U.S. patent application number 10/795743 was filed with the patent office on 2005-05-26 for device and method for preventing magnetic-resonance imaging induced damage.
This patent application is currently assigned to Biophan Technologies, Inc.. Invention is credited to Connelly, Patrick R., Helfer, Jeffrey L., Miller, Victor R., Weiner, Michael L..
Application Number | 20050113676 10/795743 |
Document ID | / |
Family ID | 33097031 |
Filed Date | 2005-05-26 |
United States Patent
Application |
20050113676 |
Kind Code |
A1 |
Weiner, Michael L. ; et
al. |
May 26, 2005 |
Device and method for preventing magnetic-resonance imaging induced
damage
Abstract
An electromagnetic shield has a first patterned or apertured
layer having non-conductive materials and conductive material and a
second patterned or apertured layer having non-conductive materials
and conductive material. The conductive material may be a metal, a
carbon composite, or a polymer composite. The non-conductive
materials in the first patterned or apertured layer may be randomly
located or located in a predetermined segmented pattern such that
the non-conductive materials in the first patterned or apertured
layer are located in a predetermined segmented pattern with respect
to locations of the non-conductive materials in the second
patterned or apertured layer.
Inventors: |
Weiner, Michael L.;
(Webster, NY) ; Miller, Victor R.; (Clarence,
NY) ; Connelly, Patrick R.; (Rochester, NY) ;
Helfer, Jeffrey L.; (Webster, NY) |
Correspondence
Address: |
BASCH & NICKERSON LLP
1777 PENFIELD ROAD
PENFIELD
NY
14526
US
|
Assignee: |
Biophan Technologies, Inc.
West Henrietta
NY
|
Family ID: |
33097031 |
Appl. No.: |
10/795743 |
Filed: |
March 8, 2004 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
10795743 |
Mar 8, 2004 |
|
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10405154 |
Apr 2, 2003 |
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Current U.S.
Class: |
600/421 ;
324/322; 600/374; 607/119 |
Current CPC
Class: |
G01R 33/288 20130101;
A61N 1/086 20170801; Y10T 428/24322 20150115; G01R 33/285 20130101;
G01R 33/4804 20130101; Y10T 428/24331 20150115 |
Class at
Publication: |
600/421 ;
324/322; 607/119; 600/374 |
International
Class: |
A61B 005/05 |
Claims
What is claimed is:
1. A device for amplifying an electrical signal of physiological
significance in a magnetic-resonance imaging environment,
comprising: at least two electrodes with associated input leads for
coupling to a patient; an amplifier having a zero-signal reference
terminal for detecting and amplifying the desired physiological
signal; and a filter, connected to said input leads and coupling
said at least two electrodes to said amplifier, to attenuate any
induced RF signal produced in the magnetic-resonance imaging
environment and passing the lower frequency desired electrical
physiological signal; said filter including a shield enclosing said
filter; said shield enclosing said filter being patterned with
non-conductive materials and conductive materials.
2. The device as claimed in claim 1, wherein said filter further
includes a low pass filter connected to one of said input
leads.
3. The device as claimed in claim 1, wherein said conductive
material is a metal.
4. The device as claimed in claim 1, wherein said conductive
material includes metal-coated carbon filaments.
5. The device as claimed in claim 1, wherein said conductive
material is a polymer composite.
6. The device as claimed in claim 1, wherein said shield is
constructed of two patterned sheaths, each patterned sheath having
non-conductive materials and conductive materials.
7. The device as claimed in claim 6, wherein said conductive
materials in one of said two patterned sheaths are randomly
located.
8. The device as claimed in claim 6, wherein said conductive
materials in said two patterned sheaths are located in a
predetermined segmented pattern.
9. The device as claimed in claim 1, further comprising: an input
lead shield to shield said input leads.
10. The device as claimed in claim 9, wherein said input lead
shield is patterned with non-conductive materials and conductive
materials.
11. The device as claimed in claim 10, wherein said conductive
material of said input lead shield is a metal.
12. The device as claimed in claim 10, wherein said conductive
material of said input lead shield includes metal-coated carbon
filaments.
13. The device as claimed in claim 10, wherein said conductive
material of said input lead shield is a polymer composite.
14. The device as claimed in claim 10, wherein said input lead
shield is constructed of two patterned sheaths, each patterned
sheath having non-conductive materials and conductive
materials.
15. The device as claimed in claim 14, wherein said conductive
materials in one of said two patterned sheaths are randomly
located.
16. The device as claimed in claim 14, wherein said conductive
materials in said two patterned sheaths are located in a
predetermined segmented pattern.
17. The device as claimed in claim 14, wherein said conductive
materials in said two patterned sheaths are located in a
predetermined segmented pattern with respect to a direction
substantially parallel with an axis of said input leads.
18. The device as claimed in claim 14, wherein said conductive
materials in said two patterned sheaths are located in a
predetermined segmented pattern with respect to a direction
substantially perpendicular with an axis of said input leads.
19. The device as claimed in claim 17, wherein said conductive
materials in said two patterned sheaths are located in a
predetermined segmented pattern with respect to a direction
substantially perpendicular with an axis of said input leads.
20. A device for amplifying an electrical signal of physiological
significance in a magnetic-resonance imaging environment,
comprising: at least two electrodes with associated input leads for
coupling to a patient; an amplifier having a zero-signal reference
terminal for detecting and amplifying the desired physiological
signal; and a filter, connected to said input leads and coupling
said at least two electrodes to said amplifier, to attenuate any
induced RF signal produced in the magnetic-resonance imaging
environment and passing the lower frequency desired electrical
physiological signal, said filter including a shield enclosing said
filter; said shield being an apertured conductive material having a
maximum aperture dimension of 0.01 millimeters to 10
millimeters.
21. The device as claimed in claim 20, wherein said filter further
includes a low pass filter connected to one of said input
leads.
22. The device as claimed in claim 20, wherein said conductive
material is a metal.
23. The device as claimed in claim 20, wherein said conductive
material includes metal-coated carbon filaments.
24. The device as claimed in claim 20, wherein said conductive
material is a polymer composite.
25. The device as claimed in claim 20, wherein said conductive
material comprises nanotubes.
Description
CROSS-REFERENCE TO RELATED US PATENT APPLICATIONS
[0001] This application is a divisional of co-pending U.S. patent
application Ser. No. 10/405,154, filed on Apr. 2, 2003. The entire
contents of U.S. patent application Ser. No. 10/405,154 are hereby
incorporated by reference.
CROSS REFERENCE TO RELATED PATENT APPLICATIONS
[0002] The subject matter of the present application is related to
the subject matter of co-pending U.S. patent application Ser. No.
09/885,867, filed on Jun. 20, 2001, entitled "Controllable,
Wearable magnetic-resonance imaging-Compatible Cardiac Pacemaker
With Pulse Carrying Photonic Catheter And VOO Functionality";
co-pending U.S. patent application Ser. No. 09/885,868, filed on
Jun. 20, 2001, entitled "Controllable, Wearable magnetic-resonance
imaging-Compatible Cardiac Pacemaker With Power Carrying Photonic
Catheter And VOO Functionality"; co-pending U.S. patent application
Ser. No. 10/037,513, filed on Jan. 4, 2002, entitled "Optical Pulse
Generator For Battery Powered Photonic Pacemakers And Other Light
Driven Medical Stimulation Equipment"; co-pending U.S. patent
application Ser. No. 10/037,720, filed on Jan. 4, 2002, entitled
"Opto-Electric Coupling Device For Photonic Pacemakers And Other
Opto-Electric Medical Stimulation Equipment"; co-pending U.S.
patent application Ser. No. 09/943,216, filed on Aug. 30, 2001,
entitled "Pulse Width Cardiac Pacing Apparatus"; co-pending U.S.
patent application Ser. No. 09/964,095, filed on Sep. 26, 2001,
entitled "Process for Converting Light"; co-pending U.S. patent
application Ser. No. 09/921,066, filed on Aug. 2, 2001, entitled
"magnetic-resonance imaging-Resistant Implantable Device";
co-pending U.S. patent application Ser. No. 10/077,842, filed on
Feb. 19, 2002, entitled "An Electromagnetic Interference Immune
Tissue Invasive System"; co-pending U.S. patent application Ser.
No. 10/077,823, filed on Feb. 19, 2002, entitled "An
Electromagnetic Interference Immune Tissue Invasive System";
co-pending U.S. patent application Ser. No. 10/077,887, filed on
Feb. 19, 2002, entitled "An Electromagnetic Interference Immune
-Tissue Invasive System"; co-pending U.S. patent application Ser.
No. 10/077,883, filed on Feb. 19, 2002, entitled "An
Electromagnetic Interference Immune Tissue Invasive System"; and
co-pending U.S. patent application Ser. No. 10/077,958, filed on
Feb. 19, 2002, entitled "An Electromagnetic Interference Immune
Tissue Invasive System".
[0003] The entire content of each of the above noted co-pending
U.S. patent applications (Ser. Nos. 09/885,867; 09/885,868;
10/037,513; 10/037,720; 09/943,216; 09/964,095; 09/921,066;
10/077,842; 10/077,823; 10/077,887; 10/077,883; and 10/077,958) is
hereby incorporated by reference.
FIELD OF THE PRESENT INVENTION
[0004] The present invention relates generally to a device and
method for preventing magnetic-resonance imaging induced damage.
More particularly, the present invention is directed to medical
assist systems, which may include leads and other implantable or
non-implantable components, that are shilded by segmented shielding
to hardened or immune the systems from electromagnetic interference
or insult, namely electromagnetic interference or insult in a
magnetic-resonance imaging environment and to a modifiable
magnetic-resonance imaging, which is, automatically or manually,
responsive to sensed tissue temperature changes or known localized
specific energy absorption ratios.
BACKGROUND OF THE PRESENT INVENTION
[0005] Magnetic-resonance imaging ("magnetic-resonance imaging")
has been developed as an imaging technique adapted to obtain both
images of anatomical features of human patients as well as some
aspects of the functional activities of biological tissue. These
images have medical diagnostic value in determining the state of
the health of the tissue examined.
[0006] In a magnetic-resonance imaging process, a patient is
typically aligned to place the portion of the patient's anatomy to
be examined in the imaging volume of the magnetic-resonance imaging
apparatus. Such an magnetic-resonance imaging apparatus typically
comprises a primary magnet for supplying a constant magnetic field
(B.sub.0) which, by convention, is along the z-axis and is
substantially homogeneous over the imaging volume and secondary
magnets that can provide linear magnetic field gradients along each
of three principal Cartesian axes in space (generally x, y, and z,
or x.sub.1, x.sub.2 and x.sub.3, respectively). The apparatus also
comprises one or more radio-frequency coils that provide excitation
and detection of the magnetic-resonance imaging signal.
[0007] The use of the magnetic-resonance imaging process with
patients who have implanted or non-implanted medical assist
devices; such as, but not limited to, cardiac assist devices,
implanted insulin pumps, catheter guide wires, leads for
neurostimulation probes, intraluminal coils, guided catheters,
temporary cardiac pacemakers, temporary esophageal pacemakers;
often presents problems.
[0008] For a specific example, as is known to those skilled in the
art, implantable devices (such as implantable pulse generators
(fPGs) and cardioverter/defibrillator/pacemakers (CDPs)) are
sensitive to a variety of forms of electromagnetic interference
(EMI) because these enumerated devices include sensing and logic
systems that respond to low-level electrical signals emanating from
the monitored tissue region of the patient and that these devices
may also have metal wire leads, which can act as antenna and
provide a path for the induced energy to travel to and possibly
damage power sensitive circuitry.
[0009] Since the sensing systems and conductive elements of these
medical assist devices are responsive to changes in local
electromagnetic fields, the medical assist devices are vulnerable
to external sources of severe electromagnetic noise, and in
particular, to electromagnetic fields emitted during the
magnetic-resonance imaging (magnetic-resonance imaging) procedure.
Thus, patients with medical assist devices are generally advised
not to undergo magnetic-resonance imaging (magnetic-resonance
imaging) procedures.
[0010] To more appreciate the problem using a specific
illustration, namely the use of implantable cardiac assist devices
during a magnetic-resonance imaging process, will be briefly
discussed.
[0011] The human heart may suffer from two classes of rhythmic
disorders or arrhythmias: bradycardia and tachyarrhythmia.
Bradycardia occurs when the heart beats too slowly, and may be
treated by a common implantable pacemaker delivering low voltage
(about 3 V) pacing pulses.
[0012] The common implantable pacemaker is usually contained within
a hermetically sealed enclosure, in order to protect the
operational components of the device from the aqueous environment
of the body, as well as to protect the body from the device.
[0013] The common implantable pacemaker operates in conjunction
with one or more electrically conductive leads, adapted to conduct
electrical stimulating pulses to sites within the patient's heart
and to communicate sensed signals from those sites back to the
implanted device.
