Low resistance implantable electrical leads

The present invention provides an implantable electrical lead having relatively low electrical resistance. The lead comprises: a wire core formed in a helical coil having pre-compression, and having distal and proximal ends; a layer of an electrically conductive material formed around the wire core such that there is electrical continuity between the wire core and the metal layer; a biocompatible, electrically insulating sheath covering the wire core; a first lead connector electrically connected to the proximal end of the wire core; and an electrode electrically connected to the distal end of the wire core. The wire core may have various cross-sectional configurations which increase the current conducting area of the wire core without increasing its outside diameter. In another embodiment, the electrical lead includes a wire core having a cross-sectional area which differs over the length of the core to enhance the fatigue resistance of the electrical lead. In yet another embodiment, the wire core may be wound in a helix having different pitches in different sections of the core.

BACKGROUND OF THE INVENTION 
The present invention relates to the field of electrical leads suitable for 
being implanted within living tissues, and more particularly to 
implantable electrical leads having relatively low electrical resistance 
which may be used in conjunction with cardiac pulse generators, neural 
stimulators, implantable sensors, and the like. 
An implantable cardiac pulse generator, referred to generically as a 
cardiac pacer, or pacemaker, is a small, sealed electronic pulse generator 
that is used to treat irregular heart rhythms. In general, such pacers 
provide minute electrical stimuli to a heart when needed to speed up 
unnaturally slow heart rates. 
An implantable cardiac defibrillator is a moderately sized, electronic 
pulse generator that is used to treat patients that are at risk from 
suffering lethal arrhythmias, most notably ventricular fibrillation. 
Ventricular fibrillation is a heart rhythm that typically results in death 
within several minutes. The defibrillator is used to provide large 
electrical stimuli when needed to interrupt the lethal arrhythmia and 
re-establish a life sustaining heart rhythm. Such pulse generators are 
typically packaged in sealed containers that are usually implanted 
subcutaneously in the thorax or abdomen of the heart patient. These 
devices monitor cardiac activity and deliver electrical pulses of 
appropriate intensity whenever needed. The energy supplied by a pulse 
generator is conducted along an electrically conductive cardiac lead from 
the pulse generator directly to the heart. 
The pulse generator is commonly powered by a battery located inside the 
sealed container which is not intended to be replaceable. The amount of 
electrical energy stored in the battery generally determines the 
operational life of the pulse generator. Although the battery is very 
efficient at storing electrical energy, the battery life, and hence the 
operational life of the pulse generator, is usually less than ten years. 
The battery depletion is in part due to energy delivered to the heart, to 
energy consumed by the resistance of the electronic circuitry of the pulse 
generator and cardiac leads, and to self-discharge of the battery over 
time. 
Each time an electrical pulse is delivered to the heart, some of the energy 
output of the battery is consumed by the cardiac leads as I.sup.2 R heat, 
where "I" represents the current through the cardiac lead, and "R" 
represents the electrical resistance of the cardiac lead. The I.sup.2 R 
losses represent wasted energy which provides no useful purpose. In an 
attempt to maximize the service life of the pulse generator, the lead 
materials and the geometries of the lead materials are chosen to minimize 
the electrical resistance of the cardiac leads while providing a lead that 
can withstand the rigors of exposure to repetitive stress. 
The lead which electrically connects the pulse generator to the heart may 
be attached to the inner surface of the heart, the endocardium, or to the 
outer surface of the heart, the epicardium. Regardless of where the lead 
is attached to the heart, the lead is mechanically flexed with every heart 
beat. Every flexure of the lead creates stress within it. Since a typical 
heart rate is 60 beats per minute, the heart beats millions of times in a 
single year, and the lead is stressed with each beat. 
Unlike a pulse generator, which is replaced when the battery is depleted, 
the lead is not normally replaced. A youthful patient who receives an 
implant may hopefully use the same lead or leads for decades. For this 
reason the materials comprising the lead should have excellent mechanical 
fatigue resistance. 
Materials known as having low electrical resistance, such as copper or 
silver, are not well suited for use as a conductor in a cardiac lead 
because they cause tissue reactions and readily form oxides which may 
ultimately result in the fracture of a lead constructed with such 
material. A further disadvantage of copper and silver is that they have 
very poor resistance to repeated stress. Core copper wire, multi-stranded 
copper wire and even tinsel copper wire would poorly withstand the 
repeated, reversing stresses the heart would impose on cardiac leads 
comprised of copper wire. Leads of such constructions tend to fail after 
only a relatively small number of flexions, much as a paper clip breaks 
after being bent a few times. 
