STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

Not applicable.

BACKGROUND

The control of bleeding during surgery accounts for a major portion of the time involved in an operation. In particular, bleeding that occurs when tissue is incised obscures the surgeon's vision, delays the operation, and reduces the precision of cutting. In addition, blood loss from the patient during surgery must be minimized to reduce or eliminate the need for supplementary blood transfusions.

Conventional surgical procedures are carried out utilizing a sequence of surgical instruments or tools. At the outset of a given procedure, sharp mechanical devices, such as a scalpel, are employed to part the skin layers so as to provide external access to the body cavity. Bleeding during such initial stages may be controlled through the use of ties, clamps, blotting procedures, and the like. As the body cavity is accessed, tissue not only is cut but also is manipulated to the extent that blunt counterparts often supplant mechanically sharp instruments.

These blunt counterparts electrically perform cutting and blood coagulating functions on demand and require the passage of high frequency electrical current through the tissue being cut and/or coagulated. Such technology, known as monopolar electrosurgery, has been available to surgeons for decades. For instance, a physicist, William T. Bovie, first developed a monopolar electrosurgical device over sixty years ago. This early device is described, for example, in U.S. Pat. No. 1,813,902, issued Jul. 14, 1931, entitled “Electrosurgical Apparatus”, has met with acceptance over the years within the surgical community to the extent that current versions are referred to as the “Bovie”. Such devices typically consist of a handle having a first or “active” electrode extending from one end. The other end of the handle is electrically coupled, via a cable, to an electrosurgical generator that provides a high frequency electric current in either a continuous high-frequency alternative current cutting mode or a pulsed coagulating mode. A remote control switch is attached to the generator and commonly is present in the form of a switch on the handle and/or as a foot switch located in proximity to the operating theater. During an operation, a second or “return” electrode, having a much larger surface area than the active electrode, will be positioned in contact with the skin of the patient. To remove tissue, the surgeon brings the active electrode into proximity with the tissue to be cut or coagulated. This is a starting condition where the instrument has not touched tissue. At this point in time, an electrical switch is actuated whereupon the active electrode is brought into close proximity with the tissue to be cut or coagulated. Electrical current then arcs from the active electrode to the adjacent tissue and flows through deeper tissue to reach the larger return electrode. In a cutting mode, the electrical arcing and corresponding current flow results in a highly intense, but localized heating, which causes cell destruction and tissue severance. Following a short cutting routine, the instrument again is elevated in still air away from the tissue for two or three seconds. In general, the device can be switched to a pulsed, higher voltage input to perform in a coagulating mode wherein a shower of arcs impinge on the adjacent tissue to effect the sealing of smaller transected blood vessels. The cutting and coagulation effect depends on the formation of small arcs between the active electrode and the adjacent tissue requiring that a small air gap be maintained to support the formation of essential electrical arcs. Disadvantageously, the application of tamponade by pressing the surgical instrument against the tissue containing the transected vessel(s) to temporarily interrupt the flow of blood from vessels (e.g., vessels having a lumen size greater than about 1 mm) is not possible since such applied pressure would prevent the essential arc formation.

Another common modality for electrosurgery is referred to as bipolar electrosurgery. With this approach, no large return electrode is in contact with the patient. Instead, a bipolar electrosurgical instrument is made having first and second electrodes arranged in close mutual proximity. The device is utilized with a dedicated bipolar cable that is inserted in appropriate bipolar outlets of an electrosurgical generator, a device found essentially in all major health care facilities. When switch activated, the bipolar device provides an electrical current that is conducted through intervening tissue located between a first electrode and a second electrode. Tissue disposed between the electrodes is heated by the flow of electrical current and the intervening tissue is coagulated. However, the current intensity is generally insufficient to enable the cutting of tissue and no electrical arcs are formed as in the case of the monopolar electrosurgery modality described above. In general, surgeons are trained in the use of both bipolar and monopolar modalities; however, particularly in conjunction with endoscopic applications, bipolar devices are becoming more accepted in view of safety considerations. In the latter regard, the bipolar approach overcomes certain of the more undesirable characteristics of monopolar instruments in that excessive necrosis is reduced and current is not passed extensively through the body of the patient. Since current arcs between adjacent electrodes, blood vessels are readily cauterized. Bipolar devices, however, generally require other auxiliary means for cutting the tissue being coagulated.

