DETAILED DESCRIPTION OF THE INVENTION

[0082] 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. The jaw assembly 100 A is carried at the distal end 104 of an introducer 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. The introducer 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. The introducer 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.

[0083] As can be seen in FIGS. 3A and 3B , the jaw assembly 100 A has first (lower) jaw element 112 A and second (upper) jaw element 112 B 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 of FIGS. 3A and 3B , both jaws are moveable relative to the introducer portion 106 .

[0084] Of particular interest, the opening-closing mechanism of the jaw assembly 100 A is capable of applying very high compressive forces on tissue on the basis of cam mechanisms with a reciprocating member 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.

[0085] In the embodiment of FIGS. 3A and 3B , a reciprocating member 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 of member 140 . The proximal end 141 and medial portion of member 140 are dimensioned to reciprocate within bore 108 of introducer sleeve 106 . The distal portion 142 of reciprocating member 140 carries first (lower) and second (upper) laterally-extending flange elements 144 A and 144 B that are coupled by an intermediate transverse element 145 . The transverse element further is adapted to transect tissue captured between the jaws with a leading edge 146 ( FIG. 3A ) that can be a blade or a cutting electrode. The transverse element 145 is adapted to slide within a channels 148 a and 148 b in the paired first and second jaws to thereby open and close the jaws. The camming action of the reciprocating 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.

[0086] In FIGS. 3A and 3B , the first and second jaws 112 A and 112 B close about an engagement plane 150 and define tissue-engaging surface layers 155 A and 155 B that contact and deliver energy to engaged tissues from electrical energy means as will be described below. The jaws can have any suitable length with teeth or serrations 156 for gripping tissue. One preferred embodiment of FIGS. 3A and 3B provides such serrations 156 at an inner portion of the jaws along channels 148 a and 148 b thus allowing for substantially smooth engagement surface layers 155 A and 155 B laterally outward of the tissue-gripping elements. The axial length of jaws 112 A and 112 B indicated at L can be any suitable length depending on the anatomic structure targeted for transection and sealing and typically will range from about 10 mm. to 50 mm. The jaw assembly can apply very high compression over much longer lengths, for example up to about 200 mm., for resecting and sealing organs such as a lung or liver. The scope of the invention also covers jaw assemblies for an instrument used in micro-surgeries wherein the jaw length can be about 5.0 mm or less.

[0087] In the exemplary embodiment of FIGS. 3A and 3B , the engagement surface 155 A of the lower jaw 112 A is adapted to deliver energy to tissue, at least in part, through a conductive-resistive matrix CM corresponding to the invention. The tissue-contacting surface 155 B of upper jaw 112 B 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. 10/032,867 filed Oct. 22, 2001 (Docket No. SRX-011) titled Electrosurgical Jaw Structure for Controlled Energy Delivery and U.S. patent application Ser. No. 10/308,362 filed Dec. 3, 2002 (Docket No. SRX-012A) titled Electrosurgical Jaw Structure for Controlled Energy Delivery, both of which are incorporated herein by reference.

[0088] Referring now to FIG. 4 , an alternative jaw structure 100 B is shown with lower and upper jaws having similar reference numerals 112 A- 112 B. The simple scissor-action of the jaws in FIG. 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 comparing FIGS. 3B and 4 , the jaws of either embodiment 100 A or 100 B can carry the same energy delivery components, which is described next.

[0089] 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″.

[0090] 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 155 A and 155 B in jaws 112 A and 112 B. It should be appreciated that the engagement surface layers 155 A and 155 B are shown in a scissors-type jaw (cf. FIG. 4 ) for convenience, and the conductive-resistive matrix system would be identical in each side of a transecting jaw structure as shown in FIGS. 3 A- 3 B.

[0091] In FIG. 5 , it can be seen that the lower jaw 112 A carries a component described herein as a conductive-resistive matrix CM that is at least partly exposed to an engagement plane 150 that is defined as the interface between tissue and a jaw engagement surface layer, 155 A or 155 B. More in particular, the conductive-resistive matrix CM comprises a first portion 160 a and a second portion 160 b . The first portion is preferably an electrically nonconductive 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, the first portion 160 a of the matrix is an elastomer, for example a medical grade silicone. The first 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.

