Note: Descriptions are shown in the official language in which they were submitted.
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TI88DE HEATINQ AND ABLATION BY8TEM8 AND METHODS
U8IN3 POROUS ELECTRODE STRUCTURES
Field of the Invention
The invention generally relates to
15 electrode structures deployed in interior regions
of the body. In a more specific sense, the
invention relates to electrode structures
deployable into the heart for diagnosis and
treatment of cardiac conditions.
20 Hackcround of the Invention
The treatment of cardiac arrhythmias
requires electrodes capable of creating tissue
lesions having a diversity of different geometries
and characteristics, depending upon the particular
physiology of the arrhythmia to be treated.
For example, a conventional 8F
diameter/4mm long cardiac ablation electrode can
transmit radio frequency energy to create lesions
in myocardial tissue with a depth of about 0.5 cm
and a width of-about 10 mm, with a lesion volume
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of up to 0.2 cm3. These small and shallow lesions
are desired in the sinus node for sinus node
modifications, or along the A-V groove for various
a
accessory pathway ablations, or along the slow
zone of the tricuspid isthmus for atrial flutter
(AFL) or AV node slow pathways ablations.
However, the elimination of ventricular
tachycardia (VT) substrates is thought to require
significantly larger and deeper lesions, with a
penetration depth greater than 1.5 cm, a width of
more than 2.0 cm, with a lesion volume of at least
1 cm3.
There also remains the need to create
lesions having relatively large surface areas with
shallow depths.
One proposed solution to the creation of
diverse lesion characteristics is to use different
forms of ablation energy. However, technologies
surrounding microwave, laser, ultrasound, and
chemical ablation are largely unproven for this
purpose.
The use of active cooling in association
with the transmission of DC or radio frequency
ablation energy is known to force the
electrode-tissue interface to lower temperature
values, As a result, the hottest tissue
temperature region is shifted deeper into the
tissue, which, in turn, shifts the boundary of the
tissue rendered nonviable by ablation deeper into
the tissue. An electrode that is actively cooled
can be used to transmit more ablation energy into ,
the tissue, compared to the same electrode that is
not actively cooled. However, control of active
cooling is required to keep maximum tissue
temperatures safely below about 100° C, at which
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tissue desiccation or tissue boiling is known to occur.
Another proposed solution to the creation of larger
lesions, either in surface area and/or depth, is the use of
substantially larger electrodes than those commercially
available. Yet, larger electrodes themselves pose problems of
size and maneuverability, which weigh against a safe and easy
introduction of large electrodes through a vein or artery into
the heart.
A need exists for multi-purpose cardiac ablation
electrodes that can selectively create lesions of different
geometries and characteristics. Multi-purpose electrodes would
possess the requisite flexibility and maneuverability permitting
safe and easy introduction into the heart. Once deployed
inside the heart, these electrodes would possess the capability
to emit energy sufficient to create, in a controlled fashion,
either large and deep lesions, or small and shallow lesions,
or large and shallow lesions, depending upon the therapy
required.
Summary of the Invention
The invention provides various porous electrode
assemblies for tissue heating and ablation systems and methods,
which enable ionic transpart of electrical energy to occur
substantially free of liquid perfusion.
The invention provides an electrode assembly for a
catheter, comprising: a wall having an interior surface and an
exterior surface, said wall surrounding an interior area and
capable of holding a medium containing ions under pressure;
and an electrode element disposed in said interior area, said
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electrode element configured for coupling the medium to a source
of electrical energy, wherein at least a portion of the wall
comprises a section of porous material sized to pass ions
contained in the medium without substantial liquid perfusion
through the porous material, thereby enabling ionic transport of
electrical energy from the electrode element to the exterior
surface of said wall, via the medium.
The wall thereby enables ionic transport of electrical
energy through the porous material to the exterior of the wall,
without substantial perfusion of liquid through the wall.
At least a portion of the wall may comprise a hydro-
philic porous material sized to pass ions contained in the
medium to thereby enable ionic transport of electrical energy
through the porous material to the exterior of the wall. The
porous material has a bubble point value greater than the
internal pressure, whereby ionic
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transport occurs substantially free of liquid
perfusion through the porous material.
Other features and advantages of the in-
ventions are set forth in the following
Description and Drawings.
Brief Description of the Drawings
Fig. 1 is a plan view of a system for
ablating heart tissue, which includes an
expandable porous electrode structure that
embodies the features of the invention;
Fig. 2 is an enlarged side elevation
view, with portions broken away, of a porous
electrode structure usable in association with the
system shown in Fig. l, with the electrode
structure shown in its expanded geometry;
Fig. 3 is an enlarged side elevation view
of the porous electrode structure shown in Fig. 2,
with the electrode structure shown in its
collapsed geometry;
Fig. 4 is a further enlarged, somewhat
diagrammatic side view, with portions broken away,
of the porous electrode structure shown in Fig. 2;
Fig. 5 is an enlarged side elevation
view, with portions broken away, of a porous
electrode structure usable in association with the
system Shawn in Fig. l, with the electrode
structure shown in its expanded geometry due to
the presence of an interior spline support
structure;
Fig. 6 is an enlarged side section view
of the porous electrode structure shown in Fig. 5,
with the electrode structure shown in its
collapsed geometry due to the manipulation of an
exterior sliding sheath;
Fig.-7 is an enlarged side elevation
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view, with portions broken away, of a porous
electrode structure usable in association with the
system shown in Fig. l, with the electrode
structure shown in its expanded geometry due to
the presence of an interior interwoven mesh
support structure;
Fig. 8 is an enlarged, somewhat
diagrammatic enlarged view, taken generally along
line 8-8 of Fig. 4, showing the ionic current
densities across the pores of the electrode body
shown in Fig. 4;
Fig. 9 is a graph showing the
relationship between sensed impedance and ionic
transport through the pores of the electrode body
shown in Fig. 4;
Fig. 10 is an enlarged side elevation
view, with portions broken away, of an alternative
porous electrode structure usable in association
with the system shown in Fig. l, with the electrode
structure comprising a porous foam body shown in
its expanded geometry;
Fig. Z1 is an enlarged side view of a
porous electrode structure usable in association
with the system shown in Fig.i, with the pores of
the structure arranged in a bulls eye pattern on
the distal end of the body;
Fig. 12 is an enlarged side view of a
porous electrode structure usable in association
with the system shown in Fig.l, with the pores of
the structure arranged in circumferentially spaced
segments along the side of the body;
Fig. 13 is a side view, with portions
broken away, showing the use of multiple chambers
to convey liquid to the segmented pore regions
shown in Fig.-12;
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Fig. 14 is an enlarged side elevation
view, with portions broken away, of a porous
electrode structure usable in association with the
system shown in Fig. 1, which also carries
nonporous electrode elements;
Fig. 15 is an enlarged side section view
of a porous electrode structure, which also
carries electrode elements formed by wire snaked
through the body of the structure;
Fig. 16 is an enlarged side elevation
view, with portions broken away, of a porous
electrode structure with interior pacing/sensing
electrodes; Figs. 17 and 18 are diagrammatic
representations of the tissue temperature profiles
associated with a porous electrode structure when
operated under different conditions;
Fig. 19 is a somewhat diagrammatic view
of a fixture and mandrel for forming a
hemispherical geometry for the distal end of a
porous electrode body from a flat sheet of porous
material;
Fig. 20 is a side sectional view of the
fixture and mandrel shown in Fig. 19 in the
process of forming the hemispherical distal end
geometry in a flat sheet of porous material;
Fig. 21 is an enlarged side section view
of the sheet of porous material after formation of
the hemispherical distal end geometry;
Fig. 22 is a somewhat diagrammatic view
of a finishing fixture for forming the
hemispherical geometry for the proximal end of the
porous electrode body from the preformed sheet
shown in Fig. 21;
Fig. 23 is an elevation view of the
porous electrode body after having been formed by
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the devices shown in Figs. 19 to 22;
Fig. 24 is a somewhat diagrammatic view
of an expandable finishing fixture that can be
used instead of the finishing fixture shown in
Fig. 22 for forming the hemispherical geometry for
the proximal end of the porous electrode body from
the preformed sheet shown in Fig. 21;
Fig. 25 is an elevation view of the
porous electrode body after having been formed by
l0 the expandable fixture shown in Fig. 24;
Fig. 26 is a somewhat diagrammatic view
of two preformed hemispherical body sections of
porous electrode body before being joined together
into a composite porous electrode body;
Fig. 27 is a side elevation view of the
composite porous electrode body formed by joining
the two hemispherical sections shown in Fig. 26
together along a circumferential seam;
Fig. 28A is a side section view showing
the eversion of the porous electrode body shown in
Fig. 27 to place the circumferential seam on the
inside of the body, away from direct tissue
contact;
Fig. 28B is a side section view of the
porous electrode body shown in Fig. 27 after
having been everted to place the circumferential
seam on the inside of the body;
Fig. 29A is a side elevation view of a
porous electrode body formed by joining two
hemispherical sections along an axial seam and
after eversion to place the axial seam on the _
inside of the body;
Fig. 29B is a top view of the porous
electrode body with the everted axial seam shown
in Fig. 29A;
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Fig. 30 is a top view of a porous
_ electrode body formed by joining two hemispherical
sections along a main axial seam, with additional
intermediate axial seams to segment the body,
T
after eversion to place the axial seams on the
inside of the body;
Fig. 31A is an enlarged side sectional
view of a seam joining two sheets of porous
material together to form an electrode body, with
1o a temperature sensing element encapsulated within
the seam, and before eversion of the body;
Fig. 31B is a side elevation view of the
seamed body, shown partially in Fig. 31A, with
temperature sensing elements encapsulated in the
seam, and before eversion of the body;
Fig. 31C is a side section view of the
body shown in Fig. 31B after eversion, placing the
seam and the signal wires of the temperature
sensing elements inside the body;
Fig. 32A is a somewhat diagrammatic view
of a porous electrode body being formed from a
regenerated cellulose material by dipping using an
expandable fixture;
Fig. 32B is the dip-formed body shown
being formed in Fig. 32A, after removal of the
expandable fixture and attachment of a fixture
with steering assembly to the distal end of the
body, and before eversion;
Fig.' 32C is the dip-formed body with
distal fixture and steering assembly, shown in
Fig. 32B, after eversion;
Fig. 33 is an exemplary porous body
formed in an elongated, cylindrical geometry with
changing radii along its length, forming the
distal and proximal neck regions;
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Fig. 34A is another exemplary porous body
formed as a tube in an elongated, cylindrical
geometry with constant radii along its length; and
Fig. 34B is the tube shown in Fig. 34A,
with its distal end closed by a seam, and a port
tube sealed to its proximal end for attachment to
a catheter tube.
The invention may be embodied in several
forms without departing from its spirit or
20 essential characteristics. The scope of the
invention is defined in the appended claims,
rather than in the specific description preceding
them. All embodiments that fall within the
meaning and range of equivalency of the claims are
therefore intended to be embraced by the claims.
