US20080045942A1 - Electrosurgical instrument and method of use - Google Patents
Electrosurgical instrument and method of use Download PDFInfo
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- US20080045942A1 US20080045942A1 US11/925,092 US92509207A US2008045942A1 US 20080045942 A1 US20080045942 A1 US 20080045942A1 US 92509207 A US92509207 A US 92509207A US 2008045942 A1 US2008045942 A1 US 2008045942A1
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B18/04—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body by heating
- A61B18/12—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body by heating by passing a current through the tissue to be heated, e.g. high-frequency current
- A61B18/14—Probes or electrodes therefor
- A61B18/1442—Probes having pivoting end effectors, e.g. forceps
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00053—Mechanical features of the instrument of device
- A61B2018/00059—Material properties
- A61B2018/00071—Electrical conductivity
- A61B2018/00077—Electrical conductivity high, i.e. electrically conducting
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00053—Mechanical features of the instrument of device
- A61B2018/00059—Material properties
- A61B2018/00071—Electrical conductivity
- A61B2018/00083—Electrical conductivity low, i.e. electrically insulating
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00053—Mechanical features of the instrument of device
- A61B2018/00107—Coatings on the energy applicator
- A61B2018/00148—Coatings on the energy applicator with metal
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00571—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body for achieving a particular surgical effect
- A61B2018/0063—Sealing
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00636—Sensing and controlling the application of energy
- A61B2018/00642—Sensing and controlling the application of energy with feedback, i.e. closed loop control
- A61B2018/00654—Sensing and controlling the application of energy with feedback, i.e. closed loop control with individual control of each of a plurality of energy emitting elements
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00636—Sensing and controlling the application of energy
- A61B2018/00773—Sensed parameters
- A61B2018/00791—Temperature
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00636—Sensing and controlling the application of energy
- A61B2018/00773—Sensed parameters
- A61B2018/00791—Temperature
- A61B2018/00797—Temperature measured by multiple temperature sensors
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B2018/00636—Sensing and controlling the application of energy
- A61B2018/00773—Sensed parameters
- A61B2018/00791—Temperature
- A61B2018/00809—Temperature measured thermochromatically
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B18/00—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body
- A61B18/04—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body by heating
- A61B18/12—Surgical instruments, devices or methods for transferring non-mechanical forms of energy to or from the body by heating by passing a current through the tissue to be heated, e.g. high-frequency current
- A61B18/14—Probes or electrodes therefor
- A61B2018/1467—Probes or electrodes therefor using more than two electrodes on a single probe
Abstract
An electrosurgical medical device and method for creating thermal welds in engaged tissue. In one embodiment, at least one jaw of the instrument defines a tissue engagement plane carrying a variable resistive body of a positive temperature coefficient material that has a selected decreased electrical conductance at each selected increased temperature thereof over a targeted treatment range. The variable resistive body can be engineered to bracket a targeted thermal treatment range, for example about 60° C. to 80° C., at which tissue welding can be accomplished. In one mode of operation, the engagement plane will automatically modulate and spatially localize ohmic heating within the engaged tissue from Rf energy application across micron-scale portions of the engagement surface. In another mode of operation, a variable resistive body will focus conductive heating in a selected portion of the engagement surface.
Description
- This application is a divisional of U.S. application Ser. No. 10/448,478 (Attorney Docket No. 021447-000510US), filed on May 30, 2003, which claimed benefit from Provisional U.S. Patent Application No. 60/384,496 (Docket No. SRX-018), filed Nov. 19, 2002, and was also a Continuation-in-Part of U.S. application Ser. No. 10/032,867 (Attorney Docket No.: 021447-000500US.SRX-011) filed Oct. 22, 2001, (now U.S. Pat. No. 6,929,644), the full disclosures all of which are incorporated herein by this reference.
- 1. Field of the Invention
- This invention relates to medical devices and techniques and more particularly relates to a working end of an electro surgical instrument that causes controlled ohmic heating of tissue across an engagement surface that modulates Rf power levels across localized micro-scale portions of the engagement surface, the system further adapted to focus conductive heating of tissue across the engagement surface.
- 2. Description of the Related Art
- In the prior art, various energy sources such as radiofrequency (Rf) sources, ultrasound sources and lasers have been developed to coagulate, seal or join together tissues volumes in open and laparoscopic surgeries. The most important surgical application relates to sealing blood vessels which contain considerable fluid pressure therein. In general, no instrument working ends using any energy source have proven reliable in creating a “tissue weld” or “tissue fusion” that has very high strength immediately post-treatment. For this reason, the commercially available instruments, typically powered by Rf or ultrasound, are mostly limited to use in sealing small blood vessels and tissues masses with microvasculature therein. The prior art Rf devices also fail to provide seals with substantial strength in anatomic structures having walls with irregular or thick fibrous content, in bundles of disparate anatomic structures, in substantially thick anatomic structures, or in tissues with thick fascia layers (e.g., large diameter blood vessels).
- In a basic bi-polar Rf jaw arrangement, each face of opposing first and second jaws comprises an electrode and Rf current flows across the captured tissue between the opposing polarity electrodes. Such prior art Rf jaws that engage opposing sides of tissue typically cannot cause uniform thermal effects in the tissue—whether the captured tissue is thin or substantially thick. As Rf energy density in tissue increases, the tissue surface becomes desiccated and resistant to additional ohmic heating. Localized tissue desiccation and charring can occur almost instantly as tissue impedance rises, which then can result in a non-uniform seal in the tissue. The typical prior art Rf jaws can cause further undesirable effects by propagating Rf density laterally from the engaged tissue thus causing unwanted collateral thermal damage.
- The commercially available Rf sealing instruments typically use one of two approaches to “control” Rf energy delivery in tissue. In a first “power adjustment” approach, the Rf system controller can rapidly adjust the level of total power delivered to the jaws' engagement surfaces in response to feedback circuitry coupled to the active electrodes that measures tissue impedance or electrode temperature. In a second “current-path directing” approach, the instrument jaws carry an electrode arrangement in which opposing polarity electrodes are spaced apart by an insulator material--which may cause current to flow within an extended path through captured tissue rather that simply between surfaces of the first and second jaws. Electrosurgical grasping instruments having jaws with electrically-isolated electrode arrangements in cooperating jaws faces were proposed by Yates et al. in U.S. Pat. Nos. 5,403,312; 5,735,848 and 5,833,690.
- The illustrations of the wall of a blood vessel in
FIGS. 1A-1D are useful in understanding the limitations of prior art Rf working ends for sealing tissue.FIG. 1B provides a graphic illustration of the opposingvessel walls portions - Now turning to
FIG. 1C , it is reasonable to ask whether the “power adjustment” approach to energy delivery is likely to cause a uniform temperature within every micron-scale tissue volume in the grid simultaneously—and maintain that temperature for a selected time interval.FIG. 1C shows theopposing vessel walls FIG. 1C graphically depicts current “paths” p in the tissue at an arbitrary time interval that can be microseconds (μs) apart. Such current paths p would be random and constantly in flux—along transient most conductive pathways through the tissue between the opposing polarity electrodes. The thickness of the “paths” is intended to represent the constantly adjusting power levels. If one assumes that the duration of energy density along any current path p is within the microsecond range before finding a new conductive path—and the thermal relaxation time of tissue is the millisecond (ms) range, then what is the likelihood that such entirely random current paths will revisit and maintain each discrete micron-scale tissue volume at the targeted temperature before thermal relaxation? Since the hydration of tissue is constantly reduced during ohmic heating—any regions of more desiccated tissue will necessarily lose its ohmic heating and will be unable to be “welded” to adjacent tissue volumes. The “power adjustment” approach probably is useful in preventing rapid overall tissue desiccation. However, it is postulated that any approach that relies on entirely “random” current paths p in tissue—no matter the power level—cannot cause contemporaneous denaturation of tissue constituents in all engaged tissue volumes and thus cannot create an effective high-strength “weld” in tissue. - Now referring to
FIG. 1D , it is possible to evaluate the second “current—path directing” approach to energy delivery in a jaw structure.FIG. 1D depictsvessel walls insulator 10 in a jaw engagement surface is that no ohmic heating of tissue can be delivered directly to the tissue volume engaged by the insulator 10 (seeFIG. 1D ). The tissue that directly contacts theinsulator 10 will only be ohmically heated when a current path p extends through the tissue between the spaced apart electrodes.FIG. 1D graphically depicts current paths p at any arbitrary time interval, for example in the μs range. Again, such current paths p will be random and in constant flux along transient conductive pathways. - This type of random, transient Rf energy density in paths p through tissue, when any path may occur only for a microsecond interval, is not likely to uniformly denature proteins in the entire engaged tissue volume. It is believed that the “current-path directing” approach for tissue sealing can only accomplish tissue coagulation or seals with limited strength.
- Now turning to
FIG. 2 , it can be conceptually understood that the key requirements for thermally-induced tissue welding relate to: (i) means for “non-random spatial localization” of energy densities in the engaged tissue et, (ii) means for “controlled, timed intervals” of power application of such spatially localized of energy densities, and (iii) means for “modulating the power level” of any such localized, time-controlled applications of energy. -
FIG. 2 illustrates a hypothetical tissue volume with a lower jaw's engagement surface 15 backed away from the tissue. The tissue is engaged under very high compression which is indicated by arrows inFIG. 2 . The engagement surface 15 is shown as divided into a hypothetical grid of “pixels” or micron-dimensionedsurface areas 20. Thus,FIG. 2 graphically illustrates that to create an effective tissue weld, the delivery of energy should be controlled and non-randomly spatially localized relative to eachpixel 20 of the engagement surface 15. - Still referring to
FIG. 2 , it can be understood that there are two modalities in which spatially localized, time-controlled energy applications can create a uniform energy density in tissue for protein denaturation. In a first modality, all cubic microns of the engaged tissue (FIG. 2 ) can be elevated to the required energy density and temperature contemporaneously to create a weld. In a second modality, a “wave” of the required energy density can sweep across the engaged tissue et that can thereby leave welded tissue in its wake. The authors have investigated, developed and integrated Rf systems for accomplishing both such modalities—which are summarized in the next Section. - The systems and methods corresponding to invention relate to creating thermal “welds” or “fusion” within native tissue volumes. The alternative terms of tissue “welding” and tissue “fusion” are used interchangeably herein to describe thermal treatments of a targeted tissue volume that result in a substantially uniform fused-together tissue mass that provides substantial tensile strength immediately post-treatment. Such tensile strength (no matter how measured) is particularly important (i) for welding blood vessels in vessel transection procedures, (ii) for welding organ margins in resection procedures, (iii) for welding other anatomic ducts wherein permanent closure is required, and also (iv) for vessel anastomosis, vessel closure or other procedures that join together anatomic structures or portions thereof.