[0092] Referring to FIG. 5 , the second portion 160 b of the matrix CM is a material that is electrically conductive and that is distributed within the first portion 160 a . In FIG. 5 , the second portion 160 b is represented (not-to-scale) as spherical elements 162 that are intermixed within the elastomer first portion 160 a of matrix CM. The elements 162 can have any regular or irregular shape, and also can be elongated elements or can comprise conductive filaments. The dimensions of elements 162 can range from nanoparticles having a scale of about 1 nm. to 2 nm. across a principal axis thereof to much larger cross-sections of about 100 microns in a typical jaw structure. In a very large jaw, the elements 162 in matrix CM can have a greater dimension that 100 microns in a generally spherical form. Also, the matrix CM can carry a second portion 160 b in the form of an intertwined filament (or filaments) akin to the form of steel wool embedded within an elastomeric first portion 160 a and fall within the scope the invention. Thus, the second portion 160 b can be of any form that distributes an electrically conductive mass within the overall volume of the matrix CM.

[0093] In the lower jaw 112 A of FIG. 5 , the matrix CM is carried in a support structure or body 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, the support structure 158 carries an insulative coating 159 to prevent electrical current flow to tissues about the exterior of the jaw assembly and between support structure 158 and the matrix CM and a conductive element 165 therein.

[0094] Of particular interest, the combination of first and second portions 160 a and 160 b provide a matrix CM that is variably resistive (in ohms-centimeters) in response to temperature changes therein. The matrix composition with the temperature-dependent resistance is alternatively described herein as a temperature coefficient material. In one embodiment, by selecting the volume proportion of first portion 160 a of the non-conductive elastomer relative to the volume proportion of second portion 160 b of the conductive nanoparticles or elements 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.

[0095] 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 in FIG. 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, see FIG. 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 in FIG. 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.

[0096] 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.

[0097] 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 ceramic first portion 160 a in combination with compressible-particle second portion 160 b of a such a metallized polymer bead to create the effects illustrated in FIGS. 8 A- 8 B below).

[0098] One aspect of the invention relates to the use of a matrix CM as illustrated schematically in FIG. 5 in a jaw's engagement surface layer 155 A with a selected treatment range between a first temperature (TE ₁ ) and a second temperature (TE ₂ ) that approximates the targeted tissue temperature for tissue welding (see FIG. 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.

[0099] 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 thererough 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.

[0100] In the exemplary jaw 112 A of FIG. 5 , the entire surface area of engagement surface layer 155 A 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 md ₁ and md ₂ in FIG. 5 , and preferably such a cross-section comprises a significant fractional volume of the jaw relative to support structure 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.

[0101] As can be seen in FIG. 5 , the interior of jaw 112 A carries a conductive element (or electrode) indicated at 165 that interfaces with an interior surface 166 of the matrix CM. The conductive element 165 is coupled by an electrical lead 109 a to a voltage (Rf) source 180 and optional controller 182 ( FIG. 4 ). Thus, the Rf source 180 can apply electrical potential (of a first polarity) to the matrix CM through conductor 165 —and thereafter to the engagement plane 150 through matrix CM. The opposing second jaw 112 B in FIG. 5 has a conductive material (electrode) indicated at 185 coupled to source 180 by lead 109 b that is exposed within the upper engagement surface 155 B.

[0102] In a first mode of operation, referring to FIG. 5 , electrical potential of a first polarity applied to conductor 165 will result in current flow through the matrix CM and the engaged tissue et to the opposing polarity 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 the engagement surface 155 A can modulate the delivery of energy to tissue et similar to the hypothetical engagement surface of FIG. 2 . Consider that the small sections of engagement surfaces represent the micron-sized, surface areas (or pixels) of the illustration of FIG. 2 (note that the jaws are not in a fully closed position in FIG. 5 ). The preferred membrane-thick engagement gap g is graphically represented in FIG. 5 .