[0014] Furthermore, the common implantable pacemaker typically has
a metal case and a connector block mounted to the metal case that
includes receptacles for leads which may be used for electrical
stimulation or which may be used for sensing of physiological
signals. The battery and the circuitry associated with the common
implantable pacemaker are hermetically sealed within the case.
Electrical interfaces are employed to connect the leads outside the
metal case with the medical device circuitry and the battery inside
the metal case.
[0015] Electrical interfaces serve the purpose of providing an
electrical circuit path extending from the interior of a
hermetically sealed metal case to an external point outside the
case while maintaining the hermetic seal of the case. A conductive
path is provided through the interface by a conductive pin that is
electrically insulated from the case itself.
[0016] Such interfaces typically include a ferrule that permits
attachment of the interface to the case, the conductive pin, and a
hermetic glass or ceramic seal that supports the pin within the
ferrule and isolates the pin from the metal case.
[0017] A common implantable pacemaker can, under some
circumstances, be susceptible to electrical interference such that
the desired functionality of the pacemaker is impaired. For
example, common implantable pacemaker requires protection against
electrical interference from electromagnetic interference (EMI),
defibrillation pulses, electrostatic discharge, or other generally
large voltages or currents generated by other devices external to
the medical device. As noted above, more recently, it has become
crucial that cardiac assist systems be protected from
magnetic-resonance imaging sources.
[0018] Such electrical interference can damage the circuitry of the
cardiac assist systems or cause interference in the proper
operation or functionality of the cardiac assist systems. For
example, damage may occur due to high voltages or excessive
currents introduced into the cardiac assist system by voltages or
currents induced in the cardiac assist system circuitry or on the
wire leads leading to and from the cardiac assist system
circuitry.
[0019] Therefore, it is required that such voltages and currents be
limited at the input of such cardiac assist systems, e.g., at the
interface. Protection from such voltages and currents has typically
been provided at the input of a cardiac assist system by the use of
one or more zener diodes and one or more filter capacitors.
[0020] For example, one or more zener diodes may be connected
between the circuitry to be protected, e.g., pacemaker circuitry,
and the metal case of the medical device in a manner which grounds
voltage surges and current surges through the diode(s). Such zener
diodes and capacitors used for such applications may be in the form
of discrete components mounted relative to circuitry at the input
of a connector block where various leads are connected to the
implantable medical device, e.g., at the interfaces for such
leads.
[0021] However, such protection, provided by zener diodes and
capacitors placed at the input of the medical device, increases the
congestion of the medical device circuits, at least one zener diode
and one capacitor per input/output connection or interface. This is
contrary to the desire for increased miniaturization of implantable
medical devices.
[0022] Further, when such protection is provided, interconnect wire
length for connecting such protection circuitry and pins of the
interfaces to the medical device circuitry that performs desired
functions for the medical device tends to be undesirably long. The
excessive wire length may lead to signal loss and undesirable
inductive effects. The wire length can also act as an antenna that
conducts undesirable electrical interference signals to sensitive
CMOS circuits within the medical device to be protected.
[0023] Additionally, the radio frequency (radio-frequency) energy
that is inductively coupled into the wire causes intense heating
along the length of the wire, and at the electrodes that are
attached to the heart wall. This heating may be sufficient to
ablate the interior surface of the blood vessel through which the
wire lead is placed, and may be sufficient to cause scarring at the
point where the electrodes contact the heart. A further result of
this ablation and scarring is that the sensitive node that the
electrode is intended to pace with low voltage signals becomes
desensitized, so that pacing the patient's heart becomes less
reliable, and in some cases fails altogether. Additionally, the
switching of the gradient magnetic fields may also induce unwanted
voltages causing problems with the circuitry and potential pacing
of the heart.
[0024] Another conventional solution for protecting the implantable
medical device from electromagnetic interference is illustrated in
FIG. 1. FIG. 1 is a schematic view of an implantable medical device
12 embodying protection against electrical interference. At least
one lead 14 is connected to the implantable medical device 12 in
connector block region 13 using an interface.
[0025] In the case where implantable medical device 12 is a
pacemaker implanted in a body 10, the pacemaker 12 includes at
least one or both of pacing and sensing leads represented generally
as leads 14 to sense electrical signals attendant to the
depolarization and repolarization of the heart 16, and to provide
pacing pulses for causing depolarization of cardiac tissue in the
vicinity of the distal ends thereof
[0026] FIG. 2 more particularly illustrates the circuit that is
used conventionally to protect from electromagnetic interference.
As shown in FIG. 2, protection circuitry 15 is provided using a
diode array component 30. The diode array consists of five zener
diode triggered semiconductor controlled rectifiers (SCRs) with
anti-parallel diodes arranged in an array with one common
connection. This allows for a small footprint despite the large
currents that may be carried through the device during
defibrillation, e.g., 10 amps. The SCRs 20-24 turn on and limit the
voltage across the device when excessive voltage and current surges
occur.
[0027] Some of the zener diode triggered SCRs may be connected to
an electrically conductive pin, with each electrically conductive
pin being connected to a medical device contact region to be wire
bonded to pads of a printed circuit board. The diode array
component 30 of FIG. 2 may be connected to the electrically
conductive pins via die contact regions along with other electrical
conductive traces of the printed circuit board.
[0028] Other attempts have been made to protect medical assist
devices from magnetic-resonance imaging fields. For example, U.S.
Pat. No. 5,968,083 (to Ciciarelli et al.) describes a device
adapted to switch between low and high impedance modes of operation
in response to EMI insult. Furthermore, U.S. Pat. No. 6,188,926 (to
Vock) discloses a control unit for adjusting a cardiac pacing rate
of a pacing unit to an interference backup rate when heart activity
cannot be sensed due to EMI.
[0029] Another problem associated with magnetic-resonance imaging
is the temperature change in tissue regions caused by using
conventional magnetic-resonance imaging techniques. When a
substance such as human tissue is subjected to a static magnetic
field, the individual magnetic moments of the spins in the tissue
align in a parallel and anti-parallel direction with the static
magnetic field. This direction along the static magnetic field can
be termed as the longitudinal direction. In magnetic-resonance
imaging, the radio frequency polarizing field used for spin
manipulation is constantly changing and thus, the individual
magnetic moments of the spins in the tissue attempt to align with
the polarizing field. The constant changing of alignment of the
magnetic moments of the spins in the tissue causes the tissue's
temperature to increase, thereby exposing the tissue to possible
magnetic-resonance imaging induced thermal damage.
[0030] Although, conventional medical assist devices provide some
means for protection against electromagnetic interference, these
conventional medical assist devices require much circuitry and fail
to provide fail-safe protection against radiation produced by
magnetic-resonance imaging procedures. Moreover, the conventional
medical assist devices fail to address the possible damage that can
be done at the tissue interface due to radio-frequency-induced
heating. Furthermore, the conventional medical assist devices fail
to address the unwanted tissue region stimulation that may result
from radio-frequency-induced electrical currents. Lastly,
conventional magnetic-resonance imaging processes fail to provide a
proper safeguard against potential magnetic-resonance imaging
induced thermal damage due to the tissue's exposure to the
switching magnetic field gradients and the circularly polarized
Radio Frequency Field of the magnetic-resonance imaging
process.
[0031] Thus, it is desirable to provide protection against
electromagnetic interference, without requiring much circuitry and
to provide fail-safe protection against radiation produced by
magnetic-resonance imaging procedures. Moreover, it is desirable to
provide medical assist devices that prevent the possible tissue
damage that can be done at the tissue interface due to induced
electrical signals. Furthermore, it is desirable to provide an
effective means for transferring energy from one point in the body
to another point without having the energy causing a detrimental
effect upon the body. Lastly, it is desirable to implement a
magnetic-resonance imaging process, which can be modified,
automatically or manually, in response to sensed tissue temperature
changes or known localized specific energy absorption rates, so as
to prevent possible magnetic-resonance imaging induced thermal
damage.
SUMMARY OF THE PRESENT INVENTION
[0032] A first aspect of the present invention is directed to a
medical assist system. The medical assist system includes a primary
device housing, the primary device housing having a control circuit
therein; a lead system to provide an electrical path between a
tissue region and the primary device housing; and a shielding
formed around the lead system to shield the lead system from
electromagnetic interference. The shielding is patterned with
non-conductive materials and conductive material.
[0033] A second aspect of the present invention is directed to a
medical assist device. The medical assist device includes a primary
device housing, the primary device housing having a control circuit
therein, and a shielding formed around the primary device housing
to shield the primary device housing and any circuits therein from
electromagnetic interference. The shielding is patterned with
non-conductive materials and conductive material.
[0034] A third aspect of the present invention is directed to an
electromagnetic shielded implantable lead. The electromagnetic
shielded implantable lead includes an electrical lead and a
shielding formed around the electrical lead to shield the
electrical lead from electromagnetic interference. The shielding is
patterned with non-conductive materials and conductive
material.
[0035] A fourth aspect of the present invention is directed to an
electromagnetic shield. The electromagnetic shield includes a first
patterned layer having non-conductive materials and conductive
material and a second patterned layer having non-conductive
materials and conductive material.
[0036] A fifth aspect of the present invention is a
magnetic-resonance imaging process. The magnetic-resonance imaging
process images a tissue region according to an image acquisition
sequence and adjusts the image acquisition sequence, in response to
a predetermined parameter, to allow for cooling of the imaged
tissue region.
[0037] A sixth aspect of the present invention is a
magnetic-resonance imaging process. The magnetic-resonance imaging
process images a tissue region according to an image acquisition
sequence; adjusts the image acquisition sequence, in response to a
predetermined parameter, to allow for cooling of the imaged tissue
region; and shields components of a medical assist device with a
shield patterned with non-conductive materials and conductive
material.
[0038] A seventh aspect of the present invention is a device for
amplifying an electrical signal of physiological significance in a
magnetic-resonance imaging environment. The device includes at
least two electrodes with associated input leads for coupling to a
patient; an amplifier having a zero-signal reference terminal for
detecting and amplifying the desired physiological signal; and a
filter, connected to the input leads and coupling the at least two
electrodes to the amplifier, to attenuate any induced
radio-frequency signal produced in the magnetic-resonance imaging
environment and passing the lower frequency desired electrical
physiological signal. The filter includes a shield enclosing the
filter, a common reference line coupled to the shield at one point
and to the zero-signal reference terminal of the amplifier, and at
least two low pass filters, each connected to one of the input
leads and each of the at least two low pass filters referenced to
the common reference line. The shield enclosing the filter is
patterned with non-conductive materials and conductive
material.
[0039] An eighth aspect of the present invention is a medical
assist system. The medical assist device includes a primary device
housing; the primary device housing having a control circuit
therein; a lead system to provide an electrical path between a
tissue region and the primary device housing; a shielding formed
around the lead system to shield the lead system from
electromagnetic interference; and a filter, connected to the lead
system, to attenuate any induced radio-frequency signal and passing
a desired electrical physiological signal. The filter includes a
shield enclosing the filter and low pass filters connected to the
lead system. The shielding is patterned with non-conductive
materials and conductive material.
[0040] A ninth aspect of the present invention is a medical assist
system. The medical assist system includes a primary device
housing, the primary device housing having a control circuit
therein; a lead system to provide an electrical path between a
tissue region and the primary device housing; and a shielding
formed around the lead system to shield the lead system from
electromagnetic interference. The shielding is an apertured
conductive material having a maximum aperture dimension of 0.01
millimeters to 10 millimeters.
[0041] A tenth aspect of the present invention is a medical assist
device. The medical assist device includes a primary device
housing, the primary device housing having a control circuit
therein; and a shielding formed around the primary device housing
to shield the primary device housing and any circuits therein from
electromagnetic interference. The shielding is an apertured
conductive material having a maximum aperture dimension of 0.01
millimeters to 10 millimeters.
[0042] An eleventh aspect of the present invention is an
electromagnetic shielded lead. The electromagnetic shielded lead
includes an electrical lead and a shielding formed around the
electrical lead to shield the electrical lead from electromagnetic
interference. The shielding is an apertured conductive material
having a maximum aperture dimension of 0.01 millimeters to 10
millimeters.
[0043] A twelfth aspect of the present invention is an
electromagnetic shield. The electromagnetic shield includes a first
apertured conductive material having a maximum aperture dimension
of 0.01 millimeters to 10 millimeters and a second apertured
conductive material having a maximum aperture dimension of 0.01
millimeters to 10 millimeters.
[0044] A further aspect of the present invention is a
magnetic-resonance imaging process. The magnetic-resonance imaging
process images a tissue region according to an image acquisition
sequence; adjusts the image acquisition sequence, in response to a
predetermined parameter, to allow for cooling of the imaged tissue
region; and shields components of a medical assist device with a
shield being an apertured conductive material and having a maximum
aperture dimension of 0.01 millimeters to 10 millimeters.
