An extractable lead and method for chronic blood contacting use. The new lead contains a hydrogel coating having a thickness increase greater than 10% when hydrated. A thick coating is used to provide a shear layer so that the coating tears during extraction, either at the coating/lead interface, between layers of the coating itself, or at the coating/tissue interface. Furthermore, because of the flexibility of such a thick coating, contracture of any fibrous capsule that may have formed is not a problem during extraction, since instead of contracting onto the lead, it contracts onto the flexible coating which can be extracted out of the tight capsule.

FIELD OF THE INVENTION 
The present invention relates in general to implantable medical structures, 
and more particularly to leads and catheters implanted in biological 
tissues such as the heart. This invention is more specifically directed 
toward improving the insertion and removal processes of cardiac pacing and 
defibrillation leads. 
BACKGROUND OF THE INVENTION 
Various types of transvenous pacing and cardioversion/defibrillation leads 
have been developed for endocardial introduction into different chambers 
of a patient's heart, typically the right ventricle (RV) or right atrial 
(RA) appendage, as well as the coronary sinus (CS). These flexible leads 
are usually constructed with an outer polymer insulator sheath, such as a 
flexible silicone or polyurethane tube or coating for encasing one or more 
coiled or cabled wire electrical conductors. One such conductor is 
typically attached at its distal tip to the shank portion of a tip 
electrode. In bipolar or multipolar leads, one or more coiled wire 
conductors are provided in a coaxial or co-linear relation to a first 
coiled wire conductor, and are connected to electrodes situated along the 
lead body. The proximal ends of the conductors are coupled to a connector 
which includes a single pin in unipolar leads and additional pins or rings 
in bipolar and multipolar leads. 
The tip electrode is usually placed in contact with the myocardial tissue 
by passage through venous access, often the subclavian or cephalic vein or 
one of its branches, which leads to the endocardial surface of the heart 
chambers. The tip electrode may be held in place passively by silicone or 
polyurethane tines within the trabeculae of the RV apex, as described, for 
example, in U.S. Pat. No. 3,902,501 to Citron et al. The tines or fins 
extend outwardly and are usually molded separately and bonded onto the 
distal end of the lead proximal to the tip electrode. Alternatively, the 
tip electrode may be held in place actively through the use of a 
manipulated anchor or screw that penetrates the myocardium, as described, 
for example, in U.S. Pat. No. 3,974,834 to Kane. These fixation mechanisms 
help to prevent dislodgment of the tip, thus maintaining consistent 
sensing and pacing characteristics over time. 
Although the state of the art of implantable pulse generators and 
endocardial lead technology has advanced considerably, endocardial leads 
nevertheless occasionally fail for a variety of reasons, such as 
insulation failure, sensing failure, wire conductor fracture, or an 
increase in electrode resistance beyond a desirable level. Also, in some 
instances, it may be desirable to add one or more leads to stimulate 
different portions of the heart than are presently being stimulated with 
leads already in place. Many patients have one or more, and sometimes as 
many as four or five previously used (abandoned) and currently used leads 
in their veins and heart. 
The risks of removing leads or introducing additional leads in the heart 
and venous system include infection, physiological complications, 
obstruction to blood flow, and formation of blood clots which may embolize 
to the lung and produce severe complications and even death. In addition, 
extra leads in the heart can interfere with tricuspid valve and mechanical 
function, and can cause considerable difficulty in the positioning and 
attachment of new endocardial leads in the heart. 
Typically in the first few months of lead implant, fibrotic tissue 
encapsulates the lead, especially in areas touching biological tissue, 
such as the endothelial layer of veins, valves, heart wall, and 
trabeculae. When small diameter veins through which the lead passes become 
occluded with fibrotic tissue, separating the lead from the vein becomes 
difficult and may severely damage the vein or tissue attached to the lead, 
making lead explant very risky, difficult and oftentimes dangerous. 