Typically, the ubiquitous electrosurgical generators exhibit outputs with frequencies ranging from about 350 KHz to 1 MHz. Such higher radiofrequency frequencies serve to avoid tissue stimulation that would otherwise occur at lower frequencies.

Investigators also have considered the implementation of resistive heating to carry out coagulation and cutting in surgery. One of the early devices known as an electrocautery device employed a very fine wire formed as a loop or extending linearly between spaced mounting points. Formed, for example, of platinum, the thinness or small diameter of the heated wire was required in order to gain a high enough resistance to develop correspondingly high enough temperature levels in conjunction with practical current levels. The requisite thinness of the wire resulted in marginal strength or rigidity, thus restricting applications of such instruments to spot coagulation with the application of a minimum level of pressure to the targeted vessel.

Surgical blades have been developed with mechanically sharp edges and side mounted electrical heating elements. With these instruments, cutting is achieved at the mechanically sharp facet edge of the blade and the coagulation or hemostasis is intended to develop as a result of contact of the sides of the blade and blade facets with the cut tissue. This, unfortunately, represents an attempt to stop bleeding after cutting, as opposed to a more desirable procedure for simultaneous cutting with coagulation. Prior art devices have employed resistance feedback control to maintain the heater temperature at a substantially constant user selected temperature under conditions ranging from operation in air (requiring minimum power deliver to heating element) to direct contact with vascular tissue (requiring maximum power delivery to heating element). In such prior art devices, the user selects the desired blade operating temperature in the range from about 70 C to 300 C. In surgical use with the blade in contact with tissue, the temperature at the cutting edge and adjacent blade facets is lower than the user selected operating temperature due to thermal impedances in the pathway between the heating element and the cutting edge and adjacent blade facets. In addition, to minimize adherence of blood coagulum and tissue during surgical use, it is essential that the lateral surfaces and facets of the blade that contact tissue are coated with a non-stick coating, such as polytetrafluoroethylene. Such non-stick coating unavoidably introduces an additional thermal impedance between the heated cutting blade and the tissue owing to the very low thermal conductivity of available non-stick coating materials. In this regard, see U.S. Pat. Nos. 4,481,057, 4,485,810, 5,308,311, 8,142,425 and 8,475,444.

Unlike monopolar electrosurgical instruments that cut tissue with electrical arcs that form between a blunt edge of the monopolar electrode “blade” and the adjacent tissue being cut or coagulated, prior art resistively heated devices have also been described that employ blunt cutting portions (i.e., unsharpened edges) that provide both thermally induced incision as well as coagulation without the passage of electrical current through the tissue being cut and/or coagulated. Thermal incision or “cutting” of tissue is achieved by transferring a sufficient heat flux to the tissue via the conduction heat transfer mode to induce rapid vaporization of cellular water within the tissue. The rapid vaporization of the cellular water that comprises up to 70% of all tissue further induces fragmentation of the cells with the result of weakening the tissue structure and consequent breaking apart of tissue in a manner that has the effect of tissue “cutting”. Such prior art resistively heated devices utilize a defined thermal cutting portion emulating surgical blades and other implements, often with edge facets having an included angle ranging from about 20 to 40 degrees. The minimum temperature required to induce the thermal cutting of tissue has been previously determined to be about 400 to 500 C. Equally important, at tissue contacting blade surface temperatures above about 400 C, tissue and blood coagulum does not adhere thereby eliminating the need for the use of a non-stick coating and the need for removal of tissue and coagulum debris from the tissue contacting portions of the blade. In this regard, one of the first such devices for the thermal cutting of tissue with a resistively heated blunt blade is seen in Eggers' U.S. Pat. No. 5,591,719 issued Jun. 15, 1999, entitled “Resistively Heating Cutting and Coagulating Surgical Instrument”.

Some resistively heated devices achieve a thermally induced cutting effect wherein the heating method employs a self-regulating temperature characteristic. Self-regulating (also known as auto-regulating) resistively heated devices maintain the cutting surface of the surgical device within a pre-selected elevated temperature range. An approach for attaining self-regulation has been to employ a ferromagnetic material in constructing the tissue-contacting end (i.e., the heating element) of the surgical instrument. When radiofrequency current is passed through a ferromagnetic material, the current density is concentrated near its outer surface. This current density attenuates exponentially as the distance into the material from the surface increases, a phenomenon known as the “skin effect”.