- The welding or fusion of tissue as disclosed herein is to be distinguished from “coagulation”, “sealing”, “hemostasis” and other similar descriptive terms that generally relate to the collapse and occlusion of blood flow within small blood vessels or vascularized tissue. For example, any surface application of thermal energy can cause coagulation or hemostasis—but does not fall into the category of “welding” as the term is used herein. Such surface coagulation does not create a weld that provides any substantial strength in the affected tissue.
- At the molecular level, the phenomena of truly “welding” tissue as disclosed herein may not be fully understood. However, the authors have identified the parameters at which tissue welding can be accomplished. An effective “weld” as disclosed herein results from the thermally-induced denaturation of collagen, elastin and other protein molecules in a targeted tissue volume to create a transient liquid or gel-like proteinaceous amalgam. A selected energy density is provided in the targeted tissue to cause hydrothermal breakdown of intra- and intermolecular hydrogen cross-links in collagen and other proteins. The denatured amalgam is maintained at a selected level of hydration—without desiccation—for a selected time interval which can be very brief. The targeted tissue volume is maintained under a selected very high level of mechanical compression to insure that the unwound strands of the denatured proteins are in close proximity to allow their intertwining and entanglement. Upon thermal relaxation, the intermixed amalgam results in “protein entanglement” as re-cross-linking or renaturation occurs to thereby cause a uniform fused-together mass.
- To better appreciate the scale at which thermally-induced protein denaturation occurs—and at which the desired protein entanglement and re-cross-linking follows—consider that a collagen molecule in its native state has a diameter of about 15 Angstroms. The collagen molecule consists of a triple helix of peptide stands about 1000 Angstroms in length (see
FIG. 2 ). In other words—a single μm3 (cubic micrometer) of tissue that is targeted for welding will contain 10's of thousands of such collagen molecules. InFIG. 2 , each tissue volume in the grid represents an arbitrary size from about 1 μm to 5 μm (microns). Elastin and other molecules fro denaturation are believed to be similar in dimension to collagen. - To weld tissue, or more specifically to thermally-induce protein denaturation, and subsequent entanglement and re-cross-linking in a targeted tissue volume, it has been learned that the following interlinked parameters must be controlled:
- (i) Temperature of thermal denaturation. The targeted tissue volume must be elevated to the temperature of thermal denaturation, Td, which ranges from about 50° C. to 90° C., and more specifically is from about 60° C. to 80° C. The optimal Td within the larger temperature range is further dependent on the duration of thermal effects and level of pressure applied to the engaged tissue.
- (ii) Duration of treatment. The thermal treatment must extend over a selected time duration, which depending on the engaged tissue volume, can range from less than 0.1 second to about 5 seconds. As will be described below, the system of the in invention utilizes a thermal treatment duration ranging from about 500 ms second to about 3000 ms. Since the objectives of protein entanglement occur at Td which can be achieved in ms (or even microseconds)—this disclosure will generally describe the treatment duration in ms.
- (iii) Ramp-up in temperature: uniformity of temperature profile. There is no limit to the speed at which temperature can be ramped up within the targeted tissue. However, it is of utmost importance to maintain a very uniform temperature across the targeted tissue volume so that “all” proteins are denatured within the same microsecond interval. Only thermal relaxation from a uniform temperature Td can result in complete protein entanglement and re-cross-linking across an entire tissue volume. Without such uniformity of temperature ramp-up and relaxation, the treated tissue will not become a fused-together tissue mass—and thus will not have the desired strength.
- Stated another way, it is necessary to deposit enough energy into the targeted volume to elevate it to the desired temperature Td before it diffuses into adjacent tissue volumes. The process of heat diffusion describes a process of conduction and convection and defines a targeted volume's thermal relaxation time (often defined as the time over which the temperature is reduced by one-half). Such thermal relaxation time scales with the square of the diameter of the treated volume in a spherical volume, decreasing as the diameter decreases. In general, tissue is considered to have a thermal relaxation time in the range of 1 ms. In a non-compressed tissue volume, or lightly compressed tissue volume, the thermal relaxation of tissue in an Rf application typically will prevent a uniform weld since the random current paths result in very uneven ohmic heating (see
FIGS. 1C-1D ). - (iv) Instrument engagement surfaces. The instrument's engagement surface(s) must have characteristics that insure that every square micron of the instrument surface is in contact with tissue during Rf energy application. Any air gap between an engagement surface and tissue can cause an arc of electrical energy across the insulative gap thus resulting in charring of tissue. Such charring (desiccation) will entirely prevent welding of the localized tissue volume and result in further collateral effects that will weaken any attempted weld. For this reason, the engagement surfaces corresponding to the invention are (i) substantially smooth at a macroscale, and (ii) at least partly of an elastomeric matrix that can conform to the tissue surface dynamically during treatment. The jaw structure of the invention typically has gripping elements that are lateral from the energy-delivering engagement surfaces. Gripping serrations otherwise can cause unwanted “gaps” and microscale trapped air pockets between the tissue and the engagement surfaces.
- (v) Pressure. It has been found that very high external mechanical pressures on a targeted tissue volume are critical in welding tissue—for example, between the engagement surfaces of a jaw structure. In one aspect, as described above, the high compressive forces can cause the denatured proteins to be crushed together thereby facilitating the intermixing or intercalation of denatured protein stands which ultimately will result in a high degree of cross-linking upon thermal relaxation.
- In another aspect, the proposed high compressive forces (it is believed) can increase the thermal relaxation time of the engaged tissue practically by an infinite amount. With the engaged tissue highly compressed to the dimension of a membrane between opposing engagement surfaces, for example to a thickness of about 0.001″, there is effectively little “captured” tissue within which thermal diffusion can take place. Further, the very thin tissue cross-section at the margins of the engaged tissue prevents heat conduction to tissue volumes outside the jaw structure.
- In yet another aspect, the high compressive forces at first cause the lateral migration of fluids from the engaged tissue which assists in the subsequent welding process. It has been found that highly hydrated tissues are not necessary in tissue welding. What is important is maintaining the targeted tissue at a selected level without desiccation as is typical in the prior art. Further, the very high compressive forces cause an even distribution of hydration across the engaged tissue volume prior to energy delivery.
- In yet another aspect, the high compressive forces insure that the engagement planes of the jaws are in complete contact with the surfaces of the targeted tissues, thus preventing any possibility of an arc of electrical energy a cross a “gap” would cause tissue charring, as described previously.
- One exemplary embodiment disclosed herein is particularly adapted for, in effect, independent spatial localization and modulation of Rf energy application across micron-scale “pixels” of an engagement surface. The jaw structure of the instrument defines opposing engagement planes that apply high mechanical compression to the engaged tissue. At least one engagement plane has a surface layer that comprises first and second portions of a conductive-resistive matrix—preferably including an elastomer such as silicone (first portion) and conductive particles (second portion) distributed therein. An electrical source is coupled to the working end such that the combination of the conductive-resistive matrix and the engaged tissue are intermediate opposing conductors that define first and second polarities of the electrical source coupled thereto. The conductive-resistive matrix is designed to exhibit unique resistance vs. temperature characteristics, wherein the matrix maintains a low base resistance over a selected temperature range with a dramatically increasing resistance above a selected narrow temperature range.
- In operation, it can be understood that current flow through the conductive-resistive matrix and engagement plane will apply active Rf energy (ohmic heating) to the engaged tissue until the point in time that any portion of the matrix is heated to a range that substantially reduces its conductance. This effect will occur across the surface of the matrix thus allowing each matrix portion to deliver an independent level of power therethrough. This instant, localized reduction of Rf energy application can be relied on to prevent any substantial dehydration of tissue proximate to the engagement plane. The system eliminates the possibility of desiccation thus meeting another of the several parameters described above.
- The conductive-resistive matrix and jaw body corresponding to the invention further can provides a suitable cross-section and mass for providing substantial heat capacity. Thus, when the matrix is elevated in temperature to the selected thermal treatment range, the retained heat of the matrix volume can effectively apply thermal energy to the engaged tissue volume by means of conduction and convection. In operation, the working end can automatically modulate the application of energy to tissue between active Rf heating and passive conductive heating of the targeted tissue to maintain a targeted temperature level.
- Of particular interest, another system embodiment disclosed herein is adapted for causing a “wave” of ohmic heating to sweep across tissue to denature tissue constituents in its wake. This embodiment again utilizes at least one engagement plane in a jaw structure that carries a conductive-resistive matrix as described previously. At least one of the opposing polarity conductors has a portion thereof exposed in the engagement plane. The conductive-resistive matrix again is intermediate the opposing polarity conductors. When power delivery is initiated, the matrix defines an “interface” therein where microcurrents are most intense about the interface of the two polarities—since the matrix is not a simple conductor. The engaged tissue, in effect, becomes an extension of the interface of microcurrents created by the matrix—which thus localizes ohmic heating across the tissue proximate the interface. The interface of polarities and microcurrents within the matrix will be in flux due to lesser conductance about the interface as the matrix is elevated in temperature. Thus, a “wave-like” zone of microcurrents between the polarities will propagate across the matrix—and across the engaged tissue. By this means of engaging tissue with a conductive-resistive matrix, a wave of energy density can be caused to sweep across tissue to uniformly denature proteins which will then re-cross-link to create a uniquely strong weld.
- In general, the system of conductive-resistive matrices for Rf energy delivery advantageously provides means for spatial-localization and modulation of energy application from selected, discrete locations across a single energy-emitting surface coupled to a single energy source
- The system of conductive-resistive matrices for Rf energy delivery provides means for causing a dynamic wave of ohmic heating in tissue to propagate across engaged tissue.