Several attempts have been made to alleviate the lead explantability 
problem by using device or tool-assisted methods for removing the lead. A 
few exemplary lead removal techniques are described in the following 
publications: 
U.S. Pat. No. 4,574,800 to Peers-Trevarton teaches a lead extractor which 
is inserted into the lumen of an implanted lead and wedged against its 
inner structure (typically a coil) at a distal location, such as near an 
electrode implanted in the atrium or ventricle. This wedging condition 
permits a pulling force to be transmitted along the length of the 
extractor to the implanted electrode location. 
U.S. Pat. No. 5,231,996 to Bardy et al. describes an endocardial lead 
having a structure that strengthens the lead body to enable its removal by 
traction after a period of chronic implant. One or more relaxed, 
nonextensible filaments are loosely contained within the insulating 
sheath, and have their proximal and distal ends mechanically connected to 
the connector and electrode shank of the lead. These filaments operate as 
means for allowing the lead body to reach a stretched length exceeding the 
relaxed length by an amount sufficient to allow the lead to be stretched 
without breaking during removal by traction. 
U.S. Pat. No. 5,207,683 to Goode et al. teaches the use of a flexible 
stylet wire with an expandable wire coil attached to the distal end for 
engaging the coiled structure of the lead. 
A potential problem with the above intraluminal devices is that they do not 
address the possible encapsulating fibrous tissue attached to the lead, 
either along the lead body or at the distal end, especially with tines. 
When using these intraluminal extraction devices, the tissue may tear 
instead of releasing from the lead, causing atrial or ventricular 
avulsion, tears in the vein or heart, tamponade, and/or hemothorax. 
Immediate surgical intervention requiring a thoracotomy may be necessary 
to prevent death. 
Other extraction devices have been developed that are meant to fit over the 
lead to separate the lead from tissue attachments. Such devices are meant 
to work either alone or in conjunction with intraluminal devices, such as 
the locking stylets described above. One company that makes such lead 
extraction tools is Cook Pacemaker Corporation (Leechburg, Pa). However, 
in many cases, these sheaths are incapable of easily separating calcified 
tissue from the lead. Often, a femoral removal approach is needed, 
especially as lead implant duration increases. It is preferred that the 
leads be removable by the venous implant site rather than femorally, if 
possible. These extraction sheaths, used alone or in combination with 
locking stylet devices, do not overcome the risks described above. 
In addition to providing special tools to aid in lead extraction, the lead 
itself may be designed to be more easily extracted, such as by adding 
coatings, discouraging tissue ingrowth, and making the lead isodiametric. 
Coated leads and catheters, as well as coating materials are described in 
the following publications, all of which are incorporated herein by 
reference: 
U.S. Pat. No. 4,487,808 to Lambert describes the process of coating a 
polymer surface with a hydrophilic coating with low friction under wet 
conditions. The process comprises applying to the polymer surface a 
solution containing between 0.05 and 40% of a compound which comprises at 
least two unreacted isocyanate groups per molecule, evaporating the 
solvent, applying a solution containing between 0.5 and 50% of 
polyethylene oxide to the thus treated polymer surface, evaporating the 
solvent of the latter solution, and curing the coating at elevated 
temperature. This patent aims at facilitating the insertion of medical 
instruments inside a body cavity by decreasing the coefficient of friction 
of the surface of the device or lead. 
U.S. Pat. No. 5,041,100 to Rowland et al. describes a catheter to which a 
friction-reducing coating may be applied, to reduce catheter friction 
particularly when the coating is hydrated. The coating includes a mixture 
of a structural plastic material and high molecular weight polyethylene 
oxide, for facilitating the insertion of the catheter into a patient's 
body. 
U.S. Pat. No. 5,077,352 to Elton describes still other abrasion resistant, 
hydrophilic, lubricious organic coatings for application to the outer 
surfaces of inorganic materials or organic polymeric medical devices, to 
facilitate the introduction of these devices inside the patient's body. 
However, while the foregoing coating processes may reduce friction to aid 
in implanting the leads, none of them were developed for nor adequately 
addresses the problem of explantation of leads should adhesions develop on 
the surface and become firmly attached. It appears that in prior art only 
a thin, surface coating is applied, with no significant dimensional 
changes. Only a thin coating is necessary to reduce friction to facilitate 
ease of insertion. 