I?ssGriQtion of the Preferred Embodiments
Fig. 1 shows a tissue ablation system 20
that embodies the features of the invention.
The system 10 includes a flexible
catheter tube 12 with a proximal end 14 and a
distal end 16. The proximal end 14 carries a
handle 18. The distal end 16 carries an electrode
structure 20, which embodies features of the
invention. The purpose of the electrode structure
20 is to transmit ablation energy.
As Figs. 2 and 3 best show, the electrode
structure 20 includes an expandable-collapsible
body 22. The geometry of the body 22 can be
altered between a collapsed geometry (Fig. 3) and
an enlarged, or expanded, geometry (Fig. 2). In
the illustrated and preferred embodiment, liquid ,
pressure is used to inflate and maintain the
expandable-collapsible body 22 in the expanded
geometry.
In this arrangement (see Fig. 2), the
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catheter tube 12 carries an interior lumen 34
along its length. The distal end of the lumen 34
opens into the hollow interior of the expandable-
collapsible body 22. The proximal end of the
lumen 34 communicates with a port 36 (see Fig. 1)
on the handle 18. The liquid inflation medium
{arrows 38 in Fig. 2) is conveyed under positive
pressure through the port 36 and into the lumen
34. The liquid medium 38 fills the interior of
the expandable-collapsible body 22. The liquid
medium 38 exerts interior pressure to urge the
expandable-collapsible body 22 from its collapsed
geometry to the enlarged geometry.
This characteristic allows the
expandable-collapsible body 22 to assume a
collapsed, low profile (ideally, less than 8
French diameter, i.e., less than about 0.267 cm)
when introduced into the vasculature. Once located
in the desired position, the expandable-
collapsible body 22 can be urged into a
significantly expanded geometry of, for example,
approximately 7 to 20 mm.
As Figs. 5 to 7 show, the structure 20
can include, if desired, a normally open, yet
collapsible, interior support structure 54 to
apply internal force to augment or replace the
force of liquid medium pressure to maintain the
body 22 in the expanded geometry. The form of the
interior support structure 54 can vary. It can,
for example, comprise an assemblage of flexible
spline elements 24, as shown in Fig. 5, or an
interior porous, interwoven mesh or an open porous
foam structure 26, as shown in Fig. 7.
In these arrangements (see Fig. 6), the
internally supported expandable-collapsible body
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22 is brought to a collapsed geometry, after the
removal of the inflation medium, by outside
compression applied by an outer sheath 28 (see
Fig. 6), which slides along the catheter tube 12.
As Fig. 6 shows, forward movement of the sheath 28
advances it over the expanded expandable-
collapsible body 22. The expandable-collapsible
body 22 collapses into its low profile geometry
within the sheath 28. Rearward movement of the
sheath 28 (see Figs. 5 or 7) retracts it away from
the expandable-collapsible body 22. Free from the
confines of the sheath 48, the interior support
structure 54 springs open to return the
expandable-collapsible body 22 to its expanded
geometry to receive the liquid medium.
As Fig. 4 best shows, the structure 20
further includes an interior electrode 30 formed
of an electrically conductive material carried
within the interior of the body 22. The material
of the interior electrode 30 has both a relatively
high electrical conductivity and a relatively high
thermal conductivity. Materials possessing these
characteristics include gold, platinum,
platinum/iridium, among others. Noble metals are
preferred.
An insulated signal wire 32 is coupled to
the electrode 30. The signal wire 32 extends from
the electrode 30, through the catheter tube 12, to
an external connector 38 on the handle 18 (see
3o Fig. 1). The connector 38 electrically couples the
electrode 30 to a radio frequency generator 40. .
In the preferred and illustrated
embodiment (see Fig. 1), a controller 42 is ,
associated with the generator 40, either as an
integrated unit or as a separate interface box.
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The controller 42 governs the delivery of radio
frequency ablation energy to the electrode 30
s
according to preestablished criteria. Further
details of this aspect of the system 10 will be
described later.
According to the invention, the liquid
medium 38 used to fill the body 22 includes an
electrically conductive liquid. The liquid 38
establishes an electrically conductive path, which
conveys radio frequency energy from the electrode
30. In conjunction, the body 22 comprises an
electrically non-conductive thermoplastic or
elastomeric material that contains pores 44 on at
least a portion of its surface. The pores 44 of
the porous body 22 (shown diagrammatically in
enlarged form in Fig. 4 for the purpose of
illustration) establishes ionic transport of
ablation energy from the electrode 30, through the
electrically conductive medium 38, to tissue
outside the body. Preferably, the liquid 38
possesses a low resistivity to decrease ohmic
loses, and thus ohmic heating effects, within the
' body 22. In the illustrated and preferred
embodiment, the liquid 38 also serves the
additional function as the inflation medium for
the body, at least in part.
The composition of the electrically
conductive liquid 38 can vary. In the illustrated
and preferred embodiment, the liquid 38 comprises
3o a hypertonic saline solution, having a sodium
y chloride concentration at or near saturation,
which is about 9~ weight by volume. Hypertonic
saline solution has a low resistivity of only
about 5 ohmcm, compared to blood resistivity of
about 150 ohmcm and myocardial tissue resistivity
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of about 500 ohm~cm.
Alternatively, the composition of the
electrically conductive liquid medium 38 can
comprise a hypertonic potassium chloride solution.
This medium, while promoting the desired ionic
transfer, requires closer monitoring of rate at
which ionic transport occurs through the pores 44,
to prevent potassium overload. When hypertonic
potassium chloride solution is used, it is
preferred keep the ionic transport rate below
about 10 mEq/min.
The system 10 as just described is
ideally suited for ablating myocardial tissue
within the heart. In this environment, a physician
moves the catheter tube 12 through a main vein or
artery into a heart chamber, while the expandable-
collapsible body 22 of the electrode structure 20
is in its low profile geometry. Once inside the
desired heart chamber, the expandable-collapsible
body 22 is enlarged into its expanded geometry and
the region containing pores 44 is placed into
contact with the targeted region of endocardial
tissue.
Due largely to mass concentration
differentials across the pores 44, ions in the
medium 38 will pass into the pores 44, because of
concentration differential-driven diffusion. Ion
diffusion through the pores 44 will continue as
long as a concentration gradient is maintained
across the body 22. The ions contained in the
pores 44 provide the means to conduct current
across the body 22.
Radio frequency energy is conveyed from .
the generator 40 to the electrode 30, as governed
by the controller 42. When radio frequency (RF)
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voltage is applied to the electrode 30, electric
current is carried by the ions within the pores
44. The RF currents provided by the ions result in
no net diffusion of ions, as would occur if a DC
a
voltage were applied, although the ions do move
slightly back and forth during the RF frequency
application. This ionic movement (and current
flow) in response to the applied RF field does not
require perfusion of liquid in the medium 38
through the pores 44.
The ions convey radio frequency energy
through the pores 44 into tissue to a return
electrode, which is typically an external patch
electrode (forming a unipolar arrangement).
Alternatively, the transmitted energy can pass
through tissue to an adjacent electrode in the
heart chamber (forming a bipolar arrangement). The
radio frequency energy heats the tissue, mostly
ohmically, forming a lesion.
The electrical resistivity of the body 22
has a significant influence on the lesion geometry
and controllability. It has been discovered that
ablation with devices that have a low-resistivity
body 22 requires more RF power and results in
deeper lesions. On the other hand, devices that
have a high-resistivity body 22 generate more
uniform heating, therefore, improve the
controllability of the lesion. Because of the
additional heat generated by the increased body
resistivity, less RF power is required to reach
similar tissue temperatures after the same
interval of time. Consequently, lesions generated
with high-resistivity bodies 22 usually have
smaller depth.
Generally speaking, lower resistivities
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values for the body 22 below about 500 ohm~cm
result in deeper lesion geometries. Likewise,
higher resistivities for the body 22 at or above
about 500 ohm~cm result in more shallow lesion
geometries.
The electrical resistivity of the body 22
can be controlled by specifying the pore size of
the material, the porosity of the material, and
the water adsorption characteristics (hydrophilic
versus hydrophobic) of the material.
St~eci.fya.ng Pore Size
The size of the pores 44 in the body 22
can vary. Pore diameters smaller than about 0.1
~.cm, typically used for blood oxygenation,
dialysis, or ultrafiltration, can be used for
ionic transfer according to the invention. These
small pores, which can be visualized by high-
energy electron microscopes, retain
macromolecules, but allow ionic transfer through
the pores in response to the applied RF field, as
above described. With smaller pore diameters,
pressure driven liquid perfusion through the pores
44 is less likely to accompany the ionic
transport, unless relatively high pressure
conditions develop within the body 22.
Larger pore diameters, typically used for
blood microfiltration, can also be used for ionic
transfer according to the invention. These larger
pores, which can be seen by light microscopy,
retain blood cells, but permit passage of ions in
response to the applied RF field. Generally .
speaking, pore sizes below 8 ~m will block most
blood cells from crossing the membrane. With ,
larger pore diameters, pressure driven liquid
perfusion, and the attendant transport of
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macromolecules through the pores 44, is also more
likely to occur at normal inflation pressures for
c
the body 22.
Still larger pore sizes can be used,
9
capable of accommodating formed blood cell
elements. However, considerations of overall
porosity, perfusion rates, and lodgment of blood
cells within the pores of the body 22 must be
taken more into account as pore size increase.
Conventional porous, biocompatible
membrane materials used for blood oxygenation,
dialysis, blood filtration such as plasmapheresis
can serve as the porous body 22. Such membrane
materials can be made from, for example,
regenerated cellulose, nylon, polycarbonate,
polyvinylidene fluoride (PTFE), polyethersulfone,
modified acrylic copolymers, and cellulose
acetate.
Alternatively, porous or microporous
materials may be fabricated by weaving a material
(such as nylon, polyester, polyethylene,
polypropylene, fluorocarbon, fine diameter
stainless steel, or other fiber) into a mesh
having the desired pore size and porosity. The use
of woven materials is advantageous, because woven
materials are very flexible as small diameter
fibers can be used to weave the mesh. By using
woven materials, uniformity and consistency in
pore size also can be achieved.
Spectrum Medical Industries, Inc.
y (Houston, Texas) commercially supplies nylon and
polyester woven materials with pore sizes as small
as 5 ~m with porosities of 2%. Stainless steel
woven materials with pore sizes as small as 30 um
with porosities of 30% can also be obtained from
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Spectrum Medical Industries, Inc. Manufacturers,
such as Tstko, also produce woven materials
meeting the desired specifications. Woven
materials with smaller pore sizes may be achieved
depending on the material.
