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BACKGROUND OF THE INVENTION
The present invention relates to field of electrical defibrillation,
including cardioversion, and more particularly to the structure for an
electrode used in implantable defibrillation systems. The term
"defibrillation", as used herein, includes cardioversion which is another
technique involving relatively high energy delivery, as compared to
pacing, as well as other aspects of defibrillation therapy such as the
monitoring of cardiac electrical activity (sensing) when not delivering
high energy impulses.
Defibrillation is a technique employed to counter arrhythmic heart
conditions including some tachycardias, flutter and fibrillation in the
atria and/or ventricles. Typically, electrodes are employed to stimulate
the heart with electrical impulses or shocks, of a magnitude substantially
greater than pulses used in cardiac pacing. One defibrillation approach
involves placing electrically conductive paddle electrodes against the
chest of the patient. During cardiac surgery, such paddles can be placed
directly against the heart to apply the necessary electrical energy.
More recent defibrillation systems include body implantable electrodes.
Such electrodes can be in the form of patches applied directly to
epicardial tissue, or at the distal end regions of intravascular
catheters, inserted into a selected cardiac chamber. U.S. Pat. No.
4,603,705 (Speicher et al), for example, discloses an intravascular
catheter with multiple electrodes, employed either alone or in combination
with an epicardial patch electrode. Compliant epicardial defibrillator
electrodes are disclosed in U.S. Pat. No. 4,567,900 (Moore).
Epicardial electrodes are considered the most efficient, in the sense that
less energy is required for defibrillation as compared to either chest
contact paddles or intravascular catheter electrodes. However epicardial
electrode implantation is highly invasive, major surgery, since it is
necessary to enter the chest cavity, which typically involves spreading of
adjacent ribs or splitting of the sternum. This procedure presents a risk
of infection. Further, implantation and attachment place physical
constraints upon the nature of electrode. These electrodes must be either
quite small, or extremely compliant and resistant to fatigue, as they
maintain conformal fit to the contracting heart.
Generally, larger defibrillation electrodes are considered more desirable,
since they reduce the impedance at or near the electrode. Sensing
artifacts also are reduced for larger electrodes. However, larger
electrodes are difficult to attach to the epicardium, as they must conform
to the heart during the contractions associated with normal cardiac
activity. Subcutaneous electrodes are more easily implanted, at less risk
to the patient. In a defibrillation electrode or any other implanted
device, however, increasing the size generally increases discomfort and
surgical risk to the patient.
Increasing the size of a defibrillation electrode affects its electrical
performance. Conventional electrodes are subject to "edge effects" arising
from the non-uniform distribution of electrical energy when the electrode
receives the pulse. In particular, current densities are greater at the
edges of the electrode than at interior regions of the electrode. An
attempt to counter the edge effect is disclosed in U.S. Pat. No. 4,291,707
(Heilman et al). A series of circular openings, through an insulative
layer framing a conductive screen, are said to substantially eliminate the
edge effect by the additional exposure of the screen. Another problem
encountered in larger electrodes is the resistance across the length
(largest linear dimension) of the electrode, leading to unwanted voltage
gradients across the electrode which can degrade electrode performance.
Therefore, it is an object of the present invention to provide an
implantable defibrillation electrode with a large effective surface area
to lower the impedance at or near the electrode, without causing undue
patient discomfort.
Another object is to provide a defibrillation electrode that has a large
effective area, yet is easier to implant and readily conforms to the
contours of its implant location.
A further object is to provide a defibrillation electrode structure
enabling a relatively large size while reducing the non-uniform field
distribution associated with conventional electrodes.
Yet another object is to provide defibrillation electrodes of sufficient
size and effectiveness to enable transthoracic delivery of defibrillation
pulses, with an implanted system.