Spring materials made of non-oxidizing, corrosion resistant alloys having 
good fatigue resistance perform much better mechanically as cardiac lead 
conductors than do conductors comprised of copper, silver, or their 
alloys. That is why for decades lead manufacturers have been using spring 
materials for the conductors in their leads. Examples of suitable 
conducting materials include stainless steel, such as Elgin Wire Co., 
Elgiloy, MP-35N, and titanium and titanium alloys. 
Generally, in the construction of cardiac leads, the conducting wire core 
is coiled to form a tight helix composed of many individual coils, similar 
to an extension spring. The helical construction greatly lowers the 
mechanical stress to which the material comprising the wire core would 
otherwise be subjected by the beating heart. Though this construction 
provides long lasting leads, the electrical resistance of leads 
manufactured of spring steel or titanium is relatively high primarily due 
to the resistivity of the material comprising the wire core. 
Low resistance leads are important both for pacing and defibrillation. DBS 
(drawn brazed strands) and DFT (drawn filled tubing) provide a cardiac 
lead having both reasonable fatigue resistance and electrical resistance. 
DBS and DFT are examples of two structures which combine low resistance, 
poor mechanical materials with the high resistance, spring materials of 
conventional leads. An example of DBS wire includes six strands of wire 
made of MP-35N that are brazed by a central silver core. An example of 
drawn filled tubing may include tubing fabricated of MP-35N and which is 
filled with silver. The silver, copper or other electrical conductor 
significantly reduces the lead resistance, but the silver or copper 
included in these leads also present several drawbacks. Chief among them 
are the toxicity and low fatigue life of copper and silver. There are 
several materials, as previously mentioned that are well suited to be 
pacing or defibrillation lead conductors. However, these materials 
unfortunately have high resistivities of about 100 micro-ohm-cm, as 
opposed to resistivities of about 1 micro-ohm-cm for copper or silver. 
The electrical resistance of a typical cardiac lead is about 100 ohms. 
Cardiac leads with larger diameter conductors, high helix pitches (a 
measure of the number of coils per unit of axial length of the lead), 
multiple conductors, or smaller helical internal diameters, may have 
electrical resistances of about 10 ohms. However, such leads also tend to 
have reduced flex life. 
A goal in the field of cardiac lead technology is to provide a cardiac lead 
having an electrical resistance of less than 1 ohm. Low resistance cardiac 
leads would provide a pulse generator with an increased service life 
because such leads would reduce I.sup.2 R losses. Thus, more energy stored 
in the battery which powers the pulse generator would be able to be 
delivered to the heart because less energy would be wasted as heat. 
Therefore, it may be appreciated that there is a need for a cardiac lead 
having low electrical resistance, as well as good resistance to the 
repeated stresses to which a cardiac lead is exposed. 
SUMMARY 
The present invention provides an implantable electrical lead having 
relatively low electrical resistance and good mechanical resistance to 
cyclical stresses. The lead comprises a wire core formed into a helical 
coil having pre-compression, and having distal and proximal ends; a layer 
of an electrically conductive material formed around the wire core such 
that there is electrical continuity between the wire core and the 
electrically conductive layer; a biocompatible, electrically insulating 
sheath covering the helical coil; a first lead connector electrically 
connected to the proximal end of the wire core; and an electrode 
electrically connected to the distal end of the wire core. The term 
"biocompatible" refers to a material that is compatible with living 
tissue. Such material does not substantially react with the tissue nor 
cause inflammation or infection. Furthermore, biocompatible material is 
not rejected by living tissue. 
The wire core may have an elongated cross-sectional area which is longer in 
the axial direction of the lead than in the radial direction. The 
elongated area provides a larger cross-sectional area having less 
resistance to electrical current flow per unit length through the core 
than would a wire core having a circular cross-sectional area with a 
diameter of the same length as the radial dimension of the elongated area. 