Woven meshes may be fabricated by
conventional techniques, including square mesh or
twill mesh. Square mesh is formed by conventional
"over and under" methods. Twill mesh is formed by
sending two fibers over and under. The materials
may be woven into a 3-dimensional structure, such
as a tube or a sphere. Alternatively, the
materials may be woven into a flat, 2-dimensional
sheet and formed (heat forming, thermal bonding,
mechanical deformation, ultrasonic welding etc.)
into the desired 3-dimensional geometry of the
body 22.
Pore size can be specified using bubble
point measurements. The bubble point value is
defined as the pressure required to force liquid
through the membrane, which is a function mainly
of pore size (given the same water adsorption
characteristic).
Pore size correlates with the expected
liquid flow resistance of the membrane. As a '
general proposition, larger pores allow more
liquid to flow through the pores and at higher
flow rates. Likewise, smaller pores limit the
volume and rate of liquid perfusion through the
pores. At a point, a pore will be small enough to
effectively block liquid perfusion, except at very
high pressure, while nevertheless enabling ionic
transport to occur in the manner described above.
Low or essentially no liquid perfusion
*Trade-mark
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through the pores 44 is preferred. Limited or
essentially no liquid perfusion through the pores
44 is beneficial for several reasons. First, it
limits salt or water overloading, caused by
S transport of the hypertonic solution into the
blood pool. This is especially true, should the
hypertonic solution include potassium chloride, as
observed above.
Furthermore, limited or essentially no
liquid perfusion through the pores 44 allows ionic
transport to occur without disruption. When
undisturbed by attendant liquid perfusion, ionic
transport creates a continuous virtual electrode
48 (see Fig. 8) at the body 22-tissue interface.
The virtual electrode 48 efficiently transfers RF
energy without need for an electrically conductive
metal surface.
The bubble point value in psi for a given
porous material also aids in specifying the nature
of ionic transport the porous material supports,
thereby indicating its suitability for tissue
ablation.
When the bubble point value for a given
porous material exceeds the pressure required to
inflate the body 22 (i.e., inflation pressure), it
is possible to pressure inflate the body 22 into
its expanded geometry, without promoting pressure-
driven liquid perfusion through the pores 44 of
the material. Specifying a material with a bubble
point value greater than body inflation pressure
. assures that ionic transfer through the pores 44
occurs without attendant liquid perfusion through
the pores 44.
A bubble point value that is
significantly less than the body inflation
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pressure also indicates that the body 22
containing the porous material may never reach its
intended expanded geometry, because of excessive
liquid perfusion through the pores 44.
On the other hand, the bubble point value "
of the porous material should not exceed the
tensile strength of the porous material. By
specifying this relationship between bubble point
value and tensile strength, liquid perfusion will
occur before abnormally high pressures develop,
lessening the chance that the body 22 will
rupture.
The bubble point value specification
mediates against the use of larger pore size
materials. Larger pore size materials pose
problems of inflation and excessive fluid
perfusion through the membrane.
~pecifying~ Porosity
The placement of the pores 44 and the
size of the pores 44 determine the porosity of the
body 22. The porosity represents the space on the
body 22 that does not contain material, or is
empty, or is composed of pores 44. Expressed as a
percentage, porosity represents the percent volume
of the body 22 that is not occupied by the body
material.
For materials having a porosity greater
than about 10~, porosity P (in ~) can be
determined as follows:
P= 100 (1 - pb)
pm
where:
p~ is the density of the body 22 as
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determined by its weight and volume, and
pm is the density of the material from
which the body 22 is made.
To derive porosity for materials having a
porosity of less than about 10~, a scanning
electron microscope can be used to obtain the
number of pores and their average diameter.
Porosity P (in ~) is then derived as follows:
z
P =Nn( 4 )
where:
to N is the pore density and equals (pn/a),
p~ is the number of pores in the body 22,
a is the total porous area of the body 22
( in cm2) , and
~r is the constant 3.1416 ...,
d is the average diameter of the pores
(in cm) .
The magnitude of the porosity affects the
liquid flow resistance of the body 22, as
discussed above. The equivalent electrical
resistivity of the body 22 also depends on its
porosity. Low-porosity materials have high
electrical resistivity, whereas high-porosity
materials have low electrical resistivity. For
example, a material with 3~ porosity, when exposed
to 9~ hypertonic solution (resistivity of 5
ohm-cm), may have an electrical resistivity
_ comparable to that of blood or tissue (between 150
and 450 ohm~cm).
The distribution of pores 44 for a given
porosity also affects the efficiency of ionic
transport. Given a porosity value, an array of
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numerous smaller pores 44 is preferred, instead of
an array of fewer but larger pores. The presence
of numerous small pores 44 distributes current
density so that the current density at each pore
44 is less. With current density lessened, the
ionic flow of electrical energy to tissue occurs
with minimal diminution due to resistive heat
loss.
An array of numerous smaller pores 44 is
to also preferred, instead of an array of fewer but
larger pores, because it further helps to impose
favorably large liquid flow resistance. The
presence of numerous small pores 44 limits the
rate at which liquid perfusion occurs through
each pore 44.
A dynamic change in resistance across a
body 22 can be brought about by changing the
diameter of the body 22 made from a porous elastic
material, such as silicone. In this arrangement,
the elastic body 22 is made porous by drilling
pores of the same size in the elastic material
when in a relaxed state, creating a given
porosity. As the elastic body 22 is inflated, its
porosity remains essentially constant, but the
wall thickness of the body 22 will decrease.
Thus, with increasing diameter of the body 22, the
resistance across the body 22 decreases, due to
decreasing wall thickness and increasing surface
area of the body 22. As the surface area of the
body 22 increases by a factor of two, the
thickness of the body 22 will decrease by a factor _
of two, resulting in a decrease in resistance by a
factor of four.
As a result, the desired lesion geometry
may be specified according to the geometry of the
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body 22. This enables use of the same porous body
22 to form small lesions, shallow and wide
lesions, or wide and deep lesions, by controlling
the geometry of the body 22.
Preferably, the porous body 22 should
possess consistent pore size and porosity
throughout the desired ablation region. Without
consistent pore size and porosity, difference in
electrical resistance of the body 22 throughout
the ablation region can cause localized regions of
higher current density and, as a result, higher
temperature. If the difference in electrical
resistance is high enough, the lesion may not be
therapeutic, because it may not extend to the
desired depth or length. Furthermore, nonuniform
areas of low porosity in the body 22 can
themselves experience physical damage as a result
of the localized heating effects.
8peaif~inq Water Adsorption
Characteristics
The porous material for the body 22 may
be either hydrophobic or hydrophilic. However,
the water adsorption characteristics of the porous
material also affect the electrical resistivity of
the material.
For materials of the same pore size and
porosity, materials that are hydrophilic possess
greater capacity to provide ionic transfer of
radiofrequency energy without significant liquid
flow through the material. Ions suspended in the
medium are more likely to fully occupy the pores
of a hydrophilic material in the absence of a
driving pressure exceeding the bubble point value
of the material, compared to hydrophobic
materials. The presence of these ions within the
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pores in the hydrophilic materials provides the
capacity o! ionic current flow with no need for
liquid perfusion through the pores. As a result,
pore sizes may be decreased more readily with
hydrophilic materials, thereby raieinc~ the bubble
point value to minimize liquid perfusion, without
adversely affecting desired ionic current-carrying
capacities. Furthermore, the relationship between
porosity and resistivity is more direct in the
case of hydrophilic materials than with
hydrophobic materials.
Some forms of nylon (e.g., nylon 6 or
nylon 6/6) are examples of hydrophilic materials
having high water adsorption suitable for use as a
porous electrode. The nylon sample identified in
Example 3 below has 4.0~ to 4.2~ moisture
adsorption at 65~ relative humidity and a
temperature of 20~ C.
Nevertheless, conventional medical grade
"balloon" materials, such as PET and PeBa~t; are
hydrophobic. Ions in the medium are less likely
to occupy the pore of a hydrophobic membrane,
absent a driving pressure exceeding the bubble
point value of the material, compared to a
hydrophilic material.' As a result, hydrophobic
materials are more likely to require liquid flow
through the pores to carry ions into the pores, to
thereby enable transmission of electrical energy
across the porous material. With such materials,
3o the inflation pressure of the body should exceed
the bubble point value to enable effective ionic
transport.
Furthermore, due to the higher surface
tension of hydrophobic material, which tends to
restrict ion flow into the pores, hydrophobic
*Trade-mark
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materials also exhibit a greater tendency to cause
material breakdown at the pores, compared with
hydrophilic materials. The large potential
differences across each pore in a hydrophobic
material may cause dissociation of water
molecules, dielectric breakdown of the membrane
material, and localized overheating. The
breakdown is associated with high temperature
effects and, depending on the material, can open
up the pores, burn the material surrounding the
pores, and generally degrade the material. In
addition, material breakdown can produce hazardous
tissue effects similar to DC ablation, such as
tissue charring.
Therefore, changing the water adsorption
characteristics of a porous material from more
hydrophobic to more hydrophilic can offset
undesired electrical characteristics, without
changing pore size or porosity. For example, the
incidence of material breakdown due to high
current densities and potential drops at the pores
can be reduced by increasing the porosity of the
material. However, the incidence of material
breakdown can be reduced or eliminated without
altering the porosity, by selecting a material
that is hydrophilic; for example, materials such
as regenerated cellulose, nylon 6, and nylon 6/6,
which typically have high water adsorption.
Alternatively, coatings or surface treatments may
3o be applied to a less hydrophilic material making
it more hydrophilic. For example, some materials
can be dipped into a specially formulated
hydrophilic coating and exposed to ultraviolet
light to bind the coating to the material surface.
This approach-is especially advantageous when
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conventional "balloon" materials are used for the
body 22, provided the coating withstands ablation
temperatures without degrading.
Other measures can be employed to offset
other undesired electrical characteristics due to
pore size or porosity or water adsorption
properties. For example, for larger pore
materials, or when porous hydrophilic materials
are used, the perfusion rate can be controlled by
controlling fluid pressures across the body 22.
alternatively, for larger pore materials,
or when porous hydrophilic materials are used, a
material can be added to the hypertonic solution
to increase its viscosity, thereby decreasing its
perfusion rate. Examples of materials that can be
added to increase viscosity include ionic contrast
(radiopaque) substances or nonionic glycerol or
concentrated mannitol solutions.
For example, the electrical performance
of woven materials having larger pore sizes may be
aided by the addition of an ionic radiopaque
contrast material like Renografin~ -76. By adding a
radiopaque material to the aqueous solution, the
body 22 may be seen under fluoroscopy (or
echocardiography, depending on the contrast
material). The flow resistance of the porous
material will effectively increase, due to the
increased viscosity of the medium.
The use of ionic materials to increase
viscosity need not excessively increase the
resistivity of the membrane, depending on the _
concentration of the ionic material. The following
Table 1 summarizes the results of .zn vitro
experiments, using an ionic radiopaque material
with a woven nylon 13.0 mm disk probe.