Hydrogel coatings have been proposed for other uses in pacing and 
defibrillation leads. The following are such examples, and are 
incorporated herein by reference: 
European patent application No. 057,450 to Cahalan et al. relates to a body 
implantable lead having a polymer-based gel electrode. European patent 
application No. 057,451 to Juncker et al. relates to a body implantable 
lead having a pressure-cushioned electrode. In both cases, the pacing tip 
electrode is coated with a hydrogel, thus separating the solid electrode 
from the excitable tissue and increasing the effective electrode area. 
While increasing effective electrode area is desirable for defibrillation 
in which the goal is field stimulation, it is desirable to keep effective 
pacing electrode area small to minimize pacing thresholds. Therefore, 
increased pacing thresholds would be a disadvantage for the Cahalan et al. 
and Juncker et al. inventions. 
1995 NASPE abstract 452, entitled "A New Surgical Temporary Pacing Lead: 
Easy to be Fixed and Easy to be Removed," by Yokoyama et al. describes the 
use of a material composed of absorbable polyglycolic acid felt (PGA felt) 
for facilitating the extraction of temporary pacing leads seven to ten 
days after implantation. However, there is no clear indication that this 
material can be useful to resolve fibrotic encapsulation problems 
resulting from the long term implantation of the lead, since the PGA felt 
is absorbed by the tissue after ten days, and tissue encapsulation may 
then begin. Additionally, the use of the PGA felt seems to be limited to 
the distal end of the lead, and consequently neglects the fibrotic growth 
on the defibrillation electrode and lead body. 
U.S. Pat. No. 5,020,544 to Dahl et al. describes a defibrillation patch 
electrode having a hydrogel incorporated in the porous conductive screen 
for preventing tissue ingrowth. The hydrogel can serve as a drug reservoir 
for antibiotics, antiseptics, antiarrhythmics, or anti-inflammatory 
steroids. 
Polymer hydrogels and other coatings have also been recommended for use on 
implantable devices to stimulate the attachment of endothelial cells for 
improving thromboresistance. U.S. Pat. No. 4,836,884 to McAuslan describes 
such hydrogels. 
U.S. Pat. No. 5,090,422 to Dahl et al. describes a porous implantable 
enclosure that covers and isolates an electrode in a way which allows 
electrical coductivity via bodily fluid which passes through but separates 
the electrode from the adjacent tissue in the manner of a dissection plane 
which substantially prevents tissue ingrowth. In addition to the improved 
explantability, the porous covering is also intended to reduce tissue 
burning and edema. However, the porous implantable enclosures in Dahl are 
not described as expandable, so that the lead would have to be implanted 
with the covering already at the final thickness required to provide the 
desired effects, which is disclosed as 10 to 100 mils. This would increase 
the overall diameter of the lead by 20 to 200 mils (up to about 5 mm), 
which would substantially increase the introducer size needed and reduce 
maneuverability through the vein to the implant sight. 
While the foregoing coating processes may aid in preventing tissue ingrowth 
into the interior of the leads, none of them adequately addresses the 
problem of explantation of leads should adhesions develop on the surface 
and become firmly attached. It appears that only a thin, surface coating 
is applied, with no significant dimensional changes. 
None of the above solutions completely satisfies the need for a new method 
and lead structure which is easy to implant and improves the rate of 
successful lead explantation, particularly after the encapsulated tissue 
has rendered various traction removal methods impractical. 
SUMMARY OF THE INVENTION 
The present invention is directed toward optimizing the construction of 
implantable leads and catheters for chronic blood contacting use. In 
particular, the present invention is directed toward improving the 
extractability of a permanently implanted cardiac lead. 
It is a further object of the present invention to provide a new lead that 
can be withdrawn from any encapsulating tissue that may form with 
relatively little or no damage to either the lead or the neighboring body 
tissue. 