- The system of conductive-resistive matrices for Rf energy delivery allows for opposing electrical potentials to be exposed in a single engagement surface with a conductive matrix therebetween to allow 100% of the engagement surface to emit energy to tissue.
- The system of conductive-resistive matrices for Rf energy application to tissue allows for bi-polar electrical potential to be exposed in a single engagement surface without an intermediate insulator portion.
- The system of conductive-resistive matrices for energy delivery allows for the automatic modulation of active ohmic heating and passive heating by conduction and convection to treat tissue.
- The system of conductive-resistive matrices for energy application to tissue advantageously allows for the creation of “welds” in tissue within about 500 ms to 2 seconds.
- The system of conductive-resistive matrices for energy application to tissue provides “welds” in blood vessels that have very high strength.
- Additional objects and advantages of the invention will be apparent from the following description, the accompanying drawings and the appended claims.
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FIG. 1A is a view of a blood vessel targeted for welding. -
FIG. 1B is a greatly enlarged sectional view of opposing wall portions of the blood vessel ofFIG. 1A taken alongline 1B-1B ofFIG. 1A . -
FIG. 1C is a graphic representation of opposing walls of a blood vessel engaged by prior art electrosurgical jaws showing random paths of current (causing ohmic heating) across the engaged tissue between opposing polarity electrodes. -
FIG. 1D is a graphic representation of a blood vessel engaged by prior art electrosurgical jaws with an insulator between opposing polarity electrodes on each side of the tissue showing random paths of current (ohmic heating). -
FIG. 2 graphically represents a blood vessel engaged by hypothetical electrosurgical jaws under very high compression with an energy-delivery surface proximate to the tissue. -
FIG. 3A is a perspective view of a jaw structure of tissue-transecting and welding instrument that carries a Type “A” conductive-resistive matrix system corresponding to the invention. -
FIG. 3B is a sectional view of the jaw structure ofFIG. 3A taken along line 3B-3B ofFIG. 3A showing the location of conductive-resistive matrices. -
FIG. 4 is a perspective view of another exemplary surgical instrument that carries a Type “A” conductive-resistive matrix system for welding tissue. -
FIG. 5 is a sectional view of the jaw structure ofFIG. 4 taken along line 5-5 ofFIG. 4 showing details of the conductive-resistive matrix. -
FIG. 6 is a graph showing (i) temperature-resistance profiles of alternative conductive-resistive matrices that can be carried in the jaw ofFIG. 5 , (ii) the impedance of tissue, and (iii) the combined resistance of the matrix and tissue as measured by a system controller. -
FIG. 7A is an enlarged view of a portion of the conductive-resistive matrix and jaw body ofFIG. 5 showing a first portion of an elastomer and a second portion of conductive particles at a resting temperature. -
FIG. 7B is another view the conductive-resistive matrix and jaw body ofFIG. 7A after a portion is elevated to a higher temperature to modulate microcurrent flow therethrough thus depicting a method of the invention in spatially localizing and modulating Rf energy application from a conductive-resistive matrix that engages tissue. -
FIG. 8A is a further enlarged view of the conductive-resistive matrix ofFIG. 7A showing the first portion (elastomer) and the second portion (conductive elements) and paths of microcurrents therethrough. -
FIG. 8B is a further enlarged view of matrix ofFIG. 7B showing the effect of increased temperature and the manner in which resistance to microcurrent flow is caused in the method of spatially localizing and modulating Rf energy application. -
FIG. 9 is an enlarged view of an alternative conductive-resistive matrix similar to that ofFIG. 7A that is additionally doped with thermally conductive, electrically non-conductive particles. -
FIG. 10 is an alternative jaw structure similar to that ofFIGS. 5 and 7 A except carrying conductive-resistive matrices in the engagement surfaces of both opposing jaws. -
FIG. 11 is a greatly enlarged sectional view of the jaws ofFIG. 10 taken along line 11-11 ofFIG. 10 . -
FIG. 12 is a sectional view of another exemplary jaw structure that carries a Type “B” conductive-resistive matrix system for welding tissue that utilizes opposing polarity electrodes with an intermediate conductive-resistive matrix in an engagement surface. -
FIG. 13A is a sectional view of alternative Type “B” jaw with a plurality of opposing polarity electrodes with intermediate conductive-resistive matrices in the engagement surface. -
FIG. 13B is a sectional view of a Type “B” jaw similar to that ofFIG. 13A with a plurality of opposing polarity electrodes with intermediate conductive-resistive matrices in the engagement surface in a different angular orientation. -
FIG. 13C is a sectional view of another Type “B” jaw similar to that ofFIGS. 13A-13B with a plurality of opposing polarity electrodes with intermediate matrices in another angular orientation. -
FIGS. 14A-14C graphically illustrate a method of the invention in causing a wave of Rf energy density to propagate across and engaged tissue membrane to denature tissue constituents: -
FIG. 14A being the engagement surface ofFIG. 12 engaging tissue membrane at the time that energy delivery is initiated causing localized microcurrents and ohmic tissue heating; -
FIG. 14B being the engagement surface ofFIG. 12 after an arbitrary millisecond or microsecond time interval depicting the propagation of a wavefronts of energy outward from the initial localized microcurrents as the localized temperature and resistance of the matrix is increased; and -
FIG. 14C being the engagement surface ofFIG. 12 after another very brief interval depicting the propagation of the wavefronts of energy density outwardly in the tissue due to increase temperature and resistance of matrix portions. -
FIG. 15 is an enlarged sectional view of the exemplary jaw structure ofFIG. 13A with a plurality of opposing polarity conductors on either side of conductive-resistive matrix portions. -
FIG. 16 is a sectional view of a jaw structure similar to that ofFIG. 15 with a plurality of opposing polarity conductors that float within an elastomeric conductive-resistive matrix portions. -
FIG. 17 is a sectional view of a jaw structure similar to that ofFIG. 16 with a single central conductor that floats on a convex elastomeric conductive-resistive matrix with opposing polarity conductors in outboard locations. -
FIGS. 18A-18C provide simplified graphic views of the method of causing a wave of Rf energy density in the embodiment ofFIG. 17 , similar to the method shown in FIGS. 14A-14C: -
FIG. 18A corresponding to the view ofFIG. 14A showing initiation of energy delivery; -
FIG. 18B corresponding to the view ofFIG. 14B showing the propagation of the wavefronts of energy density outwardly; and -
FIG. 18C corresponding to the view ofFIG. 14C showing the further outward propagation of the wavefronts of energy density to thereby weld tissue. -
FIG. 19 is a sectional view of another exemplary jaw structure that carries two conductive-resistive matrix portions, each having a different durometer and a different temperature coefficient profile. -
FIG. 20 is a sectional view of a jaw assembly having the engagement plane ofFIG. 17 carried in a transecting-type jaws similar to that ofFIGS. 3A-3B . -
FIG. 21 is a sectional view of an alternative jaw structure similar with a fully metallized engagement surface coupled to first and second polarity leads in adjacent portions thereof. -
FIG. 22 is an enlarged view of the fully metallized engagement surface ofFIG. 21 showing the first and second polarity leads that are coupled to the metal film layer. -
FIG. 23 is a sectional view of a Type “D” working end corresponding to the invention with a conductive-resistive matrix having tapered cross-sectional dimension for focusing passive heating in the center of the engagement surface. -
FIG. 24 is a sectional view of an alternative Type “D” working end with a gradient matrix for focusing passive heating in the center of the engagement surface. -
FIG. 25 is a sectional view of an alternative Type “D” working end similar to that ofFIG. 23 a with conductive-resistive matrix in the upper jaw. -
FIG. 26A is a perspective view of an alternative Type “D” instrument with an extendable blade member. -
FIG. 26B is a sectional view of the jaw structure ofFIG. 26A taken alongline 26B-26B ofFIG. 26A . -
FIG. 27 is a sectional view of an alternative jaw structure similar to that ofFIG. 26B with additional electrodes in the engagement surface. -
FIG. 28 is a sectional view of an alternative jaw structure similar to that ofFIG. 27 . -
FIG. 29 is a graph showing a power-impedance curve of a voltage source plotted against an impedance-temperature curve of a conductive-resistive matrix corresponding to the invention. - 1. Exemplary jaw structures for welding tissue.