[0045] A still further aspect of the present invention is a device
for amplifying an electrical signal of physiological significance
in a magnetic-resonance imaging environment. The device includes at
least two electrodes with associated input leads for coupling to a
patient; an amplifier having a zero-signal reference terminal for
detecting and amplifying the desired physiological signal; and a
filter, connected to the input leads and coupling the at least two
electrodes to the amplifier, to attenuate any induced RF signal
produced in the magnetic-resonance imaging environment and passing
the lower frequency desired electrical physiological signal. The
filter includes a shield enclosing the filter, a common reference
line coupled to the shield at one point and to the zero-signal
reference terminal of the amplifier, and at least two low pass
filters, each connected to one of the input leads and each of the
at least two low pass filters referenced to the common reference
line. The shield is an apertured conductive material having a
maximum aperture dimension of 0.01 millimeters to 10
millimeters.
[0046] Another aspect of the present invention is a medical assist
system. The medical assist system includes a primary device
housing, the primary device housing having a control circuit
therein; a lead system to provide an electrical path between a
tissue region and the primary device housing; a shielding formed
around the lead system to shield the lead system from
electromagnetic interference; and a filter, connected to the lead
system, to attenuate any induced RF signal and passing a desired
electrical physiological signal. The filter includes a shield
enclosing the filter and low pass filters connected to the lead
system. The shielding is an apertured conductive material having a
maximum aperture dimension of 0.01 millimeters to 10
millimeters.
BRIEF DESCRIPTION OF THE DRAWINGS
[0047] The present invention may take form in various components
and arrangements of components, and in various steps and
arrangements of steps. The drawings are only for purposes of
illustrating a preferred embodiment and are not to be construed as
limiting the present invention, wherein:
[0048] FIGS. 1 and 2 are illustrations of conventional techniques
used to protect against electromagnetic interference;
[0049] FIG. 3 is a cross-sectional view of one embodiment of
segmented shielding of a wire lead according to the concepts of the
present invention;
[0050] FIG. 4 is a cross-sectional view of another embodiment of
segmented shielding of a wire lead according to the concepts of the
present invention;
[0051] FIG. 5 is a top view of a layer of an embodiment of the
shielding, having a predetermined segmented pattern of conductive
materials, for a wire lead shielding according to the concepts of
the present invention;
[0052] FIG. 6 is a top view of a layer of another embodiment of the
shielding, having a predetermined segmented pattern of conductive
materials, for a wire lead according to the concepts of the present
invention;
[0053] FIG. 7 is a top view of a layer of another embodiment of the
shielding, having a random pattern of conductive materials, for a
wire lead according to the concepts of the present invention;
[0054] FIG. 8 is a flowchart illustrating the altering of a
magnetic-resonance imaging acquisition process in response to a
measured temperature of the tissue region being imaged, according
to the concepts of the present invention;
[0055] FIG. 9 is a flowchart illustrating the altering of a
magnetic-resonance imaging acquisition process in response to a
measured localized specific absorption ratio of the tissue region
being imaged, according to the concepts of the present
invention;
[0056] FIG. 10 is a block diagram of an ECG amplifier capable of
operating in the high static magnetic field, radio-frequency field,
and gradient field environment produced in a magnetic-resonance
imaging system, according to the concepts of the present
invention;
[0057] FIG. 11 is a block diagram of a pacemaker capable of
operating in the high static magnetic fields, radio-frequency
field, and gradient field environment produced in a
magnetic-resonance imaging system, according to the concepts of the
present invention;
[0058] FIG. 12 is a block diagram of an implantable pacemaker,
designed for external programming, which operates safely in the
environment produced in a magnetic-resonance imaging system,
according to the concepts of the present invention;
[0059] FIGS. 13 through 18 are schematic drawings of various ECG
lead and/or harness configurations designed for operation in the
environment produced by an magnetic-resonance imaging system,
according to the concepts of the present invention;
[0060] FIGS. 19 and 20 are block diagrams of ECG amplifiers with a
one-stage radio-frequency filter design, according to the concepts
of the present invention, for use with dual unshielded leads;
[0061] FIGS. 21 and 22 are block diagrams of ECG amplifiers with a
one-stage radio-frequency filter design, according to the concepts
of the present invention, for use with dual shielded leads;
[0062] FIGS. 23 and 24 are block diagrams of ECG amplifiers with a
multistage radio-frequency filter design, according to the concepts
of the present invention, for use with dual unshielded leads;
[0063] FIGS. 25 and 26 are block diagrams of ECG amplifiers with a
multistage radio-frequency filter design, according to the concepts
of the present invention, for use with dual shielded leads;
[0064] FIGS. 27 and 28 are block diagrams of ECG amplifiers with a
multistage radio-frequency filter design, according to the concepts
of the present invention, for use with a multi-lead ECG
harness;
[0065] FIG. 29 is a block diagram of a magnetic-resonance imaging
safe implantable pacemaker in accordance with the present
invention;
[0066] FIG. 30 is a block diagram of a magnetic-resonance imaging
safe implantable pacemaker using a multistage radio-frequency
filter, according to the concepts of the present invention;
[0067] FIG. 31 is a block diagram of an implantable stimulator
capable of operating in a magnetic-resonance imaging system, in
accordance with the present invention;
[0068] FIG. 32 is a cross-sectional view of one embodiment of an
apertured shielding of a wire lead according to the concepts of the
present invention;
[0069] FIG. 33 is a cross-sectional view of another embodiment of
an apertured shielding of a wire lead according to the concepts of
the present invention;
[0070] FIG. 34 is a top view of a layer of an embodiment of the
apertured shielding, having a predetermined apertured pattern of
non-conductive materials, for a wire lead shielding according to
the concepts of the present invention;
[0071] FIG. 35 is a top view of a layer of another embodiment of
the apertured shielding, having a predetermined apertured pattern
of non-conductive materials, for a wire lead according to the
concepts of the present invention; and
[0072] FIG. 36 is a top view of a layer of another embodiment of
the apertured shielding, having a random apertured pattern of
non-conductive materials, for a wire lead according to the concepts
of the present invention.
DETAILED DESCRIPTION OF THE PRESENT INVENTION
[0073] As noted above, the present invention is directed to an
implantable device, such as a medical assist device, that is immune
or hardened to electromagnetic insult or interference.
[0074] For a general understanding of the present invention,
reference is made to the drawings. In the drawings, like reference
have been used throughout to designate identical elements. In
describing the present invention, the following term(s) have been
used in the description.
[0075] For the purposes of the description below and the appended
claims, the term, medical assist device/system or tissue invasive
device/system, refers to any device/system that may enable
monitoring of living tissue(s) or living system(s) wherein the
monitoring may be, but not limited to an EKG signal, an ECG signal,
a glucose level, hormone level, cholesterol level, or magnetic
resonance image. The medical assist device/system or tissue
invasive device/system may also enable stimulus intervention to
provide assistance to living tissue(s) or living system(s) so that
the stimulus causes the selected body tissue or system to function
as desired. The stimulus may be, but not limited to, a cardiac
stimulating substance or electrical pulse, a blood thinning
substance, insulin, estrogen, progesterone, or testosterone.
[0076] The medical assist device/system or tissue invasive
device/system may also provide therapeutic treatment to the living
tissue(s) or living system(s) so that the treatment causes the
selected body tissue or system to function as desired. The stimulus
may be, but not limited to, radio frequency or laser ablation.
[0077] Furthermore, the medical assist device/system or tissue
invasive device/system may be implanted in a body cavity of a
living organism, either temporarily or permanently, or
subcutaneously implanted into a living organism either temporarily
or permanently. Moreover, the medical assist device/system or
tissue invasive device/system may be located external to the living
organism. Examples of medical assist devices/systems or tissue
invasive devices/systems are, but not limited to, wearable or
implantable cardiac pacers (such as pacemakers), implantable pulse
generators (IPGs), cardioverter/defibrillator/pacemaker- s (CDPs),
cardiac monitoring systems, insulin pump controllers, brain
monitoring systems, cardiac assist devices, implanted insulin
pumps, catheter guide wires, leads for neurostimulation probes,
intraluminal coils, guided catheters, temporary cardiac pacemakers,
temporary esophageal pacemakers, etc.
[0078] Although the thrust of the description of one embodiment of
the present invention, below, will use a cardiac assist system as
an example of the workings of the concepts of the present
invention, the concepts are directed to any medical assist
device/system, tissue invasive device/system or tissue interactive
device/system which includes electrical components that require
shielding from electromagnetic interference or insult such that the
interference or insult could cause damage to the electrical
components or damage to the surrounding tissue.
[0079] As noted above, a typical medical device may be a pacemaker.
The pacemaker may include at least one or both of pacing and
sensing leads represented generally as leads to sense electrical
signals attendant to the depolarization and repolarization of the
heart and to provide pacing pulses for causing depolarization of
cardiac tissue in the vicinity of the distal ends thereof.
[0080] The pacemaker or cardiac assist device may include a primary
device housing. The primary device housing may include a control
circuit, such as a microprocessor integrated circuit for
controlling the operations of the cardiac assist system. The
control circuit may select a mode of operation for the cardiac
assist system based on predetermined sensed parameters. The primary
device housing may also include circuitry to detect and isolate
cross talk between device pulsing operations and device sensing
operations. The control circuit may isolate physiological signals
using a noise filtering circuit or a digital noise filtering.
[0081] The control circuit can be programmable from a source
external of the primary device housing or the control circuit can
provide physiological diagnostics to a source external of the
primary device housing.
[0082] The primary device housing may include a power source. The
power source may be a battery power source in combination with a
battery power source measuring circuit. Further, the control
circuit may automatically adjust a value for determining an
elective replacement indication condition of a battery power source
such that the value is automatically adjusted by the control
circuit in response to a measured level of a state of the battery
power source, the measured level generated by the battery power
source measuring circuit connected to the battery power source.
[0083] As noted above, both the primary housing and the leads of
the medical assist device/system require protection or shielding
from electromagnetic interference or insult. The present invention
provides a shielding that substantially prevents electromagnetic
interference from damaging any, contained therein, electrical
components or damage to the tissue surrounding the medical assist
device/system's primary housing or leads. The shielding is a
compilation of patterned layers having non-conductive materials and
conductive materials contained in each patterned layer.
[0084] As is known is well known, conductive material provides a
shield or block for electromagnetic radiation, especially radio
frequency radiation ("radio-frequency") or magnetic radiation
("magnetic-resonance imaging") used in magnetic-resonance imaging
so that the magnetic-resonance imaging radiation cannot penetrate
the conductive materials to do damage to any electrical components
located on the other side of the shield.
[0085] However, although conductive material may provide a block to
the magnetic-resonance imaging radiation so as to prevent
penetration of the magnetic-resonance imaging radiation, the
blocking of the magnetic-resonance imaging radiation can cause eddy
currents to be induced in the conductive material by the changing
magnetic fields in the magnetic-resonance imaging radiation. The
eddy currents may cause heating of the conductive material, thereby
heating the enclosed electrical components so as to cause damage or
heating the surrounding tissue so as to cause tissue damage.
[0086] To prevent the eddy currents from being induced to a level,
which may cause damage from excessive heat, the conductive material
in the shielding of the present invention is patterned, apertured,
or segmented. More specifically, magnetic-resonance imaging
radiation shields generally contain thin layers of conductive
cladding sometimes separated by insulators or dielectrics. In the
present invention, the conductive material in a layer is patterned,
apertured, or segmented so that the conductive material is
literally broken up to limit the build up of eddy currents induced
by the changing magnetic fields produced during magnetic-resonance
imaging.
[0087] An example of such a patterned or segmented shielding for
use with an electrical lead is illustrated in FIG. 3. As shown in
FIG. 3, an electrical lead 60 is initially coated with an
insulation layer 62 so as to electrically insulate the electrical
lead from its surroundings. Upon the insulation layer 62, a first
patterned or segmented layer 80 of shielding is placed or formed
thereon. The first patterned or segmented layer 80 includes
conductive materials 64 and non-conductive material(s) 66.
[0088] Upon the first patterned or segmented layer 80 of shielding,
a second patterned or segmented layer 82 of shielding is placed or
formed thereon. The second patterned or segmented layer 82 includes
conductive materials 64 and non-conductive material(s) 66.
[0089] It is noted that the non-conductive material 66 may be
formed of a single integral piece of non-conductive material or be
formed from a multitude of pieces of non-conductive material, the
multitude of pieces being connected together in such amanner to
function as a single integral piece of non-conductive material.
[0090] Upon the second patterned or segmented layer 82 of
shielding, a third patterned or segmented layer 84 of shielding is
placed or formed thereon. The third patterned or segmented layer 84
includes conductive materials 64 and non-conductive material(s)
66.