In another embodiment of the invention, the electrical lead includes a wire 
core having a cross-sectional area which gradually tapers, or varies, over 
the length of the wire core from a small cross-sectional area to a larger 
cross-sectional area. One benefit of a such a varying cross-sectional area 
is that the section of the core having the smaller cross-sectional area is 
capable of withstanding the high repetitive loads from the beating heart 
while the section of the lead not subject to such loads has a larger 
cross-sectional area, and therefore, relatively less electrical resistance 
per unit length. 
In yet another embodiment, the wire core may be wound in a helix having 
different pitches in different sections of the core. Sections of the lead 
having a high helix pitch are more resistant to repeated loads from the 
heart. Sections of the lead having a low helix pitch have less electrical 
resistance per unit length of the lead. 
In a further embodiment, the electrical lead of the present invention 
includes a wire core wound into a helix having individual coils with 
pre-compression. The pre-compression assures that the coils contact one 
another at the interfaces between the coils. The two coils at every mth 
interface are welded together to promote axial current flow through those 
coil pairs, where m represents a positive integer. Such welding reduces 
the electrical resistance of the electrical lead.

DESCRIPTION OF THE PREFERRED EMBODIMENTS 
The following description is of the best mode presently contemplated for 
carrying out the invention. This description is not to be taken in a 
limiting sense, but is made for the purpose of describing the general 
principles of the invention. The scope of the invention should be 
determined with reference to the claims. 
Conventional cardiac pacing and defibrillation leads commonly employ 
corrosion resistant spring metal, such as spring steel, Elgiloy, MP-35N, 
titanium or titanium alloy, as a conducting material because such material 
has excellent resistance to metal fatigue. A wire core made of the 
conducting material is wound into a helix to form an electrical lead 
comprising a series of individual coils. The types of materials which 
commonly comprise the wire core generally exhibit thin metal oxide layers 
having high electrical resistance which impede the axial flow of current 
through the lead, thereby limiting the current to flow in a helical path 
through the cross-sectional area of the wire core. Although it is possible 
to remove the oxide, it is difficult to prevent the oxide from reforming 
in-vivo. 
The present invention provides an electrical lead having a much lower 
electrical resistance than do conventional leads, as well as good 
resistance to systematic stress. In general, an electrical lead 12 
embodying various features of the present invention is shown in FIG. 2 to 
include a wire core 13 coated with an electrically insulating, 
biocompatible sheath 14. The wire core is wound into a helix having a 
helix diameter, D.sub.H, which is to be distinguished from the diameter of 
the circular cross-section of the wire core. A lead connector 15 is 
electrically connected to the proximal end of the end of the wire core 13 
and may, for example, be designed to connect the lead to the output 
connector of a pacemaker or defibrillator. By way of example, the lead 
connector 15 may be a VS-1 bipolar or unipolar pacemaker lead connector, 
as set forth in Calfree, R.V., and Saulson, S.H., "A Voluntary Standard 
For 3.2 MM Unipolar And Bipolar Pacemaker Leads And Connectors," E, 
Vol. 9, Part II, November-December 1986. An electrode 16 for electrically 
coupling the electrical lead to a heart is electrically connected to the 
distal end of the wire core 13. There are many types of electrodes which 
may be employed in conjunction with the present invention. Such electrodes 
may include ring electrodes, hemispherical tip electrodes, and such 
electrodes as are described in U.S. Pat. Nos. 3,974,834, 4,643,201, and 
4,998,975, incorporated herein by reference. 
Although the electrical lead of the present invention is generally 
described below with reference to applications involving cardiac pulse 
generators, it is to be understood that the electrical leads of the 
invention may also be used in other applications in which it is desirable 
to implant low resistance electrical leads within a living body, as for 
example, in that of a mammal, and more particularly, in a human. For 
example, electrical leads of the present invention may be used in 
conjunction with neural stimulators or sensors implanted within a living 
body. 
More specifically, a short length of an electrical lead 20 embodying 
various features of the present invention is shown in cross-section in 
FIG. 3. Referring now to FIG. 3, lead 20 is comprised of a wire core 22 
wound in a helix and comprised of a material having good resistance to 
mechanical fatigue. By way of example, the wire core may be made from 
non-corrosive, bioresistant spring metal, such as spring steel, Elgiloy, 
MP-35N, titanium or titanium alloy, having a diameter which may be 0.1 mm 
and which may be coated with a biocompatible, electrically conductive 
layer 24 having a thickness of about 1 micron. The conductive layer may be 
composed of electrically conductive noble metals or their alloys, 
electrically conductive polymers, or carbon. The conductive layer 24 need 
only be sufficiently thick so as to not be abraded through to the 
underlying wire core 22 in response to any motion of the lead 20. An 
electrically insulating, tubular sheath 26 is placed over the helical coil 
formed of the coated, helically wound wire core 22 to electrically insulate 
the helical coil from surrounding tissues. The sheath 26 is preferably 
fabricated from a biocompatible, electrically insulating elastomeric 
material, such as polyurethane or silicone having a wall thickness of 
about 0.4 mm. 
