Patent Description:
Cardiac arrhythmia, such as atrial fibrillation, occurs when regions of cardiac tissue abnormally conduct electric signals to adjacent tissue, thereby disrupting the normal cardiac cycle and causing asynchronous rhythm. Important sources of undesired signals are located in various tissue regions in or near the heart, for example, the ventricles, the atria and/or and adjacent structures such as areas of the pulmonary veins. Regardless of the sources, unwanted signals are conducted abnormally through heart tissue where they can initiate and/or maintain arrhythmia.

Procedures for treating arrhythmia include surgically disrupting the origin of the signals causing the arrhythmia, as well as disrupting the conducting pathways for such signals. More recently, it has been found that by mapping the electrical properties of the heart muscle in conjunction with the heart anatomy, and selectively ablating cardiac tissue by application of energy, it is possible to cease or modify the propagation of unwanted electrical signals from one portion of the heart to another. The ablation process destroys the unwanted electrical pathways by formation of non-conducting lesions.

In this two-step procedure--mapping followed by ablation--electrical activity at points in the heart is typically sensed and measured by advancing a catheter containing one or more electrical sensors into the heart, and acquiring data at a multiplicity of points. These data are then utilized to select the target areas at which ablation is to be performed.

A typical ablation procedure involves the insertion of a catheter having a tip electrode at its distal end into a heart chamber. A reference electrode is provided, generally taped to the patient's skin. Radio frequency (RF) current is applied to the tip electrode, and flows through the surrounding media, i.e., blood and tissue, toward the reference electrode. The distribution of current depends on the amount of electrode surface in contact with the tissue, as compared to blood which has a higher conductivity than the tissue. Heating of the tissue occurs due to its electrical resistivity. If the tissue is heated sufficiently, cellular and other protein destruction ensues; this in turn forms a lesion within the heart muscle which is electrically non-conductive. During this process, heating of the electrode also occurs as a result of conduction from the heated tissue to the electrode itself. If the electrode temperature becomes sufficiently high, possibly above <NUM> degree C, blood clot could form on the surface of the electrode. If the temperature continues to rise, more blood clot is formed while dehydration ensues.

The tip temperature increase and the associated clot formation have two consequences: increased electrical impedance and increased probability for stroke. The former relates to clot dehydration. Because dehydrated biological material has a higher electrical resistance than heart tissue, impedance to the flow of electrical energy into the tissue also increases. Increased impedance leads to sub-optimal energy delivery to the tissue which results in inadequate lesion formation, reduced ablation efficiency and eventually to sub-optimal clinical outcome. The latter, a safety hazard, is due to possible dislodgment of the formed clot and relocation in the brain vasculature. It is therefore beneficial from a safety perspective as well as ablation efficiency to minimize the tip temperature increase and clot formation. This should be accomplished without compromising the formation of lesions of appropriate sizes.

In a typical application of RF current to the endocardium, circulating blood provides some cooling of the ablation electrode. However, there is typically a stagnant area between the electrode and tissue which is susceptible to the formation of dehydrated proteins and coagulum. As power and/or ablation time increases, the likelihood of an impedance rise also increases. As a result of this process, there has been a natural upper bound on the amount of energy which can be delivered to cardiac tissue and therefore the size of RF lesions. In clinical practice, it is desirable to reduce or eliminate impedance rises and, for certain cardiac arrhythmias, to create larger lesions. One method for accomplishing this is to monitor the temperature of the ablation electrode and to control the RF current delivered to the ablation electrode based on this temperature. If the temperature rises above a pre-selected value, the current is reduced until the temperature drops below this value. This method has reduced the number of impedance rises during cardiac ablations but has not significantly increased lesion dimensions. The results are not significantly different because this method continues to rely on the cooling effect of the blood which is dependent on the location within the heart and the orientation of the catheter to the endocardial surface.

Another method is to irrigate the ablation electrode, e.g., with physiologic saline at room temperature, to actively cool the ablation electrode instead of relying on the more passive physiological cooling provided by the blood. Additionally, due to the irrigation-mediated dilution of blood around the tip, the probability for clot creation is further reduced. Thus, irrigation tip cooling and blood dilution allow for safer increase of applied RF power. This results in lesions which tend to be larger usually measuring about <NUM> to <NUM> in depth.

The clinical effectiveness of irrigating the ablation electrode is dependent upon the distribution of flow within and around the surface of the tip electrode structure as well as the rate of irrigation flow through the tip. Effectiveness is achieved by reducing the overall electrode temperature and eliminating hot spots in the ablation electrode which can initiate coagulum formation. More channels and higher flows are more effective in reducing overall temperature and temperature variations, i.e., hot spots. Irrigation is utilized during the entire time the catheter resides inside the patient's body. Higher flow rate is used during ablation while lower-maintenance-flow rate is required in order to prevent back flow of blood into the coolant passages during non-ablation time. The coolant flow rate must be balanced against the amount of fluid that can be safely injected into the patient. Thus, reducing coolant flow by utilizing it as efficiently as possible is a desirable design objective.

