RADIOFREQUENCY ABLATION WITH PHASE SENSITIVE POWER DETECTION

Abstract: Systems for ablating tissue control radiofrequency power to an ablation electrode (16) by relying upon actual phase sensitive power measurements (72), unaffected by phase shifts between radiofrequency voltage and current. The systems also detect these phase differences (70), if they develop and integrate this factor (30) in making their control decisions.

Description: RADIOFREQUENCY ABLATION WITH PHASE SENSITIVE POWER DETECTION Field of the Invention The invention generally relates to catheters and associated power sources. In a more specific sense, the invention relates to ablation catheters that, once steered and manipulated within interior regions of the body, transmit energy to form lesions for therapeutic purposes. Background of the Invention Physicians make use of catheters today in medical procedures to gain access into interior regions of the body to ablate targeted tissue areas. It is important for the physician to control carefully and precisely the emission of energy within the body used to ablate the tissue. The need for careful and precise control over the catheter is especially critical during procedures that ablate tissue within the heart. These procedures, called electrophysiological therapy, are becoming more widespread for treating cardiac rhythm disturbances. During these procedures, a physician steers a catheter through a main vein or artery (which is typically the femoral artery) into the interior region of the heart that is to be treated. The physician then further manipulates a steering mechanism to place the electrode carried on the distal tip of the cathe- ter into direct contact with the tissue that is to be ablated. The physician directs radio frequency energy from the electrode tip through tissue to an indiffer¬ ent electrode to ablate the tissue and form a lesion. Cardiac ablation especially requires the ability to precisely monitor and control the emission of energy from the ablation electrode. Conventional systems control radiofrequency power to the ablation electrode based upon root mean squared derivations of the rapidly fluctuating radiofrequency voltage and current. These root mean squared derivations represent actual power only if the radiofrequency voltage is in phase with the radiofrequency current. When a phase shift develops , the product of the root mean squared voltage and cur- rent differ from the actual power by a factor of the cosine of the phase angle. They cease having direct relevance to the control decisions that seek to main¬ tain desired power levels at the ablation electrode. In fact, once large phase shifts develop, root mean squared derivations can lead to erroneous control decisions, gmnina-py of the Invention The invention provides an improved system for ablating tissue using radiofrequency energy. The system controls the delivery of radiofrequency energy to the ablation electrode based upon measurements of the actual phase sensitive power being delivery. The system also detects when a phase shift occurs between radiofrequency voltage and radiofrequency current and takes appropriate action. In one embodiment, the system includes a source of radiofrequency energy. An electrode assem¬ bly emits radiofrequency energy delivered by the source at the ablation site. The system measures t- 5 radiofrequency current delivered to the electrode as¬ sembly and generates a measured radiofrequency current signal. The system also measures radiofrequency volt¬ age at the electrode assembly and generates a measured radiofrequency voltage signal. 10 According to one aspect of the invention, the system includes a power monitor that multiplies the measured current signal by the measured voltage signal. The monitor thereby derives a phase sensitive power signal that represents the actual phase sensi- 15 tive power transmitted to the electrode means. The system includes a controller that performs a function based upon the actual power signal. In one arrangement, the controller displays the value of the phase sensitive power signal in a 20 user readable format. In another arrangement, the controller in¬ cludes a feedback loop that controls the voltage out¬ put of the source based upon the phase sensitive power signal. 25 According to another aspect of the inven¬ tion, the system includes a radiofrequency phase shift detector. The detector derives root mean squared cur¬ rent and voltage signal based upon the measured radiofrequency current and voltage signals. The de- 30 tector multiplies the root mean squared current signal * by the root mean squared voltage signal, thereby de¬ riving a root mean squared power signal. This signal represents the apparent power transmitted the elec¬ trode means. 