Briefly, the foregoing and other objects of the present invention are 
realized by a new hydrogel-coated lead as described below. In the coatings 
of the present invention, preferably, hydrogels having a thickness 
increase greater than 10% are used. A thicker coating is used to provide a 
shear layer so that the coating tears during extraction, either at the 
coating/lead interface, between layers of the coating itself, or at the 
coating/tissue interface. Furthermore, because of the flexibility of such 
a thick coating, contracture of any fibrous capsule that may have formed 
is not a problem during extraction, since instead of contracting onto the 
lead body itself, it contracts onto the flexible, slippery coating which 
can be extracted out of the tight capsule.

DETAILED DESCRIPTION OF THE INVENTION 
FIG. 1 illustrates a heart 10 and a superior vena cava (SVC) 12 through 
which an SVC lead 14 and a right ventricular (RV) defibrillation lead 15 
are implanted. The RV lead 15 passes through the SVC 12 into the RV 17 of 
the heart 10. The distal end 16 of the RV lead 15 includes a pacing tip 
electrode 19 for electrically stimulating the heart 10. The distal end 16 
is secured to trabeculae within the apex 18 of RV 17 by means of a 
plurality of tines 20, which in time become securely attached by fibrotic 
tissue forming around tines 20. RV lead 15 further includes a 
defibrillation electrode 22 which may rest against the heart wall and may 
mechanically stimulate the growth of fibrotic tissue around it. 
Defibrillation electrode 22 may also serve as a sensing electrode and/or 
as the return electrode for bipolar pacing. Alternatively or additionally, 
lead 15 may include one or more electrodes dedicated for sensing and/or 
pacing return electrodes (not shown). 
The SVC lead 12 includes an electrode 23 which also may become secured to 
the vein wall through fibrotic growth and encapsulation. A lead for 
implantation in the CS may also be constructed similarly to SVC lead 14, 
and be preferably 6 French or less in diameter during insertion, prior to 
hydration. A lead having good extractability may be even more important 
for the CS than for other locations because of proximity to tissue, 
smaller space possibly more prone to occlusion, weakness of venous tissue 
as compared to myocardium, a more tortuous path, and more complex lead 
configuration. For an example of a complex CS lead configuration, see U.S. 
Pat. No. 5,476,498 to Ayers which is incorporated herein by reference. 
FIGS. 2A, 2B, and 2C provide overall views of three embodiments of RV lead 
15, which include a lead body 21, that is attached to a connector 25 and 
insulated by insulation 24. As it will be described later in greater 
detail, selected parts of RV lead 15 are coated with a thin hydrophilic 
polymer layer. The hydrogel layer forms a coating 30 over tines 20, a 
coating 32 over defibrillation electrode 22, and a coating 34 over lead 
insulation 24. 
It is not necessary to apply the hydrogel to the connectors, the pacing tip 
electrode, or the region of the lead that will be coiled into the 
subcutaneous area. If the pacing tip electrode were coated, tissue would 
not grow into the pores, thereby not providing firm stabilization of the 
lead at the tip. This in turn would lead to mechanical rubbing between the 
tip and the cardiac tissue, leading to increased scar tissue formation. 
This increase in nonstimulatable tissue layer thickness, coupled with the 
thickness of the coating itself, would mean that the virtual electrode 
area for pacing would be increased to an undesirable level and require 
more energy. 
In FIG. 2A, the conductive electrode material of defibrillation electrode 
22 is made up of small diameter coils 45 wrapped around lead body 21. 
Details of this type of construction are given in U.S. Pat. No. 5,439,485 
to Mar et al. which is assigned to the assignee of the present invention 
and is incorporated herein by reference. In FIG. 2B, the conductive 
electrode material is made up of wire 46 wrapped around and secured to 
lead body 21. In FIG. 2C, the conductive electrode material is a 
conductive polymer 47. 
FIG. 3 illustrates a more detailed view of the distal end 16, and 
illustrates the coating 30 prior to implantation (in solid line) and after 
implantation (in dashed lines). Before implant, the hydrogel layer is a 
thin coating on the surface of lead 15. At implant, the hydrogel layer 
absorbs body fluids and expands in volume. This feature is illustrated by 
way of example, by the expanded coating 30a shown in dashed lines around 
tines 20. As the dry coating 30 becomes hydrated, its structure changes: 
its volume increases, and it becomes softer and mechanically weak, but not 
so weak as to be susceptible to ordinary biological and mechanical 
degradation mechanisms to which leads are typically subjected. As a 
result, the hydrogel layer acts as a mechanical buffer between lead 15 and 
the encapsulating fibrotic tissue, and prevents the encapsulating fibrotic 
tissue from attaching directly onto the polymeric surface, e.g., silicone, 
of tines 20, defibrillation electrode 22, and lead insulation 24. 
