other uses of iontophoresis are described in U.S. Pat. No.
4,383,529 entitled Iontophoretic Electrode Device, Method and Gel
Insert and U.S. Pat. No. 4,416,274 entitled Ion Mobility Limiting
Iontophoretic Bioelectrode. These generally describe iontophoretic
apparatus for localized transdermal drug delivery. Catheters
adapted to provide localized iontophoretic drug delivery are
described in U.S. Pat. No. 4,411,648 entitled Iontophoretic
Catheter Device, and U.S. Pat. No. 5,499,971 entitled Method for
Iontophoretically Delivering Drug Adjacent To A Heart. These
discuss specific problems that are most readily addressed via
localized drug delivery, including treatment of vascular regions to
reduce restenosis following PTCA, drug delivery to tumor sites and
techniques for iontophoretically delivering drugs in the vicinity
of the heart without inducing arrhythmia due to electrical
stimulation of heart muscles and nerves. In one embodiment, this is
effected together with provision of electrical fields effective in
providing drug transport by chopping a DC potential difference at a
rate of between 5 and 15 kHz or by providing an asymmetric AC
waveform that is in this frequency range. These techniques are
necessary because the current being used for iontophoresis travels
through a significant and somewhat unpredictable amount of body
tissue that may well include muscles and nerves associated with the
heart.
[0197] These concepts become more powerful when combined with the
implantable transducers 44-46 of FIGS. 4 through 6 for providing
the energy to locally deliver or locally activate the medications.
An example of an iontophoretic transducer is described in
conjunction with FIG. 27 below.
[0198] FIG. 27 illustrates an embodiment of an iontophoretic system
320 for local drug delivery in the vicinity of an implanted stent
322. The iontophoretic system 320 includes an implantable control
IC 324, which may be coupled to the stent 322. The implantable
control IC 324 is coupled via wires 326 to a first electrode 328
and to a second electrode 330. The first 328 and second 330
electrodes are insulated from the stent 322 when the stent 322
comprises conductive material, unless the stent 322 comprises one
of the electrodes 328 and 330. The first 328 and second 330
electrodes comprising an iontophoretic transducer may be disposed
on the exterior of the stent 322 (as illustrated), on the interior
of the stent 322, or may be disposed such that one is inside the
stent 322 and the other is external to the stent 322.
[0199] A potential difference is established between the first 328
and second 330 electrodes by the implantable control IC 324 in
response to signals coupled from outside the patient's body, via a
RF coupling coil (not illustrated) as discussed above. The
potential difference causes some types of drugs to migrate from one
of the electrodes 328 and 330 towards the other, according to the
polarity of the potential difference and the specific nature of the
drug. This effect may be used to provide localized drug therapy,
for example, to the wall of the vessel (not illustrated) into which
the stent 322 is implanted. For example, systemically-administered
drugs may be selectively transported from the blood into the
vasculature surrounding a stent 322 to provide increased local
concentrations of antistenotic agents.
[0200] One advantage of this technique is that the currents
produced by the iontophoretic system 320 are extremely localized,
i.e., are substantially confined to the area between the electrodes
328 and 330 and immediately surrounding tissues. This obviates some
of the problems that have been encountered with iontophoretic
systems that use a reference electrode that is placed at a body
location remote from the drug delivery electrode, e.g., a
catheterized drug delivery electrode used in conjunction with an
externally applied reference electrode. Accordingly, the
iontophoretic system 320 may employ a DC voltage to effect
iontophoretic drug delivery to parts of the body that cannot safely
be treated via a catheterized system using DC for iontophoretic
drug delivery. This is advantageous in improving the efficiency of
drug delivery and in reducing exposure of other portions of the
body to the electrical currents being employed for iontophoresis.
One area where this may provide advantages, depending on stent
placement and other factors, is in treating restenosis of cardiac
blood vessels following stent insertion as a part of a PTCA
treatment. A stent 322 intended for this purpose may also include
sensors providing signals indicative of blood flow through the
stent and therefore capable of providing data indicative of
blockage as it develops. Additionally, the stent 322 including
iontophoretic electrodes 328 and 330 may also be used to enhance
localized delivery of drugs that are activated via therapeutic
transducers coupled to the stent 322 or that are included in the
vasculature upstream of the stent 322.
[0201] Another method for localized drug activation uses light
supplied by an optical transducer, where the light is of the
appropriate wavelength and intensity to break precursor molecules
down into drugs. U.S. Pat. No. 5,445,608 entitled Method And
Apparatus For Providing Light-Activated Therapy, describes a
photodynamic therapy achieved by photoactivation of suitable
optically active drugs. As described in this patent, the drugs are
activated via catheterized light emitters inserted at the site to
be treated and providing light at the wavelength required in order
to activate the drugs and at the location where the activated drugs
are needed for therapeutic purposes. Examples of precursor
substances that can be optically activated by being broken down
into drug molecules include long-chain cyanine dyes, dimers of
phthalocyanine dyes and porphyrin compounds. A wide selection of
solid state light sources including laser diodes and light emitting
diodes is commercially available from a variety of vendors,
including Motorola of Phoenix, Ariz. Laser diodes or light emitting
diodes may be employed as transducers 44-46 in any of the systems
shown in FIGS. 1 through 6 to provide light for photoactivation of
drugs within a patient's body via signals from the implantable
electronic circuit in response to signals transmitted from the
power supply and patient monitoring console 101 of FIG. 12.