[0091] The conductive materials 64 of this embodiment of the
present invention may be a metal, a carbon composite, carbon
nanotubes, metal-coated carbon filaments (wherein the metal may be
one of the following metals: nickel, copper, cobalt, silver, gold,
tin, or zinc), metal-coated ceramic filaments (wherein the metal
may be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc), a composite of metal-coated carbon filaments
and a polymer (wherein the polymer may be one of the following:
polyether sulfone, silicone, polymide, polyvinylidene fluoride, or
epoxy), a composite of metal-coated ceramic filaments and a polymer
(wherein the polymer may be one of the following: polyether
sulfone, silicone, polymide, polyvinylidene fluoride, or epoxy), a
composite of metal-coated carbon filaments and a ceramic (wherein
the ceramic may be one of the following: cement, silicates,
phosphates, silicon carbide, silicon nitride, aluminum nitride, or
titanium diboride), a composite of metal-coated ceramic filaments
and a ceramic (wherein the ceramic may be one of the following:
cement, silicates, phosphates, silicon carbide, silicon nitride,
aluminum nitride, or titanium diboride), or a composite of
metal-coated (carbon or ceramic) filaments (wherein the metal may
be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc) and a polymer/ceramic combination (wherein the
polymer may be one of the following: polyether sulfone, silicone,
polymide, polyvinylidene fluoride, or epoxy and the ceramic may be
one of the following: cement, silicates, phosphates, silicon
carbide, silicon nitride, aluminum nitride, or titanium
diboride).
[0092] A more detailed description of the coated filaments is found
in U.S. Pat. No. 5,827,997, entitled "Metal Filaments for
Electromagnetic Interference Shielding." The entire content of U.S.
Pat. No. 5,827,997 is hereby incorporated by reference.
[0093] The non-conductive material(s) 66 of this embodiment of the
present invention may be a ceramic; glass; mica; anodized copper;
metallic oxides; natural or synthetic rubbers; or resins, such as
natural resins, epoxy resins, or silicones.
[0094] Over the shielding, a biocompatible layer 68 may be placed
or formed thereon. Preferably, the biocompatible layer is a
non-permeable diffusion resistant biocompatible material.
[0095] As illustrated in FIG. 3, the conductive materials 64 are
segmented or patterned in an x-direction, wherein the x-direction
is a direction substantially parallel to an axis of the lead 60. In
other words, the conductive materials 64 are broken up in this
direction such that one conductive material 64 is physically
separated from a neighboring conductive material 64 in the same
layer.
[0096] Moreover, as illustrated in FIG. 3, the conductive materials
64 of different layers are segmented or patterned in a y-direction,
wherein the y-direction is a direction substantially perpendicular
to an axis of the lead 60. In other words, the conductive materials
64 of different layers are broken up in this direction such that
one conductive material 64 is electrically isolated from a
neighboring conductive material 64 in an immediate adjacent layer;
i.e., a conductive material 64 in the first patterned or segmented
layer 80 is electrically isolated from a neighboring conductive
material 64 in the second patterned or segmented layer 82.
[0097] Preferably, the x-directional gap between non co-layer
immediately x-directional adjacent conductive materials 64 is much
smaller than the wavelength of the electromagnetic pulse that the
shield must attenuate to prevent the incident pulse from passing
unattenuated through the shield.
[0098] Lastly, the conductive materials 64 may be segmented or
patterned in a z-direction, wherein the z-direction represents the
planar surface of a layer as it is wrapped around the lead 60.
[0099] Another example of such a patterned or segmented shielding
for use with an electrical lead is illustrated in FIG. 4. As shown
in FIG. 4, an electrical lead 60 is initially coated with an
insulation layer 62 so as to electrically insulate the electrical
lead from its surroundings. Upon the insulation layer 62, a first
patterned or segmented layer 80 of shielding is placed or formed
thereon. The first patterned or segmented layer 80 includes
conductive materials 64 and non-conductive material(s) 66.
[0100] As noted above, the non-conductive material 66 may be formed
of a single integral piece of non-conductive material or be formed
from a multitude of pieces of non-conductive material, the
multitude of pieces being connected together in such a manner to
function as a single integral piece of non-conductive material.
[0101] Upon the first patterned or segmented layer 80 of shielding,
a layer 70 of non-conductive material is placed or formed thereon.
Upon the layer 70 of non-conductive material, a second patterned or
segmented layer 82 of shielding is placed or formed thereon. The
second patterned or segmented layer 82 includes conductive
materials 64 and non-conductive material(s) 66.
[0102] The conductive materials 64 of this embodiment of the
present invention may be a metal, a carbon composite, carbon
nanotubes, metal-coated carbon filaments (wherein the metal may be
one of the following metals: nickel, copper, cobalt, silver, gold,
tin, or zinc), metal-coated ceramic filaments (wherein the metal
may be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc), a composite of metal-coated carbon filaments
and a polymer (wherein the polymer may be one of the following:
polyether sulfone, silicone, polymide, polyvinylidene fluoride, or
epoxy), a composite of metal-coated ceramic filaments and a polymer
(wherein the polymer may be one of the following: polyether
sulfone, silicone, polymide, polyvinylidene fluoride, or epoxy), a
composite of metal-coated carbon filaments and a ceramic (wherein
the ceramic may be one of the following: cement, silicates,
phosphates, silicon carbide, silicon nitride, aluminum nitride, or
titanium diboride), a composite of metal-coated ceramic filaments
and a ceramic (wherein the ceramic may be one of the following:
cement, silicates, phosphates, silicon carbide, silicon nitride,
aluminum nitride, or titanium diboride), or a composite of
metal-coated (carbon or ceramic) filaments (wherein the metal may
be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc) and a polymer/ceramic combination (wherein the
polymer may be one of the following: polyether sulfone, silicone,
polymide, polyvinylidene fluoride, or epoxy and the ceramic may be
one of the following: cement, silicates, phosphates, silicon
carbide, silicon nitride, aluminum nitride, or titanium diboride).
Over the shielding, a biocompatible layer 68 may be placed or
formed thereon. Preferably, the biocompatible layer is a
non-permeable diffusion resistant biocompatible material.
[0103] The non-conductive materials 66 of this embodiment of the
present invention may be a ceramic; glass; mica; anodized copper;
metallic oxides; natural or synthetic rubbers; or resins, such as
natural resins, epoxy resins, or silicones.
[0104] As illustrated in FIG. 4, the conductive materials 64 are
segmented or patterned in an x-direction, wherein the x-direction
is a direction substantially parallel to an axis of the lead 60. In
other words, the conductive materials 64 are broken up in this
direction.
[0105] Moreover, as illustrated in FIG. 4, the conductive materials
64 are patterned or segmented in a y-direction, wherein the
y-direction is a direction substantially perpendicular to an axis
of the lead 60. In other words, the conductive materials 64 are
broken up in this direction such that one conductive material 64 is
electrically isolated from a neighboring conductive material 64 in
an immediate adjacent layer; i.e., a conductive material 64 in the
first patterned or segmented layer 80 is electrically isolated from
a neighboring conductive material 64 in the second patterned or
segmented layer 82.
[0106] Preferably, the x-directional gap between non co-layer
immediately x-directional adjacent conductive materials 64 is much
smaller than the wavelength of the electromagnetic pulse that the
shield must attenuate to prevent the incident pulse from passing
unattenuated through the shield.
[0107] Lastly, the conductive materials 64 may be segmented in a
z-direction, wherein the z-direction represents the planar surface
of a layer as it is wrapped around the lead 60.
[0108] As noted above, the conductive material 64 is patterned or
segmented to limit the build up of eddy currents. Examples of the
patterning of the conductive materials 64 are illustrated in FIGS.
5-7.
[0109] FIG. 5 illustrated a top view of a fabricated layer of the
shielding. As illustrated in FIG. 5, the conductive materials 64
are patterned or segmented in such a way that a width of the
conductive material 64 in the x-direction, the x-direction being
substantially parallel to an axis of a lead, is substantially
equal. More specifically, the conductive materials 64 is patterned
or segmented in such a way that conductive materials 64 spaced in a
predetermined manner; e.g., the spacing between the conductive
materials 64, in the x-direction, may be equal or set in a
predetermined manner. Lastly, the spacing between conductive
materials 64 in a z-direction, the z-direction representing the
planar surface of a layer as it is wrapped around a lead, is not
necessarily equal or set in a predetermined manner; however the
spacing in the z-direction may be equal or set in a predetermined
manner.
[0110] FIG. 6 illustrated a top view of a fabricated layer of the
shielding. As illustrated in FIG. 6, the conductive materials 64
are patterned or segmented in such a way that a width of the
conductive material 64 in the z-direction, the z-direction
representing the planar surface of a layer as it is wrapped around
a lead, is substantially equal. More specifically, the conductive
materials 64 are patterned or segmented in such a way that
conductive materials 64 spaced in a predetermined manner; e.g., the
spacing between the conductive materials 64, in the z-direction,
may be equal or set in a predetermined manner. Lastly, the spacing
between conductive materials 64 in a x-direction, the x-direction
being substantially parallel to an axis of a lead, is not
necessarily equal or set in a predetermined manner; however the
spacing in the x-direction may be equal or set in a predetermined
manner.
[0111] FIG. 7 illustrated a top view of a fabricated layer of the
shielding. As illustrated in FIG. 7, the conductive material 64 are
patterned or segmented in such a way that the widths of the
conductive materials 64 in the z-direction, the z-direction
representing the planar surface of a layer as it is wrapped around
a lead, are random. More specifically, the conductive materials 64
are patterned or segmented in such a way that conductive materials
64 are spaced in a random manner. Lastly, the spacing between
conductive materials 64 in an x-direction, the x-direction being
substantially parallel to an axis of a lead, is also random.
[0112] It is noted that the although FIGS. 3-7 show a shielding
with respect to a lead, the shielding can be used to shield other
shaped devices such as electrical component housings or other
components of a medical assist device/system.
[0113] It is further noted that the actual dimensions of the
non-conductive materials can be set to suppress the radiation and
eddy currents induced from specific frequencies; e.g., the lengths
of the non-conductive material can be varied to suppress or block
specific or undesired radiation frequencies. The dimensions of the
non-conductive material would be on the order of the wavelength of
the radiation so as to suppress, block, or shield the
radiation.
[0114] The actual layers may be formed using a photoresist/masking
process such as used in integrated circuit fabrication and thin
film deposition. By using a mask/photoresist process, the
conductive/insulative layers in the vertical direction can be
easily formed so that the conductive/insulative layers alternate
(radially outward from the axis of the lead), and the
conductive/insulative layers in the horizontal direction also
alternate (substantially parallel to an axis of the lead).
[0115] As noted above, the shielding may also be apertured. More
specifically, implantable medical devices, such as pacemaker leads,
typically demonstrate heating of 10-100.degree. C., whereas body
tissues can be irreversibly damaged by increases in temperature of
as little as 2-4.degree. C. Therefore, up to a 50.times. reduction
in the amount of energy delivered to the implanted medical device
is required to make the device safe, 100.times. if a reasonable
design margin of safety is desired. This reduction in the amount of
energy delivered to the implanted medical device corresponds to 20
dB of attenuation.
[0116] The performance of a shield or shield effectiveness (SE)
with a circular opening has been shown to be:
SE(dB)=20 Log [.lambda./2.pi.r]
[0117] where .lambda. is the incident wavelength and r is the
radius of the circular opening.
[0118] For holes of other shapes (e.g. an annulus) the factor
2.pi.r may be replaced by 2L where L is the largest dimension of
the opening (i.e. the circumference):
SE(dB)=20 Log [.lambda./2L]
[0119] This formula indicates that a slot or annulus with a
greatest dimension equal to {fraction (1/20)} of the wavelength is
required to provide a shielding effectiveness (SE) against heating
effects of 20 dB.
[0120] The wavelength at a particular frequency in air is given by
the equation .lambda.=c/f, where c is the speed of light
(3.times.10.sup.8 meters/second) and f is the frequency. However,
the wavelength in a particular material, such as body tissues, is
given by the following equation:
.lambda..sub.m=.lambda..sub.o/[.epsilon..sub.rel].sup.1/2
[0121] where .epsilon..sub.rel is the relative dielectric
constant.
[0122] At a frequency of 64 MHz and .epsilon..sub.rel=80 (a typical
value for body tissue), .lambda..sub.o is equal to 4.7 meters, and
.lambda..sub.m is equal to 0.5 meters. At a frequency of 64 MHz, a
shield effectiveness of 20 dB requires that L, the maximum aperture
dimension (i.e. annulus circumference), not exceed 0.025 meters or
25 millimeters (L=0.5 meters/20).
[0123] However, it is well known that shielding effectiveness is
reduced as the number of openings N is increased. Shield
effectiveness scales as the square root on N.
[0124] In the present invention, N may range from 10 to 1000.
Therefore, to maintain a minimum shield effectiveness against
heating effects of 20 dB, the maximum aperture dimension should not
exceed approximately 10 millimeters to 1 millimeter, depending upon
how many apertures or openings are incorporated into the
shield-design. At higher frequencies such as 128 MHz (the frequency
of emerging MRI facilities), the maximum aperture dimension should
not exceed 5 millimeters to 0.5 millimeters.