SUMMARY OF THE INVENTION
To achieve these and other objects, there is provided a body implantable
tissue stimulating electrode. The electrode includes a plurality of
flexible, electrically conductive electrode segments having a nominal
width and a length at least five times the nominal width. A means is
provided for mechanically coupling the electrode segments with respect to
one another whereby each of the segments, over the majority of its length,
is spaced apart from each one of the other segments by a distance of at
least 1.5 cm. A means is provided for electrically coupling the electrode
segments for substantially simultaneous reception of the tissue
stimulating electrical pulses from a pulse generating means. Consequently
the electrode segments, when receiving the tissue stimulating pulses,
cooperate to define an effective electrode area incorporating the
electrode segments and having a width of at least 1.5 cm.
In one preferred configuration, the electrode segments are linear and in
parallel spaced apart relation, all extending in a longitudinal direction.
The mechanical and electrical coupling means can be a transversely
extended distal portion of an elongate, electrically conductive lead. The
lead is connected to each of the respective first end portions of the
electrode segments along its distal region, and connected at its proximal
end to a pulse generating means. Preferably an electrically insulative
layer covers the lead, leaving the electrode segments exposed, to define a
substantially rectangular "phantom" area or effective electrode area.
Alternatively, the electrode segments can radiate outwardly from a common
junction, typically at the distal end of the lead or conductive coupling
wire from the pulse generating means. While the coupling wire is covered
with an insulative material over the majority of its length, a distal end
portion of the coupling wire can be left exposed, to provide one of the
electrode segments.
Yet another approach involves a single electrically conductive wire or
path, with portions of the path providing the spaced apart segments. As an
example, the path can be arranged in a serpentine configuration in which
segments are parallel to and aligned with one another, side by side.
Alternatively, the conductive path is formed as a spiral. In either event,
adjacent segments are spaced apart from one another a distance
substantially greater than their width, preferably by an order of
magnitude or more.
In a preferred example, elongate electrode segments about 30 cm long and
with a nominal width of 0.5 mm extend longitudinally, aligned with one
another and spaced apart from one another by about 3 cm. One end of each
electrode segment is mounted to the distal end portion of a conductive
lead to a pulse generator. At the opposite, free end of each segment is an
enlargement such as a loop or flared end, formed to minimize local high
current densities due to the previously described edge effects. The
combination of a large phantom area with multiple conductive segments
reduces non-uniform current distributions.
The best results are achieved with highly conductive electrode segments.
Accordingly, the segments are preferably formed of low resistance
composite conductors including drawn braised strands (DBS), drawn filled
tubes (DFT) and the like, coated with platinum or another metal from the
platinum group, e.g. iridium, ruthenium or palladium, or alternatively
with an alloy of one of these metals. The strands can be formed of
titanium or platinum. A suitable filled tubular conductor is composed of a
silver core within a stainless steel tube. The electrode segments can be
formed of single wires, pluralities of wires in a braided or twisted
configuration, helically wound coils, or a woven mesh or screen. In some
embodiments, particularly those employing the woven screen, it is further
desirable to include an insulative backing to more positively position the
electrode segments with respect to one another.
It has been found that highly conductive electrode segments reduce any
voltage gradient across the electrode, with the separate segments
simultaneously receiving a defibrillation or other stimulation pulse. The
separate segments thus cooperate to act as a single "patch" electrode,
having an effective surface area equal to that of a rectangle or other
polygon containing all of the segments. As an example, an electrode formed
as a row of five parallel electrode segments spaced apart from one another
by 3 cm, each segment being 10 cm long, would have a rectangular phantom
or effective area slightly larger than 120 (twelve times ten) square cm.
Yet, as compared to a continuous rectangular patch electrode measuring ten
by twelve cm, the branched segment electrode in accordance with the
present invention is easier to implant, reduces the high current density
regions, and more easily conforms to the thorax or other surface to which
it is attached. In fact, branched arrangements of segments can provide
effective defibrillation electrode areas in the range of from 100 to 200
square cm, while enabling easy implantation.