[0202] FIG. 28 illustrates an embodiment wherein light emitting or
optical transducers 338 are coupled to a stent 336. The light
emitting transducers 338 may comprise light emitting diodes having
an appropriate wavelength or may comprise diode lasers. The light
emitting transducers 338 may be coupled in series via lines 340, as
shown in FIG. 28; or may be coupled in parallel. When the light
emitting transducers 338 are coupled in series, one disadvantage is
that catastrophic failure of one of the light emitting transducers
338 that causes the failed light emitting transducer 338 to fail to
pass enough current for light emission may also prevent the
remainder of the light emitting transducers 338 from operating.
[0203] The light emitting transducers 338 are coupled via lines 342
to an implantable control IC 344, which is in turn coupled to a RF
coupling coil (not illustrated in FIG. 28) that provides energy and
control signals. The light emitting transducers 338 may be disposed
on the outside of the stent 336, as shown, or on the inside of the
stent 336 or both as required for a given application.
[0204] The transducers 44-46 of FIGS. 1 through 6 may concentrate
or activate medications by supplying heat, via resistive processes
or insonification, or may employ light, magnetic fields or
electrical fields for localized drug delivery or activation. The
ultrasonic transducer 290 of FIG. 24 is, among other ultrasonic
transducers, also suited to increasing drug penetration of drugs
via sonophoresis into, e.g., tumors or vascular walls via an
implantable electronic circuit such as any of those shown in FIGS.
4 through 6.
[0205] An example of an application for the systems described above
occurs in the situation where a stent is implanted to correct a
stenosis or to repair an aneurysm in a blood vessel. Over time,
tissue ingrowth at the ends of the stent can lead to stenosis,
which can lead to thrombus formation. Thrombosis threatens the
viability of the stent, and may require aggressive intervention
using surgery or drugs. It is very undesirable to have to
surgically resolve this situation if there is a viable alternative
approach for relieving the blockage. One approach is to infuse the
patient with thrombolytic drugs. This may lead to hemorrhagic
consequences in other parts of the body, especially if the patient
has, for example, recently had surgery. One approach to reducing
the amount of thrombolytic drugs required to resolve thromboses in
vitro is described in Prototype Therapeutic Ultrasound Emitting
Catheter For Accelerating Thrombolysis. J. Ultrasound Med. 16, pp.
529-535 (1997). In this study, urokinase alone as a fibrinolytic
agent was compared to urokinase in the presence of ultrasonic
energy, with the latter showing marked improvement in the degree of
fibrinolysis of artificial blood clots in glass tubes.
[0206] When, however, the stent includes a transducer, such as an
ultrasonic transducer, coupled to the implantable electronic
circuit of any of FIGS. 1 through 6, the introduction of a
thrombolytic drug into the bloodstream of the patient can be
followed by generation of ultrasound within the stent via the
transducer and under the control of an attending physician. This
allows the thrombolytic drug, e.g., urokinase, streptokinase or
tissue plasminogen activator, to be activated at the site of the
thrombus and under the control of the attending physician, reducing
the probability of hemorrhagic consequences at portions of the
patient's body remote from the site being treated. It also enables
rapid onset of treatment, which can be critical in some situations,
e.g., in the event of heart attack or stroke induced via
thrombolysis, and may obviate invasive surgery in the event that
the therapeutic transducer has already been implanted in a prior
procedure.
[0207] Additionally, when flow or pressure sensors such as are
described with respect to FIGS. 13 through 16 or 18 are also
included with the stent when the stent is implanted and these are
also coupled to the implantable electronic circuits of any of FIGS.
2 through 6, the attending physician may be able to obtain
information that is indicative of graft condition. This can allow
the physician to more readily determine if the condition is
treatable without resorting to invasive evaluation and
intervention. Monitoring during non-invasive treatment, e.g., local
drug activation, accomplished through use of an implanted blood
velocity or blood pressure transducer, may allow assessment of the
progress of thrombolysis that may, in turn, permit successful
noninvasive treatment without incurring undue risk to the
patient.
[0208] Further, when stents are implanted to relieve stenosis,
restenosis due to tissue ingrowth tends to occur within the first 6
months following angioplasty, with the greatest loss of luminal
diameter occurring between the first and third month. Detection of
tissue growth can be determined via pressure sensors as described
above or via incorporation of the dielectric sensing filaments 234
and the implantable IC sensor 220B of FIGS. 21A and 21B. When the
ultrasonic transducers, such as those of any of FIG. 13-18, 23 or
24, are included in the upstream side of an implanted stent,
precursor drugs activated sonodynamically may locally provide
antistenotic agents such as colchicine, heparin, methotrexate,
angiopeptin or hirudin to relieve or reduce restenosis without
requiring systemic administration of the drugs. Alternatively,
delivery vehicles ruptured via ultrasound may provide localized
delivery of antistenotic agents. This provides a way of controlling
restenosis on an as-needed basis as determined via the benefit of
diagnostic data, under the control of a physician, and without
requiring anesthesia or surgery. An advantage associated with at
least some of the therapeutic transducers described herein is that
they are not necessarily specific to one drug or condition. For
example, ultrasonically activated therapy provides advantages in
treatment of both restenosis and thromboses, either of which may
threaten viability of an implanted stent.
[0209] From the foregoing it will be appreciated that, although
specific embodiments of the invention have been described herein
for purposes of illustration, various modifications may be made
without deviating from the spirit and scope of the invention.
Accordingly, the invention is to be understood broadly and is not
limited except as by the appended claims.
* * * * *