The depth of the skin effect, i.e., the distance of penetrating current density into the ferromagnetic material, is defined as the depth at which current is reduced to approximately 37% of its surface value. This depth may be expressed mathematically as follows:
Skin Depth,d=C×√[ρ/(μ*f)]  (Equation 1)
where d is the skin depth measured in centimeters, ρ is the electrical resistivity in ohm-centimeters, μ is electrical relative magnetic permeability of the ferromagnetic material, C is a constant, viz. 503, and f is frequency of the applied alternating electrical potential in Hertz.

In ferromagnetic materials, such as iron, nickel, cobalt, and respective alloys, adjacent atoms and molecules couple their magnetic moments together in rigid parallelism (an interaction known as exchange coupling) in spite of the randomizing tendency of the thermal motion of atoms. If the temperature of such material is raised above a “Curie” temperature, specific for each ferromagnetic material composition, the noted exchange coupling suddenly disappears. As a result, these materials exhibit large changes in relative permeability as the temperature of the ferromagnetic material transitions to its Curie temperature. As seen from equation (1), since the relative permeability is known to change in response to the temperature of the material, the associated skin depth also will change. This relationship between skin depth and temperature enables ferromagnetic material based instruments to achieve temperature auto-regulation.

The heating elements of surgical devices have been constructed incorporating a ferromagnetic material that is selected to have a Curie temperature at or near the auto-regulation temperature desired for a particular surgical application. As a radiofrequency current passes through the ferromagnetic material, the heating element will resistively heat to approximately the Curie temperature. Once the cutting edge contacts tissue, both it and the area surrounding it will cool to a level below the Curie temperature. In response to this Curie transition, the skin depth will decrease which, in turn, results in an increased resistance of the cooled region (the resistance being a function of the ferromagnetic material's electrical resistivity multiplied by the current flow path length and divided by the current flow path area). A corresponding increase in the level of power supplied will accompany this increase in resistance. The temperature will then tend to again increase due to the increased level of resistive heating with the heating element increasing toward the Curie temperature. Thus, auto-regulation of the surgical component around the Curie temperature is achieved. See, for example, Eggers U.S. Pat. No. 5,480,398, issued Jan. 2, 1996, entitled “Endoscopic Instrument with Disposable Auto-Regulating Heater”; and Eggers, et al., U.S. Pat. No. 5,480,397, issued Jan. 2, 1996, entitled “Surgical Instrument with Auto-Regulating Heater and Method of Using Same”.

Other radiofrequency powered, resistively heated surgical devices achieve a thermally induced cutting effect wherein the heating element comprises a thin coating of ferromagnetic material deposited on a non-ferromagnetic metal substrate, such as, a round or flat hairpin-shaped loop of metal. The thickness of the ferromagnetic coating is on the order of the several skin depths, as specified above in Equation 1. In this ferromagnetic heating element design, the ferromagnetic material and operating frequency is selected such that, in those distal portions of the hairpin loop coated with the ferromagnetic material, the electrical current predominantly flows within the electrically resistive ferromagnetic coating and only minimally within the non-ferromagnetic blade substrate that exhibits a relative low electrical impedance with respect to current flow. In those proximal regions of the hairpin loop not coated with a ferromagnetic material, the electrical current flows though the non-ferromagnetic support portions of the hairpin loop. The Curie temperature of the selected ferromagnetic material coating serves to establish a maximum upper temperature limit during surgical use and the selected level of constant current determines the operating temperature range of the tissue-contacting blade (e.g., ferromagnetic material coated Beryllium-Copper hairpin-type blade) as a function of the tissue effect being attempted (e.g., tissue cutting, tissue ablation, tissue desiccation). In this regard, see U.S. Pat. Nos. 8,292,879, 8,419,724, 9,549,774, and U.S. Publication Number US 2015/0327907.

A disadvantage associated with prior art resistively heated thermal cutting devices that employ a ferromagnetic blade substrate or ferromagnetic coatings on a non-ferromagnetic hairpin-shaped round or flat wire substrate, is concerned with a lack of sufficient localization of heat at the blunt thermal cutting edge. In this regard, the entire heating element, including the support for its thermal cutting edge, is heated to a temperature of at least 400 to 500 C required to effect thermal cutting of tissue. In addition, the temperature of the ferromagnetic heating element can increase to as high as the Curie temperature of the ferromagnetic heating element (e.g., 600 to 700 C). Such elevated temperatures poses a risk that the support portions of the heating element that are proximal to those portions intended for tissue contact and cutting may contact tissue or organs not selected for incision causing unwanted thermal injury to the tissue or organs. Additionally, the time period required for the ferromagnetic heating element disposed in the cutting region to cool down to safe levels posing no threat of thermal injury can be quite significant. This time period, for example, may be ten seconds or more, an interval, which in a surgical environment is considered excessive, several seconds or less being considered acceptable, a starting condition interval to which surgeons are accustomed.