35 The detector compares the root mean squared power signal to the phase sensitive power signal gen¬ erated by the power monitor. The detector generates a deviation signal that indicates the magnitude a phase shift between the measured radiofrequency voltage sig- nal and the measured radiofrequency current signal. The detector performs a function based upon the phase shift deviation signal. In one arrangement, the phase detector in¬ terrupts the conduction of radiofrequency energy when the phase shift deviation signal exceeds a predetermined value. In a preferred arrangement, the source power feedback loop controls the voltage output of the source based upon the phase sensitive power signal when the phase shift deviation signal remains within an prescribed range. If the phase shift deviation signal exceeds this range, the feedback loop reduces the output voltage of the source. If the phase shift deviation signal exceeds a prescribed maximum amount, the feedback loops interrupts power to the electrode assembly. According to the invention, the system con¬ trols radiofrequency power to the ablation electrode relying upon actual phase sensitive power measure- ments. It also detects phase differences between the radiofrequency voltage and current and integrates this factor in making its control decisions. Brief Description of the Drawings Fig. 1 is a perspective view of a system for ablating tissue that embodies the features of the in¬ vention; Fig. 2 is a schematic view of the generator and associated monitor and control circuits for the system; Fig. 3 is a schematic view of the power mon- itor and control circuit for the system; Fig. 4 is a schematic view of the tissue impedance monitor and control circuit for the system; Figs. 5A and 5B is a schematic view of the tissue temperature monitor and control circuit for the system; Figs. 6A to C are views of an electrode with thermally insulated temperature sensing element that can be used in associated with the system to measure tissue temperature; Figs. 7A to C are views of an electrode with multiple thermally insulated temperature sensing ele¬ ments that can be used in associated with the system to measure tissue temperature; and Figs. 8A to C are views of an electrode spe¬ cially shaped to use in heart valve regions and having multiple thermally insulated temperature sensing ele¬ ments that can be used in association with the system to measure tissue temperature. Description of the Pre erred Embodiments Fig. 1 shows a system 10 for performing ab¬ lation on human tissue that embodies the features of the invention. The system 10 includes a radiofrequency generator 12 that delivers radiofre- quency energy. The system 10 also includes a steerable catheter 14 carrying a radiofrequency emitting tip electrode 16. In the illustrated embodiment, the system 10 operates in a monopolar mode. In this arrangement, the system 10 includes a skin patch electrode that serves as an indifferent second electrode 18. In use, the indifferent electrode 18 attaches to the patient's back or other exterior skin area. Alternatively, the system 10 can be operated in a bipolar mode. In this mode, the catheter 14 car- ries both electrodes. In the illustrated embodiment, the ablation electrode 16 and indifferent electrodes 18 are made of platinum. The system 10 can be used in many different environments. This specification describes the system 10 when used to provide cardiac ablation therapy. When used for this purpose, a physician steers the catheter 14 through a main vein or artery (typically the femoral artery) into the interior re¬ gion of the heart that is to be treated. The physician then further manipulates the catheter 14 to place the tip electrode 16 into contact with the tis¬ sue within the heart that is targeted for ablation. The user directs radio frequency energy from the generator 12 into the tip electrode 16 to form a le¬ sion on the contacted tissue. In the embodiment shown in Fig.l, the cathe¬ ter 14 includes a handle 20, a guide tube 22, and a tip 24, which carries the tip electrode 16 (which also will be called the ablation electrode) . The handle 20 encloses a steering mechanism 26 for the catheter tip 24. A cable 28 extending from the rear of the handle 20 has plugs (not shown) . The plugs connect the cath- eter 14 to the generator 12 for conveying radiofrequency energy to the ablation electrode 16. The radiofrequency heats the tissue to form the lesion. Left and right steering wires (not shown) extend through the guide tube 22 to interconnect the steering mechanism 26 to the left and right sides of the tip 24. Rotating the steering mechanism 26 to the left pulls on the left steering wire, causing the tip 24 to bend to the left. Also, rotating the steering mechanism 26 to the right pulls on the right steering wire, causing the tip 24 to bend to the right. In this way, the physician steers the ablation electrode 16 into contact with the tissue to be ablated. The generator 12 includes a radiofrequency power source 30 connected through a main isolation transformer 32 to first and second conducting lines 34 and 36. In the illustrated environment, the power source 30 delivers up to 50 watts of power at a fre- quency of 500 kHz. The first conducting line 34 leads to the ablation electrode 16. The second conducting line 36 leads to the indifferent patch electrode 18. Monitoring Actual and Apparent Radiofrequency Power As Figs. 2 and 3 show, the system 10 includes first monitoring means 38 for measuring the radiofrequency current and radiofrequency voltage de¬ livered by the generator 12 to the patient. The first monitoring means 38 also derives control signals in- dicative of RMS (root mean squared) voltage (in volts) , RMS current (in amps) , and actual phase sensi¬ tive power (in watts) to support other control func¬ tions of the generator 12. The first monitoring means 38 may be variously configured and constructed. In the il¬ lustrated embodiment, the first monitoring means 38 includes current monitoring means 40 for measuring the radiofrequency