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TAHLE 1
Effects of Ion Conuast Material on Ablation
with a Woven hfylon Disk
Fluid Set AverageAverage Lesion Lesion
Medium TemperaturePower ImpedanceDepth Length
996 NaCI 90 o C 23 W fib f2 9.9 21.0 mm
mm
5096-996 90 o C 13 W 85 f2 8.4 16.6 mm
NaCI mm
5096-Contrast
Contrast 90 o C 14 W 120 f2 8.6 17.9 mm
mm
Material
rvv i c: sail Lesion a~mensions are oases on ine o~ o v., aiscoiorauon
cnaraciensuc.
Table 1 shows that the ionic contrast medium
can reduce the power required to achieve
equivalent ablation results and still create
desired lesions.
For porous materials, either hydrophilic or
hydrophobic, the system 10 can include a device to
sense impedance proximate to the body-tissue
interface. As Fig. 9 shows, impedance decreases
with increasing liquid perfusion flow rate, until
a limit point is reached, at which impedance
stabilizes despite increasing perfusion rates. By
sensing impedance, it is possible to control
perfusate flow between a minimum flow rate RMtN (at
which impedance is too high) and a maximum flow
rate RMAx (above which potential salt or water
overload conditions come into existence).
The surface area of the electrode 30 bathed
in the electrically conductive medium within the
body can be increased to enhance ionic transfer.
However, the desired characteristics of small
geometry collaspsibility and overall flexibility
of the body impose practical constraints upon
electrode size.
The proximity of the electrode 30 to the
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pores 44 of the body 22 also enhances the
efficiency of ionic transfer through the
electrically conductive medium. Again, the
structural characteristics of presenting a
flexible, small collapsed profile create practical
constraints upon this consideration.
Formixict the l3od~ ~2
The expandable-collapsible body 22 can be
formed about the exterior or interior of a glass
mold. In this arrangement, the external dimensions
of the mold match the desired expanded geometry of
the expandable-collapsible body 22. The mold is
dipped in a desired sequence into a solution of
the body material until the desired wall thickness
is achieved. The mold is then etched away, leaving
the formed expandable-collapsible body 22.
Alternatively, the expandable-collapsible
body 22 may also be blow molded from extruded
tubing. In this arrangement, the body 22 is
sealed at one end using adhesive or thermal
fusion. The opposite end of the body 22 is left
open. The sealed expandable-collapsible body 22
is placed inside the mold. An inflation medium,
such as high pressure gas or liquid, is introduced
through the open tube end. The mold is exposed to
heat as the tube body 22 is inflated to assume the
mold geometry. The formed expandable-collapsible
body 22 is then pulled from the mold.
The porosity of the body 22 can be imparted
either before or after molding by COZ laser, eximer
laser, YAG laser, high power YAG laser, electronic
particle bombardment, and the like.
As earlier discussed, coatings or surface
treatments may also be applied to make the surface
more hydrophilic to improve the electrical
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properties of the body 22 for tissue ablation.
Commercially available porous materials can
also be formed into the body 22. For those
materials having poor bonding properties that are
' S formed by chemical processes, such as the
regenerated cellulose, the material may be
chemically formed into a three-dimensional
geometry by a dipping process (as generally shown
in Fig. 32A and as will be described later),
injection molding, or by varying the diameter and
geometry during extrusion.
For those materials that can be thermally
bonded, laser welded, ultrasonically welded, and
adhesively bonded, there are various ways that
make use of these bonding or welding techniques to
form a three-dimensional geometry from a sheet of
the material may be employed. Fixtures and
mandrels can be used to form the body 22 in
conjunction with heat and pressure.
Figs. 19 to 23 show a preferred way for
forming a sheet of porous material 200 into the
desired three dimensional geometry of an ablation
body 22. As Fig. 19 shows, the sheet 200 is placed
over a forming cavity 202 on a fixture 204. The
geometry of the forming cavity 202 corresponds to
the geometry desired for the distal end of the
body 22. In the illustrated embodiment, the
geometry is generally hemispherical.
As Fig. 20 shows, a forming mandrel 206
presses a section 208 of the sheet 200 into the
forming cavity 202. The geometry of the forming
mandrel 206 matches the hemispherical geometry of
the forming cavity 202. The mandrel 206 nests
within the cavity 202, sandwiching the material
section 208 between them. This sets by pressure
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the desired hemispherical shape to the material
section 208. Either the forming mandrel 206 or the
forming cavity 202, or both, may be heated to
provide an additional thermal set to the material
section 208 within the cavity 202. Pressure and,
optionally, heat within the cavity 202 shape the
material section 208 from a planar geometry into
the desired hemispherical geometry (see Fig. 21).
The sheet with the preformed section 208 is
removed from the fixture 204 and mounted upon a
finishing fixture 210 (see Fig. 22). The finishing
fixture 210 includes a distal end 212 having a
geometry that matches the geometry of the
preformed section 208. The section 208 fits on
the distal fixture end 212.
The finishing fixture 21o includes a proximal
end 214 that has the geometry desired for the
proximal end of the body 22, which in the
illustrated embodiment is hemispherical, too. The
sheet 200 is snugly draped about the proximal end
214 of the fixture 210.
The finishing fixture 210 includes a base
region 216, about which the remaining material of
the sheet 200 is gathered in overlapping pleats
218. The sheet 200 thereby tightly conforms to the
entire geometry of the fixture 210.
The finishing fixture 210 may be heated to
aid in providing an additional thermal set to the
sheet 200 in the desired geometry of the body 22.
A clam shell mold (not shown) may also be fastened
about the fixture 210 to facilitate the shaping ,
process.
The sheet material, now shaped as the porous
body 22 (see Fig. 23) is slipped from the fixture
210. The material pleats 218 that had been
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Bathared about the base region 216 of the fixture
210 are bonded together, for example, by thermal
bonding or ultrasonically welding. This forms a
reduced diameter neck region 220 in the body 22 to
lacilitata attachment of the body 22 to the distal
end of a catheter tube.
Sefors pleating, the sheet ends 21? may ba
cut into sections to minimiz~ the amount of
material which accumulates during the pleating
proces~. After pleating, the excess material may
be back-folded and bonded to the neck region 220
and/or otherwise trimmed to fore a smooth
transition between the neck region 220 and the
distal region 208.
Alternatively, after removal from the fixture
204, the sheet 200 with preformed section 208, can
ba wrapped about an expandable fixture 222 (see
Fig. 24). The proximal ands 217 of the sheet 200
are snugly tied about the neck of the fixture 222
by a tie member 223.
The fixture 222 comprises a balloon (made,
for example, from a Teflon material) or the like,
which can be expanded using gas or liquid into the
geometry desired for the body 22. fhe sheet 200
25 is thereby shaped by the expanding fixture 222 to
take the desired geometry.
Before or during expansion of the fixture
222, heat may be appiisa to the enas 2i~ of the
sheet 200 to soften the material to aid the
shaping process. External pressure may also be
applied to the proximal ends of the sheet 200 to
aid in creating the neck region 220 having the
desired reduced diameter. This also helps to
prevent "bunching" of material at the proxiiaal
°nds 2l7.
*Trade-mark
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The fixture 222 itself may also be heated by
using heated gas or liquid to expand the fixture.
The heat provides an additional thermal set to the
sheet 200 in the desired geometry of the body 22.
An external clam shell mold (not shown) may also
be fastened about the fixture 222 to facilitate
the shaping process. Alternatively. an external
shell of a material such as glass, which may be
etched away, may be used to impart the desired
final geometry.
Throughout any heating process used in
forming the body 22 using either fixture 204 or
222, a heat sink (not shown} may be used to cool
the preformed distal section 208 so that the pore
sizes do not change significantly during a heating
process.
Alternatively, the heating effects on the
pore size may be estimated and accounted for in
forming the sheet 200 in the first instance,
before shaping into the body 22. For example, if
the pores open during shaping, the pores may be
formed during manufacture proportionally smaller,
to take into account the increase in size during
shaping. Thus, the desired pore size is
ultimately achieved while shaping the sheet into
the body 22.
After the shaping process, the fixture 222 is
deflated and withdrawn (see Fig. 25). The formed
body 22 remains.
In yet another alternative process (see Figs.
26 and 27}, the body 22 can be formed by joining ,
two preformed sections 225 along a circumferential
seam 224. In the illustrated embodiment, the
sections 225 are formed as hemispheres in the
manner shown in Figs. 19 to 21, with excess
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material about the periphery of the section 225
cut away. The sections 225 could likewise be
preformed by a molding process, depending upon the
properties of the material.
The seam 224 joining the two sections 225 is
formed through thermal bonding, ultrasonically
welding, laser welding, adhesive bonding, sewing,
or the like, depending upon the properties of the
material. The bonding or sewing method employed
20 is selected to assure that the seam forms an
air-tight and liquid-sealed region. The tensile
strength of the seam 224 should also exceed the
bubble point value of the porous material.
Alternatively, two generally circular planar
sections of porous material, cut to size from a
sheet, can be joined about their peripheries by a
seam, without prior shaping. This creates a
normally collapsed disk enclosing an open
interior. The introduction of air or liquid into
the open interior during use causes the disk to
expand into the geometry desired for the body 22.
The disk could also enclose an interior support
structure 54 (as generally shown in Figs. 5 to 7),
which shapes the disk to the. desired geometry.
Preferably, after joining the hemispherical
or planar sections 225 at the seam 224, excess
material extending beyond to the seam 224 is cut
away. Still, as contact between tissue and the
somewhat roughened surface region of the seam 224
could cause trauma, the joined sections 225 are
preferably everted (see Fig. 28A). Eversion
locates the seam 224 within the interior of the
body 22 (as Fig. 28B shows), away from direct
contact with tissue.
As Fig. ~8A shows, the joined sections 225
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can be everted by creating a small hole 25o at one
end 252, inserting a pull-wire 254 and attaching
it to the opposite end 256, then pulling the
opposite end 256 through the hole 250. This turns
the attached hemispherical sections 225 inside
out.
In the foregoing embodiments, the
circumferential seam 224 extends about the axis of
the body 22. Alternatively (as Figs. 29A and 29B
show), seams 226 can extend along the axis of the
body 22 to join two or more sections 228, either
planar or preformed into a three dimensional
shape, into the body 22. Mating fixtures (not
shown) can be used, each carrying a body section
228, to hold the sections 228 stationary while
heat or ultrasonic energy is applied to create the
seam 226.
As Fig.30 shows, other axially extending
seams 230 may also be placed within a sheet of
porous material, not to join the sheet to another
sheet, but rather to segment the sheet. Further
details about segmented porous electrodes will be
discussed later. For the purpose of illustration,
Fig. 30 somewhat exaggerates the hemispherical
protrusion of the segments along the seams 226 and
230.
Preferably, the resulting body 22 is everted,
as just described, to place the axially extending
seams 226 or 230 inside the body 22.