One method for designing an ablation electrode which efficiently utilizes coolant flow is the use of a porous material structure. Such designs have the advantage of distributing the coolant evenly across the entire electrode structure. This balanced cooling results in a) eradication of possible surface or interior hot spots, and b) uniform dilution of blood at the vicinity of the electrode, thus further minimizing the chance for clot formation. Such designs are described in <CIT> and <CIT> Moaddeb describes the use of sintered metal particles to create a porous tip electrode. In addition, Moaddeb uses a non-conductive insert implanted into the porous tip electrode for mounting a thermocouple, lead wire and/or irrigation tube within the porous tip electrode. However, during irrigation the sintered metal particles can disintegrate and break away from the electrode structure. This-undesirable-particle dislodgement may be further facilitated during ablation. Additionally, (and in the context of our magnetic resonance imaging (MRI) compatibility claims-see below) the metallic material proposed for such porous tip is not optimal for MRI imaging. Furthermore, the proposed tip does not allow for high density mapping-a highly desired feature for accurate arrhythmia diagnosis. Consequently, a desire arises for a porous electrode having increased structural integrity, being compatible with the MRI environment, and allowing for high mapping density.

A porous tip electrode catheter is also described in <CIT>. The porous tip electrode comprises a porous material through which fluid can pass. The porous tip electrode is covered with a thin coating of conductive metal having openings (pores) through which fluids can pass. However, the porosity of such thin conductive coating is not easily controlled leading to inconsistent pore size and distribution. Therefore, distribution of irrigation fluid around the tip electrode may not be even or uniform. Furthermore, in this design, RF power delivery is achieved via direct connection (e.g. by soldering or other similar technique) of the RF power line to the tip's outer conductive coat. Thus, the presence of the generally non-uniform porous coating is necessary in order to establish electrical contact of the tip to the heart tissue.

Safe and efficacious ablation depends not only on optimal irrigation arrangement for the tip but also on accurate mapping of the electrophysiological behavior of the heart, which would allow for accurate diagnosis and appropriate tissue targeting. The greater the accuracy of the mapping the more accurate the diagnosis and thus the effectiveness of treatment. Improved (high resolution) cardiac mapping requires the use of a multitude of electrodes in close proximity to sense electrical activity within a small area, for example, a square centimeter or less.

Metallization of ceramics is a well-established technique and is widely used in a multitude of electronics and engineering disciplines, including fabrication of RF electronic circuits. Metallization involves the application of metal on ceramic substrates, including the formation of conductive regions, such as metallized conductor patterns or uniform metal layers on surfaces of ceramic substrates. Common ceramic substrates include aluminum oxide, beryllium oxide, ferrite, barium titanate, as well as quartz or borosilicate. Generally, ceramic metallization processes fall into three categories: thin-film, thick-film, and co-firing techniques. In the thin film approach, a thin layer of metal is deposited by vacuum processes such as sputtering, evaporation, chemical vapor deposition, and laser ablation. Electroless and electrolytic plating are also frequently grouped in the thin film category. To enhance adhesion, a preliminary adhesion-promoting layer, such as chromium or titanium, is often deposited. Thick film methods involve printing metal pastes, typically metal powders mixed with glass frits and organic binders onto ceramic substrates. The printed substrates are fired to form conductive paths on the ceramic. In the co-firing approach, unfired "green" ceramic surfaces are coated with patterned metal paste lines. The printed green ceramic is fired both to sinter the material and form the conductive metal patterns. Metallization processes are described, for example, in <CIT>; <CIT>; <CIT>; and <CIT>. Metallization depending on the type of metallization process and the substrate may include gold, platinum, or other biocompatible metals suitable for intracardial signal acquisitions.

While ablation has revolutionized the treatment of cardiac arrhythmias, ablation can be improved where physicians can assess lesions in real time. The use of magnetic resonance imaging (MRI) during an ablation procedure could enable physicians to assess lesions in real time. ; However, the ablation catheter and other associated accessory equipment can interfere with the imaging process, causing local distortions in the MRI scans. Use of appropriate MRI compatible materials is necessary to minimize these image distortions. Safety experts have cleared some metals for use during MRIs, including titanium, cobalt-chromium, copper, selected stainless steel alloys. Non-ferromagnetic metals are also MRI compatible. Such materials include copper, brass, silver, gold, aluminum, lead, magnesium, platinum and tungsten. Ceramic materials as well as other thermoplastic polymers are non-metallic and as such are highly desirable as MRI compatible materials. They not only present minimal image distortion but being electrical insulators they present no heating effects due to absence of internally induced electrical currents. Ceramic materials of porous construction are proposed in the current invention as materials for the construction of the catheter's tip.

In view of the foregoing, it is desirable to provide a catheter with a dome tip electrode made of a porous substrate for more uniform irrigation, where the dome tip electrode incorporates surface electrodes made via a metallization, printing or other process for any desirable surface electrode pattern that provides multiple electrodes in close proximity for high density mapping. It is also desirable to provide a catheter where the substrate and the surface electrodes are MRI compatible so that the physician can conduct lesion assessment in real time during an ablation procedure. <CIT> describes systems and methods for heating or ablating body tissue placed in contact with an electrode, which includes an exterior wall peripherally surrounding an interior area. A lumen conveys a medium containing ions into the interior area. At least a portion of the exterior wall of the electrode comprises a porous material sized to pass ions contained in the medium. The systems and methods transmit electrical ablation energy to the medium for ionic transport through the porous material to tissue.

Whilst no claim is directed to surgical methods per se, the catheter of the present invention is able of being used and is intended to be used in such methods.