Another disadvantage associated with prior art resistively heated thermal cutting devices that employ a ferromagnetic blade substrate or a ferromagnetic coating on a non-ferromagnetic substrate is the operating temperature of the blade can extend over a wide range from the minimum allowed temperature for thermal cutting of tissue (viz., 400 to 500 C) to the Curie temperature of the ferromagnetic material, which can be as high as 600 to 700 C. In practice, the skin depth of a ferromagnetic material actually increases only gradually as the ferromagnetic heating element approaches the Curie temperature. As a consequence, maintaining the temperature of a ferromagnetic heating element above the minimum temperature of 500 C to ensure effective thermal cutting under all heat dissipation conditions in tissue requires a Curie temperature that is at least about 100 C greater than the minimum desired temperature. Conditions during surgery that can significantly affect the rate of heat dissipation include the vascularity of tissue, the rate of advancement of cutting blade through tissue and the length of the blade in contact with tissue. This wide operating temperature range of the ferromagnetic heating element above the minimum required thermal cutting temperature of 500 C produces smoke that obscures visibility and necessitates the use of auxiliary smoke evacuation apparatus and methods, further complicating the surgical procedure.

Another disadvantage associated with prior art ferromagnetic resistively heated devices for thermal cutting of tissue is concerned with the requirement that they must be powered by a specially designed or dedicated radiofrequency power supply operating at frequencies ranging from about 300 kHz to over 20 MHz. These dedicated power supply systems generally are configured to be unique to the properties of a particular heating element and are not of a universal nature, such that they would be usable with different surgical implements. In order to maximize the auto-regulation effect, the energy source used to apply power to the heating element preferably operates at a substantially constant current.

Yet another disadvantage of some prior art ferromagnetic resistively heated devices for thermal cutting of tissue is concerned with their inability to apply tamponade with the heated blade member to [a] interrupt the flow of blood long enough to allow the heated blade to effect the sealing of the transected blood vessels or [b] apply heat to pre-seal larger blood vessels before they are transected.

As a consequence of the foregoing considerations, practitioners have found it necessary to provide device-dedicated radiofrequency energy sources for powering thermal cutting surgical devices as well as auxiliary smoke evacuation systems. Of course, such added equipment requirements pose budgetary concerns to health care institutions.

BRIEF SUMMARY

The present invention is addressed to surgical instruments employing the thermal cutting of tissue. As used herein, the thermal cutting of tissue refers to the severing of tissue by means of heating the blade and the tissue contacting surface of a surgical instrument to a sufficiently high temperature of at least 400 to 500 C to thermally weaken the tissue structure to the point that division of the tissue can be accomplished with an otherwise blunt, unsharpened edge. The mechanism of “cutting” employed in the thermal cutting of tissue is distinct from that of a conventional scalpel wherein a surgically sharp cutting edge is employed to mechanically sever and divide tissue without the need for elevating the temperature of the cutting edge.

In contrast to radiofrequency-powered devices for the thermal cutting of tissue incorporating heating elements based on the skin effect, an effect which concentrates electrical current flow in a confined depth of a ferromagnetic conductor or coating on a non-ferromagnetic conductor operating below its Curie temperature, the instruments now presented are heated by a resistive heating element disposed on a dielectric coating bonded to a blunt blade member. The level of heating power generated in heating elements based on ferromagnetic materials depends on both the temperature-dependent skin depth and the applied current level resulting in heating element temperatures that can vary over a broad range of 100 C or more. This broad range of temperature variation is dependent on the conditions during surgery that affect the rate of heat dissipation, the conditions including the vascularity of tissue, the rate of advancement of cutting blade through tissue and the length of the blade in contact with tissue.