FIGS. 3A and 3B illustrate a working end of a surgical grasping instrument corresponding to the invention that is adapted for transecting captured tissue and for contemporaneously welding the captured tissue margins with controlled application of Rf energy. Thejaw assembly 100A is carried at the distal end 104 of anintroducer sleeve member 106 that can have a diameter ranging from about 2 mm. to 20 mm. for cooperating with cannulae in endoscopic surgeries or for use in open surgical procedures. Theintroducer portion 106 extends from a proximal handle (not shown). The handle can be any type of pistol-grip or other type of handle known in the art that carries actuator levers, triggers or sliders for actuating the jaws and need not be described in further detail. Theintroducer sleeve portion 106 has a bore 108 extending therethrough for carrying actuator mechanisms for actuating the jaws and for carrying electrical leads 109 a-109 b for delivery of electrical energy to electrosurgical components of the working end. - As can be seen in
FIGS. 3A and 3B , thejaw assembly 100A has first (lower)jaw element 112A and second (upper)jaw element 112B that are adapted to close or approximate about axis 115. The jaw elements can both be moveable or a single jaw can rotate to provide the jaw-open and jaw-closed positions. In the exemplary embodiment ofFIGS. 3A and 3B , both jaws are moveable relative to theintroducer portion 106. - Of particular interest, the opening-closing mechanism of the
jaw assembly 100A is capable of applying very high compressive forces on tissue on the basis of cam mechanisms with a reciprocatingmember 140. The engagement surfaces further provide a positive engagement of camming surfaces (i) for moving the jaw assembly to the (second) closed position to apply very high compressive forces, and (ii) for moving the jaws toward the (first) open position to apply substantially high opening forces for “dissecting” tissue. This important feature allows the surgeon to insert the tip of the closed jaws into a dissectable tissue plane—and thereafter open the jaws to apply such dissecting forces against tissues. Prior art instruments are spring-loaded toward the open position which is not useful for dissecting tissue. - In the embodiment of
FIGS. 3A and 3B , a reciprocatingmember 140 is actuatable from the handle of the instrument by any suitable mechanism, such as a lever arm, that is coupled to a proximal end 141 ofmember 140. The proximal end 141 and medial portion ofmember 140 are dimensioned to reciprocate within bore 108 ofintroducer sleeve 106. The distal portion 142 of reciprocatingmember 140 carries first (lower) and second (upper) laterally-extendingflange elements FIG. 3A ) that can be a blade or a cutting electrode. The transverse element 145 is adapted to slide within achannels member 140 and jaw surfaces is described in complete detail in co-pending Provisional U.S. patent application Ser. No. 60/347,382 filed Jan. 11, 2002 (Docket No. SRX-013) titled Jaw Structure for Electrosurgical Instrument and Method of Use, which is incorporated herein by reference. - In
FIGS. 3A and 3B , the first andsecond jaws engagement plane 150 and define tissue-engagingsurface layers serrations 156 for gripping tissue. One preferred embodiment ofFIGS. 3A and 3B providessuch serrations 156 at an inner portion of the jaws alongchannels jaws - In the exemplary embodiment of
FIGS. 3A and 3B , theengagement surface 155A of thelower jaw 112A is adapted to deliver energy to tissue, at least in part, through a conductive-resistive matrix CM corresponding to the invention. The tissue-contactingsurface 155B ofupper jaw 112B preferably carries a similar conductive-resistive matrix, or the surface can be a conductive electrode or and insulative layer as will be described below. Alternatively, the engagement surfaces of the jaws can carry any of the energy delivery components disclosed in co-pending U.S. patent application Ser. No. 09/032,867 filed Oct. 22, 2001 (Docket No. SRX-011) titled Electrosurgical Jaw Structure for Controlled Energy Delivery and U.S. Prov. Patent Application Ser. No. 60/337,695 filed Dec. 3, 2001 (Docket No. SRX-012) titled Electrosurgical Jaw Structure for Controlled Energy Delivery, both of which are incorporated herein by reference. - Referring now to
FIG. 4 , analternative jaw structure 100B is shown with lower and upper jaws havingsimilar reference numerals 112A-112B. The simple scissor-action of the jaws inFIG. 4 has been found to be useful for welding tissues in procedures that do not require tissue transection. The scissor-action of the jaws can apply high compressive forces against tissue captured between the jaws to perform the method corresponding to the invention. As can be seen by comparingFIGS. 3B and 4 , the jaws of eitherembodiment - It has been found that very high compression of tissue combined with controlled Rf energy delivery is optimal for welding the engaged tissue volume contemporaneous with transection of the tissue. Preferably, the engagement gap g between the engagement planes ranges from about 0.0005″ to about 0.050″ for reduce the engaged tissue to the thickness of a membrane. More preferably, the gap g between the engagement planes ranges from about 0.001″ to about 0.005″.
- 2. Type “A” conductive-resistive matrix system for controlled energy delivery in tissue welding.
FIG. 5 illustrates an enlarged schematic sectional view of a jaw structure that carries engagement surface layers 155A and 155B injaws FIG. 4 ) for convenience, and the conductive-resistive matrix, or variable resistive body, would be identical in each side of a transecting jaw structure as shown inFIGS. 3A-3B . - In
FIG. 5 , it can be seen that thelower jaw 112A carries a component described herein as a conductive-resistive matrix CM that is at least partly exposed to anengagement plane 150 that is defined as the interface between tissue and a jaw engagement surface layer, 155A or 155B. More in particular, the conductive-resistive matrix CM comprises afirst portion 160 a and asecond portion 160 b. The first portion is preferably an electrically non-conductive material that has a selected coefficient of expansion that is typically greater than the coefficient of expansion of the material of the second portion. In one preferred embodiment, thefirst portion 160 a of the matrix is an elastomer, for example a medical grade silicone. Thefirst portion 160 a of the matrix also is preferably not a good thermal conductor. Other thermoplastic elastomers fall within the scope of the invention, as do ceramics having a thermal coefficient of expansion with the parameters further described below. - Referring to
FIG. 5 , thesecond portion 160 b of the matrix CM is a material that is electrically conductive and that is distributed within thefirst portion 160 a. InFIG. 5 , thesecond portion 160 b is represented (not-to-scale) asspherical elements 162 that are intermixed within the elastomerfirst portion 160 a of matrix CM. Theelements 162 can have any regular or irregular shape, and also can be elongated elements or can comprise conductive filaments. The dimensions ofparticles 162 can having a scale ranging from about 1 nm to 100 microns across a principal axis thereof. Preferably, theparticles 162 range between about 1 nm and 10 microns. More preferably, theparticles 162 range between about 1 nm and 1 microns in cross-section. Also, the matrix CM can carry aconductive portion 160 b in the form of separates filaments or an intertwined filament akin to the form of steel wool embedded within an elastomericfirst portion 160 a and fall within the scope the invention. Thus, thesecond portion 160 b can be of any form that distributes an electrically conductive mass within the overall volume of the matrix CM. - In the
lower jaw 112A ofFIG. 5 , the matrix CM is carried in a support structure orbody portion 158 that can be of any suitable metal or other material having sufficient strength to apply high compressive forces to the engaged tissue. Typically, thesupport structure 158 carries aninsulative coating 159 to prevent electrical current flow to tissues about the exterior of the jaw assembly and betweensupport structure 158 and the matrix CM and aconductive element 165 therein. - Of particular interest, the combination of first and
second portions first portion 160 a of the non-conductive elastomer relative to the volume proportion ofsecond portion 160 b of the conductive nanoparticles orelements 162, the matrix CM can be engineered to exhibit very large changes in resistance with a small change in matrix temperature. In other words, the change of resistance with a change in temperature results in a “positive” temperature coefficient of resistance. - In a first preferred embodiment, the matrix CM is engineered to exhibit unique resistance vs. temperature characteristics that is represented by a positively sloped temperature-resistance curve (see
FIG. 6 ). More in particular, the first exemplary matrix CM indicated inFIG. 6 maintains a low base resistance over a selected base temperature range with a dramatically increasing resistance above a selected narrow temperature range of the material (sometimes referred to herein as a switching range, seeFIG. 6 ). For example, the base resistance can be low, or the electrical conductivity high, between about 37° C. and 65° C, with the resistance increasing greatly between about 65° C. and 75° C. to substantially limit conduction therethrough (at typically utilized power levels in electrosurgery). In a second exemplary matrix embodiment described inFIG. 6 , the matrix CM is characterized by a more continuously positively sloped temperature-resistance over the range of 50° C. to about 80° C. Thus, the scope of the invention includes any specially engineered matrix CM with such a positive slope that is suitable for welding tissue as described below. - In one preferred embodiment, the matrix CM has a
first portion 160 a fabricated from a medical grade silicone that is doped with a selected volume of conductive particles, for example carbon particles in sub-micron dimensions as described above. By weight, the ration of silicone-to-carbon can range from about 10/90 to about 70/30 (silicone/carbon) to provide the selected range at which the inventive composition functions to substantially limit electrical conductance therethrough. More preferably, the carbon percentage in the matrix CM is from about 40% to 80% with the balance being silicone. In fabricating a matrix CM in this manner, it is preferable to use a carbon type that has single molecular bonds. It is less preferable to use a carbon type with double bonds that has the potential of breaking down when used in a small cross-section matrix, thus creating the potential of a permanent conductive path within deteriorated particles of the matrix CM that fuse together. One preferred composition has been developed to provide a thermal treatment range of about 75° C. to 80° C. with the matrix having about 50-60 percent carbon with the balance being silicone. The matrix CM corresponding to the invention thus becomes reversibly resistant to electric current flow at the selected higher temperature range, and returns to be substantially conductive within the base temperature range. In one preferred embodiment, the hardness of the silicone-based matrix CM is within the range of about Shore A range of less than about 95. More preferably, an exemplary silicone-based matrix CM has Shore A range of from about 20-80. The preferred hardness of the silicone-based matrix CM is about 150 or lower in the Shore D scale. As will be described below, some embodiments have jaws that carry cooperating matrix portions having at least two different hardness ratings. - In another embodiment, the particles or
elements 162 can be a polymer bead with a thin conductive coating. A metallic coating can be deposited by electroless plating processes or other vapor deposition process known in the art, and the coating can comprise any suitable thin-film deposition, such as gold, platinum, silver, palladium, tin, titanium, tantalum, copper or combinations or alloys of such metals, or varied layers of such materials. One preferred manner of depositing a metallic coating on such polymer elements comprises an electroless plating process provided by Micro Plating, Inc., 8110 Hawthorne Dr., Erie, Pa. 16509-4654. The thickness of the metallic coating can range from about 0.00001″ to 0.005″. (A suitable conductive-resistive matrix CM can comprise a ceramicfirst portion 160 a in combination with compressible-particlesecond portion 160 b of a such a metallized polymer bead to create the effects illustrated inFIGS. 8A-8B below). - One aspect of the invention relates to the use of a matrix CM as illustrated schematically in
FIG. 5 in a jaw'sengagement surface layer 155A with a selected treatment range between a first temperature (TE1) and a second temperature (TE2) that approximates the targeted tissue temperature for tissue welding (seeFIG. 6 ). The selected switching range of the matrix as defined above, for example, can be any substantially narrow 1°-10° C. range that is about the maximum of the treatment range that is optimal for tissue welding. For another thermotherpy, the switching range can fall within any larger tissue treatment range of about 50°-200° C. - No matter the character of the slope of the temperature-resistance curve of the matrix CM (see