Furthermore, in the event that encapsulating fibrotic tissue forms on the 
hydrogel, because the hydrogel is compliant and has a low coefficient of 
friction, the lead can be pulled out more easily. 
Because the hydrogel prevents fibrotic tissue from attaching directly to 
the polymeric or metallic surfaces of the lead, and because the hydrated 
hydrogel layer is weak, in the event that lead explantation is required, 
the hydrogel layer can act as a shear layer. Consequently, lead 15 may be 
easily removed from the encapsulating fibrotic sheath with minimal or no 
damage to neighboring body tissue. While it would be preferable to remove 
the entire hydrogel layer during explantation, part of the hydrogel layer 
may be left in the heart. The residual layer will not cause damage to the 
heart, since it is biocompatible, being composed mostly of water with 
similar ionic content to that of blood. The residual layer will occopy a 
small volume and is mechanically compliant and soft to prevent mechanical 
irritation. 
As used herein "hydrogel" or "hydrogel layers" includes a polymer that 
expands in volume when it hydrates. The following are some exemplary 
hydrogels: Poly 1-hydroxyethyl methacrylate, polymethacrylic acid, 
poly(N,N, dimethyl-aminoethyl methacrylate), polyacrylamide, poly(N-vinyl 
pyrrolidone), polyvinyl alcohol, polyethylene oxides, hydrolyzed 
polyacrylonitrile, polyelectrolyte complexes, polymethacrylic acid and 
polyacrylonitrile, anionic and cationic hydrogels or composites or 
copolymers of one or more of these hydrogels and other suitable 
biocompatible materials such as silicone, polymers, and polyurethane. 
In one embodiment the hydrogel can be used as a vehicle for drug delivery, 
by loading the polymeric hydrogel precursor with a blood soluble or 
insoluble drug, or by chemically binding a drug such as antibiotic, 
antiseptic, antiarrhythmic, antiinflammatory steroid or other agents, to 
the polymer network. Upon expansion of the hydrogel layer, the drug is 
released to the neighboring tissue. 
The hydrogel can also be loaded with an additive, such as Na.sup.+ Cl.sup.- 
or other electrolyte, to enhance its electrical conductivity. It is 
preferred that the hydrated coating consist mostly of water with an ionic 
content at least that of blood. Examples of loading methods include 
dissolving a salt in a monomer-water-initiator solution and polymerizing, 
or simply soaking the unloaded, coated lead in an electrolyte solution. 
The hydrogel may also be blended with a non-hydrogel polymer that increases 
the physical strength and/or helps control the degree of swell, while the 
hydrogel carries the electric current. Methods for preparing 
hydrogel/non-hydrogel polymer blends are known in the art, with an example 
being described in U.S. Pat. No. 4,883,699 to Aniuk et al., which is 
incorporated herein by reference. A mixture of the components can be 
prepared using, for example, a two-roll mill, an internal mixer, such as a 
Brabender or Banbury mixer, and extruder, for example a twin-screw 
extruder. 
Depending on the thickness of the hydrogel, the expanded hydrogel layer may 
reduce stimulation edge effects. It may further allow a more uniform 
distribution of the stimulation energy for decreasing the chance of tissue 
damage. The biocompatibility and dimensional and chemical stability of 
hydrogels render them well-suited for long-term implantation. The expanded 
hydrogel layer may also inhibit tissue adhesion and thrombus formation on 
the lead. 
In one embodiment, the hydrogel layer is formed on lead 15 by dip coating 
predetermined parts of lead 15 in a hydrogel solution. Other exemplary 
coating techniques include molding, spraying, and vapor deposition. 