Unlike thermal cutting surgical devices that incorporate ferromagnetic heating elements, the thermal cutting surgical instruments now presented utilize resistance-feedback control to regulate the temperature of the heating element to within about 5 C of the optimum temperature for the thermal cutting of tissue (e.g., an optimum preselected temperature of 500 C). The resistance-feedback control of the heating element, wherein the heating element is disposed on an electrically insulative dielectric layer interposed between the heating element and the blade substrate, enables a nearly constant heating element temperature to be maintained over the entire range of possible heat dissipation conditions. Such heat dissipation conditions range from operation of the heated blade in still air to exposure of the full length of the heated portion of the blade to highly vascular tissue during passage through the tissue at the maximum rate of advancement.

The thermal cutting surgical instruments of the present disclosure incorporate a blade comprising two distinct portions, a distal portion constituting a heated portion of the blade and a proximal portion constituting a support member portion of the blade. The heated portion of blade includes a substrate having relatively high thermal conductance in order to minimize thermal gradients across the length or the width of the heated portion of the blade. Such gradients can otherwise exist when the only a small portion of the entire heated length of the blade is in contact with tissue with the resulting heat dissipation that occurs at the blade/tissue interface limited to only a small fraction of the length of the heater.

By way of example, a first substrate supporting the resistance heater in the heated portion of the blade may be a high thermal conductivity material, such as, for example, silver, silver alloy, copper or a copper alloy. By way of another example, the substrate supporting the resistance heater in the heated portion of the blade may comprise a laminate structure produced by conventional roll bonding of dissimilar metals comprising a core of copper or silver surrounded on either side by stainless steel claddings on both sides of the copper or silver core (e.g., stainless steel Type 430 or 304). Such a three-layer laminate increases the stiffness of the heated portion of the blade by the buttressing effect of the stainless steel owing to its much higher modulus of elasticity than that of the high thermal conductivity copper or silver core material. The stainless steel claddings of equal thickness on either side of a copper or silver or silver core also serve to limit warpage due to any difference in the thermal expansion coefficient of the copper or silver core and stainless steel cladding materials. In addition, the stainless steel cladding disposed on the copper or silver core enables an improved level of chemical bonding between the electrically insulative dielectric layer and the stainless steel cladding present in the heated portion of the blade. Preferably, the high thermal conductivity metal selected for first substrate or core of first substrate is greater than 2 watt/cm-C, more preferably greater than 3 watt/cm-C.

The use of a known biocompatible material, such as silver, in the core of the three layer laminate structure is advantageous since a biocompatible coating is not required over the exposed edge portions of the first substrate involving a copper core. A biocompatible coating will otherwise be required since copper is known to be cytotoxic with respect to human cells and, therefore, can not be used in contact with human tissue or blood during surgical procedures. In this regard, see Cortizo, M., et al., Cytotoxicity of Copper Ions Released from Metal. Biological Trace Elements Research, December 2004; 102(1-3); 129-141. The biocompatible coatings may include materials such as, for example, titanium nitride, titanium aluminum nitride, chromium nitride, zirconium nitride, gold, silver and, if the maximum operating temperature is below 500 C, Parylene HT® (Parylene HT® being a registered trademark of Specialty Coating Systems, Inc., Indianapolis Ind.).

The support member portion of the blade incorporates a substrate having a relatively low thermal conductance in order to minimize heat conduction from the heated portion of the blade through the support member portion of the blade and into the handpiece that supports the blade. By way of example, a second substrate comprising the support member portion of the blade may be a low thermal conductivity material such as stainless steel Type 304 or 430. Preferably, the thermal conductivity of metal selected for support member portion of blade is less than 0.6 watt/cm-C, more preferably less than 0.3 watt/cm-C.

The first substrate exhibiting high thermal conductance in the heated portion of the blade having a first thickness is joined by a welding process to a second substrate exhibiting low thermal conductance in a support member portion of the blade having a second thickness. The first and second substrates may advantageously have equal first and second thicknesses so that a composite of the two substrates is planar and having a uniform thickness. The composite of the first and second substrates having a uniform thickness enables subsequent thick-film printing processes including the printing and firing of one or more electrically insulative dielectric layers on the upper surfaces of the first and second substrates including the weld zone located at the boundary between the first and second substrates.

Alternatively, a slot may be formed in a strip (e.g., by grinding) of stainless steel substrate material to receive an inlay strip of a high thermal conductivity material (e.g., silver or copper). The strip of stainless steel substrate with inlay strip of high thermal conductivity material is next roll bonded to achieve a metallurgical bond between the inlay strip of high thermal conductivity material and the stainless steel substrate, typically with an overall thickness reduction of at least 50%. Next, the length of the roll bonded composite stainless steel strip incorporating an inlay of high thermal conductivity material is cut in half and the two composite strips are positioned together such that the outer edges are aligned and the inlays of high thermal conductivity material on both strips are facing each other. These two composite strips then are roll bonded together such that the inlay of high thermal conductivity is located within outer layers of the stainless steel substrate. This alternative approach eliminates the need for welding a three-layer laminate strip comprising stainless steel and a high thermal conductivity material to strips of stainless steel.