current passing from the first line 34 through the tissue to the second line 36 (i.e., from the ablation electrode 16 to the indifferent patch electrode 18) . The first monitoring means 38 also includes voltage monitoring means 42. The voltage monitoring means 42 measures the radiofrequency voltage generated between the first and second conducting lines 34 and 36 (i.e., between the ablation electrode 16 and the indifferent patch electrode 18) . The first monitoring means 38 includes three control outputs 44, 46, and 48. The first control output 44-carries a signal representative of RMS current conducted by the ablation electrode 16. The second control output 46 carries a sig¬ nal representative of the RMS voltage between the ablation electrode 16 and the indifferent patch electrode 18. The third control output 48 carries a signal representative of actual phase sensitive power transmitted by the ablation electrode 16. In the illustrated embodiment (as Figs. 2 and 3 show) , the current monitoring means 40 includes an isolated current sensing transformer 50 connected in the second conducting line 36. In this arran¬ gement, the current sensing transformer 50 directly measures the radiofrequency current passing through the ablation electrode 16 to the indifferent patch electrode 18. The measured value is a radiofrequency sig¬ nal varying at the selected rate, which in the il- lustrated embodiment is 500 kHz. The current sensing transformer 50 is con¬ nected to the first control output 44, which derives RMS current. The first control output 44 includes an integrated circuit RMS converter 52 to do this function. The RMS current converter first squares the radiofrequency current input signal from the current sensing transformer 50, and then averages the squared signal over a user prescribed period (which in the illustrated embodiment is about once every 0.01 sec- ond) . The RMS current converter 52 then takes the square root of the average squared value. The result¬ ing output represents RMS current. The RMS current signal takes the form of a relatively slowly varying signal, compared with the rapidly varying radiofrequency current input signal. As Figs. 2 and 3 show, the voltage monitoring means 42 includes an isolated voltage sens¬ ing transformer 54 that is connected between the first and second conducting lines. In this arrangement, the voltage sensing transformer 54 directly measures the radiofrequency voltage across the body tissue between the ablation electrode 16 and the indifferent patch electrode 18. Like the value measured by the current sens- ing transformer 50, the measured voltage value is a radiofrequency signal varying at the selected 500 kHz rate. The voltage sensing transformer 54 is con¬ nected to the second control output 46, which derives RMS voltage. The second control output 46 includes an integrated circuit RMS converter 56 to do this function. The RMS voltage converter 56 squares the radiofrequency voltage input signal and then averages it over the same user prescribed period used by the current converter 52. The RMS voltage converter 56 then takes the square root of the average squared voltage value. The resulting RMS voltage signal (like the RMS current signal) takes the form of a relatively slowly varying signal. The voltage sensing transformer 54 is also connected to the third control output 48, which derives actual phase sensitive power. The third con¬ trol output 48 includes an analog multiplier in- tegrated circuit 58 to do this function. The multi- plier circuit 58 receives as one input the radiofre¬ quency input current signal directly from the current sensing transformer 50. The multiplier circuit 58 also receives as a second input the radiofrequency input voltage signal directly from the voltage sensing transformer 54. The output of the multiplier circuit 58 is the product of these two inputs, which represents the actual radiofrequency power transmitted by the abla- tion electrode 16. The power value is (like its component cur¬ rent and voltage inputs) a radiofrequency signal vary¬ ing at a relatively high radiofrequency rate. The third control output 48 also includes a low pass filter 60. In the illustrated embodiment, which operates with a radiofrequency rate of 500 kHz, the cut off frequency of the filter 60 selected is about 100 Hz. The rapidly varying measured input pow¬ er value, is low pass filtered by the filter 60 into a relatively slowly varying signal. This signal represents the actual phase sen¬ sitive power signal of the radiofrequency energy that the ablation electrode 16 delivers to the targeted tissue. The first, second, and third control outputs 44, 46, and 48 each includes appropriate inline scal¬ ing circuits 62. The scaling circuits 62 scale the RMS current signal, the RMS voltage signal, and the actual phase sensitive power signal to a specified voltage range that can be usable by the remainder of generator 12 circuitry. In the illustrated em¬ bodiment, the scaled range is 0.0 to 5.0 volts. The first monitoring means 38 also includes an analog to digital converter 64. The converter 64 digitizes a selected one or more of the analog RMS current output signal, RMS voltage output signal, and the actual phase sensitive power signal. The digital output(s) of the converter 64 can be used to display measurement results. In the illustrated embodiment, the system 10 includes a first digital