Assuming electrical current flow through the lead 20 is in the direction of 
the arrow 21, then axial current flow between successive coils 25 is 
represented by the arrows 27 shown in FIG. 3. Current also flows in a 
helical path through the wire core 22 such that the symbols ".sym." 
represent current flow in a direction into the FIG. 3, and the symbols " " 
represent helical current flow through the cross-sectional area of the wire 
core 22 in a direction coming out of FIG. 3. Current flows serially from 
one coil 25 to an adjacent coil 25 through the electrically conductive 
layer 24 surrounding the wire core 22. Current flow through the 
electrically conductive layer 24 is parallel to the current flowing in the 
helical path through the wire 22. Therefore, a thin layer of the 
electrically conductive layer 24 achieves a significant reduction in the 
electrical resistance of the lead 20. The conductive coating 24 should be 
composed of a material which remains conductive on its surface even after 
decades of being implanted within the body. Such conductive materials 
include the noble metals, conductive polymers, and conductive carbon. An 
example of a particularly suitable noble metal is platinum. 
The electrical resistance of a material is directly related to the product 
of the intrinsic resistivity of the material multiplied by the length of 
the current path through the material, and is inversely related to the 
cross-sectional area of the material. The electrical lead 20 achieves the 
beneficial result of low electrical resistance by effectively reducing the 
length of the circuit path through the wire core 22. This result is 
achieved by allowing electrical current to flow from coil to coil, axially 
along the length of the lead, as well as in a helical path through the 
cross-sectional area of the wire core 22. 
If local yielding or cracking of the electrically conductive layer 24 
occurs at one location along the lead 20 with no effect on the underlying 
wire core 22, electrical current still conducts from coil to coil until a 
crack is encountered. Then, the electrical current simply conducts around 
such crack through the electrically conductive layer 24 and/or through the 
core material to the adjacent coil. 
Platinum or platinum alloys are preferred materials for the electrically 
conductive layer 24 because they are excellent electrical conductors, have 
excellent immunity to oxidation, and are biocompatible. However, other 
metals such as gold or gold alloys may also be employed to comprise the 
conducting layer 24. If such metals are not biocompatible, then in 
addition to electrically insulating the wire core from the surrounding 
tissues, the sheath 26 also provides the important function of preventing 
contact between the electrically conductive layer 24 and the surrounding 
tissues, and/or bodily fluids. 
In the case where the electrically conductive layer 24 is a metal, the 
metal layer may be formed on the wire core 22 by any of a number of well 
known manufacturing methods, e.g., sputtering, plating, condensation, etc. 
Generally, the electrically conductive layer 24, whether formed of metal, 
polymers, or carbon, should be formed on the wire core 22 before the wire 
core 22 is wound into a helical coil in order to ensure that the wire core 
is completely coated with the conductive layer. Alternatively, the 
electrically conductive layer 24 may be applied to the wire core 22 after 
the wire core 22 is wound into the helical coil while the coil is 
maintained elastically separated. After the electrically conductive layer 
24 is applied, the coil is allowed to return to its compressed condition 
for use in the lead. Any electrically insulating oxides or films that may 
be present on the surface of the wire core 22 must be substantially 
removed before forming the electrically conductive layer 24 on the wire 
core 22 in order to provide good electrical continuity between the wire 
core and the electrically conductive layer 24. An example of one technique 
for removing such oxides or films from the surface of the wire core 22 
involves etching the wire core with an acid in an inert atmosphere, such 
as argon or nitrogen, before forming the conductive material thereon. The 
removal of such films and oxides also promotes mechanical adhesion between 
the conductive layer 24 and the surface of the wire core 22. 