It should be appreciated that the sections
225 or 228 shown in Figs. 26 to 30, whether planar
or preforming in three dimensional geometries,
need not be made of the same material. Materials
of different porous characteristics can be joined
by seams in the manner just described.
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Alternatively, porous materials may be joined by
seams to nonporous materials, which can themselves
be either electrically conductive or electrically
insulating. Or, still alternatively, electrically
conductive materials can be joined by seams to
insulating materials, to provide double sided
electrode bodies, one (electrically conductive)
for contacting tissue, and the other (electrically
insulating) for exposure to the blood pool.
Virtually any flexible material suitable for use
in association with an electrode body can be
combined using seams according to this aspect of
the invention. Also, it should be realized that
the number of sections joined together by seams to
form a composite electrode body can vary.
Various specific geometries, of course, can
be selected, as well. The preferred geometry is
essentially spherical and symmetric, with a distal
spherical contour, as Fig. 2 shows. However,
nonsymmetric or nonspherical geometries can be
used. For example, the expandable-collapsible body
22 may be formed with a flattened distal contour,
which gradually curves or necks inwardly for
attachment with the catheter tube 12. Elongated,
cylindrical geometries can also be used, such as
shown in Figs. 33 and 34B, which will be discussed
later.
Fig. 10 shows an alternative expandable-
collapsible porous body 50. In this embodiment,
the body 50 comprises open cell foam molded to
normally assume the shape of the expanded
geometry. The electrode 30 is encapsulated within
the foam body 50. The hypertonic liquid medium 38
is introduced into the foam body 50, filling the
open cells, to enable the desired ionic transport
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of ablation energy, as already described. The
transport of ions using the foam body 50 will also
occur if the body 50 includes an outer porous skin
51(as the right side of Fig. 10 shows), which can
provide a porosity less than the porosity of the
foam body 50 to control the perfusion rata.
In this arrangement, a sliding sheath (as
previously described) can be advanced along the
catheter tube 12 to compress the foam body 50 into
the collapsed geometry. Likewise, retraction of
the sheath removes the compression force. The
foam body 50, free of the sheath, springs open to
return the expandable-collapsible body 50 back to
the expanded geometry.
In the illustrated and preferred embodiment,
a distal steering mechanism 52 (see Fig. 1)
enhances the manipulation of the porous electrode
structure 20 or 50, both during and after
deployment. The steering mechanism 52 can vary.
In the illustrated embodiment (see Fig. 1), the
steering mechanism 52 includes a rotating cam
wheel 56 coupled to an external steering lever 58
carried by the handle 18. The cam wheel 56 holds
the proximal ends of right and left steering wires
60. The wires 60 pass'with the ablation energy
signal wires 32 through the catheter tube 12 and
connect to the left and right sides of a resilient
bendable wire or leaf spring (not shown) adjacent
the distal tube end 16. Further details of this
and other types of steering mechanisms are shows
in Lundquist and Thompson U.S. Patent 5,254,088.
As shown in Fig. 1, the leaf spring of the
steering mechanism 52 is carried within in the
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distal end 16 of the catheter tube 12. As Fig. 1
shows, forward movement of the steering lever 58
pulls on one steering wire 60 to flex or curve the
leaf spring, and, with it, the distal catheter end
16 and the electrode structure 20, in one
direction. Rearward movement of the steering
lever 58 pulls on the other steering wire 60 to
flex or curve the leaf spring 62, and, with it,
the distal catheter end 16 and the electrode
structure 20, in the opposite direction.
Alternatively, as Fig. 32C shows, a steerable
leaf spring 268 is part of a distal fixture 270,
which is itself attached to the distal end of the
porous body 22. In this arrangement, the leaf
spring 268 extends beyond the distal catheter end
16 within a tube 272 inside the porous body 22.
The steering wires 60 and 62 attached to the leaf
spring 268 also pass through the tube 272. The
proximal end of the leaf spring 268 is secured to
a hub 274 attached to the distal catheter end 16.
In this arrangement, forward and rearward
movement of the steering lever 58 on the handle 18
bends the leaf spring 268 in opposite directions
within the body 22. The leaf spring 268 moves the
distal fixture 270 and deforms the porous body 22
in the direction that the leaf spring 268 bends.
In either arrangement, the steering mechanism
54 is usable whether the expandable-collapsible
body is in its collapsed geometry or in its
expanded geometry.
Figs. 32A and 32B show a preferred way of
securing the distal fixture 270 and leaf spring
268 to a porous body 22. In Fig. 32A, the porous
body 22 is formed by dipping an expandable fixture
276 having a desired geometry into solution of
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regenerated cellulose 278. The details of such an
expandable fixture 276 have already been described
in another context and are shown in Figs. 24 and
25. It should be appreciated that the porous body
22 can be formed in various other ways, as already
described.
As Fig. 32B shows, the fixture forms a dip-
formed porous body 22 having a proximal neck
region 280 and a distal neck region 282. After
molding the body 22, the expandable fixture 276 is
collapsed and withdrawn, as Fig. 32B also shows.
As Fig. 32B shows, the distal neck region 282
is secured about the distal fixture 270, for
example using adhesive or a sleeve 288 that is
secured by adhesive bonding, thermal bonding,
mechanical bonding, screws, winding, or a
combination of any of these.
The distal fixture 270 has, preattached to
it, the leaf spring 268 and associated components,
already described. When initially secured to the
fixture 270, the proximal neck region 280 of the
body 22 is oriented in a direction opposite to the
leaf spring 268.
After securing the distal neck region 282 to
the fixture 270, the body 22 is everted about the
distal fixture 270 over the leaf spring 268, as
Fig. 32C shows. The proximal end of the leaf
spring 268 is secured to the hub 274 carried by
the distal catheter end 16. The evened proximal
neck region 280 is then secured to the distal
catheter end by use of a sleeve 286. The sleeve _
286 can be secured about the catheter tube in
various ways, including adhesive bonding, thermal
bonding, mechanical bonding, screws, winding, or a
combination of any of these.
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8
As will be described in greater detail later,
the distal fixture 270 can also serve as a
nonporous electrically co#~ductive region on the
porous body 22. Similar fixtures 270 can be
located elsewhere on the porous body 22 for the
same purpose.
A stilette (not shown) may also be attached
to the distal fixture 2?0 instead of or in
combination with the leaf spring 268. From there,
the ctilette extends inside the body 22 (following
aversion), through the catheter tube 12, to a
suitable push-pull controller on the handle 18
(not shown). The stilette is movable along the
axis of the catheter tube 12 to push and pull
axially upon the distal fixture 270, thereby
elongating or shortening the body 22.
30 Figs. 33 and 34A/34H show exemplary electrode
bodies having elongated, cylindrical geometries,
which can be associated with various distal
fixtures in the manner shown and attached to
distal catheter ends 16 in the manner shown in
Figs. 32H and 32C.
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In Fig. 33, the body 290 is formed by
extrusion, dipping, or molding into an elongated
geometry with varying radii to form the distal and
proximal neck regions 280 and 282. A suitable
distal fixture 270 (shown in phantom lines) can be
secured within the distal neck region 282 and the
elongated body 22 everted to complete the
assembly, in the manner shown in Figs. 32B and
32C. The proximal neck region 280 can then be
secured to a distal catheter end 16 in the manner
shown in Fig. 32C.
In Figs. 34A and 34B, the body 22 is formed
from a tube 292 of material formed by extrusion,
molding, or dipping with a uniform radius (shown
in Fig. 34A). In this arrangement (see Fig. 34B),
a seam 294, formed the manner previously
disclosed, closes the distal end of the tube 292.
The proximal end of the tube 292 is sealed about a
tubular port 296, for attachment to the distal
catheter end 16. Alternatively, the distal end of
the tube 292 can be sealed about a distal fixture
270 (shown in phantom lines in Fig. 34B). In the
latter case, the tube 292 is everted about the
distal fixture 270 before attachment to the
catheter distal end 16.
The pattern of pores 44 that define the
porous region of the body may vary. Preferably,
as generally shown in Figs. 2 and 3, the region of
at least the proximal 1/3rd surface of the
expandable-collapsible body 22 is free of pores
44.
The absence of pores 44 on the at least
proximal 1/3rd surface of the expandable-
collapsible body 22 is desirable for several
reasons. This region is not normally in contact
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- 41 -
with tissue, so the presence of the virtual
electrode boundary serves no purpose.
Furthermore, this region also presents the
smallest diameter. If electrically conductive,
this region would possess the greatest current
density, which is not desirable. Keeping the
proximal region of smallest diameter, which is
usually free of tissue contact, free of pores 44
assures that the maximum current density will be
distributed at or near the distal region of the
expandable-collapsible body 22, which will be in
tissue contact.
When it is expected that ablation will occur
with the distal region of body 22 oriented in end-
on contact with tissue, the porous region should,
or course, be oriented about the distal tip of the
expandable-collapsible body 22. For this end-on
orientation, the porous region may comprise a
continuous cap deposited upon the distal 1/3rd to
'~ of the body 22, as Figs. 2 and 3 show. However,
when distal contact with tissue is contemplated,
the preferred embodiment (see Fig. 11) segments
the electrically conductive porous region into
separate energy transmission zones 62 arranged in
a concentric "bulls eye" pattern about the distal
tip of the body 22.
When it is expected that ablation will occur
with the side region of the body 22 oriented in
contact with tissue, the porous region is
preferably segmented into axially elongated energy
transmission zones 62 (see Fig. 12), which are
circumferentially spaced about the distal 1/3rd to
of the body.
When the porous region comprises segmented
zones 62 on the body 22, an interior group of
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- 42 -
sealed bladders 64 (see Fig. 13) can be used to
individually convey liquid 38 to each porous
region segment 62. Each bladder 64 individually
communicates with a lumen 66 to receive the
electrically conductive liquid for the one porous
region 62 it services. The multiple lumens pass
through the catheter tube 12. The multiple
bladders 64 also provide the ability to more
particularly control the geometry of the expanded
body 22, by selectively inflating with the liquid
some but not all the bladders 64.
The bladders 64 may be separately formed and
inserted into the body 22, or they may be
- integrally formed during molding the main
expandable-collapsible body 22.
As Fig. 12 shows, segmented porous zones 62
are also well suited for use in association with
folding expandable-collapsible bodies 22. In this
arrangement, the regions that are free of pores
comprise creased or folding regions 68. To create
these regions 68, the mold for the body 22 has a
preformed surface geometry such that the
expandable-collapsible material would be formed
slightly thinner, indented, or ribbed along the
desired regions 68. The expandable-collapsible
body 22 collapses about these creased regions 68,
causing the body 22 to circumferentially fold upon
itself in a consistent, uniform fashion. The
resulting collapsed geometry can thus be made more
uniform and compact.
It should be appreciated that the foldable _
body 22 shown in Fig. 12 can also be used for
other patterns of porous regions. The creased
regions 68 can also be provided with pores, if
desired.