FIG. 6 ), a preferred embodiment has a matrix CM that is engineered to have a selected resistance to current flow across its selected dimensions in the jaw assembly, when at 37° C., that ranges from about 0.0001 ohms to 1000 ohms. More preferably, the matrix CM has a designed resistance across its selected dimensions at 37° C. that ranges from about 1.0 ohm to 1000 ohms. Still more preferably, the matrix CM has with a designed resistance across its selected dimensions at 37° C. that ranges from about 25 ohms to 150 ohms. In any event, the selected resistance across the matrix CM in an exemplary jaw at 37° C. matches or slightly exceeds the resistance of the tissue or body structure that is engaged. The matrix CM further is engineered to have a selected conductance that substantially limits current flow therethrough corresponding to a selected temperature that constitutes the high end (maximum) of the targeted thermal treatment range. As generally described above, such a maximum temperature for tissue welding can be a selected temperature between about 50° C. and 90° C. More preferably, the selected temperature at which the matrix's selected conductance substantially limits current flow occurs at between about 60° C. and 80° C. - In the
exemplary jaw 112A ofFIG. 5 , the entire surface area ofengagement surface layer 155A comprises the conductive-resistive matrix CM, wherein the engagement surface is defined as the tissue-contacting portion that can apply electrical potential to tissue. Preferably, any instrument's engagement surface has a matrix CM that comprises at least 5% of its surface area. More preferably, the matrix CM comprises at least 10% of the surface area of engagement surface. Still more preferably, the matrix CM comprises at least 20% of the surface area of the jaw's engagement surface. The matrix CM can have any suitable cross-sectional dimensions, indicated generally at md1 and md2 inFIG. 5 , and preferably such a cross-section comprises a significant fracational volume of the jaw relative to supportstructure 158. As will be described below, in some embodiments, it is desirable to provide a thermal mass for optimizing passive conduction of heat to engaged tissue. - As can be seen in
FIG. 5 , the interior ofjaw 112A carries a conductive element (or electrode) indicated at 165 that interfaces with an interior surface 166 of the matrix CM. Theconductive element 165 is coupled by an electrical lead 109 a to a voltage (Rf)source 180 and optional controller 182 (FIG. 4 ). Thus, theRf source 180 can apply electrical potential (of a first polarity) to the matrix CM throughconductor 165—and thereafter to theengagement plane 150 through matrix CM. The opposingsecond jaw 112B inFIG. 5 has a conductive material (electrode) indicated at 185 coupled tosource 180 by lead 109 b that is exposed within theupper engagement surface 155B. - In a first mode of operation, referring to
FIG. 5 , electrical potential of a first polarity applied toconductor 165 will result in current flow through the matrix CM and the engaged tissue et to the opposingpolarity conductor 185. As described previously, the resistance of the matrix CM at 37° C. is engineered to approximate, or slightly exceed, that of the engaged tissue et. It can now be described how theengagement surface 155A can modulate the delivery of energy to tissue et similar to the hypothetical engagement surface ofFIG. 2 . Consider that the small sections of engagement surfaces represent the micron-sized surface areas (or pixels) of the illustration ofFIG. 2 (note that the jaws are not in a fully closed position inFIG. 5 ). The preferred membrane-thick engagement gap g is graphically represented inFIG. 5 . -
FIGS. 7A and 8A illustrate enlarged schematic sectional views ofjaws conductor 165 will cause current flow within and about theelements 162 ofsecond portion 160 b along any conductive path toward the opposingpolarity conductor 185.FIG. 8A more particularly shows a graphic representation of paths of microcurrents mcm within the matrix wherein theconductive elements 162 are in substantial contact.FIG. 7A also graphically illustrates paths of microcurrents met in the engaged tissue across gap g. The current paths in the tissue (across conductive sodium, potassium, chlorine ions etc.) thus results in ohmic heating of the tissue engaged betweenjaws elements 162 in the surface of theengagement layer 155A, which can act like surface asperities or sharp edges to induce current flow therefrom. - Consider that ohmic heating (or active heating) of the shaded
portion 188 of engaged tissue et inFIGS. 7B and 8B elevates its temperature to a selected temperature at the maximum of the targeted range. Heat will be conducted back to the matrix portion CM proximate to the heated tissue. At the selected temperature, the matrix CM will substantially reduce current flow therethrough and thus will contribute less and less to ohmic tissue heating, which is represented inFIGS. 7B and 8B . InFIGS. 7B and 8B , the thermal coefficient of expansion of the elastomer offirst matrix portion 160 a will cause slight redistribution of the secondconductive portion 160 b within the matrix—naturally resulting in lessened contacts between theconductive elements 162. It can be understood by arrows A inFIG. 8B that the elastomer will expand in directions of least resistance which is between theelements 162 since the elements are selected to be substantially resistant to compression. - Of particular interest, the small surface portion of matrix CM indicated at 190 in
FIG. 8A will function, in effect, independently to modulate power delivery to the surface of the tissue T engaged thereby. This effect will occur across the entireengagement surface layer 155A, to provide practically infinite “spatially localized” modulation of active energy density in the engaged tissue. In effect, the engagement surface can be defined as having “pixels” about its surface that are independently controlled with respect to energy application to localized tissue in contact with each pixel. Due to the high mechanical compression applied by the jaws, the engaged membrane all can be elevated to the selected temperature contemporaneously as each pixel heats adjacent tissue to the top of treatment range. As also depicted inFIG. 8B , the thermal expansion of the elastomeric matrix surface also will push into the membrane, further insuring tissue contact along theengagement plane 150 to eliminate any possibility of an energy arc across a gap. - Of particular interest, as any portion of the conductive-resistive matrix CM falls below the upper end of targeted treatment range, that matrix portion will increase its conductance and add ohmic heating to the proximate tissue via current paths through the matrix from
conductor 165. By this means of energy delivery, the mass of matrix and the jaw body will be modulated in temperature, similar to the engaged tissue, at or about the targeted treatment range. -
FIG. 9 shows another embodiment of a conductive-resistive matrix CM that is further doped withelements 192 of a material that is highly thermally conductive with a selected mass that is adapted to provide substantial heat capacity. By utilizingsuch elements 192 that may not be electrically conductive, the matrix can provide greater thermal mass and thereby increase passive conductive or convective heating of tissue when the matrix CM substantially reduces current flow to the engaged tissue. In another embodiment (not shown) the material ofelements 162 can be both substantially electrically conductive and highly thermally conductive with a high heat capacity. - The manner of utilizing the system of
FIGS. 7A-7B to perform the method of the invention can be understood as mechanically compressing the engaged tissue et to membrane thickness between the first andsecond engagement surfaces conductor 165, which potential is conducted through matrix CM to maintain a selected temperature across engaged tissue et for a selected time interval. At normal tissue temperature, the low base resistance of the matrix CM allows unimpeded Rf current flow fromvoltage source 180 thereby making 100 percent of the engagement surface an active conductor of electrical energy. It can be understood that the engaged tissue initially will have a substantially uniform impedance to electrical current flow, which will increase substantially as the engaged tissue loses moisture due to ohmic heating. Following an arbitrary time interval (in the microsecond to ms range), the impedance of the engaged tissue—reduced to membrane thickness—will be elevated in temperature and conduct heat to the matrix CM. In turn, the matrix CM will constantly adjust microcurrent flow therethrough—with each square micron of surface area effectively delivering its own selected level of power depending on the spatially-local temperature. This automatic reduction of localized microcurrents in tissue thus prevents any dehydration of the engaged tissue. By maintaining the desired level of moisture in tissue proximate to the engagement plane(s), the jaw assembly can insure the effective denaturation of tissue constituents to thereafter create a strong weld. - By the above-described mechanisms of causing the matrix CM to be maintained in a selected treatment range, the actual Rf energy applied to the engaged tissue et can be precisely modulated, practically pixel-by-pixel, in the terminology used above to describe
FIG. 2 . Further, theelements 192 in the matrix CM can comprise a substantial volume of the jaws' bodies and the thermal mass of the jaws, so that when elevated in temperature, the jaws can deliver energy to the engaged tissue by means of passive conductive heating—at the same time Rf energy delivery in modulated as described above. This balance of active Rf heating and passive conductive heating (or radiative, convective heating) can maintain the targeted temperature for any selected time interval. - Of particular interest, the above-described method of the invention that allows for immediate modulation of ohmic heating across the entirety of the engaged membrane is to be contrasted with prior art instruments that rely on power modulation based on feedback from a temperature sensor. In systems that rely on sensors or thermocouples, power is modulated only to an electrode in its totality. Further, the prior art temperature measurements obtained with sensors is typically made at only at a single location in a jaw structure, which cannot be optimal for each micron of the engagement surface over the length of the jaws. Such temperature sensors also suffer from a time lag. Still further, such prior art temperature sensors provide only an indirect reading of actual tissue temperature—since a typical sensor can only measure the temperature of the electrode.
- Other alternative modes of operating the conductive-resistive matrix system are possible. In one other mode of operation, the
system controller 182 coupled tovoltage source 180 can acquire data from current flow circuitry that is coupled to the first and second polarity conductors in the jaws (in any locations described previously) to measure the blended impedance of current flow between the first and second polarity conductors through the combination of (i) the engaged tissue and (ii) the matrix CM. This method of the invention can provide algorithms within thesystem controller 182 to modulate, or terminate, power delivery to the working end based on the level of the blended impedance as defined above. The method can further include controlling energy delivery by means of power-on and power-off intervals, with each such interval having a selected duration ranging from about 1 microsecond to one second. The working end andsystem controller 182 can further be provided with circuitry and working end components of the type disclosed in Provisional U.S. patent application Ser. No. 60/339,501 filed Nov. 9, 2001 (Docket No. S-BA-001) titled Electrosurgical Instrument, which is incorporated herein by reference. - In another mode of operation, the
system controller 182 can be provided with algorithms to derive the temperature of the matrix CM from measured impedance levels—which is possible since the matrix is engineered to have a selected unique resistance at each selected temperature over a temperature-resistance curve (seeFIG. 6 ). Such temperature measurements can be utilized by thesystem controller 182 to modulate, or terminate, power delivery to engagement surfaces based on the temperature of the matrix CM. This method also can control energy delivery by means of the power-on and power-off intervals as described above. -
FIGS. 10-11 illustrate a sectional views of analternative jaw structure 100C—in which both the lower andupper engagement surfaces FIGS. 7A-7B and 8A-8B to apply energy to engaged tissue. The jaw structure ofFIGS. 10-11 illustrate that the tissue is engaged on opposing sides by a conductive-resistive matrix, with each matrix CMA and CMB in contact with an opposing polarity electrode indicated at 165 and 185, respectively. It has been found that providing cooperating first and second conductive-resistive matrices in opposing first and second engagement surfaces can enhance and control both active ohmic heating and the passive conduction of thermal effects to the engaged tissue. - 3. Type “B” conductive-resistive matrix system for tissue welding.