[0103] FIGS. 7A and 8A illustrate enlarged schematic sectional views of jaws 112 A and 112 B and the matrix CM. It can be understood that the electrical potential at conductor 165 will cause current flow within and about the elements 162 of second portion 160 b along any conductive path toward the opposing polarity conductor 185 . FIG. 8A more particularly shows a graphic representation of paths of microcurrents mc _m within the matrix wherein the conductive elements 162 are in substantial contact. FIG. 7A also graphically illustrates paths of microcurrents mc _t 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 between jaws 112 A and 112 B. In fact, the flux of microcurrents mc _m within the matrix and the microcurrents mc _t within the engaged tissue will seek the most conductive paths—which will be assisted by the positioning of elements 162 in the surface of the engagement layer 155 A, which can act like surface asperities or sharp edges to induce current flow therefrom.

[0104] Consider that ohmic heating (or active heating) of the shaded portion 188 of engaged tissue et in FIGS. 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 in FIGS. 7B and 8B . In FIGS. 7B and 8B , the thermal coefficient of expansion of the elastomer of first matrix portion 160 a will cause slight redistribution of the second conductive portion 160 b within the matrix—naturally resulting in lessened contacts between the conductive elements 162 . It can be understood by arrows A in FIG. 8B that the elastomer will expand in directions of least resistance which is between the elements 162 since the elements are selected to be substantially resistant to compression.

[0105] 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 entire engagement surface layer 155 A, 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 in FIG. 8 B, the thermal expansion of the elastomeric matrix surface also will push into the membrane, further insuring tissue contact along the engagement plane 150 to eliminate any possibility of an energy arc across a gap.

[0106] 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.

[0107] FIG. 9 shows another embodiment of a conductive-resistive matrix CM that is further doped with elements 192 of a material that is highly thermally conductive with a selected mass that is adapted to provide substantial heat capacity. By utilizing such 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 of elements 162 can be both substantially electrically conductive and highly thermally conductive with a high heat capacity.

[0108] The manner of utilizing the system of FIGS. 7 A- 7 B to perform the method of the invention can be understood as mechanically compressing the engaged tissue et to membrane thickness between the first and second engagement surfaces 155 A and 155 B of opposing jaws and thereafter applying electrical potential of a frequency and power level known in electrosurgery to 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 from voltage 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.

[0109] 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, the elements 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.

[0110] 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.

[0111] Other alternative modes of operating the conductive-resistive matrix system are possible. In one other mode of operation, the system controller 182 coupled to voltage 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 the system 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 and system controller 182 can further be provided with circuitry and working end components of the type disclosed in Provisional U.S. Patent Application Serial No. 60/339,501 filed Nov. 9, 2001 (Docket No. S-BA-001) titled Electrosurgical Instrument, which is incorporated herein by reference.

[0112] 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 (see FIG. 6 ). Such temperature measurements can be utilized by the system 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.

[0113] FIGS. 10 - 11 illustrate a sectional views of an alternative jaw structure 100 C—in which both the lower and upper engagement surfaces 155 A and 155 B carry a similar conductive-resistive matrices indicated at CM _A and CM _B . It can be easily understood that both opposing engagement surfaces can function as described in FIGS. 7 A- 7 B and 8 A- 8 B to apply energy to engaged tissue. The jaw structure of FIGS. 10 - 11 illustrate that the tissue is engaged on opposing sides by a conductive-resistive matrix, with each matrix CM _A and CM _B 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.

[0114] 3. Type “B” conductive-resistive matrix system for tissue welding. FIGS. 12 and 14 A- 14 C illustrate an exemplary 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. The jaw 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. 3 A- 3 B). For convenience, the jaw assembly 200 is shown as a scissor-type instrument that allows for clarity of explanation.

[0115] The Type “A” system and method as described above in FIGS. 5 and 7 A- 7 B 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.

[0116] Of particular interest, referring to FIG. 12 , the Type “B” system according to the invention provides an engagement surface layer of at least one jaw 212 A and 212 B with a conductive-resistive matrix CM intermediate a first polarity electrode 220 having exposed surface portion 222 and second polarity electrode 225 having exposed surface portion 226 . Thus, the microcurrents within tissue during a brief interval of active heating can flow to and from said exposed surface portions 222 and 226 within the same engagement surface 255 A. By providing opposing polarity electrodes 220 and 225 in an engagement surface with an intermediate conductive-resistive matrix CM, it has been found that the dynamic “wave” of energy density (ohmic heating) can be created that proves to be a very effective means for creating a uniform temperature in a selected cross-section of tissue to thus provide very uniform protein denaturation and uniform cross-linking on thermal relaxation to create a strong weld. While the opposing polarity electrodes 220 and 225 and matrix CM can he carried in both engagement surfaces 255 A and 255 B, the method of the invention can be more clearly described using the exemplary jaws of FIG. 11 wherein the upper jaw's engagement surface 250 B is an insulator indicated at 252 .