In the preferred embodiment, the wire core 22 should be wound into a helix 
having pre-compression between the coils. Such pre-compression increases 
the contact area and provides good electrical contact between the 
individual coils 25 comprising the lead 20. 
A conventional cardiac lead is constructed to have a uniform 
cross-sectional area over most of its length which can bear the cyclical 
stresses to which a portion of it is subjected. Despite the relatively 
long length of a typical cardiac lead, there are only a limited number of 
sections along the length of the lead which are subject to fatigue 
failure. One of these sections is that part of the lead in and near the 
heart itself. This section experiences millions of reversing stresses 
generated by the beating heart. Another section of the lead that 
experiences stress is near the pulse generator where the lead wraps around 
the implanted pulse generator. In the region of the pulse generator, 
however, skeletal motion causes the lead to flex much less frequently than 
the lead flexes near the heart. 
The electrical resistance of the electrical lead 20 shown in FIG. 3 may be 
estimated in accordance with the following analysis. Assuming that an 
electrically conducting, biocompatible electrically conductive layer 24 
has been formed on the wire core 22, the contact area between the 
individual coils comprising the helically wound wire core may be 
approximated using formulae empirically derived for materials in which the 
stress in an elastic body is proportional to the strain to which the body 
is subjected by an applied load (i.e., the material behaves in accordance 
with Hooke'law). Further assuming that the wire core 22 is wound into a 
helix with a pre-compression resulting from, say for example, a 0.001 inch 
coil overlay, the initial compression between individual coils is about 
0.16 pounds, resulting in a region of contact about 0.016 micro-inches 
wide over a helix length of the wire core. 
The total electrical resistance of an implantable electrical lead 20 as 
embodied by the present invention may be estimated by calculating the 
resistances of a single coil and of the contact area between adjacent 
coils, and then adding these resistances for each coil and contact area 
comprising the lead. For a lead coil made with 0.020 inch internal 
diameter winding of a 0.005 inch diameter MP-35N wire coated with about a 
1 micron thick layer of platinum, and assuming that all current conducts 
to neighboring coils only through the contact regions, yields a predicted 
electrical resistance value of 4.2 ohms per axial meter of the lead. Much 
of the current will flow through the platinum, never entering the MP-35N, 
however, not all the current conducts from coil to coil. For a unifilar 
lead, less than 1% of the current is conducted in a helical path through 
the MP-35N cross-sectional area of the wire core in parallel with the 
axial flowing current. A portion of the current conducts through the 
platinum along the inside diameter of the helix which provides an 
electrical resistance contribution of about 6.6 ohms/m. Similarly, 
conduction along the outer surface of the helix provides an electrical 
resistance of about 8.8 ohms/m. These three parallel resistances combine 
to provide the lead 20 with an electrical resistance, R, where 
EQU 1/R=1/4.2+1/6.6+1/8.8. 
Therefore, an estimate of the total resistance, R, of the electrical lead 
20 is about 2 ohms per meter. For a typical lead length, the total 
resistance is approximately 1 ohm. 
Another embodiment of the present invention provides an electrical lead 
having reduced electrical resistance which is achieved by providing the 
wire core with a cross-sectional area which varies over the length of the 
lead. Such a lead may include a wire core having a tapered cross-sectional 
area to provide a large cross-sectional area in sections of the lead 
subject to low repetitive stresses, and a small cross-sectional area in 
sections of the lead subject to the severe repetitive stresses. The 
sections of the lead in which the wire core has the large cross-sectional 
area have relatively low electrical resistance per unit length through the 
helically wound wire core. However, these sections have limited flexibility 
making them less tolerant of repetitive, cyclical stress than the sections 
of the wire core having the smaller cross-sectional area. Therefore, a 
wire core can be configured with sections having cross-sectional areas 
well suited to the particular environment in which those sections are to 
be situated. 
Referring now to FIG. 4, an electrical lead 40 embodying various features 
of the present invention is shown to include a helically wound wire core 
42 comprised of material having good mechanical fatigue resistance, such 
as MP-35N, coated with a biocompatible, electrically insulating 
elastomeric sheath 43. The sheath 43 may for example, be comprised of a 
0.4 mm thick layer of pellethane. The wire core 42 maintains a uniformly 
round outside helix diameter, "d.sub.o ", over its entire length. However, 
the diameter, "d.sub.1," of the wire core 42 itself over its length, 
providing the lead 40 with variable stiffness appropriate for the types of 
forces to which particular sections of the lead are subjected. The wire 
core 42 is designed to minimize mechanical stress in the sections where 
the lead 40 is subject to high repetitive mechanical loads, as for 
example, near the heart, and to minimize electrical resistance in all 
other sections of the lead. It should be noted that the taper of the wire 
core may be precisely controlled so as to be as gradual as desired. 