The present invention is directed to a catheter having a multifunctional "virtual" tip electrode with a porous substrate and a multitude of surface microelectrodes. The surface microelectrodes in close proximity to each other and in a variety of configurations sense tissue for highly localized intracardiac signal detection, and high density local electrograms and mapping and the porous substrate allows for flow of conductive fluid for ablating tissue. The surface microelectrodes can be formed via a metallization process that allows for any shape or size and close proximity, and the fluid "weeping" from the porous substrate provides more uniform irrigation in the form of a thin layer of saline. The delivery of RF power to the catheter tip is based on the principle of "virtual electrode," where the conductive saline flowing through the porous tip acts as the electrical connection between the tip electrode and the heart surface.

Moreover, the substrate and the surface electrodes are constructed of MRI compatible materials so that the physician can conduct lesion assessment in real time during an ablation procedure. The surface electrodes include noble metals, including, for example, platinum, gold and combinations thereof.

In some embodiments, the catheter includes an elongated catheter body, and a distal electrode member having a porous substrate and a plurality of distinct surface microelectrodes. A plurality of lead wires are connected to the surface microelectrodes for transmitting electrical signals sensed by the microelectrodes. The porous substrate has an interior chamber adapted to receive conductive fluid which is in electrical contact with a lead wire that extends into the chamber, wherein such electrified fluid passes from the chamber to outside the substrate for distal irrigation and tissue ablation.

In some detailed embodiments, the porous substrate is comprised of a ceramic material. The substrate has a plurality of surface microelectrodes ranging between about one and <NUM>. Each surface microelectrode has a surface area ranging between <NUM><NUM> and <NUM><NUM>.

In some detailed embodiments, the porous substrate and the chamber both have a generally cylindrical shape, with a generally uniform wall thickness between the chamber and the outer surface of the substrate.

These and other features and advantages of the present invention will be better understood by reference to the following detailed description when considered in conjunction with the accompanying drawings wherein:.

In one embodiment of the invention, there is provided a steerable catheter having an irrigated tip adapted for diagnostic and/or therapeutic procedures. As shown in <FIG>, catheter <NUM> comprises an elongated catheter body <NUM> having proximal and distal ends, an intermediate deflection section <NUM> extending from a distal end of the catheter body <NUM>, a tip electrode section <NUM> extending from a distal end of the catheter body <NUM>, and a control handle <NUM> at the proximal end of the catheter body <NUM>.

With reference to <FIG> and <FIG>, the catheter body <NUM> comprises an elongated tubular construction having a single, axial or central lumen <NUM>. The catheter body <NUM> is flexible, i.e., bendable but substantially non-compressible along its length. The catheter body <NUM> can be of any suitable construction and made of any suitable material. In one embodiment, the catheter body <NUM> comprises an outer wall <NUM> made of a polyurethane or PEBAX. The outer wall <NUM> comprises an imbedded braided mesh of high-strength steel, stainless steel or the like to increase torsional stiffness of the catheter body <NUM> so that, when the control handle <NUM> is rotated axially, the rest of the catheter, including the sections <NUM> and <NUM>, also rotates axially. The outer diameter of the catheter body <NUM> is not critical, but is preferably no more than about <NUM> french (<NUM> french = <NUM>/<NUM>), more preferably about <NUM> french, still more preferably about <NUM> french. Likewise, the thickness of the outer wall <NUM> is not critical, but is thin enough so that the central lumen <NUM> can accommodate an irrigation tube, puller wire(s), lead wires, and any other wires, cables or tubes. The inner surface of the outer wall <NUM> is lined with a stiffening tube <NUM>, which can be made of any suitable material, such as polyimide or nylon. The stiffening tube <NUM>, along with the braided outer wall <NUM>, provides improved torsional and longitudinal stability. The outer diameter of the stiffening tube <NUM> is about the same as or slightly smaller than the inner diameter of the outer wall <NUM>. Polyimide tubing is presently preferred for the stiffening tube <NUM> because it may be very thin walled while still providing very good stiffness. This maximizes the diameter of the central lumen <NUM> without sacrificing strength and stiffness. A particularly preferred catheter has an outer wall <NUM> with an outer diameter of from about <NUM> inches (<NUM> inch = <NUM>) to about <NUM> inches and an inner diameter of from about <NUM> inches to about <NUM> inches and a polyimide stiffening tube <NUM> having an outer diameter of from about <NUM> inches to about <NUM> inches and an inner diameter of from about <NUM> inches to about <NUM> inches.

As shown in <FIG>, <FIG>, and <FIG>, the intermediate deflectable section <NUM> comprises a short section of tubing <NUM> having multiple lumens, including off-axis lumens <NUM>, <NUM>, <NUM> and <NUM>. The tubing <NUM> is made of a suitable non-toxic material that is preferably more flexible than the catheter body <NUM>. In one embodiment, the material for the tubing <NUM> is braided polyurethane, i.e., polyurethane with an imbedded mesh of braided high-strength steel, stainless steel or the like. The outer diameter of the deflection section <NUM>, like that of the catheter body <NUM>, is preferably no greater than about <NUM> french, more preferably about <NUM> french, still more preferably about <NUM> french. The size of the lumens is not critical. In one embodiment, the deflection section <NUM> has an outer diameter of about <NUM> french (<NUM> inches) and the second lumen <NUM> and third lumen <NUM> are generally about the same size, each having a diameter of from about <NUM> inches to about <NUM> inches, preferably about <NUM> inches, with the first and fourth lumens <NUM> and <NUM> having a slightly larger diameter of from about <NUM> inches to about <NUM> inches, preferably about <NUM> inches.