FIGS. 12 and 14 A-14C illustrate anexemplary jaw assembly 200 that carries a Type “B” conductive-resistive matrix system for (i) controlling Rf energy density and microcurrent paths in engaged tissue, and (ii) for contemporaneously controlling passive conductive heating of the engaged tissue. The system again utilizes an elastomeric conductive-resistive matrix CM although substantially rigid conductive-resistive matrices of a ceramic positive-temperature coefficient material are also described and fall within the scope of the invention. Thejaw assembly 200 is carried at the distal end of an introducer member, and can be a scissor-type structure (cf.FIG. 4 ) or a transecting-type jaw structure (cf.FIGS. 3A-3B ). For convenience, thejaw assembly 200 is shown as a scissor-type instrument that allows for clarity of explanation. - The Type “A” system and method as described above in
FIGS. 5 and 7 A-7B allowed for effective pixel-by-pixel power modulation—wherein microscale spatial locations can be considered to apply an independent power level at a localized tissue contact. The Type “B” conductive-resistive matrix system described next not only allows for spatially localized power modulation, it additionally provides for the timing and dynamic localization of Rf energy density in engaged tissues—which can thus create a “wave” or “wash” of a controlled Rf energy density across the engaged tissue reduced to membrane thickness. - Of particular interest, referring to
FIG. 12 , the Type “B” system according to the invention provides an engagement surface layer of at least onejaw first polarity electrode 220 having exposedsurface portion 222 andsecond polarity electrode 225 having exposedsurface portion 226. Thus, the microcurrents within tissue during a brief interval of active heating can flow to and from said exposedsurface portions same engagement surface 255A. By providing opposingpolarity electrodes polarity electrodes engagement surfaces FIG. 11 wherein the upper jaw's engagement surface 250B is an insulator indicated at 252. - More in particular, referring to
FIG. 12 , the first (lower)jaw 212A is shown in sectional view with a conductive-resistive matrix CM exposed in a central portion ofengagement surface 255A. Afirst polarity electrode 220 is located at one side of matrix CM with thesecond polarity electrode 225 exposed at the opposite side of the matrix CM. In the embodiment ofFIG. 12 , the body orsupport structure 258 of the jaw comprises theelectrodes insulated body portion 262. Further, the exterior of the jaw body is covered by aninsulator layer 261. The matrix CM is otherwise in contact with theinterior portions electrodes - The jaw assembly also can carry a plurality of alternating opposing
polarity electrode portions FIGS. 13A-13C . Any of these arrangements of electrodes and intermediate conductive-resistive matrix will function as described below at a reduced scale—with respect to any paired electrodes and intermediate matrix CM. -
FIGS. 14A-14C illustrate sequential views of the method of using of the engagement surface layer ofFIG. 11 to practice the method of the invention as relating to the controlled application of energy to tissue. For clarity of explanation,FIGS. 14A-14C depict exposedelectrode surface portions FIG. 14A , theupper jaw 212B and engagement surface 250B is shown in phantom view, and comprises aninsulator 252. The gap dimension g is not to scale, as described previously, and is shown with the engaged tissue having a substantial thickness for purposes of explanation. -
FIG. 14A provides a graphic illustration of the matrix CM within engagement surface layer 250A at time T1—the time at which electrical potential of a first polarity (indicated at +) is applied toelectrode 220 via an electrical lead fromvoltage source 180 andcontroller 182. InFIGS. 14A-14C , the sphericalgraphical elements 162 of the matrix are not-to-scale and are intended to represent a “region” of conductive particles within the non-conductive elastomer 164. Thegraphical elements 162 thus define a polarity at particular microsecond in time just after the initiation of power application. InFIG. 14A , the bodyportion carrying electrode 225 defines a second electrical potential (−) and is coupled tovoltage source 180 by an electrical lead. As can be seen inFIG. 14A , thegraphical elements 162 are indicated as having a transient positive (+) or negative (−) polarity in proximity to the electrical potential at the electrodes. When thegraphical elements 162 have no indicated polarity (seeFIGS. 14B & 14C ), it means that the matrix region has been elevated to a temperature at the matrix switching range wherein electrical conductance is limited, as illustrated in that positively sloped temperature-resistance curve ofFIG. 6 and the graphical representation ofFIG. 8B . - As can be seen in
FIG. 14A , the initiation of energy application at time T1 causes microcurrents mc within the central portion of the conductive matrix CM as current attempts to flow between the opposingpolarity electrodes FIG. 14A , which results in highly localized ohmic heating and denaturation effects along that interface which extends from the matrix CM into the engaged tissue. Thus,FIG. 14A provides a simplified graphical depiction of the interface or plane P that defines the “non-random” localization of ohmic heating and denaturation effects—which contrasts with all prior art methods that cause entirely random microcurrents in engaged tissue. In other words, the interface between the opposing polarities wherein active Rf heating is precisely localized can be controlled and localized by the use of the matrix CM to create initial heating at that central tissue location. - Still referring to
FIG. 14A , as the tissue is elevated in temperature in this region, the conductive-resistive matrix CM in that region is elevated in temperature to its switching range to become substantially non-conductive (seeFIG. 6 ) in that central region. -
FIG. 14B graphically illustrates the interface or plane P at time T2—an arbitrary microsecond or millisecond time interval later than time T1. The dynamic interface between the opposing polarities wherein Rf energy density is highest can best be described as planes P and P′ propagating across the conductive-resistive matrix CM and tissue that are defined by “interfaces” between substantially conductive and non-conductive portions of the matrix—which again is determined by the localized temperature of the matrix. Thus, the microcurrent mc′ in the tissue is indicated as extending through the tissue membrane with the highest Rf density at the locations of planes P and P′. Stated another way, the system creates a front or wave of Rf energy density that propagates across the tissue. At the same time that Rf density (ohmic heating) in the localized tissue is reduced by the adjacent matrix CM becoming non-conductive, the matrix CM will begin to apply substantial thermal effects to the tissue by means of passive conductive heating as described above. -
FIG. 14C illustrates the propagation of planes P and P′ at time T3—an additional arbitrary time interval later than T2. The conductive-resistive matrix CM is further elevated in temperature behind the interfaces P and P′ which again causes interior matrix portions to be substantially less conductive. The Rf energy densities thus propagate further outward in the tissue relative to theengagement surface 255A as portions of the matrix change in temperature. Again, the highest Rf energy density will occur at generally at the locations of the dynamic planes P and P′. At the same time, the lack of Rf current flow in the more central portion of matrix CM can cause its temperature to relax to thus again make that central portion electrically conductive. The increased conductivity of the central matrix portion again is indicated by (+) and (−) symbols inFIG. 14C . Thus, the propagation of waves of Rf energy density will repeat itself as depicted inFIGS. 14A-14C which can effectively weld tissue. - Using the methods described above for controlled Rf energy application with paired electrodes and a conductive-resistive matrix CM, it has been found that time intervals ranging between about 500 ms and 4000 ms can be sufficient to uniformly denature tissue constituents re-cross-link to from very strong welds in most tissues subjected to high compression. Other alternative embodiments are possible that multiply the number of cooperating opposing
polarity electrodes -
FIG. 15 depicts an enlarged view of the alternative Type “B”jaw 212A ofFIG. 13A wherein the engagement surface 250A carries a plurality of exposed conductive matrix portions CM that are intermediate a plurality of opposingpolarity electrode portions lower jaw 212A has a structural body that comprises theelectrodes insulator member 266 that provide the strength required by the jaw. Aninsulator layer 261 again is provided on outer surfaces of the jaw excepting theengagement surface 255A. The upper jaw (not shown) of the jaw assembly can comprise an insulator, a conductive-resistive matrix, an active electrode portion or a combination thereof. In operation, it can be easily understood that each region of engaged tissue between each exposedelectrode portion FIGS. 14A-14C . - The type of engagement surface 250A shown in
FIG. 15 can have electrode portions that define an interior exposed electrode width ew ranging between about 0.005″ and 0.20″ with the exposedoutboard electrode surface - In the embodiment of
FIG. 15 , theelectrode portions insulator member 266 of the jaw body thus substantially preventing flexing of the engagement surface even though the matrix CM may be a flexible silicone elastomer.FIG. 16 shows an alternative embodiment wherein theelectrode portions -
FIG. 17 illustrates an alternative Type “B” embodiment that is adapted for further increasing passive heating of engaged tissue when portions of the matrix CM are elevated above its selected switching range. Thejaws FIG. 17 illustrates that afirst polarity electrode 220 is a thin layer of metallic material that floats on the matrix CM and is bonded thereto by adhesives or any other suitable means. The thickness of floatingelectrode 220 can range from about 1 micron to 200 microns. Thesecond polarity electrode 225 has exposedportions engagement planes FIG. 17 creates controlled thermal effects in engaged tissue by several different means. First, as indicated inFIGS. 18A-18C , the dynamic waves of Rf energy density are created between the opposingpolarity electrode portions engagement surface layer 255B cause microcurrents between the engagement surface layers 255A and 255B, as well as to the outboard exposed electrode surfaces exposedportions -
FIG. 19 illustrates another Type “B” embodiment of jaws structure that again is adapted for enhanced passive heating of engaged tissue when portions of the matrix CM are elevated above its selected switching range. Thejaws engagement surface layer 255B is convex and has an elastomeric hardness ranging between about 20 and 80 in the Shore A scale and is fabricated as described previously. - Of particular interest, the embodiment of
FIG. 19 depicts afirst polarity electrode 220 that is carried in a central portion ofengagement plane 255A but the electrode does not float as in the embodiment ofFIG. 17 . Theelectrode 220 is carried in a first matrix portion CM1 that is a substantially rigid silicone or can be a ceramic positive temperature coefficient material. Further, the first matrix portion CM, preferably has a differently sloped temperature-resistance profile (cf.FIG. 6 ) that the second matrix portion CM2 that is located centrally in thejaw 212A. The first matrix portion CM1, whether silicone or ceramic, has a hardness above about 90 in the Shore A scale, whereas the second matrix portion CM2 is typically of a silicone as described previously with a hardness between about 20 and 80 in the Shore A scale. Further, the first matrix portion CM1 has a higher switching range than the second matrix portion CM2. In operation, the wave of Rf density across the engaged tissue fromelectrode 220 to outboard exposedelectrode portions FIGS. 18A-18C . The rigidity of the first matrix CM1 prevents flexing of theengagement plane 255A. During use, passive heating will be conducted in an enhanced manner to tissue fromelectrode 220 and the underlying second matrix CM2 which has a second selected lower temperature switching range, for example between about 60° C. to 70° C. This Type “B” system has been found to be very effective for rapidly welding tissue—in part because of the increased surface area of theelectrode 220 when used in small cross-section jaw assemblies (e.g., 5 mm. working ends). -
FIG. 20 shows theengagement plane 255A ofFIG. 17 carried in a transecting-type jaws assembly 200D that is similar to that ofFIGS. 3A-3B . As described previously, the Type “B” conductive-resistive matrix assemblies ofFIGS. 12-19 are shown in a simplified form. Any of the electrode-matrix arrangements ofFIGS. 12-19 can be used in the cooperating sides of a jaw with a transecting blade member—similar to the embodiment shown inFIG. 20 . - 3. Type “C” system for tissue welding.