display 66 on the generator 12 to show the user the actual phase sensitive power signal. The digital output(s) of the converter 64 also can be used to control operation of the generator 12. In the illustrated embodiment, the system 10 uses the digitized outputs in a feedback loop that main¬ tains radiofrequency output voltage within a desired range or at a constant value to control radiofrequency power at the ablation electrode 16. By controlling the power delivered by the generator 12, the physician can reproducibly form lesions of the desired depth during an ablation procedure. In this arrangement, the system 10 includes an input 68 for the user to enter an operating value desired for the actual phase sensitive power for the generator 12. The system 10 includes power control means 70 that includes comparator 71 to compare de¬ sired power with actual phase sensitive power. The output of the comparator varies the output voltage of radiofrequency power source 30 to maintain minimum er¬ ror between the measured actual power and the set point power. In the illustrated embodiment, the power control means 70 also monitors phase differences bet- ween radiofrequency voltage and current. The power control means 70 does this function by computing ap¬ parent power and by comparing the computed apparent power to the actual phase sensitive power. If the radiofrequency voltage and current signals are exactly in phase, the apparent power and actual phase sensi- tive power will be the same. However, if there is a phase difference, actual phase sensitive power will differ from the apparent power by a factor that repre¬ sents the cosine of the phase angle. In the illustrated embodiment, the power control means 70 includes a multiplier circuit 72 that obtains the product of the RMS current and RMS volt¬ age. The resulting output of the multiplier circuit 72 forms the apparent (i.e., not phase sensitive) pow- er of the system 10. The power control means 70 in¬ cludes a comparator 74 to compare the derived apparent power with the actual phase sensitive power. The mag¬ nitude of the output of the comparator 74 quantifies the amount of the phase shift. If the output of the phase shift comparator 74 exceeds a preselected amount, the power control means 70 generates a warning signal to show that a phase shift between the radiofrequency voltage and current has occurred. The system 10 may include a flashing light and audible alarm (not shown) to warn the user. The power control means 70 operates to main¬ tain a constant set power when the output of the phase shift comparator 74 remains within an allowable range above the threshold amount. The power control means 70 operates to reduce the output voltage of the source 30 when the output of the phase shift comparator 74 increases beyond this range. If the output of the phase shift comparator 74 shows a phase shift beyond a maximum threshold value, the power control means 70 generates a signal to shut off all power to the abla¬ tion electrode 16. Monitorinσ Tissue Impedance As Fig. 4 shows, the system 10 further in- eludes second monitoring means 76 for deriving the im- pedance of the tissue undergoing ablation. The second monitoring means 76 derives tissue impedance not only in absolute terms, but it also serves to record chang¬ es in tissue impedance over time. The second monitoring means 76 generates appropriate control signals based upon observed abso¬ lute values of tissue impedance as well as sensed changes according to preprogrammed criteria. The second monitoring means 76 may be variously configured and constructed. In the il¬ lustrated embodiment, the second monitoring means 76 includes a microprocessor 78. The microprocessor 78 samples the digitized outputs of the analog-to-digital converter 64 at prescribed intervals (for example, every 20 milliseconds, which represents a 50 Hz sam¬ pling rate) . The microprocessor 78 also divides the sampled digitized RMS voltage signal by the sampled digitized RMS current signal. The digital result is the tissue impedance (in ohms) for the sample. Preferably, the system 10 includes a second display 80 on the generator 12 that shows the user the sampled tissue impedance. The microprocessor 78 also maintains a re- cord of sampled tissue impedances over time. From this record, the microprocessor 78 calculates the changes in tissue impedance during a selected interval and generates appropriate control signals based upon predetermined criteria. The predetermined criteria under which the microprocessor 78 generates the control signals based upon tissue impedance can vary. Preferably, the tis¬ sue impedance control signals are used to augment the monitoring and control functions of the power control means 70 just described. In the illustrated embodiment, if measured tissue impedance falls outside a predetermined set range, the microprocessor 78 generates a command sig¬ nal to shut off power to the ablation electrode 16, whatever the actual phase sensitive power level sensed. The set range for tissue impedance for a car¬ diac ablation procedure is believed to be about 50 to 300 ohms. When tissue impedance begins in the set range and, over time, increases beyond it, the most likely cause is the coagulation of blood on the abla¬ tion electrode 16. A sudden rise in tissue impedance over the set range suggests the sudden onset of coag¬ ulation or a sudden shift in the position of the