Thus, in accordance with the present invention there is disclosed a process
for applying defibrillation pulses to a human heart, including the
following steps:
(a) implanting a first compliant electrode in a patient, proximate the
pleural cavity and the rib cage, and on a first side of the thoracic
region of the body;
(b) implanting a second compliant electrode in the body, proximate the
pleural cavity, and the rib cage, and on a second side of the thoracic
region opposite the first side, with at least a portion of the heart
between the first and second electrodes;
(c) implanting a defibrillation pulse generator; and
(d) electrically coupling the first and second electrodes to a
defibrillation pulse generator and providing defibrillation pulses from
the pulse generator across the first and second electrodes.
If desired, one or more electrodes implanted proximate the pleural cavity
and rib cage can be used in combination with one or more coil electrodes
mounted on an intravascular catheter, preferably positioned in the right
atrium and the right ventricle of the heart, with the distal end of the
catheter near the apex of the right ventricle.
As compared to the entry into the chest cavity normally associated with
implanting epicardial electrodes, transthoracic placement of subcutaneous
electrodes as outlined above is substantially less invasive, preserves the
integrity of the rib cage and the pleural cavity, and reduces risk of
infection.
Nonetheless, other implant locations, including direct attachment to
epicardial tissue, can be employed in accordance with the present
invention, to achieve relatively large effective electrode areas while
maintaining patient comfort with substantially more uniform distribution
current density.
IN THE DRAWINGS
For a further understanding of the above and other features and advantages,
reference is made to the detailed description and to the drawings, in
which:
FIG. 1 is a top plan view of a defibrillation electrode constructed in
accordance with the present invention;
FIG. 2 is a sectional view taken along the line 2--2 in FIG. 1;
FIG. 3 is a sectional view taken along the line 3--3 in FIG. 1;
FIG. 4 is a top plan view of an alternative embodiment electrode
constructed in accordance with the present invention;
FIGS. 5-9 illustrate alternative constructions for electrode segments of
the electrodes;
FIG. 10 is plan view of another alternative embodiment electrode
constructed in accordance with the present invention;
FIGS. 11-13 illustrate further alternative configurations of the electrode
of FIG. 9;
FIG. 14 is a top plan view of another alternative embodiment electrode;
FIGS. 15, 16 and 17 illustrate a further embodiment electrode;
FIG. 18 is a top plan view of yet another embodiment electrode;
FIG. 19 is a schematic representation of the electrical field between a
continuous patch electrode and an electrode having segments, but in which
the segments are too close to one another;
FIG. 20 is a schematic representation of the electrical field between two
electrodes constructed according to the present invention;
FIG. 21 is a plot of intraelectrode impedance as a function of the spacing
between adjacent segments of each of the electrodes, for electrodes with
from two to four segments; and
FIGS. 22, 23 and 24 diagrammatically illustrate alternative implantation
approaches for defibrillation systems incorporating electrodes embodying
the present invention.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
Turning now to the drawings, there is shown in FIG. 1 a defibrillation
electrode 16 including three parallel and spaced apart electrode segments
18, 20 and 22. Each of the segments has a length (L in the figure)
substantially longer than its width (W), e.g. 30 cm long with a nominal
width preferably about 0.5 mm. Generally, the width should be within the
range of from 0.25-5 mm. Adjacent segments are spaced apart a distance (D)
substantially greater than the nominal width, e.g. 3 cm. This
center-to-center spacing should be at least 1.5 cm, and preferably does
not exceed 30 cm.
Electrode segments 18, 20 and 22 are fixed at respective first ends to a
distal end portion 24 of an electrically conductive lead 26. The lead
conducts electrical pulses to the electrode segments from a pulse
generator (not shown) coupled to the proximal end of the lead. Lead 26 at
the distal end structurally supports the longitudinally extended electrode
segments in the transversely spaced apart configuration shown.
The electrically conductive portion of lead 26 is surrounded by an
electrically insulative cover or sheath 28, preferably constructed of a
body compatible polymer, e.g. a medical grade silicone rubber or
polyurethane. As seen in FIG. 2, the lead includes a composite conductor
formed of a core 30 of silver surrounded by a tube 32 of stainless steel.