[0125] Moreover, implantable medical devices, such as pacemaker
leads, must also be capable of making sensitive electrical
measurements such as intracardiac ECG signals, which are typically
1-25 millivolts. To insure the accurate measurements of such
signals, a measurement resolution capability of less than 0.1
millivolt is required. Experimental measurements have shown that
MRI procedures can induce voltages of as much as 100 volts in
implanted pacing leads. Therefore, a shield attenuation of
1,000,000 or 60 dB is required, which requires that the maximum
aperture dimension must equal {fraction (1/2000)} of the
wavelength. At 64 MHz, L=0.5 meters/2000=0.00025 meters or 0.25
millimeters. Since N may range from 10 to 1000, the maximum
aperture dimension is further reduced to approximately 0.1
millimeters to 0.01 millimeters.
[0126] Therefore, depending upon the specific application and
shield design, a maximum aperture dimension of 0.01 millimeters to
10 millimeters should be used.
[0127] An example of such an apertured shielding for use with an
electrical lead is illustrated in FIG. 32. As shown in FIG. 32, an
electrical lead 60 is initially coated with an insulation layer 62
so as to electrically insulate the electrical lead from its
surroundings. Upon the insulation layer 62, a first patterned or
apertured layer 800 of shielding is placed or formed thereon. The
first patterned or apertured layer 800 includes conductive material
64 and non-conductive materials 66.
[0128] Upon the first patterned or apertured layer 800 of
shielding, a second patterned or apertured layer 820 of shielding
is placed or formed thereon. The second patterned or apertured
layer 820 includes conductive material 64 and non-conductive
materials 66.
[0129] It is noted that the conductive material 64 may be formed of
a single integral piece of conductive material or be formed from a
multitude of pieces of conductive material, the multitude of pieces
being electrically connected together in such a manner to function
as a single integral piece of conductive material.
[0130] Upon the second patterned or apertured layer 820 of
shielding, a third patterned or apertured layer 840 of shielding is
placed or formed thereon. The third patterned or apertured layer
840 includes conductive material 64 and non-conductive materials
66.
[0131] The conductive materials 64 of this embodiment of the
present invention may be a metal, a carbon composite, carbon
nanotubes, metal-coated carbon filaments (wherein the metal may be
one of the following metals: nickel, copper, cobalt, silver, gold,
tin, or zinc), metal-coated ceramic filaments (wherein the metal
may be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc), a composite of metal-coated carbon filaments
and a polymer (wherein the polymer may be one of the following:
polyether sulfone, silicone, polymide, polyvinylidene fluoride, or
epoxy), a composite of metal-coated ceramic filaments and a polymer
(wherein the polymer may be one of the following: polyether
sulfone, silicone, polymide, polyvinylidene fluoride, or epoxy), a
composite of metal-coated carbon filaments and a ceramic (wherein
the ceramic may be one of the following: cement, silicates,
phosphates, silicon carbide, silicon nitride, aluminum nitride, or
titanium diboride), a composite of metal-coated ceramic filaments
and a ceramic (wherein the ceramic may be one of the following:
cement, silicates, phosphates, silicon carbide, silicon nitride,
aluminum nitride, or titanium diboride), or a composite of
metal-coated (carbon or ceramic) filaments (wherein the metal may
be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc) and a polymer/ceramic combination (wherein the
polymer may be one of the following: polyether sulfone, silicone,
polymide, polyvinylidene fluoride, or epoxy and the ceramic may be
one of the following: cement, silicates, phosphates, silicon
carbide, silicon nitride, aluminum nitride, or titanium
diboride).
[0132] A more detailed description of the coated filaments is found
in U.S. Pat. No. 5,827,997, entitled "Metal Filaments for
Electromagnetic Interference Shielding." The entire content of U.S.
Pat. No. 5,827,997 is hereby incorporated by reference.
[0133] The non-conductive materials 66 of this embodiment of the
present invention may be a ceramic; glass; mica, anodized copper;
metallic oxides; natural or synthetic rubbers or resins, such as
natural resins, epoxy resins, or silicones.
[0134] Over the shielding, a biocompatible layer 68 may be placed
or formed thereon. Preferably, the biocompatible layer is a
non-permeable diffusion resistant biocompatible material.
[0135] As illustrated in FIG. 32, the conductive material 64 is
patterned or apertured in an x-direction through the utilization of
non-conductive materials 66, wherein the x-direction is a direction
substantially parallel to an axis of the lead 60. In other words,
the conductive material 64 is broken up, patterned, or apertured in
this direction such that one non-conductive material 66 is
physically separated from a neighboring non-conductive material 66
in the same layer.
[0136] Moreover, as illustrated in FIG. 32, the conductive
materials 64 of different layers are patterned in a y-direction,
wherein the y-direction is a direction substantially perpendicular
to an axis of the lead 60. In other words, the conductive materials
64 of different layers are broken up in this direction such that
one conductive material 64 is electrically isolated from a
neighboring conductive material 64 in an immediate adjacent layer;
i.e., a conductive material 64 in the first patterned or apertured
layer 800 is electrically isolated from a neighboring conductive
material 64 in the second patterned or apertured layer 82.
[0137] Preferably, the distance of the overlap is much smaller than
the wavelength of the electromagnetic pulse that the shield must
attenuate to prevent the incident pulse from passing unattenuated
through the shield.
[0138] If the length of the overlapping shield dimension is
designed to exceed the thickness of the shield material, the
structure thus formed resembles a waveguide and frequencies well
below its cut-off frequency are subject to additional attenuation.
For a frequency of 64-128 MHz, this overlap corresponds to
approximately 2-3.times. the maximum aperture dimension.
[0139] Therefore, depending upon the specific application and
shield design, a minimum shield overlap dimension should be
2-3.times. the maximum aperture dimension. It is also important to
note that the heating may also increase with increased frequency.
At a static magnetic field of 1.5 Tesla the RF transmission
frequency is 63.7 MHz, at 3 T it is 127.5 MHz (factor of 42.5 MHz
for each Tesla).
[0140] Lastly, the conductive material 64 may be segmented or
apertured in a z-direction, wherein the z-direction represents the
planar surface of a layer as it is wrapped around the lead 60.
[0141] Another example of such a shielding for use with an
electrical lead is illustrated in FIG. 33. As shown in FIG. 33, an
electrical lead 60 is initially coated with an insulation layer 62
so as to electrically insulate the electrical lead from its
surroundings. Upon the insulation layer 62, a first patterned or
apertured layer 800 of shielding is placed or formed thereon. The
first patterned or apertured layer 800 includes conductive material
64 and non-conductive materials 66.
[0142] It is noted that the conductive material 64 may be formed of
a single integral piece of conductive material or be formed from a
multitude of pieces of conductive material, the multitude of pieces
being electrically connected together in such a manner to function
as a single integral piece of conductive material.
[0143] Upon the first patterned or apertured layer 800 of
shielding, a layer 70 of non-conductive material is placed or
formed thereon. Upon the layer 70 of non-conductive material, a
second patterned or apertured layer 820 of shielding is placed or
formed thereon. The second patterned or apertured layer 820
includes conductive material 64 and non-conductive materials
66.
[0144] The conductive materials 64 of this embodiment of the
present invention may be a metal, a carbon composite, carbon
nanotubes, metal-coated carbon filaments (wherein the metal may be
one of the following metals: nickel, copper, cobalt, silver, gold,
tin, or zinc), metal-coated ceramic filaments (wherein the metal
may be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc), a composite of metal-coated carbon filaments
and a polymer (wherein the polymer may be one of the following:
polyether sulfone, silicone, polymide, polyvinylidene fluoride, or
epoxy), a composite of metal-coated ceramic filaments and a polymer
(wherein the polymer may be one of the following: polyether
sulfone, silicone, polymide, polyvinylidene fluoride, or epoxy), a
composite of metal-coated carbon filaments and a ceramic (wherein
the ceramic may be one of the following: cement, silicates,
phosphates, silicon carbide, silicon nitride, aluminum nitride, or
titanium diboride), a composite of metal-coated ceramic filaments
and a ceramic (wherein the ceramic may be one of the following:
cement, silicates, phosphates, silicon carbide, silicon nitride,
aluminum nitride, or titanium diboride), or a composite of
metal-coated (carbon or ceramic) filaments (wherein the metal may
be one of the following metals: nickel, copper, cobalt, silver,
gold, tin, or zinc) and a polymer/ceramic combination (wherein the
polymer may be one of the following: polyether sulfone, silicone,
polymide, polyvinylidene fluoride, or epoxy and the ceramic may be
one of the following: cement, silicates, phosphates, silicon
carbide, silicon nitride, aluminum nitride, or titanium
diboride).
[0145] Over the shielding, a biocompatible layer 68 may be placed
or formed thereon. Preferably, the biocompatible layer is a
non-permeable diffusion resistant biocompatible material.
[0146] The non-conductive materials 66 of this embodiment of the
present invention may be a ceramic; glass; mica; anodized copper;
metallic oxides; natural or synthetic rubbers; or resins, such as
natural resins, epoxy resins, or silicones.
[0147] As illustrated in FIG. 4, the conductive materials 64 are
patterned or apertured in an x-direction, wherein the x-direction
is a direction substantially parallel to an axis of the lead 60. In
other words, the conductive materials 64 are broken up, patterned,
or apertured in this direction such that the non-conductive
materials 66 break up the conductive material 64.
[0148] Moreover, as illustrated in FIG. 33, the conductive
materials 64 are patterned in a y-direction, wherein the
y-direction is a direction substantially perpendicular to an axis
of the lead 60. In other words, the conductive materials 64 are
broken up in this direction such that one conductive material 64 is
electrically isolated from a neighboring conductive material 64 in
an immediate adjacent layer; i.e., a conductive material 64 in the
first patterned or apertured layer 800 is electrically isolated
from a neighboring conductive material 64 in the second patterned
or apertured layer 820.
[0149] Preferably, the distance of the overlap is much smaller than
the wavelength of the electromagnetic pulse that the shield must
attenuate to prevent the incident pulse from passing unattenuated
through the shield.
[0150] If the length of the overlapping shield dimension is
designed to exceed the thickness of the shield material, the
structure thus formed resembles a waveguide and frequencies well
below its cut-off frequency are subject to additional attenuation.
For a frequency of 64-128 MHz, this overlap corresponds to
approximately 2-3.times. the maximum aperture dimension.
[0151] Therefore, depending upon the specific application and
shield design, a minimum shield overlap dimension should be
2-3.times. the maximum aperture dimension.
[0152] Lastly, the conductive materials 64 may be segmented in a
z-direction, wherein the z-direction represents the planar surface
of a layer as it is wrapped around the lead 60.
[0153] As noted above, the conductive material 64 is patterned or
apertured to limit the build up of eddy currents. Examples of the
patterning or aperturing of the conductive material 64 are
illustrated in FIGS. 34-36.
[0154] FIG. 34 illustrated a top view of a fabricated layer of the
shielding. As illustrated in FIG. 34, the conductive material 64 is
patterned or apertured in such a way that a width of the conductive
material 64 in the x-direction, the x-direction being substantially
parallel to an axis of a lead, is substantially equal. More
specifically, the conductive material 64 is patterned or apertured
in such a way that non-conductive materials 66 spaced in a
predetermined manner; e.g., the spacing between the non-conductive
materials 66, in the x-direction, may be equal or set in a
predetermined manner. Lastly, the spacing between non-conductive
materials 66 in a z-direction, the z-direction representing the
planar surface of a layer as it is wrapped around a lead, is not
necessarily equal or set in a predetermined manner; however the
spacing in the z-direction may be equal or set in a predetermined
manner.
[0155] Preferably, the distance of the overlap is much smaller than
the wavelength of the. electromagnetic pulse that the shield must
attenuate to prevent the incident pulse from passing unattenuated
through the shield.
[0156] If the length of the overlapping shield dimension is
designed to exceed the thickness of the shield material, the
structure thus formed resembles a waveguide and frequencies well
below its cut-off frequency are subject to additional attenuation.
For a frequency of 64-128 MHz, this overlap corresponds to
approximately 2-3.times. the maximum aperture dimension.
[0157] Therefore, depending upon the specific application and
shield design, a minimum shield overlap dimension should be
2-3.times. the maximum aperture dimension.
[0158] FIG. 35 illustrated a top view of a fabricated layer of the
shielding. As illustrated in FIG. 35, the conductive material 64 is
patterned or apertured in such a way that a width of the conductive
material 64 in the z-direction, the z-direction representing the
planar surface of a layer as it is wrapped around a lead, is
substantially equal. More specifically, the conductive material 64
is patterned or apertured in such a way that non-conductive
materials 66 spaced in a predetermined manner; e.g., the spacing
between the non-conductive materials 66, in the z-direction, may be
equal or set in a predetermined manner. Lastly, the spacing between
non-conductive materials 66 in a x-direction, the x-direction being
substantially parallel to an axis of a lead, is not necessarily
equal or set in a predetermined manner; however the spacing in the
x-direction may be equal or set in a predetermined manner.