A means for attaching the catheter body <NUM> to the deflection section <NUM> is illustrated in <FIG> and <FIG>. The proximal end of the deflection section <NUM> comprises an outer circumferential notch <NUM> that receives the inner surface of the outer wall <NUM> of the catheter body <NUM>. The deflection section <NUM> and catheter body <NUM> are attached by adhesive (e.g. polyurethane glue) or the like. Before the deflection section <NUM> and catheter body <NUM> are attached, however, the stiffening tube <NUM> is inserted into the catheter body <NUM>. The distal end of the stiffening tube <NUM> is fixedly attached near the distal end of the catheter body <NUM> by forming a glue joint (not shown) with polyurethane glue or the like. Preferably, a small distance, e.g., about <NUM>, is provided between the distal end of the catheter body <NUM> and the distal end of the stiffening tube <NUM> to permit room for the catheter body <NUM> to receive the notch <NUM> of the deflection section <NUM>. A force is applied to the proximal end of the stiffening tube <NUM>, and, while the stiffening tube <NUM> is under compression, a first glue joint (not shown) is made between the stiffening tube <NUM> and the outer wall <NUM> by a fast drying glue, e.g. Super Glue. Thereafter, a second glue joint (not shown) is formed between the proximal ends of the stiffening tube <NUM> and outer wall <NUM> using a slower drying but stronger glue, e.g. polyurethane.

At the distal end of the deflection section <NUM> is the distal tip electrode section <NUM> having a connector tube <NUM> and a tip electrode <NUM>. In the illustrated embodiment of <FIG> and <FIG>, the connector tube <NUM> is a relative short piece of tubing, about <NUM> in length, for example, made of polyetheretherketone (PEEK). The proximal end of the connector tube <NUM> has a circumferential notch whose outer surface is surrounded by an inner surface of a circumferential notch formed in the distal end of the tubing <NUM> of the deflection section <NUM>. The ends are bonded to each other by polyurethane glue or the like.

As shown in <FIG>, the tip electrode <NUM> has a diameter about the same as the outer diameter of the tubing <NUM> and the connector tube <NUM>. The tip electrode <NUM> includes a porous substrate <NUM> and a plurality of surface electrodes <NUM>. The porous substrate <NUM> is formed by porous ceramic material or any other suitable non-conductive polymer, such as polyethylene, or Teflon. In the illustrated embodiment, the substrate <NUM> has an elongated cylindrical shape with a narrower proximal stem portion 38N. The substrate <NUM> is formed with an interior chamber <NUM> that also has a similar elongated cylindrical shape extending longitudinally in the substrate <NUM>. In one embodiment, the porous substrate <NUM> has a total length ranging from about <NUM> to about <NUM>, more preferably about <NUM>. For a <NUM> long tip electrode, each of the body form 38B and the proximal stem portion 38N may have a length of about <NUM>.

The chamber <NUM> has an opening 37P at the proximal end of the substrate <NUM> and a distal end 37D near the distal end of the substrate. It is understood that the chamber <NUM> and the substrate need not have the same general shape, and further that depending on the volume of the chamber <NUM> the thickness T of the wall between the chamber <NUM> and the outer surface of the substrate may be varied as desired or appropriate.

Abutting the proximal face of the substrate stem portion 38N, a plug member <NUM> seals the proximal face and plugs the opening thus enclosing the chamber <NUM>. As shown in <FIG>, the plug member <NUM> has a first through-hole <NUM> for lead wire <NUM> to pass through and enter into the chamber <NUM> and a second through-hole <NUM> for receiving a distal end of an irrigation tubing <NUM> which supplies fluid, e.g., saline or any electrically conductive fluid, into the chamber <NUM>.

In the embodiment of <FIG>, the proximal stem portion 38N is received in a distal end of the connector tube <NUM>. The stem portion 38N and the connector tube <NUM> are attached by polyurethane glue or the like.

The porous non-conductive material of the substrate <NUM> can be made using any conventional technique In the illustrated embodiment, the non-conductive material comprises sintered ceramic powder, or polymer particles formed from polyethylene or Teflon. As used herein, the term "sinter" refers to the process of bonding adjacent particles in a powder mass or compacting the particles by heating them to a temperature below the melting point of the main constituent at a predetermined and closely controlled time-temperature regime, including heating and cooling phases, in a protective atmosphere. The porosity of the sintered material is controlled by the amount of particle compacting in the mold or glue, the particle size, and the particle distribution. The sintered particles permit passage of a cooling fluid through the tip electrode, as described in more detail below. The final shape of the tip can be obtained with a variety of techniques including machining, grinding, etching, or molding.