This type of composite conductor is known as drawn field tube (DFT) of
MP35N (brandname) alloy available from FWM Research Products of Fort
Wayne, Ind. Further, a coating 34 of platinum is applied over the
stainless steel, preferably by sputtering or other deposition process.
While preferably platinum, coating 34 also can consist of another metal
from the platinum group (e.g. iridium, ruthenium and palladium) or an
alloy of these metals. Insulative sheath 28 is contiguous with and
surrounds the platinum layer.
As seen in FIG. 3, the construction of electrode segment 22 (and likewise
segments 18 and 20) over substantially all of its length is substantially
similar to the construction of the conductive portion of lead 26. Thus the
segments also are highly electrically conductive. Platinum coating 34
provides a further advantage for the segments, which are not covered by
the insulative sheath. In particular, the platinum coating when applied by
vapor deposition provides a microtexture which substantially increases the
reactive surface area of the electrode segments, to reduce near field
impedance of the electrode (the term "near field" impedance refers to the
voltage losses associated with the electrode due to chemical and field
effects). For a further discussion of this feature, reference is made to
U.S. Pat. No. 5,074,313, and assigned to the assignee of the present
application. The reduced interface impedance increases the ratio of bulk
impedance to the total system impedance as measured between the
stimulating electrode and the indifferent or signal return electrode.
Thus, more of the voltage drop occurs across tissue, where it is useful
for causing the desired stimulation, with proportionately less of the
voltage drop occurring at the electrodes where it is non-productive. This
enables a reduction in overall potential or pulse duration, in either
event reducing the required energy for defibrillation.
Given adequate separation between segments 20, 22 and 24, the current
distribution is made more uniform. To further counter any current density
differentials due to edge effects at the ends of segments 20, 22 and 24,
loops 36, 38 and 40 are formed at these ends, respectively. Alternatively,
the ends can be flared or otherwise enlarged, and remain substantially
free of undesirable concentrations of high current. Such enlargements also
facilitate implant, as they tend to positionally fix the electrode
segments.
Because the electrode segments are electrically common, the electrodes
receive and transmit defibrillation pulses simultaneously. The electrode
segments are sufficiently near one another to function in concert,
providing an effective area or phantom area incorporating the segments, as
indicated in broken lines at 42. In other words, electrode segments 20, 22
and 24 define a generally rectangular effective area, with substantially
greater compliance to contours and movements of body tissue, as compared
to a continuous patch electrode. In addition, the spacing between
electrodes performs an important electrical function by producing a
substantially more uniform current distribution than that of a continuous
patch electrode. Patch electrodes are known to have regions of very high
current density around their outside edges, and regions of low current
density at their centers. By using a segmented electrode, with segments
properly spaced apart from one another, much higher currents can be
delivered to the central region of the effective or phantom area because
current is able to flow between adjacent segments. This results in a more
uniform electrical field across the heart.
FIG. 4 illustrates an alternative embodiment defibrillation electrode 44
including five elongate electrode segments 46, 48, 50, 52 and 54, each
with a preferred width and substantially greater preferred length as
described in connection with electrode 16. Each of electrode segments
46-54 is part of a wire mesh pattern 55 and extends longitudinally.
Transversely extended end portions 56 and 57 of the pattern couple the
segments to a lead 58. An insulative sheath 62 surrounds lead 58 from
electrode 44 to the proximal end of the lead. An electrically insulative
backing 64 supports mesh pattern 55. The mesh pattern is covered by an
insulative layer 66. Slots 68, 70, 72 and 74 are formed in backing 64 and
layer 66 between adjacent electrode segments.
FIG. 5 illustrates an alternative form of composite conductor known as DBS
(drawn braised strand), available from FWM Research Products, Fort Wayne,
Ind. As shown, a silver core 73 is surrounded by six stainless steel wires
75. The structure is heated and drawn to braise all wires together. The
results is a solid, continuous composite conductor composed of a silver
core and a stainless steel outer shell or tube.