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Fig. 14 shows an embodiment of an expandable-
collapsibie electrode structure 70 that serves
dual functions. The structure 70 includes an
expandable-collapsible body 22, as previously
described, containing the interior electrode 30.
The body 22 contains an electrically conductive
fluid 38, and also includes one or more porous
regions 62 to enable ionic transport of electrical
energy, also as just described.
The structure 70 shown in Fig. 14 also
includes one or more nonporous, electrically
conductive regions 72 on the surface of the body
22. In one embodiment (as Fig. 14 shows), the
nonporous conductive regions 72 comprise metal,
such as gold, platinum, platinum/iridium, among
others, deposited upon the expandable-collapsible
body 22 by sputtering, vapor deposition, ion beam
deposition, electroplating over a deposited seed
layer, or a combination of these processes.
Alternatively, the nonporous conductive regions 72
can comprise thin foil affixed to the surface of
the body. Still alternatively, the nonporous
conductive regions can comprise solid fixtures
(like the distal fixture 270 shown in Fig. 32C)
carried by the porous body 22 at or more
locations. Signal wires (not shown) within the
body are electrically coupled to the nonporous
regions. The signal wires traverse the catheter
tube 12 for coupling to the connectors 38 carried
by the handle 18.
_ In the preferred embodiment (see Fig. 15),
the nonporous conductive regions 72 comprise
_ insulated signal wires 26 passed into the interior
of the body and then snaked through the body 22 at
the desired point of electrical connection. The
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electrical insulation of the distal end of the
snaked-through wire 26 is removed to exposed the
electrical conductor, which is also preferably
flattened, to serve as the conductive region 72.
The flattened region 72 is affixed by an
electrically conductive adhesive 73 to body 22.
Adhesive 73 is also preferable applied in the
region of the body 22 where the wire 26 passes to
seal it. The same signal wire 26 can be snaked
through the body 22 multiple times to establish
multiple regions 72, if desired.
The nonporous regions 72 can be used to sense
electrical activity in myocardial tissue. The
sensed electrical activity is conveyed to an
external controller, which processes the
potentials for analysis by the physician. The
processing can create a map of electrical
potentials or depolarization events for the
purpose of locating potential arrhythmia foci.
Once located with the nonporous regions 72, the
porous regions 62 can be used to convey radio
frequency energy as previously described to ablate
the foci.
Alternatively, or in combination with sensing
electrical activities, the nonporous regions 72
can be used to convey pacing signals. In this
way, the nonporous regions can carry out pace
mapping or entrainment mapping.
Preferably (see Figs. 16j, the interior
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surface of the body 22 carries electrodes 100
suitable for unipolar or bipolar sensing or
pacing. Although these electrodes 100 are located
on the interior surface of the body 22, their
ability for sensing or pacing is not impaired
because of the good electrical conductive
properties of the body 22.
Different electrode placements can be used
for unipolar or bipolar sensing or pacing. For
example, pairs of 2-mm length and 1-mm width
electrodes 100 can be deposited on the interior
surface of the body 22. Connection wires 102 can
be attached to these electrodes 100. To prevent
the hypertonic solution from electrically
short-circuiting these electrodes, they have to be
covered with an electrically insulating material
104 (e.g. epoxy, adhesive etc.). Preferably the
interelectrode distance is about 1 mm to insure
good quality bipolar electrograms. Preferred
placements of these interior electrodes are at the
distal tip and center of the structure 22. Also,
when multiple zones are used, it is desired to
have the electrodes 100 placed in between the
ablation regions.
It is also preferred to deposit opaque
markers 106 on the interior surface of the body 22
so that the physician can guide the device under
fluoroscopy to the targeted site. Any high-atomic
weight material is suitable for this purpose. For
example, platinum, platinum-iridium. can be used
__ to build the markers 106. Preferred placements of
these markers 106 are at the distal tip and center
of the structure 22.
The expandable-collapsible structure 70 shown
in Fig. 14 thereby combines the use of "solid"
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nonporous electrodes 72 with ''liquid" or porous
electrodes 62. The expandable-collapsible
structure makes possible the mapping of myocardial
tissue for therapeutic purposes using one
electrode function, and the ablation of myocardial
tissue for therapeutic purposes using a different
electrode function.
In an alternative embodiment, the nonporous
regions 72 of the structure 70 can be used in
tandem with the porous regions 62 to convey radio
frequency energy to ablate tissue. In this
arrangement, the signal wires serving the region
72 are electrically coupled to the generator 40 to
convey radio frequency energy for transmission by
one or more regions 72. At the same time, the
interior electrode 30 receives radio frequency
energy for transmission by the medium 38 through
the porous body. The ionic transport across the
porous structure surrounding the regions 72
extends the effective surface area of the ablation
electrode.
In this embodiment, the expandable-
collapsible structure 70 shown in Fig. 14 thereby
combines the use of electrodes 72 having a first
effective surface area for sensing and mapping.
The first effective surface area can be
selectively increased for ablation purposes by
ionic transport of a hypertonic liquid across a
porous structure surrounding the electrodes 72.
If liquid perfusion occurs through the pores,
an interior electrode 30 is not required to
increase the effective electrode surface area of
the regions. The liquid perfusion of the ionic
medium through the pores at the time the regions
transmit radio frequency energy is itself
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sufficient to increase the effective transmission
surface area of the regions 72. However, if ionic
transfer occurs without substantial liquid
perfusion, it is believed that it would be
advantageous in increasing the effective surface
area to also transmit radio frequency energy using
an interior electrode 30 at the same time that
radio frequency is being delivered to the exterior
regions 72 for transmission.
It should also be appreciated that, in this
embodiment, the regions 72 can themselves be made
from a porous, electrically conducting material.
In this way, ionic transport can occur across the
regions 72 themselves.
As before described (see Fig. 1), a
controller 32 preferably governs the conveyance of
radio frequency ablation energy from the generator
30 to the electrode carried within the body 22.
In the preferred embodiment (see Fig. 2), the
porous electrode structure 20 carries one or more
temperature sensing elements 104, which are
coupled to the controller 32.
The temperature sensing elements 104 can take
the form of thermistors, thermocouples, or the
equivalent. The sensing elements 104 are in
thermal conductive contact with the exterior of
the electrode structure 20 to sense conditions in
tissue outside the structure 20 during ablation.
Temperatures sensed by the temperature
sensing elements 104 are processed by the
_ controller 32. Based upon temperature input, the
controller adjusts the time and power level of
radio frequency energy transmissions by the
electrode 30, to achieve the desired lesion
patterns and other ablation objectives.
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As Figs. 31A, 318, and 31C show, temperature
sensing elements 104 can also be positioned
proximal to or within a seam 258 joining sheets
l0 260 and 262 of porous material together into a
body 22. The formation of such seams 258 has been
already described and is also shown in Figs. 26 to
30.
As shown in Figs. 31A and 31B, each
15 temperature sensing element 104 is placed on one
sheet 260, and then covered by the other sheet
262. The two sheets 260 and 262 are then seamed
together, forming the body 22. The seam 258
encapsulates the sensing elements. The signal wire
20 264 for each sensing element 104 extends.free of
the seam 258 to the exterior of the sheets 260 and
262, as Figs. 31A and 31B show.
As previously described, the body 22 is
preferably evened (see Fig. 28A). As Fig. 31C
25 shows, aversion locates both the seam 258 and the
encapsulated signal wires 264 within the interior
of the body 22. The signal wires 264 are passed
through the neck 266 for coupling to the
controller 32.
30 Instead of or in addition to the temperature
sensing elements 104, pacing electrodes or sensing
electrodes may be encapsulated within the seams
258 of averted electrode bodies 22 in the manner
shown in Figs. 31A to 31C. In such arrangements,
35 it is preferred to locate the electrodes so that,
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after eversion, they are located within the seams
close to the surface of the material where
intended contact between the body material and
tissue is to take place.
Further details of the use of multiple
ablation energy transmitters controlled using
multiple temperature sensing elements are
disclosed in U.S. Patent No. 5,769,847.
La i
In Vitro llnalysis
Three electrode configurations were analyzed:
(1) a porous expandable-collapsible
electrode structure made according to the
invention, having a 13 mm disk-shaped body
constructed from dialysis tubing made from
regenerated cellulose(manufactured by Spectra),
with a molecular weight cut off of 12,000 - 14,000
Daltons, and using 9% saline solution as the
internal liquid medium;
(2) a sputtered platinum disk-shaped
electrode body having a diameter of 13 mm; and
(3) an expandable-collapsible hemispherical
electrode structure constructed of aluminum foil
with a diameter of 10 mm.
A thermistor was embedded 0.5 mm into animal
tissue (sheep) at the region where maximum
temperature conditions existed. For electrodes (1)
and t2), the maximum temperature was at edges of
the disk-shaped body, due to edge heating effects.
For the electrode (3), the maximum temperature was
at the distal'tip of the hemispherical body, where
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current densities are greatest.
Radio frequency electromagnetic power was
regulated to maintain the thermistor temperature
at 60°C, 70°C, 80°C, or 90°C. All lesion
dimensions were measured based on the 60°C
isotherms, marking the discoloration of tissue.
The following Table 2 lists the observed in
vitro results.
TABhE 2
ElectrodeTime Average Average AverageLesion Lesion
(Sec) Power ImpedanceTamp Depth Length
(Watts) (Ohms) (' C) (mm) (mm)
1 i 20 11 75 59 4.2 15.8
1 120 21 72 68 6.4'"" 17.5
1 120 17 69 78 8.0 17.1
1 120 15 62 78 9.T"' 18.3
1 120 25 68 87 8.9T" 19.5
1 120 26 66 87 8.8'" 20.8
1 120 22 69 88 10.0'"'19.1
2 128 9 76 78 6.3 14.2
3 120 23 69 78 6.0 13.7
3 120 37 55 86 7.4 18.5
Note: TlVI indicates that the lesion was transmural, so depths were actually
larger
than measured.
The porous electrode structure (1) was
minimally affected by convective cooling compared
to normal metal ablation electrode, such as
electrode (2). This was observed by varying fluid
flow about the electrode during a particular
lesion and observing that, with a porous electrode
structure, no change in power was required to
maintain thermistor temperature.
Table 2 shows that the porous electrode
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structure created lesions at least as large as a
metal coated electrode structures when regulating
power based on tissue temperature. The porous
electrode structure also had reasonable impedance
levels compared to the metal coated electrode
structures.
This Example demonstrates that a porous
electrode structure can create lesions deeper than
1.0 cm in a controlled fashion to ablate
epicardial, endocardial, or intramural VT
substrates.
The dialysis tubing forming the porous
electrode structure has a high water adsorption
characteristic. The dialysis tubing becomes
significantly more flexible when exposed to water.