The printing and firing of the electrically insulative dielectric layer is followed by the thick-film printing and firing of an electrically resistive heating element on the surface of the electrically insulative dielectric layer within the heated portion of the blade. The thick-film printable heating element material has a temperature coefficient of electrical resistance of at least 1000 and preferably more than 3000 parts-per-million/degree C. over the range from 20 C to 600 C and a sheet resistance at 20 C of about 30 to 200 milliohms/square. Two or more electrically conductive leads are thick-film printed and fired on the electrically insulative dielectric layer within the region of the support member of the blade and extend into the region of the heated portion of the blade and overlap at the termini of the electrically resistive heating element to provide electrical communication between the leads and the heating element. The thick-film printable electrically conductive lead material has a sheet resistance of less than about 5 milliohms/square. An electrically insulative dielectric overcoat layer is thick-film printed and fired over the entire electrically resistive heater element as well as all but the most proximal portions of the electrically conductive leads. A short length of the electrically conductive leads at the proximal end of the support member portion of the blade are not covered with the electrically insulative dielectric overcoat layer to permit engagement of the proximal end of the electrically conductive leads with electrical contacts within the distal end of the handpiece.

The thick-film pastes selected for the printing of the electrically insulative dielectric layer, electrically resistive heating element, electrically conductive leads and electrically insulative dielectric overcoat layer have a firing temperature in the range from about 600 C to 900 C to enable operation of the heating element of thermal cutting surgical instrument at a temperature of at least 400 C, preferably at least 500 C. The thickness of the thick-film printed and fired electrically insulative dielectric layer (e.g., less than about 0.001 inch) is [a] thin enough to minimize the temperature gradient across the electrically insulative dielectric layer during the application of electrical power to the heating element and [b] thick enough to prevent electrical breakdown of the dielectric layer during the application of the maximum voltage differential to the heating element.

The heat capacity of the heated portion of the blade is advantageously minimized to reduce the time required to raise the temperature of the heated portion of the blade from room temperature to the preselected blade operating temperature required for the thermal cutting of tissue (e.g., about 500 C). The heat capacity of the heated portion of the blade is minimized by reducing the thickness and width of the heated portion of the blade to levels consistent with maintaining the ability of the structure of the heated portion of the blade to withstand contact pressure with tissue during the application of tamponade in combination with the heated portion of the blade to seal transected blood vessels. However, the minimum width of the heated portion of the blade must also be sufficiently wide to provide a sufficient surface area at the heating element/dielectric layer interface to maintain the maximum heat flux between the heating element and the first substrate within an acceptable level, for example not greater than about 250 watts/sq. cm. Preferably the heat capacity of the heated portion of blade is less than about 0.025 calories/C.

In a preferred embodiment of the present invention, a direct current (DC) voltage is applied to the terminals of one or more heating element segments located in the heated portion of the blade. Resistance-feedback control circuitry is incorporated within handpiece to maintain the heating element in the heated portion of the blade within a narrow band of temperature that only varies about 5 C above or below the preselected thermal cutting and blood vessel sealing temperature (e.g., 500 C).

In addition to incorporating all heating element temperature control circuitry within the handpiece, the handpiece also includes a pressure sensitive switch that energizes the heating element and maintains the heating element at the preselected thermal cutting and blood vessel sealing temperature as long as the pressure sensitive switch is depressed. Also, the handpiece incorporates display lights to indicate operating states of heated portion of blade. A first display light corresponds to a first operating state indicating that the blade ready to be heated. A second display light corresponds to a second state indicating that the blade is at the temperature required for thermal cutting of tissue and/or sealing of blood vessels. A third display light corresponds to a third state corresponding to a failed heating element requiring the operator to replace the currently used thermal cutting blade with a new sterile blade by removing the defective blade from the handpiece and inserting a new blade in its place.