FIG. 6 illustrates an alternative construction for the electrode segments
of either electrode 16 or electrode 44, involving a plurality of composite
conductors 76 in a twisted configuration. Each of the conductors can
include a silver core within a stainless steel tube coated with platinum
as previously described. Alternative composite conductors for single and
multiple wire arrangements include platinum or titanium ribbon or wire,
clad with platinum. The twisted construction enhances flexibility and
resistance to fatigue in the electrode segments. Other alternatives
include braided or knitted wires.
FIG. 7 shows another alternative construction for the electrode segments,
in the form of a woven mesh or screen 78 on an electrically insulative
backing 80. This type of electrode segment construction is particularly
well suited for epicardial positioning, e.g. with electrode 44 in FIG. 4.
Another alternative segment construction, shown in FIGS. 8 and 9, involves
a flexible, electrically insulative cylindrical core 82 of polyurethane,
medical grade silicone rubber, or other suitable body compatible material.
Core 82 is surrounded by an electrically conductive coil winding 84,
preferably a wire or composite cable such as illustrated in FIG. 2. The
helically wound coil conductor provides the greatest flexibility and
fatigue resistance of any of the arrangements discussed, and for this
reason is preferred in the case of direct epicardial attachment, or any
other implant location in which the lead segments are subject to continued
or repeated muscular contraction or other abrupt tissue movements. A
disadvantage, relative to other embodiments, is that a helical coil
electrode segment, as compared to other segments of equal length, involves
a substantially longer conductive path with less tensile strength.
All of the alternative constructions provide electrode segments which are
highly compliant, first in the sense that they readily adjust to the
contours of body tissue at the implant site when they are implanted, and
secondly over the long term, in continually conforming to the tissue
during muscular contractions and other tissue movement.
FIG. 10 illustrates a further embodiment defibrillation electrode 86
including electrode segments 88, 90 and 92 formed as branches, radiating
or extended outwardly from a common junction and stress relief area 94.
Junction 94 is positioned at the distal tip region of a lead 96 to a pulse
generator (not shown), and includes a conductive portion surrounded by an
insulative sheath 98. The conductive region of the lead and the electrode
segments can be constructed as previously described.
The stress relief portion of the electrode is electrically insulative and
covers portions of the segments, leaving exposed portions of the segments
spaced apart from one another and defining an effective or phantom area
100 shown by the broken line. As before, segments 88-92 have a nominal
width preferably about 0.5 mm, and are longer than they are wide, for
example by at least a factor of five. At the free ends of the segments are
respective masses or bodies 102, 104 and 106. The bodies are constructed
of an electrically conductive, plastically deformable material such as
platinum or gold and, as seen in FIG. 10, include slots 108 slightly wider
than the thickness of segments 88-92. Each body is applied to the free end
of its respective electrode segment by inserting the free end within the
respective slot and pinching the body to frictionally secure the body to
the electrode segment. Bodies 102-106 thus provide enlargements at the
free ends of the segments to reduce the chance for high current densities
at the free ends, and provide a means of fixation of the free ends.
FIGS. 11-13 schematically illustrate alternative configurations for
electrode 86. More particularly, FIG. 11 illustrates a clamp 110 for
electrically and mechanically coupling two intersecting cables 112 and
114. Cable 112 is part of lead 96, with a distal portion of the lead
providing center segment 90. Electrode segments 88 and 92 are opposite
portions of cable 114. An extension 116 of electrically insulative sheath
98 covers clamp 110 and portions of cables 112 and 114, leaving the
segments exposed.
In FIG. 12, segments 88, 90 and 92 extend radially from a crimping member
118 at the distal end of lead 96. Alternatively, segment 90 is the distal
end of the lead, in which case the remainder of the lead, crimping member
118 and portions of the electrode segments are provided with an insulative
covering 119.
In FIG. 13, crimping member 118 secures electrode segments 88, 90 and 92 to
the distal section 120 of lead 96. Insulative sheath 98 leaves distal
section 120 exposed, so that it functions as a fourth electrode segment.
FIG. 14 shows a further embodiment defibrillation electrode 122 including a
lead 124 havin | | |