[0117] More in particular, referring to FIG. 12 , the first (lower) jaw 212 A is shown in sectional view with a conductive-resistive matrix CM exposed in a central portion of engagement surface 255 A. A first polarity electrode 220 is located at one side of matrix CM with the second polarity electrode 225 exposed at the opposite side of the matrix CM. In the embodiment of FIG. 12 , the body or support structure 258 of the jaw comprises the electrodes 220 and 225 with the electrodes separated by insulated body portion 262 . Further, the exterior of the jaw body is covered by an insulator layer 261 . The matrix CM is otherwise in contact with the interior portions 262 and 264 of electrodes 220 and 225 , respectively.

[0118] The jaw assembly also can carry a plurality of alternating opposing polarity electrode portions 220 and 225 with intermediate conductive-resistive matrix portions CM in any longitudinal, diagonal or transverse arrangements as shown in FIGS. 13 A- 13 C. 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.

[0119] FIGS. 14 A- 14 C illustrate sequential views of the method of using of the engagement surface layer of FIG. 11 to practice the method of the invention as relating to the controlled application of energy to tissue. For clarity of explanation, FIGS. 14 A- 14 C depict exposed electrode surface portions 220 and 225 at laterally spaced apart locations with an intermediate resistive matrix CM that can create a “wave” or “front” of ohmic heating to sweep across the engaged tissue et. In FIG. 14 A, the upper jaw 212 B and engagement surface 250 B is shown in phantom view, and comprises an insulator 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.

[0120] FIG. 14A provides a graphic illustration of the matrix CM within engagement surface layer 250 A at time T ₁ —the time at which electrical potential of a first polarity (indicated at +) is applied to electrode 220 via an electrical lead from voltage source 180 and controller 182 . In FIGS. 14 A- 14 C, the spherical graphical 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 . The graphical elements 162 thus define a polarity at particular microsecond in time just after the initiation of power application. In FIG. 14 A, the body portion carrying electrode 225 defines a second electrical potential (−) and is coupled to voltage source 180 by an electrical lead. As can be seen in FIG. 14 A, the graphical elements 162 are indicated as having a transient positive (+) or negative (−) polarity in proximity to the electrical potential at the electrodes. When the graphical elements 162 have no indicated polarity (see FIGS. 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 of FIG. 6 and the graphical representation of FIG. 8B .

[0121] As can be seen in FIG. 14 A, the initiation of energy application at time T ₁ causes microcurrents me within the central portion of the conductive matrix CM as current attempts to flow between the opposing polarity electrodes 220 and 225 . The current flow within the matrix CM in turn localizes corresponding microcurrents mc′ in the adjacent engaged tissue et. Since the matrix CM is engineered to conduct electrical energy thereacross between opposing polarities at about the same rate as tissue, when both the matrix and tissue are at about 37° C., the matrix and tissue initially resemble each other, in an electrical sense. At the initiation of energy application at time T ₁ , the highest Rf energy density can be defined as an “interface” indicated graphically at plane P in FIG. 14 A, 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.

[0122] Still referring to FIG. 14 A, 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 (see FIG. 6 ) in that central region.

[0123] FIG. 14B graphically illustrates the interface or plane P at time T ₂ —an arbitrary microsecond or millisecond time interval later than time T ₁ . 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 nonconductive, the matrix CM will begin to apply substantial thermal effects to the tissue by means of passive conductive heating as described above.

[0124] FIG. 14C illustrates the propagation of planes P and P′ at time T ₃ —an additional arbitrary time interval later than T ₂ . 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 the engagement surface 255 A 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 in FIG. 14C . Thus, the propagation of waves of Rf energy density will repeat itself as depicted in FIGS. 14 A- 14 C which can effectively weld tissue.