Generally, the more gradual the taper, the less stress concentration there 
is in the tapered section of the wire core. 
The sections of the wire core 42 near the proximal end 44 of the lead 40 
preferably have a relatively large cross-sectional area to minimize 
electrical resistance in that section of the lead. However, the large 
cross-sectional area of the wire core 42 in that section of the lead 
provides less resistance to mechanical fatigue than would a smaller 
diameter wire core. The proximal end 44 is commonly electrically connected 
to a lead connector, as for example, the lead connector 15 of FIG. 2. The 
wire core 42 at the distal end 46 of the lead 40 has a relatively small 
diameter, providing the wire core with good flexibility and resistance to 
mechanical fatigue in that section. However, disadvantageously, a small 
diameter wire has higher electrical resistance per unit length than would 
larger diameter wire of the same material. Therefore, the length of the 
section having the smaller cross-sectional area is made as short as 
possible in order to minimize the electrical resistance of the lead, yet 
long enough to provide the lead with sufficient resistance to withstand 
the stresses imposed in the lead by the beating heart. The distal end 46 
of the lead 40 is generally connected to an electrode, such as the 
electrode 16 of FIG. 2. 
The wire core 42 of FIG. 4 has a uniform outside diameter (isodiametric), 
d.sub.o, and may be manufactured by winding a suitable wire around an 
appropriately tapered mandrel. As an alternative to having a isodiametric 
outside diameter, the wire core 42 may be wound so that it has a uniform 
inside helix diameter by winding the wire core 42 around an isodiametric 
mandrel, not shown. 
The wire core 42 also may be coated with a suitable electrically conductive 
layer, as for example, an electrically conductive noble metal such as 
platinum or a platinum alloy, as described herein above, or with a layer 
of carbon or an electrically conductive polymer to further reduce the 
electrical resistance of the lead 40. 
The proximal section of the lead is preferably sufficiently flexible so 
that the lead may be wound around the pulse generator implanted in the 
chest of the patient, as is common practice in the field of pacemaker 
implantation. However, the proximal end of the lead 40 is generally not 
flexed often during normal activities, nor is most of the length of the 
lead between the pulse generator and the heart frequently flexed and 
stressed. Generally, only the section of the wire core nearest the heart 
is highly and repeatedly flexed. Therefore, the lead 40 may be 
constructed, as shown by way of example in FIG. 4, whereby the wire core 
42 has a relatively large diameter for the proximal 2/3 of the overall 
length, "L", of the lead and a smaller diameter for the distal 1/3 of the 
length of the lead. The larger diameter of the wire core 42 may be 
approximately 0.2 mm, and gradually tapers over a predetermined length to 
a smaller diameter of about 0.1 ram, although other diameters suitable for 
a particular application may also be employed. 
A wire core having a varied diameter, as described above, may be 
manufactured, for example, by feeding a wire of uniform diameter through 
an etchant, such as an aqua regia, at varying speed. A length of wire 
having a smaller diameter may be manufactured by feeding the wire through 
the etchant at relatively slow speed so that the wire core is exposed to 
the etchant a sufficient time for the etchant to chemically etch the wire 
core to a predetermined diameter. The length of wire having the 
cross-sectional area which tapers from the smaller diameter to the larger 
diameter may be obtained by gradually increasing the speed of the wire 
through the etchant, thereby producing a wire core having a tapered 
diameter with no discontinuities. 
An electrical lead having relatively low electrical resistance and 
embodying various features of the present invention may include a 
helically coiled wire core having a given helix outside diameter, and an 
elongated cross-sectional area. The axial dimension of such elongated 
cross-sectional area is greater in the axial direction along the length of 
the lead than in the transverse, or radial direction. For example, as shown 
in FIG. 5, an electrical lead 50 is shown to include a wire core 52, 
having, by way of example, an oval-shaped cross-sectional area where the 
dimension of the oval in the axial direction of the lead 50 is greater 
than the radial or transverse direction along the diameter, "D," of the 
helix. The lead is surrounded by a 0.4 mm thick insulating sheath 54 such 
as polyurethane or silicone. The benefit of the oval cross-sectional area 
is that it provides the electrical lead 50 with a greater cross-sectional 
area and shorter length through which the current flows helically than 
would an implantable electrical lead comprised of a wire core having a 
circular cross-sectional area and the same helix diameter, "D." The 
combination of both the increased cross-sectional area and reduced length 
serve to provide the lead with relatively low electrical resistance. 