Alternatively, a fully crosslinked hydrogel tube may be hydrated, inserted 
over a lead, and allowed to dry onto the lead. Primer and/or adhesive may 
be added at the interface that is activated by heat or UV. 
In some applications, it is desirable to enhance the adhesion of the 
hydrogel layer to either the lead body or the electrodes. Several methods 
may be used to accomplish such result, such as ion etching, covalent 
crosslinking, thermal or chemical surface treatment, radiation (i.e., 
electron or gamma beam), and adding a primer layer between the lead and 
the hydrogel layer. 
Defibrillation electrode 22 and coating 32 may be formed as illustrated in 
FIG. 4, which is a greatly enlarged surface view of FIG. 2A. It includes 
wrapping conductive coiled wires 45 around lead body 21, and then coating 
RV electrode 22 with an insulation layer 48, such as a urethane or 
silicone layer. Insulation layer 48 defines raised portions over the 
conductive coils 45, which raised portions are then ablated for allowing 
portions 45a of the coiled coils 45 to be exposed. In one embodiment, the 
outer surface of insulation layer 48 is treated, such as by applying a 
primer, to produce active sites 49 for binding to the hydrogel. A hydrogel 
layer is then formed or deposited over insulation layer 48. Additionally, 
if hydrogel is desired onto the metal portions, the metal may be etched 
and or primed to improve adhesion of hydrogel. 
FIGS. 5 and 6 illustrate another way to form the coating of the 
defibrillation electrodes, for instance RV electrode 22 of FIG. 2B. RV 
electrode 22 is formed of one or more conductive wires 46 wound around 
lead body 21 (shown in dashed lines). A thin dry hydrogel coating 32 is 
formed on the portions of lead body 21 that are not covered by wire 46. 
Coating 32 may be formed by dipping RV electrode 22 in a hydrogel 
solution, and by allowing the hydrogel layer to dry. Coating 32 
preferentially adheres to the material of lead body 21, which is typically 
silicone, and does not adhere as well to the wire 46, which is typically 
platinum or a platinum iridium alloy. Lead 15 may be dipped beyond 
electrode 22 such that portions of lead insulation 24 (not shown) are also 
coated. FIG. 6 illustrates the swelling of coating 32 after implantation 
and hydration, to form a pliant coating 32a, that covers the conductive 
wire 46 (shown in dashed lines). Expanded coating 32a may cover the wire 
46 entirely or partially, depending on the volume swell of hydrogel 
coating 32. 
FIG. 7 is an enlarged side view of an alternative lead body 21 showing 
coatings 34 and 32 of insulation 24 and defibrillation electrode 22 in the 
form of longitudinal strips of a hydrogel coating. The strips may be 
rectangular, semicircular, or other shapes in cross section. 
FIG. 8 is a cross-sectional view of the defibrillation electrode of FIG. 7, 
taken along line 8--8. (The inner coil structure has been omitted from the 
drawing for clarity.) 
FIG. 9 is a cross-sectional view of the defibrillation electrode of FIG. 8, 
illustrating an expanded hydrated hydrogel layer 34a. 
FIG. 10 is an enlarged side view of an alternative arrangement of hydrogel 
coatings 34 and 32 of insulation 24 and defibrillation electrode 22, 
respectively, with a dot pattern of hydrogel applied. The hydrogel may be 
applied in any pattern. 
FIGS. 11 and 12 are cross-sectional views of lead body 21, taken along line 
11--11 in FIG. 2A. Lead body 21 includes a biocompatible, insulative outer 
sheath 24 which is coated with a thin dry hydrogel layer 34. FIG. 12 
illustrates the expansion of hydrogel layer 34 into a swollen pliant 
coating 34a after implant and hydration. As explained above, coating 34a 
provides a shear layer between outer sheath 24 and attached fibrous 
encapsulation, and therefore facilitates the explantation of lead 15. In 
one embodiment of lead 15, hydrogel layer 34 is not applied to the entire 
length of outer sheath 24, but rather on selected portions of outer sheath 
24 that are more likely to be encapsulated by fibrous tissue. 