[0125] 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-crosslink 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 220 and 225 and intermediate or surrounding matrix portions CM.

[0126] FIG. 15 depicts an enlarged view of the alternative Type “B” jaw 212 A of FIG. 13A wherein the engagement surface 250 A carries a plurality of exposed conductive matrix portions CM that are intermediate a plurality of opposing polarity electrode portions 220 and 225 . This lower jaw 212 A has a structural body that comprises the electrodes 220 and 225 and an insulator member 266 that provide the strength required by the jaw. An insulator layer 261 again is provided on outer surfaces of the jaw excepting the engagement surface 255 A. 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 exposed electrode portion 222 and 226 will function as described in FIGS. 14 A- 14 C.

[0127] The type of engagement surface 250 A 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 exposed outboard electrode surface 222 and 226 having any suitable dimension. Similarly, the engagement surface 250 A has resistive matrix portions that portions that define an exposed matrix width mw ranging between about 0.005″ and 0.20″.

[0128] In the embodiment of FIG. 15 , the electrode portions 220 and 225 are substantially rigid and extend into contact with the 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 the electrode portions 220 and 225 are floating within, or on, the surface layers of the matrix 250 A.

[0129] 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. The jaws 212 A and 212 B and engagement surface layers 255 A and 255 B both expose a substantial portion of matrix to the engaged tissue. The elastomeric character of the matrix can range between about 20 and 95 in the Shore A scale or above about 40 in the Shore D scale. Preferably, one or both engagement surface layers 255 A and 255 B can be “crowned” or convex to insure that the elastomeric matrices CM tend to compress the engaged tissue. The embodiment of FIG. 17 illustrates that a first 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 floating electrode 220 can range from about 1 micron to 200 microns. The second polarity electrode 225 has exposed portions 272 a and 272 b at outboard portions of the engagement planes 255 A and 255 B. In operation, the jaw structure of FIG. 17 creates controlled thermal effects in engaged tissue by several different means. First, as indicated in FIGS. 18 A- 18 C, the dynamic waves of Rf energy density are created between the opposing polarity electrode portions 220 and 225 and across the intermediate matrix CM exactly as described previously. Second, the electrically active components of the upper jaw's engagement surface layer 255 B cause microcurrents between the engagement surface layers 255 A and 255 B, as well as to the outboard exposed electrode surfaces exposed portions 272 a and 272 b , between any portions of the matrices that are below the selected switching range. Third, the substantial volume of matrix CM is each jaw provides substantial heat capacity to very rapidly cause passive heating of tissue after active tissue heating is reduced by increasing impedance in the engaged tissue et.

[0130] 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. The jaws 212 A and 212 B and engagement surface layers 255 A and 255 B again expose matrix portions to engaged tissue. The upper jaw's engagement surface layer 255 B is convex and has an elastomeric hardness ranging between about 20 and 80 in the Shore A scale and is fabricated as described previously.

[0131] Of particular interest, the embodiment of FIG. 19 depicts a first polarity electrode 220 that is carried in a central portion of engagement plane 255 A but the electrode does not float as in the embodiment of FIG. 17 . The electrode 220 is carried in a first matrix portion CM ₁ that is a substantially rigid silicone or can be a ceramic positive temperature coefficient material. Further, the first matrix portion CM ₁ preferably has a differently sloped temperature-resistance profile (cf. FIG. 6 ) that the second matrix portion CM ₂ that is located centrally in the jaw 212 A. The first matrix portion CM ₁ , whether silicone or ceramic, has a hardness above about 90 in the Shore A scale, whereas the second matrix portion CM ₂ 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 CM ₁ has a higher switching range than the second matrix portion CM ₂ . In operation, the wave of Rf density across the engaged tissue from electrode 220 to outboard exposed electrode portions 272 a and 272 b will be induced by matrix CM ₁ having a first higher temperature switching range, for example between about 70° C. to 80° C., as depicted in FIGS. 18 A- 18 C. The rigidity of the first matrix CM ₁ prevents flexing of the engagement plane 255 A. During use, passive heating will be conducted in an enhanced manner to tissue from electrode 220 and the underlying second matrix CM ₂ 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 the electrode 220 when used in small cross-section jaw assemblies (e.g., 5 mm.