[0159] Preferably, the distance of the overlap is much smaller than
the wavelength of the electromagnetic pulse that the shield must
attenuate to prevent the incident pulse from passing unattenuated
through the shield.
[0160] If the length of the overlapping shield dimension is
designed to exceed the thickness of the shield material, the
structure thus formed resembles a waveguide and frequencies well
below its cut-off frequency are subject to additional attenuation.
For a frequency of 64-128 MHz, this overlap corresponds to
approximately 2-3.times. the maximum aperture dimension.
[0161] Therefore, depending upon the specific application and
shield design, a minimum shield overlap dimension should be
2-3.times. the maximum aperture dimension.
[0162] FIG. 36 illustrated a top view of a fabricated layer of the
shielding. As illustrated in FIG. 36, the conductive material 64 is
patterned or apertured in such a way that the widths of the
non-conductive materials 66 in the z-direction, the z-direction
representing the planar surface of a layer as it is wrapped around
a lead, are random. More specifically, the conductive materials 64
are patterned or apertured in such a way that non-conductive
materials 66 spaced in a random manner. Lastly, the spacing between
non-conductive materials 66 in an x-direction, the x-direction
being substantially parallel to an axis of a lead, is also
random.
[0163] It is noted that the although FIGS. 32-36 show a patterned
or apertured shielding with respect to a lead, the patterned or
apertured shielding can be used to shield other shaped devices such
as electrical component housings or other components of a medical
assist device/system.
[0164] It is further noted that the actual dimensions of the
non-conductive materials can be set to suppress the radiation and
eddy currents induced from specific frequencies; e.g., the lengths
of the non-conductive material can be varied to suppress or block
specific or undesired radiation frequencies. The dimensions of the
non-conductive material would be on the order of the wavelength of
the radiation so as to suppress, block, or shield the
radiation.
[0165] Preferably, the distance of the overlap is much smaller than
the wavelength of the electromagnetic pulse that the shield must
attenuate to prevent the incident pulse from passing unattenuated
through the shield.
[0166] The actual layers may be formed using a photoresist/masking
process such as used in integrated circuit fabrication and thin
film deposition. By using a mask/photoresist process, the
conductive/insulative layers in the vertical direction can be
easily formed so that the conductive/insulative layers alternate
(radially outward from the axis of the lead), and the
conductive/insulative layers in the horizontal direction also
alternate (substantially parallel to an axis of the lead).
[0167] Although, as stated above, eddy currents in a conductive
structure may cause heating to the surrounding tissue, eddy
currents induced in the tissue may also cause the heating of the
surrounding tissue. As noted above, when a substance such as human
tissue is subjected to a static magnetic field, the individual
magnetic moments of the spins in the tissue align in a parallel and
anti-parallel direction with the static magnetic field. This
direction along the static magnetic field can be termed as the
longitudinal direction. In magnetic-resonance imaging, the radio
frequency polarizing field used for spin manipulation is constantly
changing and thus, the individual magnetic moments of the spins in
the tissue attempt to align with the polarizing field. The constant
changing of alignment of the magnetic moments of the spins in the
tissue causes the tissue's temperature to increase, thereby
exposing the tissue to possible magnetic-resonance imaging induced
thermal damage. Thus, this second possible cause of undesirable
heat needs to also be addressed to prevent tissue damage caused by
excessive heat.
[0168] In a further embodiment of the present invention, the heat
being accumulated in the surrounding tissue is addressed and
reduced or prevented, notwithstanding the source or cause of the
excessive heat. To better describe this embodiment, the term,
specific absorption ratio, will be used in discussing the excessive
heat.
[0169] The specific absorption ratio is defined as the amount of
power absorbed by a sample within a magnetic-resonance imaging
scanner. The sample could be a suitable phantom mimicking tissue
properties or a tissue region of patient.
[0170] As noted above, this excessive heat could be caused by
either eddy currents induced in conductive structures or eddy
currents induced in the actual tissue. It is noted that although
two distinct sources of excessive heat are identified for the
purposes of describing this embodiment of the-present invention,
the actual source of the heat may not beknown, nor is it a real
consideration in the actual functionality of this embodiment of the
present invention.
[0171] With respect to induced eddy currents, due to their
electromagnetic receptive capabilities, cardiac assist device
leads, metallic guidewires, neurostimulator leads and intraluminal
coil leads may induce a localized increase in the specific
absorption ratio, thereby resulting in an increased heating of the
surrounding tissue. Moreover, due to the individual magnetic
moments of the spins in the tissue attempting to align with the
polarizing field, the acquisition sequence of the
magnetic-resonance imaging process may also induce a localized
increase in the specific absorption ratio, thereby resulting in an
increased heating of the surrounding tissue. This tissue heating,
notwithstanding the source, may be a function of the time of
radio-frequency excitation, gradient coil switching speed, or
strength of the static magnetic field.
[0172] In this embodiment of the present invention, the localized
specific absorption ratio of tissue is reduced by altering the
actual magnetic-resonance imaging acquisition process. More
specifically, according to a preferred embodiment of the present
invention, the localized specific absorption ratio of tissue is
reduced by varying the imaging pulse sequence and/or changing the
timing parameters of magnetic-resonance imaging acquisition
process. For example, in response to increases in the localized
specific absorption ratio or during magnetic-resonance image
acquisition, the radio-frequency excitation cycle may be
interrupted to allow for cooling of the tissue. In other words, if
a threshold temperature level or a localized specific absorption
ratio is exceeded, the magnetic-resonance imaging acquisition can
be halted until the tissue temperature or localized specific
absorption ratio decreases to an accepted level.
[0173] It is noted that the increased in the localized specific
absorption ratio may or may not have resulted from placement of a
conductive structure near the tissue region.
[0174] FIG. 8 illustrates a magnetic-resonance imaging process
according to the concepts of the present invention, which alters
the actual magnetic-resonance imaging acquisition process in
response to a temperature change in the tissue region being imaged.
In this process, at Step S1, an image acquisition sequence is
started by the magnetic-resonance imaging system. At step S2, the
temperature change in the tissue region being imaged is measured.
It is noted that this temperature change may be relative or
absolute. The temperature change in the tissue region being imaged
can be measured a number of ways.
[0175] For example, in one embodiment, the relative temperature
change can be monitored by magnetic image acquisition techniques.
More specifically, relative temperature changes can be measured
using a double gradient echo sequence to acquire a baseline phase
image and a second phase image at within a single radio-frequency
excitation of one same pulse sequence. A relative spatial
temperature map of the desired region of interest can be calculated
based on the subtraction of the two phase images collected during
the double gradient echo sequence acquisition.
[0176] A more detail description of this relative temperature
change acquisition process is described in U.S. Pat. No. 5,711,300,
entitled "Real Time In Vivo Measurement Of Temperature Changes With
Magnetic-resonance Imaging." The entire contents of U.S. Pat. No.
5,711,300 are hereby incorporated by reference. It is noted that
this is only one example of collecting phase image data for
relative temperature mapping. Other methods for collecting phase
image data for relative temperature mapping may also be used.
[0177] It is further noted that absolute temperature changes can be
determined using known methods, such as fiber optic thermocouples
attached to guidewires or catheters, or non-contact thermographic
imaging techniques.
[0178] Upon calculating the temperature change of the desired
tissue region, such as described above with respect to the relative
spatial temperature map of the desired region of interest, at Step
S3, a predetermined upper temperature threshold is compared with
the calculated temperature change, or in the example described
above, compared with the relative temperature map. If the
temperature change exceeds or is equal to the predetermined upper
temperature threshold, the imaging acquisition sequence is adjusted
or halted so as to reduce the temperature change below the
predetermined upper temperature threshold, at Step S4.
[0179] The process returns to Step S2 to make another determination
of the temperature change in the tissue region of interest. If the
temperature change is below the predetermined upper temperature
threshold, Step S5 determines if the measured temperature change is
above a predetermined lower temperature threshold. If Step S5
determines that the measured temperature change is above a
predetermined lower temperature threshold, the process returns to
Step S4 and the imaging acquisition sequence remains adjusted or
halted so as to reduce the temperature of the tissue region of
interest. On the other hand, if Step S5 determines that the
measured temperature change is below or is equal to a predetermined
lower temperature threshold, the process resume image acquisition
at to Step S6.
[0180] It is to be understood that this type of adjustment can be
applied to any form of pulse sequence or imaging acquisition method
so as it meets the requirements of the user. There are an infinite
number of variations in pulse sequence design and pulse sequence
parameter determination to minimize heating. The actual variation
in pulse sequence design and pulse sequence parameter determination
will be dependent upon the image quality and the temporal
resolution of acquisition desired; however, the required tissue
temperature threshold(s) will be the ultimate determinant.
[0181] This embodiment may also include a user interactive process,
which can be programmed. An example of such a process is described
in detail in U.S. Pat. No. 6,396,266, entitled "magnetic-resonance
Imaging System With Interactive magnetic-resonance Geometry
Prescription Control." The entire contents of U.S. Pat. No.
6,396,266 are hereby incorporated by reference.
[0182] In another embodiment, the SAR coefficients of a wide range
of implantable medical devices are programmed into the
magnetic-resonance imaging system, which then uses these data to
predetermine and apply as safe magnetic-resonance imaging
sequence.
[0183] In yet another embodiment, as illustrated in FIG. 9, a
magnetic-resonance imaging process, according to the concepts of
the present invention, alters the actual magnetic-resonance imaging
acquisition process in response to a calculated specific absorption
ratio of the tissue region being imaged. In this process, at Step
S1, an image acquisition sequence is started by the
magnetic-resonance imaging system. At step S21, the specific
absorption ratio is estimated using theoretical modeling and/or
empirical results from certain conductors being imaged in tissues.
At Step S31, the calculated specific absorption ratio level(s) are
then be compared to threshold levels.
[0184] If the calculated specific absorption ratio level exceeds or
is equal to the predetermined upper specific absorption ratio level
threshold, the imaging acquisition sequence is adjusted or halted
so as to reduce the temperature of the tissue region of interest,
at Step S4.
[0185] The process returns to Step S21 to make another
determination of the specific absorption ratio level. If the
determined specific absorption ratio level is below the
predetermined upper specific absorption ratio level threshold, Step
S51 determines if the determined specific absorption ratio level is
above a predetermined lower specific absorption ratio threshold. If
Step S51 determines that the determined specific absorption ratio
level is above a predetermined lower specific absorption ratio
threshold, the process returns to Step S4 and the imaging
acquisition sequence remains adjusted or halted so as to reduce the
temperature of the tissue region of interest. On the other hand, if
Step S51 determines that the determined specific absorption ratio
level is below or is equal to a predetermined lower specific
absorption ratio threshold, the process resume image acquisition at
to Step S6.
[0186] This embodiment may also include a user interactive process
described above.
[0187] In the embodiments illustrated in FIGS. 8 and 9, specific
absorption ratio or energy deposition of radio frequency energy in
the tissue can be reduced by specific manipulations of each
individual radio frequency pulse. Moreover, these embodiments of
the present invention may manipulate the pulse waveform, pulse
duration and pulse amplitude to reduce the specific absorption
ratio. The choice of the integrated values of these three
parameters is determined upon by the requirements of the user. For
example, reduction of the amplitude and increase in the duration of
the pulse may reduce the specific absorption ratio, but this may
not be appropriate for the specified timing required to image
certain contrast parameters.
[0188] A more detail discussion of the specific manipulations that
be used to reduce specific absorption ratio is set forth in U.S.
Pat. No. 4,760,336, entitled "Variable Rate Magnetic Resonance
Selective Excitation For Reducing radio-frequency Power And
Specific Absorption Rate." The entire contents of U.S. Pat. No.
4,760,336 are hereby incorporated by reference.
[0189] It is also noted that the embodiments illustrated in FIGS. 8
and 9 can be used not only as a stand-alone solution to reduce the
increases in specific absorption ratio resulting from conductive
structures in tissue, but also in conjunction with the patterned
shielding embodiments described above.
[0190] A block diagram of an ECG acquisition and processing system
is shown in FIG. 10. The ECG monitoring system can be used as a
stand-alone ECG monitor or as part of an external or internal
pacemaker. A patient is connected to ECG monitoring circuitry 110
via surface or internal electrodes 112 and lead wires 114. The lead
wires 114 connect to the input ECG/radio-frequency filter 116.
[0191] This filter design enables the monitoring circuit to operate
in the high radio-frequency field environment generated by the
magnetic-resonance imaging system. First, the radio-frequency
filter prevents the development of excess voltage on the ECG input
amplifier and processor 118 that would otherwise damage the
amplifier, and excess voltage on the ECG leads potentially inducing
a dangerous heart rhythm. Second, it prevents the development of
excess current on the ECG electrodes 112 that would induce
thermogenic damage at the electrode-heart tissue interface that
could lead to loss of capture of pacing signals and possibly death.