In one embodiment, a sintering process involves providing ceramic, polyethylene or Teflon powder particles in a certain sieve fraction, e.g., in the range of from about <NUM> microns to about <NUM> microns. The particles are preferably in the range of from about <NUM> microns to about <NUM> microns. In a particularly preferred embodiment, at least two different sized particles can be provided. For example, particles in the range of from about <NUM> microns to about <NUM> microns, and more preferably about <NUM> microns, in combination with particles in the range of from about <NUM> microns to about <NUM> microns, and more preferably about <NUM> microns, could be used. When two different sized particles are used, preferably the larger particles have a mean diameter at least about <NUM> times greater than the mean diameter of the smaller particles, and more preferably at least about <NUM> times greater. Alternatively, a single particle size can be used, which can provide a denser packing and result in a higher pressure drop across the porous electrode. Whatever material is used, the particles are preferably rounded and more preferably spherical, so as to provide a tip electrode surface that is not rough. However, the particles can be irregularly shaped, i.e. having differing shapes, which is a low cost alternative. Tip surface irregularities could also be smoothed through secondary operations such as mechanical polishing and laser etching.

In one process, the particles are put into a mold, such as a ceramic mold, having the desired electrode shape. If desired, the particles can be mixed with a suitable binder prior to being put into the mold. When a binder is used, the mold containing the binder and particles is placed into a low temperature oven and heated to a temperature sufficient to evaporate the binder. The particles are then sintered under vacuum or air at a temperature ranging from about <NUM> degree C to about <NUM> degree C, although the temperature can vary depending on the composition of the porous polymer. However, the temperature should be below the melting point of the composition. The resulting tip electrode is then removed from the mold and assembled onto the flexible tubing of the tip section.

In the embodiment of <FIG>, the porous substrate <NUM> has a generally cylindrical shape with domed distal end. However, it is understood that the porous substrate may have different shapes, as desired or appropriate. For example, porous substrate <NUM> of <FIG> has a bulbous shape and an elongated stem portion 138N, along with a bulbous shaped chamber <NUM> with an elongated proximal portion137P. The wall thickness T is generally uniform throughout the substrate <NUM>. Moreover, it is understood that the porous substrate and the chamber may have dissimilar shapes. For example, substrate <NUM> of <FIG> is generally cylindrical with a domed distal end, but its chamber <NUM> is cylindrical with a narrowed distal end 237D which is adapted for receiving and anchoring a distal end of the electrifying lead wire. Likewise, the wall thickness may vary throughout the substrate, some portions T1 being thinner and other portions T2 being thicker.

Disposed over the surface of the porous substrate are the one or more sensing microelectrodes <NUM> in the form of individual and separate thin metal coatings, as depicted in <FIG>. It is understood that the thin metal coatings may be applied in close proximity to each other using any suitable process, including, for example, metallization, core plating, electroplating and/or <NUM>-D printing, and may involve more than one layers, with the outer most layer comprising suitable electrode material (or alloys) known in the art, such as gold, platinum, platinum/iridium. In accordance with a feature of the present invention, the one or more metal coatings are made of conductive material that is also MRI-compatible, (e.g. platinum or gold). In some embodiments, the metal coating <NUM> is made of a platinum-iridium alloy, e.g. <NUM>% Platinum/<NUM>% Iridium, applied to the surface of the porous substrate <NUM> by metallization treatment or process impregnating a thin layer of platinum-iridium alloy onto the porous substrate <NUM>, as known in the art.

The thickness of the metal coating may vary as desired. The thickness can be uniform or not uniform. For example, the metal coating may have a uniform thickness ranging from <NUM> to about <NUM>. In some embodiments, coating forming one or more microelectrodes 40X, as shown in <FIG>, has non-uniform thickness e.g. thicker towards the center and thinner towards the periphery. This allows for a protrusion configuration or a raised profile. The ratio of the central thickness H to the thickness at the periphery may range between about <NUM> and <NUM>. Such-protruding-shape allows for improved contact with the heart tissue and consequently improved electrogram quality.

The configuration of the micro-electrodes is as defined in claim <NUM>. As shown in <FIG>, the metal coatings can be of any desired plurality and of any desired configuration and/or orientation to form individual and separate surface micro-electrodes, for example, circular, oval, rectangular, elongated, ring, axial, radial, and co-centric. For example, in <FIG>, the metal coatings provide axial proximal rectangular surface micro-electrodes 40T, more distal ring surface micro-electrodes 40R, and distal tip circular surface micro-electrode 40C. For example, in <FIG>, the metal coatings provide axial proximal rectangular surface micro- electrodes 40TP and axial distal rectangular surface micro-electrodes 40TD that are axially offset from each other. For example, in <FIG>, the metal coatings provide a plurality (four) of rings surface micro-electrodes 40R and a distal tip circular surface micro-electrode 40C. For example, in <FIG>, the metal coatings provide a distal tip circular surface micro-electrode 40C, a plurality of smaller circular micro-electrodes 40C and a proximal ring surface micro-electrode 40R. For example, in <FIG>, a series of concentric circular surface micro-electrodes 40C1, 40C2 and 40C3 of different radii is shown on one side of the tip. It is understood that the size of the micro-surface electrodes in the drawings herein, including <FIG>, is not to scale and that their size is exaggerated so as to show their structure with better clarity.