The molecular weight cutoff was 12,000 to 14,000
daltons. Larger or smaller molecular weight cut-
offs are available from Spectrum. The conversion
from molecular weight cutoff to estimated pore
size for the dialysis tubing tested is 100,000
daltons equals 0.01 pm; 50,000 daltons equals
0. 004 ;Cm; 10, 000 daltons equals 0.0025 ~Cm; 5, 000
daltons equals 0.0015 ~tm.
The dialysis tubi.zg possesses a hydrophilic
nature and high porosity despite low pore sizes.
As a result, the bubble point value is extremely
high and the resistivity is low enough to not
require fluid flow during delivery of
radiofrequency energy.
EYAMPLE 2
Finite Bl~meat Analysis
A three-dimensional finite element model was
created for a porous electrode structure having a
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body with an elongated shape, with a total length
of 28.4 mm, a diameter of 6.4 mm, and a body wall
thickness of .1 mm. A .2-mm diameter metal wire
extended within the length of the body to serve as
an interior electrode connected to an RF energy
source. The body was filled with 9~ hypertonic
solution, having an electrical resistivity of 5.0
ohm~cm. The porous body of the structure was
modeled as an electric conductor. Firm contact
with cardiac tissue was assumed along the entire
length of the electrode body lying in a plane
beneath the electrode. Contact with blood was
assumed along the entire length of the electrode
body lying in a plane above the electrode. The
blood and tissue regions had resistivities of 150
and 500 ohm-cm, respectively.
Analyses were made based upon resistivities
of 1.2 k-ohm~cm and 12 k-ohm~cm for the electrode
body.
Table 3 shows the depth of the maximum tissue
temperature when RF ablation power is applied to
the porous electrode at various power levels and
at various levels of resistivity for the porous
body of the electrode.
TABLE 3
Resistivity ~epth of
of Power Time Maximum Maximum
the Porous Tissue Tissue
Body (Watts) (Sec) Temp (C) Temp(cmy
(0c-ohmcm)
3 0 1.2 58 120 96.9 1.1
1.2 58 240 97.9 1.4
12 40 120 94.4 0.8
12 40 240 95.0 1.0 -
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Fig. 17 shows the temperature profiles when
power is applied to the electrode at 58 watts for
y 240 seconds when the porous body has a resistivity
of 1.2 k-ohmcm. The depth of 50 C isotherm in
Fig. 17 is 1.4 cm.
Fig. 18 shows the temperature profiles when
power is applied to the electrode at 40 watts for
240 seconds when the porous body has a resistivity
of 12 k-ohmcm. Depth of 50 C isotherm in Fig.
18 is 1.0 cm.
In all cases, the maximal temperature is
located at the interface between tissue and the
opposite end edges of the elongated porous
structure. This dictates that the preferred
location for temperature sensing elements for an
elongated geometry of the porous body is at each
end edge of the body. Preferably, each edge
should carry at least one temperature sensing
element, and multiple sensing elements should be
located in diametrically opposite sides to assure
that at least one of them faces tissue.
The data also show that the hottest region is
not moved deep into the tissue, as would be
observed with metal surface electrodes. The
hottest region consistently resides at the
tissue-electrode body interface for direct
sensing. This feature reduces the difference
between sensed temperature and actual hottest
tissue temperature potentially to a theoretical O
C, although somewhat higher differentials may be
encountered given other aspects of the
instrumentation.
The porous electrode body with higher
resistivity body (see Fig. 18) generated more
uniform temperature profiles, compared to a porous
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body having the lower resistivity value (see Fig.
17). Due to additional heating generated at the
tissue-electrode body interface with increased
electrode body resistivity, less power was
required to reach same maximal temperature. The
consequence was that the lesion depth decreased.
As before explained, by selecting the
resistivity of the body 22, the physician can
significantly influence lesion geometry. The use
of a low-resistivity body 22 results in deeper
lesions, and vice versa. The following Table 4,
based upon empirical data, demonstrates the
relationship between body resistivity and lesion
depths.
TAHI~B 4
ResistivityPower TemperatureLesion Depth
(ohmcm) (Watts) (C) (cm) Time
isec)
850 94 97 1.2 120
1200 58 97 1.1 120
2 0 7 2,000 40 95 0.8 120
Because of the reduced thermal conductivity
of the porous electrode structure, when compared
to nonporous, metallic surface electrodes, lesion
formation is expected to be less sensitive to
dynamic blood flow conditions around the
electrode. The application of ablation energy
through porous electrodes can be more closely
controlled to obtain desired lesion
characteristics, particularly when shallow atrial
lesion are desired. -
The following Table 5, based upon empirical
data, demonstrates the reduced sensitivity of -
porous electrode structures to connective cooling
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Connective Maximum
Cooling Temperature Power (Watts)Lesion Depth
Conditions 1' C) Icm)
Normal 97 94 1.2
5096 Reduced 98 94 1.3
7596 Reduced 95 79 0.8
The use of porous electrode structures
provides structural benefits. It isolates
possible adherence problems that may be associated
with the placement of metal, electrically
conductive shells to the outside of expandable-
collapsible bodies. Porous electrode structures
also avoid potential problems that tissue sticking
to exterior conductive materials could create.
In addition to these structural benefits, the
temperature control of the ablation process is
improved. When using a conventional metal
electrode to ablate tissue, the tissue-electrode
interface is convectively cooled by surrounding
blood flow. Due to these connective cooling
effects, the region of maximum tissue temperature
is located deeper in the tissue. As a result, the
temperature conditions sensed by sensing elements
associated with metal electrode elements do not
directly reflect actual maximum tissue
temperature. In this situation, maximum tissue
temperature conditions must be inferred or
predicted from actual sensed temperatures. Using a
porous electrode structure 20 or 70, connective
cooling of the tissue-electrode interface by the
surrounding blood flow is minimized. As a result,
the region of maximum temperature is located at
conditions due to changes in blood flow rates.
TABLE 5
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the interface between tissue and the porous
electrode. As a result, the temperature
conditions sensed by sensing elements associated
with porous electrode elements will more closely
reflect actual maximum tissue.
E%AEPLE 3
In Vitro experiments were performed to
compare hydrophilic materials (Hphl) versus
hydrophobic materials (HPhb) in terms of their use
as porous tissue ablation elements. Table 6
summarizes the results.
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57 _
TABhE 6
Summary of Porous Ablation Materials
Mat'I Mfgr HPhb HPhI Pore BubbleNo Flow ImpedancMat'i Lesion
Size point Impedancea w/ BrkdwnDepth
value Flow
DialysisSpectrum ~/ .025frtnHigh 87 fl 87 fl No 13.8
mm
Tubing
Nylon Spectrum V/ 5 ErtnMad 88 f2 88 i2 No 9.9
mm
Mesh
Stain-StSpectrum 30,vm Low 87 i2 B7 f2 No 9.7
mm
Mesh
PolycarbMillipore ~ 1.2 High 78 t2 78 i2 No 11.8
~rtn 14 mm
Film psi
Polyvin-Millipore~ 5 /rtnHigh > 300 84 fl Yes 10.7
f2 mm
ylidene w/
flow
Fluoride
PTFE Millipore~ 5 ~rtnHigh > 300 > 300 N/A NONE
f2 !;2
PolyethersGelman ~ 5 /rtnMed 80 i2 80 f2 No '10.8
1- mm
ulfone 6 psi
PolyethersGelman ~ O.i High > 300 > 300 N/A NONE
~rtn f2 t2
ulfone
ModifiedGelmen ~ 10 Mad 88 f1 B8 i2 Yes 9.9
Ertn mm
Acrylic 1-6
psi
copolymer
ModifiedGelman ~/ 5 frtnHigh > 300 70 fl Yes 11.0
f2 mm
Acrylic w/
flow
copolymer
ModifiedGelman ~ 10 High > 300 61 f2 Yea 11.3
/rtn fl mm
Acrylic w/
w/ flow
3 0 backing
PTFE Pore ~ 1 ,rm~High > 300 > 300 N/A NONE
Tech f2 f2
CelluloseGoodfelio ~ Very High > 300 > 300 N/A NONE
S2 fl
Acetatew low
Note: "Mat'I Brkdwn" refers to the presence of material breakdown, as
described
above.
- Table 6 demonstrates that pore sizes may be
decreased using hydrophilic materials, thereby
minimizing or stopping liquid perfusion through
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the porous material, while still enabling ionic
transport through the membrane.
Hydrophobic porous materials make possible
the realization of high resistivity porous
. 5 electrodes. On the other hand, hydrophilic porous
materials make possible the realization of low
resistivity porous electrodes.
obtaining Desired Lesion charactsristias
As the foregoing tables demonstrate, the same
expandable-collapsible porous electrode structure
is able to selectively form lesions that are
either wide and shallow or large and deep.
Various methodologies can be used to control the
application of radio frequency energy to achieve
15 this result.
A. D5~ Function
In one representative embodiment, the
controller 42 includes an input 300 (see Fig. 1)
for receiving from the physician a desired
20 therapeutic result in terms of (i) the extent to
which the desired lesion should extend beneath the
tissue-electrode interface to a boundary depth
between viable and nonviable tissue and/or (ii) a
maximum tissue temperature developed within the
lesion between the tissue-electrode interface and
the boundary depth.
The controller 42 also includes a processing
element 302 (see Fig. 1), which retains a function
that correlates an observed relationship among
lesion boundary depth, ablation power level,
ablation time, actual sub-surface tissue _
temperature, and electrode temperature. The
processing element 302 compares the desired
therapeutic result to the function and selects an
operating condition based upon the comparison to
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achieve the desired therapeutic result without
exceeding a prescribed actual or predicted
sub-surface tissue temperature.
The operating condition selected by the
processing element 302 can control various aspects
of the ablation procedure, such as controlling the
ablation power level, limiting the ablation time
to a selected targeted ablation time, limiting the
ablation power level subject to a prescribed
maximum ablation power level, and/or the
orientation of the porous region 44 of the body
22, including prescribing a desired percentage
contact between the region 44 and tissue. The
processing element 302 can rely upon temperature
sensors carried by or otherwise associated with
the expandable-collapsible structure 20 that
penetrate the tissue to sense actual maximum
tissue temperature. Alternatively, the processing
element 302 can predict maximum tissue temperature
based upon operating conditions.
In the preferred embodiment, the electrode
structure 20 carries at least one temperature
sensing element 104 to sense instantaneous
localized temperatures (T1} of the thermal mass of
the region 44. The temperature T1 at any given
time is a function of the power supplied to the
electrode 30 by the generator 40.
The characteristic of a lesion can be
expressed in terms of the depth below the tissue
surface of the 50° C isothermal region, which will
be called DSOC- The depth DSOC is a function of the
physical characteristics of the porous region 44
(that is, its electrical and thermal
conductivities, resistivities, and size}; the
percentage of-contact between the tissue and the
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porous region 44; the localized temperature T1 of
the thermal mass of the region 44; the magnitude
of RF power {P) transmitted by the interior
electrode 30, and the time (t) the tissue is
exposed to the RF power.