Third, the radio-frequency filter 116, when attached between the
lead wires 114 and the ECG input amplifier, will not interfere with
proper sensing of the ECG signal.
[0192] The input ECG/radio-frequency filter 116 has the
characteristics of a low-pass filter with approximately a 100 dB
signal attenuation in the commonly used magnetic-resonance imaging
system radio-frequency frequency range. The frequencies contained
in the intracardiac or surface ECG signal, however, are passed
without any attenuation.
[0193] In the case where the lead wires 114 are of long length,
such as in external monitoring, there is significant stray
capacitance between the lead wires. As a result, electromagnetic
fields generated between the electrodes 112 can produce dangerously
high voltage within and current flow through the electrodes.
[0194] To limit this voltage and current, inductors 120, may be
placed in the lead wires 114, close to the electrodes. The
ECG/radio-frequency input filter 116 and lead wire inductors 120
must contain components with a resonance frequency higher than the
radio-frequency frequency of the magnetic-resonance imaging system
in use.
[0195] The ECG amplifier/processor 118 contains low-pass or band
reject filter 119 placed after the initial amplifier 117 and before
final stages 121. The low-pass or band reject filter must pass the
QRS signals and reject gradient field and radio-frequency
frequencies. The output ECG signal 122, which can take any
conventional form such as an R-wave detector output or analog ECG,
can be used for any purpose. It can be used to display the ECG on
monitor 124, or synchronize an external or internal device 126,
such as an magnetic-resonance imaging system or cardiac
pacemaker.
[0196] In another embodiment shown in FIG. 11, a pacemaker system
128 is described. This pacemaker system can be either external or
implantable. In addition to the elements shown in FIG. 10, this
embodiment also contains a pacing output stage 130 that is
controlled by the pacing logic 132.
[0197] The pacing logic is itself controlled by signals from the
ECG amplifier/processor 134. The patient is connected to the
pacemaker system 128, via surface or internal sensing electrodes
136 and lead wires 138. The lead wires are connected to the ECG
amplifier/processor 134, via the input ECG/radio-frequency filter
140. The pacing output stage 130 is connected to pacing electrodes
142 via the output radio-frequency filter 144 and pacing lead wires
146. The ECG amplifier/processor 134 contains a low-pass or band
reject filter 135 placed after the initial amplifier 133 and before
the final stages 137.
[0198] The low-pass or band reject filter must pass the QRS signal
and reject gradient field frequencies caused by the
magnetic-resonance imaging. As in the previous embodiment,
inductors 148, 150 may be placed in the lead wires 138, 146 close
to the electrodes 136, 142. In this embodiment, the sensing 136 and
pacing 142 electrodes, wires 138, 146, inductors 148, 150, and
radio-frequency filters 149, 144 are shown separately. However,
these components can be merged to create a combined ECG
sensing/pacing radio-frequency electrodes, wires, inductors, and
filters.
[0199] An implantable pacemaker 150 capable of bi-directional
communication with an external programmer 152 is shown in FIG. 12.
The implantable unit contains conventional programmable pacing
logic 154 for sensing and/or pacing the ventricle and/or atrium and
includes an ECG amplifier and processor 156, pacing output stage
158, combined input/output radio-frequency filter 160, electrodes
162 and leads 164. The ECG amplifier and processor 156 contains a
low-pass or band reject filter 157 placed after the initial
amplifier 155 and before final stages 159. The low-pass or band
reject filter must pass the QRS signals and reject gradient field
frequencies caused by the magnetic-resonance imaging.
[0200] As in the previous embodiment, inductors 166 may be placed
in the lead wires 164 close to the electrodes 162. The electrodes
162, leads 164, inductors 166, and input/output filter 160 may be
single elements performing combined input and output functions or
may be separate input and output elements. For dual chamber
pacemakers, multiple sets of sensing/pacing electrodes may be used
159, 162, which are each connected to input/output radio-frequency
filters 160.
[0201] A first electrode pair may be for atrium sensing or pacing
and a second electrode pair may be for ventricle sensing or pacing.
As is taught in the pacing art, pacing logic 154 may be programmed,
for example, to sense the atrium signal and pace via the ventricle
electrodes. In addition, a new form of circulatory support called
cardiomyoplasty may utilize other electrodes 161 and leads 167 to
pace skeletal muscle wrapped around the heart and stimulated in
synchrony with pacing the heart. In each type of pacemaker the
leads 163, 164, 167 may contain inductors 165, 166, 169 to reduce
electrical currents across the electrodes and each lead may be
connected to separate input/output radio-frequency filter 160.
[0202] In order to protect the pacemaker from the effects of the
radio-frequency field generated by the magnetic-resonance imaging
system, it is necessary to surround the components with an
radio-frequency shield 168. In a preferred embodiment of the
present invention, this shield is as described above with respect
to FIGS. 3 through 7.
[0203] More specifically, the shield may consist of patterned
layers including conductive materials and non-conductive materials.
The conductive materials of the present invention may be a metal, a
carbon composite, carbon nanotubes, metal-coated carbon filaments
(wherein the metal may be one of the following metals: nickel,
copper, cobalt, silver, gold, tin, or zinc), metal-coated ceramic
filaments (wherein the metal may be one of the following metals:
nickel, copper, cobalt, silver, gold, tin, or zinc), a composite of
metal-coated carbon filaments and a polymer (wherein the polymer
may be one of the following: polyether sulfone, silicone, polymide,
polyvinylidene fluoride, or epoxy), a composite of metal-coated
ceramic filaments and a polymer (wherein the polymer may be one of
the following: polyether sulfone, silicone, polymide,
polyvinylidene fluoride, or epoxy), a composite of metal-coated
carbon filaments and a ceramic (wherein the ceramic may be one of
the following: cement, silicates, phosphates, silicon carbide,
silicon nitride, aluminum nitride, or titanium diboride), a
composite of metal-coated ceramic filaments and a ceramic (wherein
the ceramic may be one of the following: cement, silicates,
phosphates, silicon carbide, silicon nitride, aluminum nitride, or
titanium diboride), or a composite of metal-coated (carbon or
ceramic) filaments (wherein the metal may be one of the following
metals: nickel, copper, cobalt, silver, gold, tin, or zinc) and a
polymer/ceramic combination (wherein the polymer may be one of the
following: polyether sulfone, silicone, polymide, polyvinylidene
fluoride, or epoxy and the ceramic may be one of the following:
cement, silicates, phosphates, silicon carbide, silicon nitride,
aluminum nitride, or titanium diboride).
[0204] Since the implantable pacemaker shown in FIG. 12 is enclosed
in an radio-frequency shield, it was necessary to develop a unique
method of external programming. The information received by the
pacemaker 150 is processed to separate the desired communication
signal from the radio-frequency signals produced by the
magnetic-resonance imaging system. An antenna 170 is connected to
the telemetry radio-frequency amplifier 172 via the telemetry
receiver/band pass filter 174. This filter is a band pass filter
that passes only the specific programming frequency to the
telemetry amplifier 172.
[0205] The telemetry logic circuit 176 interprets all
radio-frequency received by the telemetry antenna 170 and will only
allow programming of the pacing logic circuit 154 when a specific
telemetry enable pattern is received. The telemetry enable logic
circuit 176 will also inhibit control of the pacing output stage
158 while programming is taking place. This prevents improper and
potentially dangerous pacing parameters from controlling the pacing
stage. As a further safety measure, the telemetry enable circuit
176 enables pacing at a preset rate ("safety pacing") to provide
adequate pacing back-up for pacemaker dependent patients during
external programming.
[0206] FIGS. 13 through 18 are schematic diagrams of the inductor
elements (elements 120, 148, 150, 165, 166, and 169) that are
placed in the lead wires to reduce harmful currents. As mentioned
previously, significant stray capacitance can be produced along the
lead wires, particularly for external pacers. As a result,
[0207] Electromagnetic fields generated between the electrodes
because of the radio-frequency field can produce dangerously high
voltages across and currents flowing through the electrodes. To
limit this current, and enhance patient safety, the inductor
elements shown in FIGS. 13 through 18 are placed in the lead wires
close to the electrodes. FIG. 13 shows the inductor elements 178
when a two-wire lead is used. FIG. 14 shows the inductor elements
180 when a two-wire shielded lead is used. FIG. 15 shows
alternative inductor elements 182, which incorporate a capacitor
184, thus comprising an additional low-pass L-C filter. FIG. 16
shows the same alternative inductor elements 183 and a capacitor
185, as shown in FIG. 15, but using a shielded two-wire lead. FIG.
17 shows the inductor elements 186 used in a multi-lead shielded
harness. FIG. 18 shows alternative inductor elements 188 used in a
multi-lead shielded harness, which also includes capacitor
components 190 connected to the shielding enclosure of the
capacitors.
[0208] The leads can be used to measure ECG, EEG, or other
electrical signals of physiological significance. FIGS. 13 through
18 show twisted leads, but it is to be understood that coaxial as
well as other forms of lead wires could be used.
[0209] FIGS. 19 through 31 show various embodiments of the input
radio-frequency filter, output radio-frequency filter, and combined
input/output radio-frequency filter. According to the concepts of
the present invention, there are many possible specific
implementations of the radio-frequency filter design that will
depend on: 1) whether two leads or multiple leads are used; 2)
whether the leads are shielded; 3) whether single or multiple stage
filtering is used; and 4) whether each radio-frequency filter is
housed in a separate shielded enclosure.
[0210] It should be apparent that the filters could be placed
either at the proximal end or the distal end of the pulse
generator--lead assembly.
[0211] It will also become apparent that the same filter design
principles can be used whether the filter is acting as an input
filter for an ECG monitor, as separate input filters and output
filters for a pacemaker, or as a single combined input/output
filter for a pacemaker.
[0212] FIGS. 19 and 20 show one-stage filters for use with two
unshielded leads. FIGS. 21 and 22 show alternative ways to shield
the electronics.
[0213] FIG. 19 shows the input filter 192 in a separate shielded
enclosure; whereas FIG. 20 shows the input filter 196 located in
the same-shielded enclosure 198 as the processing electronics 1100
with an internal wall 1102 shielding the monitoring electronics
1100 from the input filter 196. (Whether one or two enclosures are
used, the filter section should be shielded from the other
circuitry.) As can be seen in FIG. 19, each lead is connected to a
separate low-pass filter. (Leads 1103 may be twisted to reduce
electromagnetic interference). A first low-pass filter is made from
inductor 1104 and capacitor 1106; and a second low-pass filter is
made from inductor 1108 and capacitor 1110. Each low-pass filter
offers approximately 100 dB signal attenuation in the
radio-frequency frequency range of the magnetic-resonance imaging
system and passes with little attenuation the desired physiological
signal.
[0214] As mentioned previously, the L-C components must be high
frequency elements retaining their desired characteristics
throughout the frequency range. The output from the input filter
192 is connected to differential amplifier 1112. In order to
properly utilize the differential amplifier 1112, the non-inverting
input of the ECG amplifier 1112 is connected to the amplifier
zero-signal reference terminal via a resistor 1114, as is well
known in the art.
[0215] Obviously, any other method of converting a two to three
conductor input may also be used. The two capacitors 1106, 1110 in
the low-pass filters are referenced to a third wire 1116 that is
connected to the shield surrounding the filter 1122 and also
connected to the zero-signal reference terminal of the ECG
amplifier 1112.
[0216] The ECG amplifier may be a single differential amplifier or
a combination of multiple differential amplifiers. Alternatively,
capacitors 1106, 1110 can be connected directly to the shield 1122
with a separate wire connecting the shield 1122 to the zero-signal
reference terminal of the ECG amplifier 1112.
[0217] To improve the characteristics of the filter, resistors
1118, 1120 may be added between the input filter 192 and the input
of the differential amplifier 1112. Just as the input filter 192 is
enclosed in a shield 1122, the ECG amplifier 1112 and processor
1124 may also be enclosed in a separate shield 1125.
[0218] The shields may consist of patterned layers including
conductive materials and non-conductive materials. The conductive
materials of the present invention may be a metal, a carbon
composite, carbon nanotubes, metal-coated carbon filaments (wherein
the metal may be one of the following metals: nickel, copper,
cobalt, silver, gold, tin, or zinc), metal-coated ceramic filaments
(wherein the metal may be one of the following metals: nickel,
copper, cobalt, silver, gold, tin, or zinc), a composite of
metal-coated carbon filaments and a polymer (wherein the polymer
may be one of the following: polyether sulfone, silicone, polymide,
polyvinylidene fluoride, or epoxy), a composite of metal-coated
ceramic filaments and a polymer (wherein the polymer may be one of
the following: polyether sulfone, silicone, polymide,
polyvinylidene fluoride, or epoxy), a composite of metal-coated
carbon filaments and a ceramic (wherein the ceramic may be one of
the following: cement, silicates, phosphates, silicon carbide,
silicon nitride, aluminum nitride, or titanium diboride), a
composite of metal-coated ceramic filaments and a ceramic (wherein
the ceramic may be one of the following: cement, silicates,
phosphates, silicon carbide, silicon nitride, aluminum nitride, or
titanium diboride), or a composite of metal-coated (carbon or
ceramic) filaments (wherein the metal may be one of the following
metals: nickel, copper, cobalt, silver, gold, tin, or zinc) and a
polymer/ceramic combination (wherein the polymer may be one of the
following: polyether sulfone, silicone, polymide, polyvinylidene
fluoride, or epoxy and the ceramic may be one of the following:
cement, silicates, phosphates, silicon carbide, silicon nitride,
aluminum nitride, or titanium diboride).