In addition to display lights to indicate the functional state of the thermal cutting surgical instrument, the handpiece also may incorporate a sound generating element that provides an audible cue to the operator that the heating element is at the preselected temperature to effect thermal cutting and coagulation of tissue. The handpiece of the preferred embodiment incorporates a flexible, two-lead cable that extends from the handpiece to a simple, low-cost DC power supply of the type commonly used with laptop computers and other portable electronic devices. The DC power supply is removably connectable to any electrical outlet (e.g., 115 volt, 60 cycle line power) available in any operating room or surgery setting and provides a constant DC voltage to the resistance-feedback control circuitry, switch function and display lights incorporated within the handpiece. The remote DC power supply may include a single on/off power switch.

The present disclosure includes methods of manufacturing an electrically heated thermal cutting surgical blade, comprising the sequence of steps of:a. providing a first substrate in the form of a laminate strip of first thickness having a core of a high thermal conductivity material having lateral faces surrounded by a pair of outer layers of metal having a high modulus of elasticity, having a coefficient of expansion that is similar to the electrically insulative dielectric layer and suitable for mechanical and/or chemical bonding to an electrically insulative dielectric layer;b. providing a second substrate strip of second thickness of a low thermal conductivity material where first thickness and second thickness are preferably substantially the same and having a coefficient of expansion that is similar to the electrically insulative dielectric layer and suitable for mechanical and/or chemical bonding to an electrically insulative dielectric layer;c. metallurgically joining, as by welding, the laminate strip of first substrate to the strip of second substrate along a longitudinal joint to form a composite strip incorporating both first and second substrates;d. providing the composite strip with a smooth surface at the longitudinal joint;e. perforating the composite strip to define at least one blade blank having a first region extracted or cut from the first substrate that functions as the heated portion of the blade and a second region extracted or cut from the second substrate that functions as the support member portion of the blade;f. depositing a thin layer of an electrically insulative dielectric material on one side of the blade blank in both the first region and second region, the electrically insulative dielectric material having a thermal expansion coefficient in the range of the thermal expansion coefficient of the outer layers of the first region and the material of the second region;g. depositing a thin layer of an electrically resistive material in a first pattern on the layer of electrically insulative dielectric material in the first region of the blade blank that functions as the heating element;h. depositing a thin layer of an electrically conductive material in a second pattern on the layer of the electrically insulative dielectric material in the second region of the blade blank that functions as the electrical leads, the second pattern overlapping the first pattern of electrically resistive material dielectric so that the terminals of the heating element comprising electrically resistive material and distal terminals of electrical leads comprising electrically conductive material are in electrical communication;i. depositing a thin electrically insulative dielectric overcoat layer over the first and second patterns of electrically resistive and electrically conductive materials except for a portion of the second pattern that is left exposed to accommodate completing an electrical contact between the electrically conductive leads and the electrical terminals of the handle andj. optionally depositing a thin coating of biocompatible material over the entire surface of the first and second region of the blade up to but distal to the portions of the electrically conductive leads not covered by the electrically insulative dielectric overcoat layer, the coating of biocompatible material not required in the event all materials of construction of the tissue contacting portions of the blade are biocompatible.

An advantage of the present invention is an improved surgical blade for the thermal cutting of tissue without the need for a surgically sharp cutting edge, and methods of making such blades. Another advantage is a durable blade having improved thermal delivery capabilities over a broad range of conditions encountered in the thermal cutting of tissue. A yet another advantage is a thermally regulated heated blade having a cutting region that is maintained at a nearly uniform temperature over the full range of surgical cutting conditions. A further advantage is a thermal cutting blade that permits sustained operation without adherence of tissue or coagulum during surgical use and without the need for a non-stick coating. A yet further advantage is an electrically heated blade that reduces conduction of heat from the heated portion of the surgical blade to the blade supporting handpiece. Another advantage is an electrically heated surgical blade having thick-film printed electrical leads characterized by a low electrical resistance to reduce Joulean heating within a blade support member by resistive losses in the electrical leads. A further advantage is a surgical instrument for the thermal cutting of tissue whose heated portion of the blade has a heat capacity that is sufficiently low to enable heat up of the blade from room temperature to a thermal cutting temperature within about one second. Another advantage is the elimination the of a costly control system external to the handpiece. A further advantage is a method of manufacturing a surgical blade to maximize temperature uniformity of the heated portion of the blade and minimize thermal conduction from the heating element and tissue cutting regions of the blade to the handpiece supporting the blade. These and other advantages will be apparent to the skilled artisan based on the present disclosure.