FIGS. 21 and 22 illustrate anexemplary jaw assembly 400 that carries a Type “C” system that optionally utilizes at least one conductive-resistive matrix CM as described previously for (i) controlling Rf energy density and microcurrent paths in engaged tissue, and (ii) for contemporaneously controlling passive conductive heating of the engaged tissue. - In
FIG. 21 , it can be seen thatjaws respective engagement surfaces 455A and 455B. Theupper jaw 412B and engagement surface 455B can be as described in the embodiment ofFIGS. 17 and 19 , or the upper engagement surface can be fully insulated as described in the embodiment ofFIGS. 14A-14C . Preferably, upper engagement surface layer 455B is convex and made of an elastomeric material as described above. Both jaws have astructural body portion body portions electrical source 180 and have exposedsurfaces portions 472 a and 472 b in the jaws' engagement planes to serve as an electrode defining a first polarity, as thesurface portions 472 a and 472 b are coupled to, and transition into, themetallic film layer 475 described next. - As can be seen in
FIG. 21 , theentire engagement surface 455A of thelower jaw 412A comprises any thin conductive metallic film layer indicated at 475. For example, the layer can be of platinum, titanium, gold, tantalum, etc. or any alloy thereof. The thin film metallization can be created by electroless plating, electroplating processes, sputtering or other vapor deposition processes known in the art, etc. The film thickness ft of themetallic layer 475 can be from about 1 micron to 100 microns. More preferably, themetallic film layer 475 is from about 5 to 50 microns. - The matrix CMA preferably is substantially rigid but otherwise operates as described above. The
metallic film layer 475 is shown as having an optional underlying conductive member indicated at 477 that is coupled toelectrical source 180 and thus comprises an electrode that defined a second polarity. - Of particular interest, referring to
FIG. 22 , it can be seen thatengagement surface 455A entirely comprises the thinmetallic film layer 475 that is coupled in spaced apartportions FIG. 22 ,intermediate portions 485 of the metallic film layer 475 (that are intermediate the central and outboard metallic film portions coupled to the opposing polarities of the electrical source) are made to have an altered resistance to current flow therethrough to thereby induce microcurrents to flow through adjacent engaged tissue rather than throughintermediate portions 485. This can be advantageous for precise control of localizing the microcurrents in engaged tissue. At the same time, the thin dimension of thefilm 475 allows for very rapid adjustment in temperature and thus allows enhanced passive conductive heating of engaged tissue when the engaged tissue is no longer moist enough for active Rf density therein. One preferred manner of fabricating theintermediate portions 485 is to provide perforations orapertures 488 therein that can range in size from about 5 microns to 200 microns. Stated another way, theintermediate portions 485 can haveapertures 488 therein that make the regions from about 1 percent to 60 percent open, no matter the size or shape of the apertures. More preferably, theintermediate portions 485 are from about 5 percent to 40 percent open. Theapertures 488 can be made in thefilm 475 by any suitable means, such as photo-resist methods. As shown inFIG. 22 , theintermediate portions 485 are not-to-scale and have a width w that ca range from about 0.005″ to 0.20″ in a typical electrosurgical jaw. - 4. Type “D” conductive-resistive matrix system for tissue welding. FIGS. 23 to 28 illustrate exemplary Type “D” jaw structures that utilize a conductive-resistive matrix CM or variable resistive body in a different manner than described previously. The Type “D” system still controls Rf energy density and microcurrent paths in engaged tissue as described previously, but also provides means for a more “focused” application of passive conductive heating of engaged tissue.
-
FIG. 23 illustrates a firstexemplary jaw assembly 500A that carries a Type “D” conductive-resistive matrix system CM that can comprise an elastomeric or non-elastomeric matrix of a positive temperature coefficient material. Thejaw assembly 500A is carried at the distal end of an introducer member, and can be a scissor-type structure (cf.FIG. 4 ) or a transecting-type jaw structure (cf.FIGS. 3A-3B ). InFIG. 23 , thejaw assembly 500A is shown as a scissor-type instrument to allow for simplified of explanation of the features corresponding to the invention. - Still referring to
FIG. 23 , thejaw assembly 500A depicts first andsecond jaws engagement surfaces jaw 512B carries a conductor or returnelectrode element 558 that is exposed inengagement surface 555B which in turn is coupled tovoltage source 180 and is indicated for purposes of explanation as having a negative (—) polarity. - The lower (first)
jaw 512A carries an opposing polarity (+)electrode element 560 that is embedded within the interior ofinsulator material 556 that makes up the exterior body ofjaw 512A. The jaw has a thinsurface conductor element 565 inengagement surfaces 555A—that is not directly coupled to thevoltage source 180. Rather, thesurface conductor element 565 is electrically/conductively coupled to thesurface conductor 565 only by an intermediate conductive-resistive matrix CM that contacts both theactive electrode 560 and thesurface conductor 565. Of particular interest, the conductive-resistive matrix CM has a cross-section that diminishes in the direction of thesurface conductor 565. InFIG. 23 , the matrix CM is shown with triangular cross-section that tapers from afirst region 580 that has an extended dimension to a secondreduced dimension region 585 that conductively contacts thecentral portion 570 ofsurface conductor 565. It can be understood theconductive material 565 only functions as an electrode to actively conduct current to engaged tissue when the matrix CM is below its switching range. At other time intervals when the matrix CM is above its switching range, thesurface conductor 565 will not provide current paths to the engaged tissue and passive heating of the matrix will be focused in thecentral portion 570 of thesurface conductor 565. - It can be understood that when the conductive-resistive matrix CM is below it switching temperature range, current will flow between the interior
active electrode 560 and thesurface conductor 565. However, when the engaged tissue is elevated in temperature, which elevates the temperature ofsurface conductor 565, the portion of the matrix CM proximate to surfaceconductor 565 will be heated to its switching range before the other portions of the matrix more proximate to the interioractive electrode 560. Thereafter, the conductive-resistive matrix CM will have a temperature that hovers about the upper end of its switching range, which also is the targeted tissue treatment range. Contemporaneously, the central matrix portion will focus its passive (conductive) heating at a selected location within theengagement surface 255B. It has been found that a centrally focused passive heating as depicted in the embodiment ofFIG. 23 is very useful in tissue welding. It should be appreciated that the scope of the invention includes the use of a positive or negative temperature coefficient material volume CM intermediate anactive electrode 560 and asurface conductor 565 that engages tissue wherein the surface area of the matrix material CM has a firstgreater surface area 580 in contact with theactive electrode 560 and secondlesser surface area 585 in contact with theconductor 565 in the engagement surface. The actual cross-section of the matrix volume CM can be any shape such as triangular, pyramid-shaped, “T”-shaped, or any curvilinear shape that tapers. In a preferred embodiment, the secondlesser surface area 585 in contact withsurface conductor 565 is less than about 50 percent of thefirst surface area 580. More preferably, thesecond surface area 585 in contact withsurface conductor 565 is less than about 25 percent of thefirst surface area 580 in contact withactive electrode 560. - As can be seen in
FIG. 23 , the jaws carry insulated projectingelements 572 that can be located anywhere in the engagement surfaces 555A and 555B or the jaw perimeters to prevent inadvertent contact between the opposing polarity electrodes when the jaws are moved toward the fully closed position. -
FIG. 24 illustrates anotherexemplary jaw structure 500B that corresponds to the invention. The positive temperature coefficient conductive-resistive matrix CM is again intermediateactive electrode 560 andsurface conductor 565 but this embodiment has a jaw that carries a plurality of matrix portions CM1 to CM3 that create a gradient in the temperature coefficients of resistance within adjacent matrix portions. The matrices CM1 to CM3 are graphically indicated (not-to-scale) to haveconductive particles 160 b of different dimensions/volumes within thenon-conductive portion 160 a of each matrix to provide varied temperature-resistance curves. It should be appreciated that any temperature coefficient material of any type can be used, or any combination of materials or material types can be used to fabricate the gradient. In this embodiment, the more outboard matrix CM3 that is exposed in theengagement surface 555A has a significantly greater resistance to current flow therethrough than the embedded matrix CM1 that defines that tapered cross-section between theinterior electrode 560 and the surface conductor. Also the overall matrix can have a substantially continuous gradient across the matrix volume and fall within the scope of the invention. - In the embodiment of
FIG. 24 , thesurface conductor 565 is directly coupled to thevoltage source 180 to define a polarity (−) therein that opposes the polarity of theinterior electrode 560. The embodiment ofFIG. 24 has its interior electrode 560 (indicated with positive (+) polarity) comprising a body portion of thelower jaw 512A with aninsulative coating 586 about most of its exterior surface. In theengagement surface 555A, the laterally outward portions indicated at 588 also are exposed portions of theinterior electrode 560 indicated as having positive (+) polarity. Theseelectrode portions 588 cooperate with theconductor 565 as described in the embodiments ofFIGS. 14A-14C andFIGS. 18A-18C . - In the embodiment of
FIG. 24 , the central matrix portion CM1 has a switching range and contact area with thesurface conductor 565 to thus cause differential passive conductive heating across the engagements surface 555A andconductor 565, with the more focused passive heating in thecentral region 570 of the jaw surface. As can be seen inFIG. 24 , the central matrix portion CM1 has a firstgreater surface area 580 in contact withinterior electrode 560 and a secondlesser surface area 585 in contact withsurface conductor 565. - The embodiment of
FIG. 24 further illustrates that matrix CM2 and matrix CM3 have surface areas that differ between the contacts with theinterior electrode 560 and thesurface conductor 565. These surface areas can be manipulated to advantage to cause “focused” active or passive heating at theengagement surface 555A, and the scope of the invention includes having any such surface area larger than the other. - Now turning to