abla- tion electrode 16. Rapid fluctuations of the tissue impedance also could suggest poor contact between the ablation electrode 16 and the targeted tissue. All require prompt response; for example, withdrawal and cleaning of the ablation electrode 16, or reposition- ing of the ablation electrode 16. The system 10 preferably includes flashing lights and an audible alarm (not shown) to transmit a warning to the user when these conditions occur. A very high tissue impedance value could suggest poor skin contact with the indifferent elec¬ trode 18, or an electrical problem in the system 10. Again, this calls for prompt corrective action. If tissue impedance remains within the set range, but rises beyond a prescribed amount within the range, the second monitoring means 76 generates a con¬ trol signal that reduces but does not interrupt, the power output. In this arrangement, a relatively tight range of tissue impedance can be established (for ex¬ ample, 80 to 150 ohms) to maintain relatively constant power within this range. Monitoring Tissue Temperature As Fig. 5 shows, the system 10 includes third monitoring means 82 for sensing the temperature of the tissue in contact with the ablation electrode 16. The third monitoring means 82 includes temperature sensing means 84 carried in the ablation electrode 16. The system 10 also includes control means 86 for the generator 12 that is responsive to the sensed tissue temperature for performing generator control functions. Thermal insulation means 88 thermally isolates the temperature sensing means 84 from the thermal mass of the ablation electrode 16. Thus, the temperature sensing means 84 does not add to or serve as a part of the thermal mass of the ablation elec¬ trode 16. It serves to show the true temperature of the tissue with which it is in contact, without adding to the thermal mass of the ablation electrode 16 and without being influenced by the temperature of the surrounding thermal mass of the ablation electrode 16. In the embodiment illustrated in Figs. 6A to C, the ablation electrode 16 includes an interior well 90 at its tip end 92. The temperature sensing means 84 occupies this well. In this arrangement, the thermal insulating means 88 isolates the temperature sensing means 84 from the interior surface of the well 90 and the rest of the thermal mass of the ablation electrode 16. In Figs. 6A to C, the temperature sensing means 84 includes a small bead thermistor 94 with two associated lead wires 96 and 98. The temperature sen¬ sing tip of the thermistor 94 is exposed at the tip end 92 of the ablation electrode 16 for tissue con¬ tact. In the illustrated embodiment (see Figs. 6A to C) , the third monitoring means 82 includes means 132 for calibrating the thermistor 94 to account for deviations in nominal resistance among different thermistors 94. During manufacture of the catheter 10, the resistance of thermistor 94 is measured at a known temperature; for example, 75 degrees C. A cali¬ bration resistor 134 equal to the measured value is incorporated into the catheter handle 20. The leads of the calibration resistor 134 are connected to the third monitoring means 82. The thermistor 94 of the type shown is com¬ mercially available from the Fenwal Co. (Ma¬ ssachusetts) under the trade designation 111-202CAK- BD1. The lead wires 96 and 98 comprise #36 AWG signal wire Cu+ clad steel (heavy insulation) . Potting compound 100 encapsulates the thermistor 94 and lead wires 96 and 98 within the electrode well. Insulating sheathes 102 also shields the encapsulated lead wires 96 and 98. Together, the compound 100 and sheathes 102 electrically insulate the thermistor 94 from the surrounding ablation elec¬ trode 16. The potting compound 100 and insulation sheathes 102 can be made with various materials. In the illustrated embodiment, loctite adhesive serves as the potting compound 100, although cyanoacrylate adhe¬ sive or RTV adhesive and the like could be used. The sheathes 102 are made from polymide material, although other conventional electrical insulating materials also can be used. In the illustrated embodiment, the thermal insulating means 88 comprises a tube 104 that en¬ velopes the encapsulated thermistor 94 and lead wires 96 and 98. The thermal insulation tube 104 is itself adhesive bonded to the interior wall of the well 90. The thermal insulating material of the tube 104 can vary. In the illustrated embodiment, it is a polymide material having a wall thickness of about .003 inch. Other thermal insulating materials like mylar or kapton could be used. The lead wires 96 and 98 for the thermistor 94 extend from the electrode well 90 through the guide tube 22 and into the catheter handle 20. There, the lead wires 96 and 98 electrically couple to the cable 28 extending from the handle 20. A cable plug (not shown) connects to the generator 12 and transmits the temperature signal from the thermistor 94 to the third monitoring means 82. Figs. 7A to C show an alternate embodiment of an ablation electrode 16 having an array of temperature sensing means 84. At least one temperature sensing means 84, and preferably all of them, are thermally isolated from the ablation electrode 16 in the manner shown in Figs 6A to C. As Figs. 7A to C show, the ablation electrode 16 in the multiple array version includes an interior core well 106. Five branch wells 108A to E extend from the core well 106. The branch