A wire core having an oval-shaped cross-sectional area may be manufactured 
by well known techniques such as extruding heated and softened material of 
which the wire is comprised through a suitably shaped die, or flattening a 
wire having a circular cross-sectional area with a roller mill. A wire 
core, such as wire core 52, may also be shaped to have other 
cross-sectional areas, such as rectangles or ellipses, oriented so that 
when the wire is coiled, the cross-sectional area of the wire core is 
longer in the axial direction than in the radial direction. 
In another embodiment of the present invention, an electrical lead 60, as 
shown in FIG. 6, having relatively low electrical resistance may include a 
helically wound wire core 62 comprised of individual coils 65 sheathed in a 
biocompatible electrical insulator sheath 64 which electrically isolates 
the wire core 62 from the surrounding tissues and bodily fluids. Adjacent 
coils 65 contact one another at interfaces 67 between coils. Every mth 
interface 67 between the coils 65 is welded so as to provide electrical 
continuity between the adjacent coils 65, where m is a positive integer, 
and the welds are represented by reference numbers 66. For example, every 
third interface 67 may be welded to the coils forming the interface along 
a weld line 66. 
An electrical lead 70 embodying various features of the present invention, 
as shown in FIG. 7, and having relatively low electrical resistance may 
include a helically wound wire core 72 comprised of coils 73 sheathed in a 
biocompatible electrical insulator sheath 74. The wire core 72 is wound 
with a helix pitch which differs over the length of the lead. The length, 
L.sub.1, of the electrical lead 70 is suitable for being positioned near 
the heart where it will be subjected to high, repetitive stresses. 
Therefore, the length L.sub.1 is wound to have many closely spaced coils, 
i.e., a high helix pitch, whereas the length, L.sub.2, of the electrical 
lead 70, not subjected to high stress levels may include coils wound with 
a lower helix pitch. The helix pitch may be defined as the number of coils 
of the wire core per unit length of the lead. The sections of the 
electrical lead having the coils 73 with the lower helix pitch provide the 
electrical lead 70 with a shorter, helical electrical path length for a 
given overall axial length of the lead than would a section of the lead 
where the coils have a higher helix pitch. Therefore, the section of the 
lead 70 having the lower helix pitch has less electrical resistance than 
does the section having the higher helix pitch. 
The electrical lead shown in FIG. 7 has variable stiffness attributable to 
the varying helix pitch over the length of the wire core. Such variable 
stiffness may be used to promote forming the lead into desired shapes. For 
example, the relatively flexible section of length L.sub.1 includes the 
distal end 76 of the lead 70. Such distal end 76 may be formed into an 
atrial "J," a common lead preform shape which maintains a stable lead 
position in the right atrial appendage of the heart. Such preform shape 
may also reduce the load that an electrode, not shown, mounted to the 
distal end of the lead exerts on the heart by reducing the inertial load 
of the lead against the heart. 
It has been shown that the present invention provides several embodiments 
of electrical leads having very low electrical resistance and excellent 
fatigue resistance. An electrical lead having low electrical resistance is 
expected to increase the operational life of an implanted pulse generator 
by reducing the expenditure of energy stored in the pulse generator 
battery that is wasted in the form of I.sup.2 R heating losses. 
Furthermore, such electrical leads may be manufactured from materials 
commonly employed in the construction of standard cardiac leads using well 
known manufacturing techniques. 
It is to be appreciated that the present invention may be practiced 
utilizing other lead constructions than those described herein such as 
multifilar, i.e., multiple wire core, construction. 
While the present invention has been described in terms of preferred 
embodiments, it is to be understood that the invention is not to be 
limited to the exact form of the apparatus or processes disclosed. 
Therefore, it is to be understood that the invention may be practiced 
other than as specifically described without departing from the scope of 
the claims.