[0219] FIG. 20 shows an alternative embodiment where the input
filter 196 and processing electronics 1100 are housed in a single
contiguous shielding enclosure 198. A separate metallic wall 1102
is used to shield the electronics 1100 from the input filter 196.
In this embodiment, the two capacitors 1126, 1128 that make up the
low-pass filters are connected to the zero-signal reference
terminal of the differential amplifiers 1130 via a third reference
wire 1132; the reference wire 1132 and zero-signal reference
terminal 1134 both connected to the pacemaker shielding case at a
single point.
[0220] Both FIGS. 19 and 20 show processing electronics 1224, 1136
that are used to further process the signal obtained from the
differential amplifiers 1112, 1130. The processing electronics may
be used to process and display the ECG signal, in the case of an
external monitor, or be used to process and control pacing in the
case of a pacemaker. It is to be understood that the filter
embodiments shown in FIGS. 19 and 20 could be used in a device that
monitors physiologically significant electrical signals other than
ECG.
[0221] FIGS. 21 and 22 show one-stage filter designs 1135 and 1137
that is preferred when the input leads are shielded. The design is
identical to that shown in FIGS. 19 and 20, except that the lead
shield must be coupled to the shield enclosure via an
radio-frequency filter. To accomplish this, the lead shield 1140 of
the leads 1138, in FIG. 21, is connected by an inductor 1146 to the
third reference wire 1148, and the reference wire is connected to
the filter's shielding enclosure 1150. Analogously, the shield 1144
of the leads 1142 in FIG. 22 is connected by an inductor 1154 to
the reference wire 1156, connected in turn to the enclosure 1160.
In order to further reduce the effects of the magnetic-resonance
imaging system radio-frequency signal, more than one stage of
filtering can be used.
[0222] In the embodiment shown in FIG. 23, two stages of filtering
are used. Each filter stage is similar to the filter described in
FIG. 19. Each input lead is connected to a low-pass filter. Each
low-pass L-C filter has a capacitor connected to a reference wire
that is connected to the shielding enclosure.
[0223] The first stage filter 1162 contains low-pass filters
comprising inductors 1164, 1166 and capacitors 1168, 1170.
Capacitors 1168, 1170 are coupled to a reference wire 1172 that is
connected to the enclosure 1174. Similarly, the second filter stage
1176 contains low-pass filters made from inductors 1178, 1180 and
capacitors 1182, 1184, with capacitors 1182, 1184 coupled to the
reference wire 1172 that is connected to the shielding enclosure
1177. All the capacitors (1168, 1170, 1182 and 1184) used in the
low-pass filters are connected to the reference wire 1172 which is
connected to the zero-signal reference terminal of the differential
amplifier 1188.
[0224] Each stage in FIG. 23 is housed in a separate shielding
enclosure and each shielding enclosure is coupled to the reference
wire 1172. It is understood that this concept can be extended to
multistage filtering. The second stage 1176 could actually
represent the Nth stage of a multistage filter, with each identical
stage coupled together as shown.
[0225] An alternative embodiment for multiple stage filtering is
shown in FIG. 24. In this embodiment each filtering stage as well
as the differential amplifier and the associated electronics are
housed in the same contiguous shielding enclosure 1192. The filter
design in this embodiment is identical to that shown in FIG. 20;
however, the reference wire is only connected to the shielding case
1192 at a single point. In effect, the low-pass filters in each
stage and the zero-signal reference terminal of the differential
amplifier 1194 are all referred to the same point on the shielding
enclosure. It is to be understood that the embodiment shown in FIG.
24 may also have three or more stages, each stage shielded and
coupled together as shown.
[0226] FIGS. 25 and 25 are schematic drawings of a two-stage filter
for use with shielded leads. These filters are similar to the
multistage filter embodiments shown in FIGS. 23 and 24, except that
the lead shields are coupled via inductors to the reference wire in
the last filter stage. FIG. 25 is a two-stage filter where each
stage is separately isolated in enclosures 1196, 1198. Again, as in
the previous embodiments, low-pass filters are connected in each
input lead and referenced to a common reference line 1200 (i.e.,
capacitors 1202, 1204, 1206, and 1208 are connected to the
reference line 1200). The common reference line 1200 is connected
to the shielding enclosure for each stage and is input to the
zero-signal reference terminal of the differential amplifier
1210.
[0227] The lead shield 1212 is connected to inductor 1214 in the
first filter stage and via line 1215 to inductor 1216 in the second
filter stage. The inductor 1216 in the second, or final filter
stage, is connected to the common reference line 1290. As discussed
previously, the inductors 1214, 1216 provide high impedance to the
undesirable high frequency radio-frequency voltages produced by the
magnetic-resonance imaging system and prevent introduction of
radio-frequency interference into the processing electronics via
the common wire.
[0228] To increase performance, an additional capacitor 1218 may be
connected between lines 1215 and 1200 in the first stage and any
subsequent filter stage other than the last stage. Inductor 1214,
coupled with capacitor 1218, provides a third low-pass filter.
[0229] FIG. 26 is a multistage filter housed in a single shielding
enclosure 1220 with separate walls 1222, 1224 isolating the
filtering stages from each other and from the processing
electronics. In the first stage 1226, each lead 1228 and the lead
shield 1230 are coupled to low-pass filters referenced by
capacitors 1230, 1232, and 1234 to a common reference line 1236. In
the last stage 1238, only the two lead wires contain low-pass
filters referenced by capacitors 1240, 1242 to the common reference
line 1236. The lead shield 1230 is coupled via inductor 1244
directly to the reference line 1236.
[0230] As in the other embodiments, the reference line 1236 is
connected to the zero-signal reference terminal of the differential
amplifier 1248. Also as in the other embodiments, when a single
case is used, the common reference line 1236 is connected to the
shielding enclosure at a single point. It is of course understood
that the two-stage filters shown in FIGS. 25 and 26 could be easily
extended to multiple stage filtering with each stage connected as
described.
[0231] FIGS. 27 and 28 show an extension of the single and multiple
stage filtering to multiple lead shielded ECG harness. In FIG. 27,
a five conductor shielded cable 1250 is filtered by a single stage
filter 1252. In FIG. 28, a five conductor shielded cable 1254 is
filtered by a multiple stage filter 1256.
[0232] Thus far, the filters have been described as input filters.
However, as shown in FIG. 29, the filters can operate as both input
and output filters. As generally true in pacemakers, the lead pair
1258 can be used to sense electrical activity of the heart as well
as to pace the heart. As shown previously, low pass filters (i.e.,
inductor 1260/capacitor 1262 pair and inductor 1264/capacitor 1266
pair) for each lead are referenced to a common reference line 1268.
The reference line 1268 then connects to the zero-signal reference
terminal of the differential amplifier and to the shielding
enclosure. A pacing amplifier 1270 in controlled by electronics
1272 to generate a pacing signal. The pacing signal is connected
via lines 1274, 1276 through the filter, to the leads 1258. Two
inductors 1278, 1280 are placed in lines 1274, 1276 to further
prevent any current generated by the magnetic-resonance imaging
system from affecting the pacing amplifier 1270; the inductors
1278, 1 provide an additional high impedance to the high frequency
radio-frequency signals produced by the magnetic-resonance imaging
system.
[0233] Connecting the pacing amplifier to the leads 1258 via the
filter can be used with each of the filter embodiments described
above. In multistage filtering embodiments, pacing amplifiers would
be preferably connected into the last filtering stage.
[0234] FIG. 30 is an exemplary embodiment, showing how the pacing
amplifier 1282 can be connected via a multi-stage filter to the
shielded pacing/sensing leads 1234. The pacing amplifier 1282 is
connected via inductors 1286 and 1288 to the lead wires in the last
stage of the filter. As discussed above, the inductors 1286, 1288
provide additional protection from the high frequency
radio-frequency signals generated by the magnetic-resonance imaging
system.
[0235] There are an unlimited number of ways in which the filters
described herein can be connected to both sensing electronic and
pacing amplifiers. In fact, the same filter design can be used in
stimulators, as shown in FIG. 31, where only a pacing signal is
produced.
[0236] The filter design is identical to those described
previously. The leads 1290 are each connected to low pass filters
(inductor 1291/capacitor 1292 pair and inductor 1293/capacitor 1294
pair) that are referred by capacitors 1292, 1294 to a common
reference line 1296. Inductors 1298, 1300 provide additional
isolation to protect the pacing amplifier. The filter design for
stimulation use only can be extended to multistage embodiments as
described above.
[0237] In order to protect the implantable pacemaker elements from
the effects of the radio-frequency signals produced by the
magnetic-resonance imaging system, it is necessary to surround the
components with an radio-frequency shield. This shielding may
consist of patterned layers including conductive materials and
non-conductive materials. The conductive materials of the present
invention may be a metal, a carbon composite, carbon nanotubes,
metal-coated carbon filaments (wherein the metal may be one of the
following metals: nickel, copper, cobalt, silver, gold, tin, or
zinc), metal-coated ceramic filaments (wherein the metal may be one
of the following metals: nickel, copper, cobalt, silver, gold, tin,
or zinc), a composite of metal-coated carbon filaments and a
polymer (wherein the polymer may be one of the following: polyether
sulfone, silicone, polymide, polyvinylidene fluoride, or epoxy), a
composite of metal-coated ceramic filaments and a polymer (wherein
the polymer may be one of the following: polyether sulfone,
silicone, polymide, polyvinylidene fluoride, or epoxy), a composite
of metal-coated carbon filaments and a ceramic (wherein the ceramic
may be one of the following: cement, silicates, phosphates, silicon
carbide, silicon nitride, aluminum nitride, or titanium diboride),
a composite of metal-coated ceramic filaments and a ceramic
(wherein the ceramic may be one of the following: cement,
silicates, phosphates, silicon carbide, silicon nitride, aluminum
nitride, or titanium diboride), or a composite of metal-coated
(carbon or ceramic) filaments (wherein the metal may be one of the
following metals: nickel, copper, cobalt, silver, gold, tin, or
zinc) and a polymer/ceramic combination (wherein the polymer may be
one of the following: polyether sulfone, silicone, polymide,
polyvinylidene fluoride, or epoxy and the ceramic may be one of the
following: cement, silicates, phosphates, silicon carbide, silicon
nitride, aluminum nitride, or titanium diboride).
[0238] In the various embodiments described above with respect to
FIGS. 10 through 31, a common reference line was described in the
context of a unipolar lead system. It is noted that the common
reference line may not be a separate line, but the common reference
line may be one of the leads in a bipolar lead system.
[0239] In summary the present invention, as described above, is an
electromagnetic shield having a first patterned layer having
non-conductive materials and conductive materials and a second
patterned layer having non-conductive materials and conductive
materials. The conductive materials in the first patterned layer
may be randomly located or located in a predetermined segmented
pattern. Moreover, the conductive materials in the first patterned
layer may be located in a predetermined segmented pattern with
respect to locations of the conductive materials in the second
patterned layer. The electromagnetic shield may further include a
non-conductive layer between the first and second patterned layers
such that the conductive materials in the first patterned layer can
overlap the conductive materials in the second patterned layer with
respect to a direction substantially perpendicular to a planar
surface of the second patterned layer.
[0240] Therefore, the present invention provides a patterned
shielding that substantially prevents undesired radiation from
entering the lead from the radial direction or electrical component
and also substantially prevents the formation or inducement of eddy
currents on the surface of the shielding, which are induced by the
pulsing of the radiation or changing magnetic field of the
radiation. More specifically, the conductive material in a
shielding layers of the present invention is patterned or segmented
so that the conductive material is literally broken up to limit the
build up of eddy currents induced by the changing magnetic
fields.
[0241] Moreover, the present invention can reduce specific
absorption ratio by altering the magnetic-resonance imaging
acquisition process by altering the imaging pulse sequence and/or
changing the timing parameters. Furthermore, the radio-frequency
excitation cycle may be interrupted to allow for cooling. Therefore
if a threshold temperature level is exceeded, the imaging
acquisition can be halted until the tissue temperature is at an
acceptable level.
[0242] While various examples and embodiments of the present
invention have been shown and described, it will be appreciated by
those skilled in the art that the spirit and scope of the present
invention are not limited to the specific description and drawings
herein, but extend to various modifications and changes all as set
forth in the following claims.
* * * * *