wells 108A to E open at the surface of the ablation electrode 16. One branch well 108A opens at the tip of the ablation electrode 16, like the single temperature sensing means 84 shown in Figs. 6A to C. The other four branch wells 100B to E extend at an angle from the core well 106 at arcuate intervals of 45 degrees. The four branch wells 108B to E open at the side of ablation electrode 16 and encircle the tip well branch 100A. A temperature sensing means 84 occupies each branch well 108A to E. In the illustrated and pre- ferred embodiment, a thermal insulating means 88 isolates each temperature sensing means 84 from the interior surface of the associated branch well 108A to E and the remaining thermal mass of the ablation elec¬ trode 16. As in the embodiment shown in Figs. 6A to C, each thermal sensing means 84 includes a small bead thermistor 94 with two associated lead wires 96 and 98. The temperature sensing tips of the thermistors 94 are exposed at the tip of the ablation electrode 16 for multiple contact with the tissue. The associated lead wires 96 and 98 are bundled within the center core well 106 and pass through the guide tube 22 to the handle 20. As in the embodiment shown in Figs. 6 to C, potting compound 100 encapsulates each thermistor 94 and its lead wires 96 and 98 within the associated branch well. Insulating sheathes 102 also shield the encapsulated lead wires 96 and 98. Together, the com¬ pound 100 and sheathes 102 electrically insulate each thermistor 94 from the surrounding ablation electrode 16. As in the embodiment shown in Figs. 6A to C, a thermal insulating tube 104 envelopes each electrically encapsulated thermistor 94 and its lead wires 96 and 98. And, as in the Figs 6A to C em¬ bodiment, adhesive bonds each thermal insulation tube 104 to the interior wall of each branch well 108A to E. Figs. 8A to C show another alternate embodi- ment of an ablation electrode 16 having multiple temperature sensing means 84. In the arrangement shown in Figs. 8 to C, the ablation electrode 16 includes a forward electrode region 110 and a rear electrode region 112. The for- ward electrode region 110 and the rear electrode re- gion 112 are generally spherical in shape. An electrical and thermal insulating sleeve 114 separates the forward electrode region 110 and the rear electrode region 112. The sleeve 114 is general- ly cylindrical in shape. The resulting "peanut" shape is well suited for use within the valve regions of the heart. In the illustrated embodiment, the forward electrode region 110 and the rear electrode region 112 are made of platinum. The sleeve 114 is made of a polysulfone material. Multiple temperature sensing means 84 occupy the surface of each forward and rear electrode region 11.0 and 112. At least one, and preferably all, the temperature sensing means 84 are thermally isolated from the remainder of the surrounding body of the as¬ sociated electrode region 110 and 112. Each electrode region 110 and 112 includes an interior core well 116 and branch wells 118 that open at the surface of the associated electrode region 110 and 112. A temperature sensing means 84 occupies each well branch. In the illustrated and preferred embodiment, thermal insulating means 88 also isolates each temperature sensing means 84 from the interior surface of the associated branch well 116 and 118 and the rest of the thermal mass of the electrode region 110 and 112. As in the embodiment shown in Figs. 6A to C, each thermal sensing means 84 includes a small bead thermistor 94 with two associated lead wires 96 and 98. The temperature sensing tips of the thermistors 94 are exposed at the surface of the associated electrode regions 110 and 112 for multiple contact with the tissue. The associated lead wires 96 and 98 are bundled in the center core well 116, passing through the guide tube 22 to the handle 20. As in the previous embodiments, potting com¬ pound 100 encapsulates each thermistor 94 and its lead wires 96 and 98 within the associated branch well 116 and 118. Insulating sheathes 102 also shield the en¬ capsulated lead wires 96 and 98. Also as in the previous embodiments, a thermal insulating tube 104 envelopes each electrically encapsulated thermistor 94 and its lead wires 96 and 98. Adhesive bonds the thermal insulation tube 104 to the interior wall of each branch well 116 and 118. The number and arrangement of possible ar¬ rays of multiple temperature sensing means 84 can, of course, vary from the specific configurations shown in Figs. 6, 7, and 8. For example, one. or more temperature sensing means 84 can occupy the side re¬ gion of the ablation electrode 16, below its tip. The branch wells holding the temperature sensing means 84 also can extend from the center well at various angles, acute, obtuse, or perpendicular. Not all the temperature sensing means 84 need be thermally isolated from the electrode 16, but preferable they all are. As Fig. 5 shows, the third monitoring means 82 can perform different display and control functions in response to sensed temperature conditions according to different prescribed criteria. Preferably, the third monitoring means 82 responds to sensed tissue temperature not only in ab- solute terms, but it also serves to record changes in tissue temperature over time and respond to these changes as well. In the illustrated embodiment, the third monitoring means 82 includes a control output 120 for