Advantageously, the surface electrodes <NUM> are sized as micro-electrodes for obtaining highly localized electrograms and providing high density mapping of heart tissue. The surface area of each surface electrode ranges between about <NUM><NUM> and <NUM><NUM>, preferably between about <NUM><NUM> and <NUM><NUM>. In that regard, it is understood that the figures herein are not necessarily to scale. The plurality of surface electrodes on the substrate may range between about one and <NUM>, preferably about two and <NUM>. Each surface electrode <NUM> is connected to a respective lead wire <NUM> whose proximal end terminates in the control handle <NUM> in an input jack (not shown) that may be plugged into an appropriate signal processor (not shown). The lead wires <NUM> extend from the control handle <NUM> and through the central lumen <NUM> of the catheter body (<FIG>), the first lumen <NUM> of tubing <NUM> of the deflection section <NUM> (<FIG>), and the lumen of the connector tube <NUM> (<FIG>). The portion of the lead wires <NUM> extending through at least the catheter body <NUM> and the deflection section <NUM> may be enclosed within a protective sheath (not shown), which can be made of any suitable material, preferably polyimide. The protective sheath may be anchored at its distal end to the proximal end of the deflection section <NUM> by gluing it in the first lumen <NUM> with polyurethane glue or the like.

The lead wires <NUM> are attached or electrically connected to the surface electrodes <NUM> through surface electrode leads <NUM> (<FIG>) which may be applied or deposited on the outer surface of the porous substrate <NUM> and the stem portion 38N in the same manner as the surface microelectrodes <NUM>, as described above. As shown in <FIG> and <FIG>, distal end portions of lead wires <NUM> pass between an inner surface of the connector tubing <NUM> and the peripheral edge of the plug member <NUM> and the outer surface of the stem portion 38N. These surfaces at and near the proximal end of the substrate <NUM> may be sealed by glue and the like. Distal ends of the lead wires <NUM> are attached to respective proximal ends of the surface electrode leads <NUM> at or near the distal end of the connector tubing <NUM>. Accordingly, electrical signals of the heart tissue sensed by the microelectrodes are transmitted proximally toward the control handle via the surface electrode leads <NUM> and the lead wires <NUM>.

As understood by one of ordinary skill in the art, selected surface electrode leads <NUM> and surface sensing microelectrodes <NUM> are insulated from each other where they overlap each other. An insulating layer may be placed in between surface electrode leads <NUM> and surface electrodes <NUM> and grooves <NUM> (<FIG>, <FIG> and <FIG>) may be formed on the outer surface of the porous substrate for underpassing surface electrode leads <NUM> so that overlying surface electrodes <NUM> can lie flat on the outer surface of the porous substrate <NUM>.

For ablation purposes, the porous substrate <NUM> is "energized" by the lead wire <NUM> which passes into the chamber <NUM> via the first through-hole <NUM> in the plug member <NUM>. When energized, the lead wire <NUM> renders the porous substrate <NUM> into a "virtual" ablation electrode by conducting the energy through the conductive- irrigation fluid, e.g., saline, delivered by the irrigation tubing <NUM> which enters the chamber <NUM> and weeps through the porous substrate <NUM> in providing a generally uniform thin layer of energized fluid throughout its exposed surfaces <NUM> (in between the surface microelectrodes <NUM>) to further improve ablation safety. Wherever the fluid is present on or flowing from the porous substrate <NUM>, ablation may be accomplished therefrom.

In the embodiment of <FIG>, the distal portion of the lead wire <NUM> in the chamber <NUM> is elongated and linear. However, it is understood that the distal portion may assume any shape as desired or appropriate. The distal portion of the lead wire <NUM> in the chamber <NUM> may be configured nonlinearly, for example, wrapped around itself (<FIG>) or coiled around a support member <NUM> (<FIG>) for increased surface area exposure and contact with the fluid in the chamber <NUM> for greater conduction between the lead wire and the fluid. A proximal end of the support member <NUM> may be affixed to and mounted on a distal face of the plug member.

The distal portion of the lead wire <NUM> may also extend linearly and deeply distally in the chamber <NUM> along the longitudinal center axis (<FIG>), spiral widely approaching the inner surface of the chamber <NUM> (<FIG>), or be wrapped or coiled around an extended distal portion of the irrigation tubing <NUM> such that both extend deeply distally in the chamber <NUM>. The irrigation tubing <NUM> may be perforated with pores <NUM> along its length (<FIG>). Such configuration improves the uniformity of irrigation within the chamber <NUM> and the tip <NUM> and allows for even greater exposure of the lead wire to the conductive fluid.

In the illustrated embodiment, the catheter includes three ring electrodes <NUM> proximal of the distal tip section <NUM>, mounted on the tubing <NUM> of the deflection section <NUM> and/or the connector tubing <NUM>, as shown in <FIG>, <FIG> and <FIG>. It is understood that the presence and number of ring electrodes <NUM> may vary as desired, likewise their function as monopolar or bipolar electrodes for local electrogram sensing and/or location referencing in relation to the location sensor <NUM> housed in the connector tubing <NUM>. Each ring electrode <NUM> is slid over the tubing <NUM> and/or <NUM> and fixed in place by glue or the like. The ring electrodes <NUM> can be made of any suitable material, and are preferably machined from platinum-iridium bar (<NUM>% platinum/<NUM>% iridium), gold, or gold alloys.