Preferably, the dry, unexpanded hydrogel coating thickness is 0.1 to 0.5 
mm, and the hydrated, expanded thickness is approximately 0.25 to 1.5 mm, 
and most preferably about 0.5 to 1 mm. Therefore, a typical dry, coated 
lead may start out with a diameter of about 3 mm; when the coating is 
expanded, the final lead diameter may be about 4 mm. When the dry lead 
becomes hydrated, the increase in the hydrogel layer thickness itself is 
preferably greater than 10% and less than 500%. 
It is preferred that hydration take place quickly enough to perform 
standard implant testing without delay following lead insertion. 
Therefore, hydration can occur any time between the time the lead leaves 
the end of the introducer and about one minute following positioning, 
preferably about 30 seconds. 
With these properties, while maintaining small size for ease of implant 
through an introducer, the electrode is effectively enlarged and 
explantability is improved. The enlargeability of the electrode diameter 
leads to reduced peak current density and more even current distribution, 
which may serve to decrease incidence and severity of burns, edema, 
necrosis, and or tissue stunning. Furthermore, if high current density 
areas are refibrillatory, by reducing them, defibrillation thresholds may 
actually be lowered, thereby allowing lower energies to be used and 
prolonging battery life. 
FIG. 13 is a side view of lead 15 of FIG. 1 shown encapsulated in a tissue 
sheath 50. FIGS. 14A, 14B, 15A, 15B, 16A, and 16B are side views of the RV 
lead of FIG. 1 being explanted from the heart. Because the tissue response 
will occur at the outer surface of hydrogel 32, the tissue capsule 50 
would be separated from electrode 22 by the hydrogel layer. This 
facilitates easy removal of the electrode from the tissue sheath. The 
hydrogel is mechanically very weak and acts as a shear layer to allow the 
electrode to easily pull away from the tissue sheath. The ultimate tensile 
strength (UTS) of fibrous capsule has been reported at 334 psi, and the 
UTS of typical silicone rubber, a major component of leads, is 900 psi. 
Cardiac muscle has a UTS of 16 psi, and veins have a UTS of 247 psi in the 
longitudinal direction. Thus, a hydrogel with an ultimate tensile strength 
of 1 to 300 psi would provide a weak link for allowing the lead to break 
free from the fibrous tissue. 
Lead 15 is extracted from the tissue capsule 50 by applying traction at a 
proximal portion of the lead, such as the connector end 25 (shown in FIG. 
2A) or at an accessible portion of the lead body 21 where it exits the 
venous access site (not shown). A twisting motion may be used to aid in 
separating the lead from the tissue capsule. A locking stylet-type device 
and/or an extraction sheath (as described herein under Background) may be 
used to aid in this separation process. 
In FIGS. 14A and 14B, lead 15 is shown separated from tissue 50 between 
hydrogel coating 30, 32 and the rest of the lead 56 (lead 15 minus the 
hydrogel coating). In FIGS. 15A and 15B, lead 15 is shown separated from 
tissue 50 between hydrogel coating 30, 32 and tissue 50. In FIGS. 16A and 
16B, lead 15 is shown separated from tissue 50 by separating hydrogel 
coating 30, 32 into layers, leaving the extracted lead with a first layer 
52 of coating and leaving a second layer 54 of hydrogel coating attached 
to tissue 50. The extraction process may involve a combination of 
separations, with portions of the encapsulated lead separating between the 
coating and the rest of the lead, other portions separating between layers 
of the coating, and still other portions separating between the coating 
and the tissue. Alternatively, the extraction process may be primarily 
limited to only one type of separation, with only a small amount of 
hydrogel coating remaining adherent in patches to either the rest of the 
lead or to the tissue. 
While the foregoing embodiments were explained in relation to RV lead 15, 
it should be understood that the same or similar inventive concepts may be 
applied to SVC (or CS) lead 14 as well as to other implantable medical 
structures in the venous system, including other defibrillation leads, 
catheters, and pacemaker leads. 
It should be understood that various alternatives to the embodiments of the 
invention described herein may be employed in practicing the invention. It 
is intended that the following claims define the scope of the invention 
and that structures and methods within the scope of these claims and their 
equivalents be covered thereby.