each temperature sensing means 84 carried by the as- sociated ablation electrode 16. The lead wires 96 and 98 for each thermistor 94 provide the input for the control outputs 120. Al¬ ternatively, when the ablation electrode 16 carries multiple thermistors 94, the number of lead wires 96 and 98 traversing the guide tube 22 can be minimized by providing an integrated circuit 122 within the third monitoring means 82 for multiplexing the input signals of the thermistors 94. In the illustrated embodiment, the third monitoring means 82 includes a converter 124 to obtain an average temperature for each grouped array of thermistors 94 during a user prescribed period (which in the illustrated embodiment is about once every 0.01 second) . The embodiment shown in Figs. 6A to C includes one thermistor 94, so the input signal and the average will be the same. The embodiment shown in Figs. 7A to C includes a single grouped array of five thermistors 94 clustered at the tip of the ablation electrode 16. For this array, the converter adds the individual in¬ put signals and divides by five. The embodiment shown in Figs. 8A to C includes two grouped arrays, one having five thermis¬ tors 94 on the forward electrode region 110, the other having four thermistors 94 on the rear electrode re¬ gion 112. The converter 124 adds the input signals for each grouped array and divides by the associated number of thermistors 94 in each grouped array to ob¬ tain an average for the forward electrode region 110 and a separate average for the rear electrode region 112. The third monitoring means 82 includes an analog to digital converter 126. The converter 126 digitizes the sensed temperature average(s) for the system 10. The converter 126 also digitizes the value of the calibration resistor 134. The thermistor re- sistance value is divided by the calibration resistor value to get a normalized resistance for the thermis¬ tor 94. This value is the input to a read only memory (ROM) 136 (see Fig. 5B) containing stored thermistor temperature data. The ROM 136 output is the actual measured tissue temperature (in degrees C) , taking into account deviations in the nominal resistance of the thermistor 94. Although not shown in the drawings, the em¬ bodiments having multiple thermistors 94 would include an equal number of calibration resistors 134, one for each thermistor 94. The digital output(s) of the converter can be used to display measurement results. In the illus¬ trated embodiment, the system 10 includes a third dig- ital display 128 on the generator 12 to show the user the average sensed temperature. If the "peanut" type electrode shown in Figs. 8A to C is used, the system 10 includes a separate display for the forward and rear electrode region 110 and 112. The digital output(s) of the converter 126 also can be used to control operation of the generator 12. Preferably, the temperature control signals of the third monitoring means 82 also are used to further augment the function of the first and second monitor¬ ing means 38 and 76 previously described. In the illustrated embodiment, the system 10 uses the digitized temperature outputs in a feedback loop that maintains radiofrequency output voltage within a desired range or at a constant value to con- trol radiofrequency power at the ablation electrode 16. By controlling the power delivered by the genera¬ tor 12 based upon temperature, the physician is able to control the size of the lesion generated. For this purpose, the system 10 includes an input 130 for the user to enter an operating value desired for the tissue temperature. If tissue temperature remains within a pre¬ set range, but deviates a prescribed amount within the range, the third monitoring means 82 generates a con¬ trol signal that either increases or reduces but does not interrupt, the power output. If the tissue temper¬ ature rises, the control signal reduces the power out¬ put. If the tissue temperature falls, the control signal increases the power output. If measured tissue temperature falls outside the preset range, the third monitoring means 82 generates a command signal to shut off power to the ablation electrode 16. A represen¬ tative set range for tissue temperature for cardiac ablation is believed to be about 40 degrees to 100 degrees C. When temperature begins in the set range and, over time, fall outside it, the most likely cause is the coagulation of blood on the ablation electrode 16, requiring withdrawal and cleaning of the ablation electrode 16. A sudden change in tissue temperature outside the set range suggests a shift in the position of the ablation electrode 16, requiring repositioning of the ablation electrode 16. The system 10 preferably includes flashing lights and an audible alarm (not shown) to transmit a warning to the user when these temperature- based con¬ ditions occur. The system 10 as described can provide pre- cise control over the ablation procedure. The moni- toring and control of actual phase sensitive power assure the effective distribution of radiofrequency to the ablation electrode 16. The monitoring and control of tissue impedance and tissue temperature, either separately or in combination, set safe physiological limits in terms of lesion size and detection of coagu¬ lation. Monitoring and control of tissue impedance and/or tissue temperature also provide information regarding orientation of the ablation electrode 16.