Connection of a lead wire <NUM> to a ring electrode <NUM> is preferably accomplished by first making a small hole through the tubing <NUM> and/or <NUM>. Such a hole can be created, for example, by inserting a needle through the tubing and heating the needle sufficiently to form a permanent hole. A lead wire <NUM> is then drawn through the hole by using a microhook or the like. The ends of the lead wire <NUM> are then stripped of any coating and soldered or welded to the underside of the ring electrode <NUM>, which is then slid into position over the hole and fixed in place with polyurethane glue or the like.

The irrigation tubing <NUM> is provided within the catheter body <NUM> for infusing fluids, e.g. saline, to electrify the porous substrate <NUM> of the tip electrode <NUM> and provide cooling during ablation. The irrigation tubing <NUM> may be made of any suitable material, and is preferably made of polyimide tubing. In one embodiment, the irrigation tubing has an outer diameter of from about <NUM> inches to about <NUM> inches and an inner diameter of from about <NUM> inches to about <NUM> inches.

The irrigation tubing <NUM> extends from the control handle <NUM> and through the central lumen <NUM> of the catheter body <NUM> (<FIG>), the lumen <NUM> of the tubing <NUM> of the deflection section <NUM> (<FIG>), and the connector tube <NUM> (<FIG>), and into the second through-hole <NUM> in the plug member <NUM> and the chamber <NUM> of the substrate <NUM> (<FIG>). The proximal end of the irrigation tubing <NUM> extends through the control handle <NUM> to a fluid source and a pump (not shown). The fluid introduced through the catheter is preferably a biologically compatible fluid such as saline, or water. In addition to, or instead of, being used to cool the tip electrode, the infused fluid also forms a buffer layer to maintain biological materials, such as blood, at a distance from the tip electrode, thereby minimizing contact of the tip electrode with the biological material. This buffer layer reduces coagulation of biological materials and regulates the impedance or resistance to energy transfer of the tissue near the tip electrode during ablation. Saline or any other conductive fluid is preferred where the tip electrode is to function as an ablative electrode.

The rate of fluid flow through the catheter may be controlled by any suitable fluid infusion pump or by pressure. A suitable infusion pump is the COOLFLOW available from Biosense Webster, Inc. (Diamond Bar, CA). The rate of fluid flow through the catheter preferably ranges from about <NUM>/min to about <NUM>/min, more preferably from about <NUM>/min to about <NUM>/min. Preferably, the fluid is maintained at about room temperature.

It is understood that a temperature sensing means is provided for the tip electrode <NUM>, as known in the art. Any conventional temperature sensing means, e.g., a thermocouple or thermistor, may be used. A suitable thermistor for use in the present invention is Model No. AB6N2-GC14KA143E/37C sold by Thermometrics (New Jersey). The temperature sensing means may also be used as a feedback system to adjust the RF power delivered to the tissue through the catheter to maintain a desired temperature at the tip electrode.

As shown in <FIG> and <FIG>, a pair of puller wires <NUM> and <NUM> extend through the catheter body <NUM> for bidirectional deflection. The puller wires <NUM> and <NUM> are anchored at their proximal end to the control handle <NUM>, and are anchored at their distal ends to the deflection section <NUM> at or near its the distal end. The puller wires are made of any suitable metal, such as stainless steel or Nitinol, and may be coated with Teflon or the like. The coating imparts lubricity to the puller wires. Each of the puller wires may have a diameter ranging from about <NUM> inches to about <NUM> inches.

A compression coil <NUM> is situated within the catheter body <NUM> in surrounding relation to each puller wire <NUM> (<FIG>). Each compression coil <NUM> extends from the proximal end of the catheter body <NUM> to about the proximal end of the deflection section <NUM>. The compression coils are made of any suitable metal, preferably stainless steel. Each compression coil <NUM> is tightly wound on itself to provide flexibility, i.e., bending, but to resist compression. The inner diameter of the compression coil is slightly larger than the diameter of the puller wire. The Teflon coating on the puller wires <NUM> and <NUM> allows them to slide freely within their respective compression coil. If desired, particularly if the lead wires <NUM> and <NUM> are not enclosed by a protective sheath, the outer surface of each compression coil can be covered by a flexible, non-conductive sheath <NUM>, e.g., made of polyimide tubing, to prevent contact between the compression coil and any other wires within the catheter body <NUM>.

Each compression coil <NUM> is anchored at its proximal end to the proximal end of the stiffening tube <NUM> in the catheter body <NUM> by a glue joint (not shown) and at its distal end to the deflection section <NUM> by glue joint <NUM> (<FIG>). Both glue joints may comprise polyurethane glue or the like. The glue may be applied by means of a needle or the like through a hole made in the side wall of the respective tubing, which needle is heated sufficiently to form a permanent hole. The glue is then introduced through the hole and wicks around the outer circumference to form a glue joint about the entire circumference of the compression coil.