FIG. 25 , another alternative jaw assembly 500C is shown that has alower jaw 512A that is identical to the embodiment ofFIG. 23 . However, theupper jaw 512B has anengagement surface 555B that carries an exposed conductive-resistive matrix CM5 that is coupled to areturn electrode 558 embedded at an interior of the upper jaw body that is fabricated of aninsulator material 556. Such an upper jaw was shown inFIGS. 10, 11 and 17. In all other respects, the working end functions as described previously. - Referring now to
FIGS. 26A-26B , anotheralternative jaw assembly 500D is shown that is a scissor-type instrument similar to the embodiment ofFIG. 4 , except that the jaws carry aslot 590 for receiving an extendable transecting blade (not shown). InFIG. 26B , it can be seen thatlower jaw 512A has left andright sides FIG. 23 . In all respects, the embodiment ofFIG. 26B functions as described above. - Referring now to
FIG. 27 , anotheralternative jaw assembly 500E is shown that again has scissor-type first andsecond jaws slot 590 in the jaws for receiving an extendable blade for transecting tissue in the embodiment ofFIG. 26B . In this embodiment, theengagement surface 555A of the lower jaw carries lateral conductive elements indicated at 592 (collectively) of an opposing (−) polarity from the spaced apart centralconductive element 565, making this embodiment function in the manner of the embodiments shown in FIGS. 17, 18A-18C, 19, 20 and 24. The material between theconductive element 565 and the lateralconductive elements 592 can be a conductive-resistive matrix (cf.FIG. 24 ) or an insulative material (seeFIG. 27 ). -
FIG. 28 illustrates analternative jaw assembly 500F that is similar to the embodiment ofFIG. 27 . The first andsecond jaws engagement surface 555A of thelower jaw 512A again carries lateralconductive elements 588 coupled to the voltage source to define an opposing polarity from the spaced apart from centralconductive element 565 which is also coupled to the voltage source. The jaw carries first and conductive-resistive matrices CM1 and CM2. An insulative layer 589 is provided about the exterior surface of the jaw. Again, the pyramidal cross-section central matrix CM1 has first and second contact areas of greater and lesser dimensions, respectively, for contacting theinterior electrode 560 and thesurface conductor 565. - Now referring to
FIG. 29 , the systems and method of the invention for controlled application of energy to tissue can be described in terms of equilibrium temperature and impedance characteristics that can be designed into the working end.FIG. 29 first illustrates a typical power output-impedance curve for a radiofrequency generator (voltage source 180).FIG. 29 also illustrates a selected temperature-impedance curve for the conductive-resistive matrix CM corresponding to the invention. It can be easily understood that thevoltage source 180 and matrix CM can be designed to provide a selected equilibrium temperature EQ which is indicated at the intersection of the curves. Thus, one preferred system of the invention comprises (i) a working end that carries a matrix as described above having a positive temperature coefficient of resistance that defines a selected temperature-impedance curve, and (ii) a voltage source that defines a selected power output-impedance curve wherein the temperature-impedance curve and power output-impedance curve define an equilibrium temperature at which the matrix dissipates power output from the voltage source to thereby maintain said equilibrium temperature within the matrix CM. Practicing the method of the invention thus consists of (i) providing the matrix CM and voltage source as described above, (ii) engaging tissue with the engagement surface, and (iii) applying electrosurgical energy to the tissue through the matrix material wherein the selected temperature-impedance curve and power output-impedance curve define the matrix's dissipation of power to thereby maintain a selected temperature in the engaged tissue. The equilibrium temperature EQ can be any temperature, and for the purposes of welding tissue can be between about 60° C. and 100° C. More preferably, the equilibrium temperature EQ is between about 65° C. and 85° C. - In another aspect of the invention, still referring to
FIG. 29 , the invention provides an electrosurgical system that insures that tissue will not be desiccated, and insures that sparks will not cross any gaps between the engagement surface and tissue which thereby prevents tissue from sticking to the engagement surface. The system provides a conductive-resistive matrix material CM that is exposed in an engagement surface that receives electrosurgical energy, or coupled to a conductor in the engagement surface, wherein the matrix material defines a positive temperature coefficient of resistance. The invention further provides a radiofrequency energy source or voltage source for generating the electrosurgical energy. Further, the matrix CM is designed so that the combined impedance of engaged tissue and the matrix material CM is such that voltage developed across any gap of a selected dimension between the engagement surface and the tissue is less than the breakdown voltage required to cross of a gap having that selected dimension. In other words, the source's power-impedance curve and the matrix's impedance-temperature curve—together with that potential impedance parameters of the engaged tissue—can be engineered to insure that no sparks will jump across a gap in the interface between the engagement surface and the tissue. - In another aspect of the invention, the invention can provide a
controller 182 coupled to thevoltage source 180 that includes algorithms that convert energy delivery from a continuous mode to a pulsed mode upon the system reaching a selected parameter such as an impedance level. For example, thecontroller 182 can alter energy delivery to a pulsed mode upon the combination of the matrix CM and the engaged tissue reaching a particular impedance level. It has been found that such a pulsed mode of energy delivery will allow moisture within the tissue to re-hydrate the engaged tissue to further prevent tissue desiccation, while still maintaining the targeted tissue temperature. - Although particular embodiments of the present invention have been described above in detail, it will be understood that this description is merely for purposes of illustration. Specific features of the invention are shown in some drawings and not in others, and this is for convenience only and any feature may be combined with another in accordance with the invention. Further variations will be apparent to one skilled in the art in light of this disclosure and are intended to fall within the scope of the appended claims.
Claims (10)
1. An electrosurgical system, comprising:
an instrument body having a proximal end and a working end that has an engagement surface for engaging tissue;
a voltage source that defines a selected power output-impedance curve; and
a variable resistive body carried in the working end that is positioned intermediate said engagement surface and an interior electrode coupled to the voltage source, the variable resistive body having a positive temperature coefficient of resistance that defines a selected temperature-impedance curve;
wherein the temperature-impedance curve and power output-impedance curve are selected to define a selected equilibrium temperature at which the variable resistive body dissipates power output from the voltage source to thereby maintain said equilibrium temperature in the matrix.
2. The electrosurgical system of claim 1 , wherein the selected equilibrium temperature is between about 60° C. and 100° C.
3. The electrosurgical system of claim 1 , wherein the selected equilibrium temperature is between about 65° C. and 85° C.
4. An electrosurgical system for controlled application of energy to tissue, comprising:
an instrument with a working end that has an engagement surface for engaging tissue;
a voltage source that defines a selected power output-impedance curve; and
a variable resistive body carried in the working end that is positioned intermediate said engagement surface and an interior electrode coupled to the voltage source, the variable resistive body having a positive temperature coefficient of resistance that defines a selected temperature-impedance curve;
wherein the temperature-impedance curve and power output-impedance curve are selected to define a breakdown voltage across a gap of a selected dimension.
5. An electrosurgical system for controlled application of energy to tissue, comprising:
an instrument with a working end that has an engagement surface for engaging tissue;
a voltage source and controller for operatively coupled to the engagement surface;
a variable resistive body carried in the working end that is positioned intermediate said engagement surface and an interior working end electrode coupled to the voltage source, the variable resistive body having a positive temperature coefficient of resistance that defines a selected temperature-impedance curve;
wherein the controller defines means for modulating energy delivery between continuous modes and pulsed modes in response to the impedance of the variable resistive body.
6. The electrosurgical system of claim 5 , wherein the controller defines means for modulating energy delivery between continuous modes and pulsed modes in response to the impedance of the combination of variable resistive body and the engaged tissue.
7. An electrosurgical method, comprising the steps of:
providing an instrument with a working end that has an engagement surface for engaging tissue;
providing a voltage source that defines a power output-impedance curve;
providing a variable resistive body that is intermediate the engagement surface and the voltage source, the variable resistive body having a positive temperature coefficient of resistance and thereby defining a selected temperature-impedance curve; and
engaging tissue with the engagement surface and applying electrosurgical energy to the tissue through the variable resistive body;
wherein the selected temperature-impedance curve and power output-impedance curve will define a point at which the variable resistive body dissipates power to maintain a selected temperature in the engaged tissue.
8. The electrosurgical method of claim 7 , wherein the selected temperature is between about 60° C. and 100° C.
9. The electrosurgical method of claim 7 , wherein the selected temperature is between about 65° C. and 85° C.
10. An electrosurgical method, comprising the steps of:
providing an instrument with a working end that has an engagement surface for engaging tissue;
providing a voltage source operatively coupled to the working end;
providing a variable resistive body within the working end that is intermediate the engagement surface and an interior electrode coupled to the voltage source, the variable resistive body defining a positive temperature coefficient of resistance;
engaging tissue with the engagement surface and applying electrosurgical energy to the tissue through the variable resistive body;
wherein the combined impedance of the tissue and the variable resistive body is such that voltage developed across any gap between the engagement surface and the tissue is less than the breakdown voltage required to cross a selected gap dimension.
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