Claims: Claims; We Claim: 1. A system for ablating tissue using radiofrequency energy comprising a source of radiofrequency energy, electrode means electrically connected to the source for emitting radiofrequency energy at the ablation site, current monitoring means for measuring radiofrequency current delivered to the electrode means and generating a measured radiofrequency current signal, voltage monitoring means for measuring radiofrequency voltage at the electrode means and gen¬ erating a measured radiofrequency voltage signal, power monitoring means for multiplying the measured current signal by the measured voltage signal and deriving a phase sensitive power signal that rep¬ resents the actual phase sensitive power transmitted the electrode means, and control means for performing a function based upon the actual power signal. 2. A system according to claim 1 wherein the control means includes display means for indicating the value of the phase sensitive power signal in a user readable format. 3. A system according to claim 1 wherein the control means includes feedback means for controlling the voltage output of the source based upon the phase sensitive power signal. 4. A system according to claim 1 wherein the feedback means includes register means for storing a desired phase sensitive power amount, comparator means for comparing the de- sired phase sensitive power amount in the register means with the phase sensitive power signal generated by the power monitoring means and generating a devia¬ tion signal, and command means for varying output vol¬ tage of the source in response to the deviation sig¬ nal. 5. A system according to claim 4 wherein the command means varies output voltage of the source to maintain a prescribed minimum deviation signal. 6. A system according to claim 1 and further including phase shift detection means comprising . means for deriving a root mean squared current signal based upon the measured radiofrequency current signal, means for deriving a root mean squared voltage signal based upon the measured radiofrequency voltage signal, means for multiplying the root mean squared current signal by the root mean squared vol¬ tage signal and deriving a root mean squared power signal that represents the apparent power transmitted the electrode means, means for comparing the root mean squared power signal with the phase sensitive power signal generated by the power monitoring means and generating a deviation signal that indicates a phase shift between the measured radiofrequency voltage sig- nal and the measured radiofrequency current signal at the electrode means, and phase shift control means for perfor¬ ming a function based upon the phase shift deviation signal. 7. A system according to claim 6 wherein the phase shift control means interrupts the conduction of radiofrequency energy when the phase shift deviation signal exceeds a predetermined value. 8. A system according to claim 6 wherein the phase shift control means generates a warning signal when the phase shift devi¬ ation signal exceeds a predetermined value. 9. A system according to claim 6 wherein the control means includes feedback means for controlling the voltage output of the source based upon the phase sensitive power signal when the phase shift deviation signal remains within an pre¬ scribed range. 10. A system according to claim 9 wherein the feedback means reduces the out¬ put voltage of the source when the phase shift devia¬ tion signal increases beyond the prescribed range. 11. A system according to claim 10 wherein the feedback means interrupts power to the electrode means when phase shift deviation sig¬ nal exceeds a prescribed maximum value. 12. A system according to claim 1 and further including tissue impedance moni¬ toring means comprising means for deriving a root mean squared current signal based upon the measured radiofrequency current signal, means for deriving a root mean squared voltage signal based upon the measured radiofrequency voltage signal, and means for dividing the root mean squared voltage signal by the root mean squared cur¬ rent signal to derive a measured tissue impedance sig- nal, and impedance control means for performing a function based upon the measured tissue impedance signal. 13. A system according to claim 12 wherein the impedance control means calcu¬ lates changes in the measured tissue impedance signals during a selected interval and generates control sig- nals based upon predetermined criteria. 14. A system according to claim 12 wherein the impedance control means generates a command signal to interrupt energy to the electrode means when the measured tissue impedance signal falls outside a predetermined range. 15. A system according to claim 12 wherein the predetermined range is about 50 to 300 ohms. 16. A system according to claim 12 wherein the impedance control means generates a control signal when the measured tissue impedance signal begins in a predetermined range and, over time, increases beyond it, suggesting coagulation of blood on the electrode means. 17. A system according to claim 12 wherein the impedance control means generates a control signal in response to an incremental increase in the measured tissue impedance signal beyond a prescribed amount, suggesting the on¬ set of coagulation or a sudden shift in the position of the electrode means. 18. A system according to claim 12 wherein the impedance control means generates a control signal when the measured tissue impedance signal exceeds a predetermined maximum amount, suggesting poor skin contact with the elec- trode means or an electrical problem in the system. 19. A system according to claim 12 wherein the impedance control means includes display means for indicating the value of the measured tissue impedance signal in a user readable format.

Inventors: EDWARDS STUART D, STERN ROGER A, EDWARDS, STUART, D., STERN, ROGER, A.
Applicants: EP TECHNOLOGIES, EP TECHNOLOGIES, INC.
Application year: 1992
Main classification: A61B17
IPC codes: A61B17, A61B18