The puller wires <NUM> and <NUM> extend into the lumens <NUM> and <NUM> (<FIG>), respectively, of the deflection section <NUM>. The puller wires are anchored at their distal end to the deflection section <NUM>. In one embodiment, an anchor is fixedly attached to the distal end of each puller wire, as depicted in <FIG>. The anchor is preferably formed by a metal tube <NUM>, e.g. a short segment of hypodermic stock, which is fixedly attached, e.g. by crimping, to the distal end of the puller wires <NUM> and <NUM>. The tubes <NUM> have a section that extends a short distance beyond the distal end of the puller wires. A cross-piece <NUM> made of a small section of stainless steel ribbon or the like is soldered or welded in a transverse arrangement to the distal end of each tube section <NUM>, which is flattened during the operation. This creates a T-bar anchor. Two notches are created in the sidewall of the deflection section <NUM>, resulting in openings into the lumens <NUM> and <NUM> into which the puller wires <NUM> and <NUM> extend. The anchors lie partially within the notches. Because the length of the ribbons forming the cross-pieces <NUM> are longer than the diameter of the openings into the lumens <NUM> and <NUM>, the anchors cannot be pulled completely into the lumens <NUM> and <NUM>. The notches are then sealed with polyurethane glue or the like to give a smooth outer surface. Within the lumens <NUM> and <NUM> of the deflection section <NUM>, each of the puller wires <NUM> and <NUM> extends through a respective plastic, preferably Teflon sheath <NUM>, which prevents the puller wires from cutting into the wall of the tubing <NUM> when the deflection section <NUM> is deflected.

Longitudinal movement of the puller wires <NUM> and <NUM> relative to the catheter body <NUM>, which results in deflection of the deflection section <NUM>, is accomplished by suitable manipulation of the control handle <NUM>. A suitable control handle for use with the present invention is described in <CIT>.

In the illustrated embodiment, an electromagnetic sensor <NUM> is provided and housed in the lumen of the connector tube <NUM>. A sensor cable <NUM> extends from the control handle <NUM>, and through the central lumen <NUM> of the catheter body <NUM> and the lumen <NUM> of the tubing <NUM> of deflection section <NUM> and the lumen of the connector tube <NUM>. The sensor cable <NUM> extends out the proximal end of the control handle <NUM> within an umbilical cord (not shown) to a sensor control module (not shown) that houses a circuit board (not shown). Alternatively, the circuit board can be housed within the control handle <NUM>, for example, as described in <CIT>. The electromagnetic sensor cable <NUM> comprises multiple wires encased within a plastic covered sheath. In the sensor control module, the wires of the electromagnetic sensor cable are connected to the circuit board. The circuit board amplifies the signal received from the electromagnetic sensor and transmits it to a computer in a form understandable by the computer by means of the sensor connector at the proximal end of the sensor control module. Also, because the catheter is designed for single use only, the circuit board preferably contains an EPROM chip which shuts down the circuit board approximately <NUM> hours after the catheter has been used. This prevents the catheter, or at least the electromagnetic sensor, from being used twice. Suitable electromagnetic sensors for use with the present invention are described, for example, in <CIT>, <CIT>,<CIT>, <CIT> and<CIT> and International Publication No. <CIT>. A preferred electromagnetic sensor <NUM> has a length of from about <NUM> to about <NUM> and a diameter of about <NUM>.

In use, a suitable guiding sheath (not shown) is inserted into the patient with its distal end positioned at or near a desired tissue location for diagnostics such as mapping and/or treatment such as ablation. An example of a suitable guiding sheath for use in connection with the present invention is the Preface Braided Guiding Sheath, commercially available from Biosense Webster, Inc. (Diamond Bar, Calif. The catheter <NUM> is passed through the guiding sheath and advanced therethrough to the desired tissue location. The guiding sheath is pulled proximally, exposing the tip electrode section <NUM> and the deflection section <NUM>.

The user actuates a thumb knob on the control handle to deflect the catheter and position the tip electrode <NUM> on tissue surface. With the multiple surface microelectrodes <NUM> in contact (or close proximity) with tissue, the catheter <NUM> is adapted for high density electrode sensing detecting electrical activity in the tissue which is transmitted through the catheter via the lead wires <NUM> for processing by a signal processor (not shown) for generating high density mapping with highly localized electrograms. If ablation is desired, the lead wire <NUM> is energized by an energy source, e.g., RF generator (not shown), whose distal end portion in the chamber <NUM> of the porous substrate <NUM> electrifies the conductive irrigation fluid delivered into the chamber <NUM> via irrigation tubing <NUM>. Passing of such electrified fluid from the chamber to the exposed surfaces of the porous substrate <NUM> renders the porous substrate <NUM> into a "virtual" ablation electrode. During and after ablation, the surface microelectrodes <NUM> on the porous substrate <NUM> can sense electrical activity at and around the ablated tissue to confirm the formation of electrically blocked tissue regions.

Claim 1:
A catheter (<NUM>) comprising:
an elongated catheter body (<NUM>);
a distal electrode member (<NUM>) having a porous substrate (<NUM>) and a plurality of surface electrodes (<NUM>) on portions of an outer surface of the porous substrate, the porous substrate having an interior chamber (<NUM>) adapted to receive conductive fluid;
a plurality of lead wires (<NUM>), each connected to a respective surface electrode; and
a lead wire (<NUM>) having a distal portion extending into the interior chamber, the lead wire adapted to electrify the conductive fluid in the chamber,
wherein the porous substrate is adapted to pass the conductive fluid from the chamber to the outer surface of the porous substrate, and characterized in that the plurality of surface electrodes comprises:
axial proximal rectangular micro-electrodes (40T), distal ring micro-electrodes (40R) and a distal tip circular micro-electrode (40C); or
a distal tip circular micro-electrode (40C), a plurality of circular micro-electrodes and a proximal ring micro-electrode (40R); or
a series of concentric circular micro electrodes (40C1, 40C2, 40C3) of different radii.