For a desired lesion depth Dsoc, additional
considerations of safety constrain the selection
of an optimal operating condition among the
operating conditions listed in the matrix. The
principal safety constraints are the maximum
tissue temperature TMAX and maximum power level
PMAX.
The maximum temperature condition TMAX lies
within a range of temperatures which are high
Z5 enough to provide deep and wide lesions {typically
between about 85° C and 95° C), but which are
safely below about 100° C, at which tissue
desiccation or tissue micro-explosions are known
to occur. Tt is recognized that TMAX will occur a
distance below the electrode-tissue interface
between the interface and DSOC
The maximum power level PMAX takes into
account the physical characteristics of the
interior electrode 30 and the power generation
capacity of the RF generator 40.
These relationships can be observed
empirically and/or by computer modeling under
controlled real and simulated conditions, as the
foregoing examples illustrate. The Dsoc function
for a given porous region 44 can be expressed in
terms of a matrix listing all or some of the .
foregoing values and their relationship derived
from empirical data and/or computer modeling.
The processing element 302 includes in memory
this matrix of operating conditions defining the
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D5~ temperature boundary function, as described
above !or t = 120 seconds and TMAX = 95~ C and for
an array of other operating conditions.
Th~ physician also uses the input 300 to
identify the characteristics of the structure 20,
using a prescribed identification code; set a
desired maximum RF power level PMAX; a desired
time t; and a desired maximum tissue temperature
TMAX.
Based upon these inputs, the processing
element 302 compares the desired therapeutic
result to the function defined in the matrix. The
generator 42 selects an operating condition to
achieve the desired therapeutic result without
exceeding the prescribed TMAX by controlling the
function variables.
This arrangement thereby permits the
physician, in effect, to "dial-a-lesion" by
specifying a desired Due.
Further details of. deriving the DS~ function
and its use in obtaining a desired lesion pattern
are found ~n U.S. Patent No. 6,056,745.
8. 8egmentad Regioas: Duty Cpole Control
Various RP' energy control schemes can also be
used in conjunction with segmented porous patterns
shown in~Fig. 11 (the axially spaced, bull s-eye
pattern of zones) and Fig. 12 (the
circumferentially spaced zones) For the purpose
of discussion, the porous zones 44 (which will
also be called electrode regions) will be
symbolically designated E(J), where J represents a
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given zone 44 (J = 1 to N}.
As before described, each electrode region
E(J} has at least one temperature sensing element
104, which will be designated S(J,K}, where J
represents the zone and K represents the number of '
temperature sensing elements on each zone (K = 1
to M}.
In this mode, the generator 40 is conditioned
through an appropriated power switch interface to
deliver RF power in multiple pulses of duty cycle
1/N.
With pulsed power delivery, the amount of
power (PE~~~} conveyed to each individual electrode
region E(J} is expressed as follows:
PE w~ cx AMP~c ~ 2 x DUTYCYC.LEEt .n
where:
AMPEC~~ is the amplitude of the RF voltage
conveyed to the electrode region E(J}, and
DUTYCYCLEE~~~ is the duty cycle of the pulse,
expressed as follows:
D UTYCYC_LE~,t ~ ~ TONE t~ ~
TONS t~~ + TOFFE ~~~
where:
TONE~~~ is the time that the electrode region
E(J} emits energy during each pulse period,
TOFFE~~~ is the time that the electrode region
E(J) does not emit energy during each pulse
period.
The expression TON~~~~ + TOFFE~~~ represents the
period of the pulse for each electrode region
E (J} .
In this mode, the generator 40 can
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collectively establish duty cycle (DUTYCYCLEE~~~) of
1/N for each electrode region (N being equal to
the number of electrode regions).
The generator 40 may sequence successive
power pulses to adjacent electrode regions so that
the end of the duty cycle for the preceding pulse
overlaps slightly with the beginning of the duty
cycle for the next pulse. This overlap in pulse
duty cycles assures that the generator 40 applies
power continuously, with no periods of
interruption caused by open circuits during pulse
switching between successive electrode regions.
In this mode, the temperature controller 42
makes individual adjustments to the amplitude of
the RF voltage for each electrode region (AMPE~~~),
thereby individually changing the power PE~~~ Of
ablating energy conveyed during the duty cycle to
each electrode region, as controlled by the
generator 40.
In this mode, the generator 40 cycles in
successive data acquisition sample periods.
During each sample period, the generator 40
selects individual sensors S(J,K), and temperature
codes TEMP(J) (highest of S(J,K)) sensed by the
sensing elements 104, as outputted by the
controller 42.
When there is more than one sensing element
104 associated with a given electrode region (for
example, when edge-located sensing elements are
used, the controller 42 registers all sensed
temperatures for the given electrode region and
selects among these the highest sensed
temperature, which constitutes TEMP(J).
In this mode, the generator 40 compares the
temperature TEMP(J) locally sensed at each
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electrode E(J) during each data acquisition period
to a set point temperature TEMPseT established by
the physician. Based upon this comparison, the
generator 40 varies the amplitude AMPEt~~ of the RF
voltage delivered to the electrode region E(J), '
while maintaining the DUTYCYCLEEt~~ for that
electrode region and all other electrode regions,
to establish and maintain TEMP(J) at the set point
temperature TEMPser .
The set point temperature TEMPSeT can vary
according to the judgment of the physician and
empirical data. A representative set point
temperature for cardiac ablation is believed to
lie in the range of 40°C to 95° C, with 70° C
being a representative preferred value.
The manner in which the generator 40 governs
AMPEt~~ can incorporate proportional control
methods, proportional integral derivative (PID)
control methods, or fuzzy logic control methods.
For example, using proportional control
methods, if the temperature sensed by the first
sensing element TEMP(1) > TEMPSET, the control
signal generated by the generator 30 individually
reduces the amplitude AMPE«> of the RF voltage
applied to the first electrode region E(1), while
keeping the duty cycle DUTYCYCLE Et~~ for the first
electrode region E(1) the same. If the temperature
sensed by the second sensing element TEMP(2) <
TEMPS~T, the control signal of the generator 30
increases the amplitude AMPEtz~ of the pulse applied
to the second electrode region E(2), while
keeping the duty cycle DUTYCYCLEEtZ~ for the
second electrode region E(2) the same as DUTYCYCLE
Etl~, and so on. If the temperature sensed by a
given sensing~element is at the set point
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temperature TEMP ~T, no change in RF voltage
amplitude is made for the associated electrode
region.
The generator 40 continuously processes
voltage difference inputs during successive data
acquisition periods to individually adjust AMPE~~~
at each electrode region E(J), while keeping the
collective duty cycle the same for all electrode
regions E(J). In this way, the mode maintains a
desired uniformity of temperature along the length
of the ablating element.
Using a proportional integral differential
(PID) control technique, the generator takes into
account not only instantaneous changes that occur
in a given sample period, but also changes that
have occurred in previous sample periods and the
rate at which these changes are varying over time.
Thus, using a PID control technique, the generator
will respond differently to a given proportionally
large instantaneous difference between TEMP (J)
and TEMP~t, depending upon whether the difference
is getting larger or smaller, compared to previous
instantaneous differences, and whether the rate at
which the difference is changing since previous
sample periods is increasing or decreasing.
Further details of individual
amplitude/collective duty cycle control for
segmented electrode regions based upon temperature
sensing are found in U.S. Patent No. 5,810,802.
C. Segmented Regions: Differential
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Temperature Disablina~
In this control mode, the controller 42
selects at the end of each data acquisition phase
the sensed temperature that is the greatest for
that phase (TEMPs~X). The controller 42 also
selects for that phase the sensed temperature that
is the lowest (TEMPSMt~) .
The generator compares the selected hottest
sensed temperature TEMPs~X to a selected high set
point temperature TEMPHiser ~ The comparison
generates a control signal that collectively
adjusts the amplitude of the RF voltage for all
electrode regions using proportional, PID, or
fuzzy logic control techniques.
In a proportion control implementation
scheme:
(i) If TEMPs~X > TEMPN1SET ~ the control
signal collectively decreases the
amplitude of the RF voltage
delivered to all regions;
(ii) If TEMPs~ < TEMPHiser , the control
signal collectively increases the
amplitude of the RF voltage
delivered to all regions:
(iii) If TEMPs~ = TEMPHiser , no change in
the amplitude of the RF voltage
delivered to all regions.
It should be appreciated that the generator
can select for amplitude control purposes any one
of the sensed temperatures TEMPs~AX, TEMPsMiN, or
temperatures in between, and compare this ,
temperature condition to a preselected temperature
condition.
The generator governs the delivery of power to
the regions based upon difference between a given
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local temperature TEMP (J) and TEMP~IW. This
implementation computes the difference between
local sensed temperature TE1KP (J) and TEMP~~w and
compares this difference to a selected set point
temperature difference oTF,~iPmT. The comparison
generates a control signal that governs the
delivery of power to the electrode regions.
If the local sensed temperature TEIdP(J) for a
given electrode region E(J) exceeds the lowest
1o sensed temperature TEMPmi~ by as much as or more
than ~TEMP~ (that is, if TEMP(J) - TEMP~~w
oTEMP$ET), the generator turns the given region
E(J) off. The generator turns the given region
E (J) back on when TEMP (J) - TEMP~t~ < dTEMP~.
Alternatively, instead of comparing TEMP(J)
and TEMP~tw, the generator can compare TE~iPand
TEMP~IN. When the difference between TEMPS and
TE'MP~~w equals or exceeds a predetermined amount
~TEMPs~, the generator turns all regions off,
except the region where TEMP~Iw exists. The
generator 30 turns these regions back on when the
temperature dif f erence between TEMPS and TEt~P~~w
is less than ~TEMPmr.
Further details of the use of differential
temperature disabling are found in
U.S. Patent No. 5,769,847.
D. BID Control
With porous electrode structures, the minimal
effects of connective cooling by the blood pool
enables the use of actual sensed temperature
conditions as~maximum tissue temperature TMAX ,
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instead of predicted temperatures. Because of
this, such structures also lend themselves to the
use of a proportional integral differential (PID)
control technique. An illustrative PID control
techniques usable in association with,these
electrode structures are disclosed in copending
U.S. Patent No. 5,755,715.
Finally, it should be appreciated that
interior electrodes 30 located within porous
expandable-collapsible bodies can be used for
mapping myocardial tissue within the heart. In
this use, the interior electrodes sense electrical
activity in the heart, which can take the form,
for example, of electrical potentials or tissue
resistivity: The sensed electrical activity is
conveyed to an external controller, which
processes the sensed activities for analysis by
the physician.
It should further be appreciated that interior
electrodes 30 located within porous expandable-
collapsible bodies can be used alternatively, or
in combination with sensing electrical activities,
to convey pacing signals. In this way, the
interior electrodes 30 can carry out pace mapping
or entrainment mapping.
Various features of the invention are set
forth in the following claims.
