Method and system for generating a respiratory signal for use in electrophysiological treatment

Abstract

Within the electroanatomical mapping system, a respiratory signal may be generated from the non-driven impedance signals received from the plurality of patch electrodes. The non-driven impedance signals may be used to determine a reference respiratory signal. Each subset of the non-driven impedance signals may then be compared to the reference respiratory signal to determine a polarity value. A scale factor may also be calculated to normalize the non-driven impedance signals. The polarity value and scale factor may be applied to the non-driven impedance signals to generate weighted non-driven impedance signals that may be summed as a composite respiratory signal. The polarity value and scale factor may be applied to the composite respiratory signal for use in real time (e.g., gating data collection, respiratory compensation, detecting irregular breathing, etc.).

Description

Background Art

[0001] (Cross - reference to related applications) This application claims the benefit of U.S. Provisional Application No. 63 / 237,269, filed Aug. 26, 2021, which is hereby incorporated by reference herein as if fully set forth herein.

[0002] This disclosure generally relates to electrophysiological procedures, such as cardiac diagnostic and therapeutic procedures including electrophysiological mapping and cardiac ablation. In particular, this disclosure relates to the generation of composite respiratory signals for use in such electrophysiological procedures.

[0003] It is known to use respiratory signals in various electrophysiological procedures. For example, respiratory signals can be used for gating, detection of irregular breathing, and correction of certain movements during electrophysiological mapping (e.g., as described in U.S. Patent No. 7,263,397, which is hereby incorporated by reference herein as if fully set forth herein).

[0004] In many existing electroanatomical mapping systems, respiratory signals are obtained from impedance measurements at one or more body surface (e.g., patch) electrodes. However, there are variations in the placement of such electrodes, and further variations in environmental and physiological conditions, so these impedance signals may have drawbacks such as baseline fluctuations, low amplitude, and polarity reversals. These drawbacks can lead to inaccurate gating, false positive or false negative detection of irregular breathing, and errors in respiratory motion correction algorithms. This can, in turn, lead to motion artifacts in the electrophysiological catheter.

Summary of the Invention

Means for Solving the Problems

[0005] The present disclosure provides a method for generating a respiratory signal within an electroanatomical mapping system. According to aspects of the present disclosure, the method includes the electroanatomical mapping system receiving a plurality of non-driven impedance signals from a plurality of patch electrodes, the electroanatomical mapping system determining a reference respiratory signal using the plurality of non-driven impedance signals, and for each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, the electroanatomical mapping system calculating a polarity value of the non-driven impedance signal, calculating a scale factor of the non-driven impedance signal, and the electroanatomical mapping system calculating a composite respiratory signal from the subset of the plurality of non-driven impedance signals.

[0006] In an embodiment of the present disclosure, calculating the polarity value of the non-driven impedance signal includes calculating a correlation coefficient between the non-driven impedance signal and the reference respiratory signal, and calculating the polarity value based on the sign of the correlation coefficient.

[0007] In another embodiment of the present disclosure, calculating the scale factor of the non-driven impedance signal includes normalizing the non-driven impedance signal, such as by dividing the non-driven impedance signal by its signal range.

[0008] A plurality of weighted non-driven impedance signals can be calculated by multiplying the polarity value and the scale factor corresponding to each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, and a composite respiratory signal can be calculated from the subset of the plurality of non-driven impedance signals by summing the plurality of weighted non-driven impedance signals.

[0009] It is also conceivable to determine the polarity value of the composite respiration signal. For example, the polarity value of the composite respiration signal may be determined so that the polarity of the composite respiration signal corresponds to the polarity of the PRS-A signal. Alternatively, the polarity value of the composite respiration signal is determined so as to correspond to the polarity of the PRS-A signal only when the correlation coefficient between the composite respiration signal and the PRS-A signal exceeds a preset threshold, for example, about 75%.

[0010] In still another embodiment of the present disclosure, the polarity value of the composite respiration signal may be determined according to the assumption that the expiration duration of the composite respiration signal exceeds the inspiration duration of the composite respiration signal. For example, the polarity value of the composite respiration signal may be determined so that the minimum value of the composite respiration signal is closer to the average value of the composite respiration signal than the maximum value of the composite respiration signal.

[0011] As another example, the polarity value of the composite respiration signal may be determined so that the average value of the troughs of the composite respiration signal is closer to the average value of the composite respiration signal than the average value of the peaks of the composite respiration signal.

[0012] As yet another example, the polarity value of the composite respiration signal may be determined so that the time interval between when the composite respiration signal first crosses zero downward and then crosses zero upward is longer than the time interval between when the composite respiration signal first crosses zero upward and then crosses zero downward.

[0013] A composite scale factor of the composite respiration signal for normalizing the composite respiration signal may also be calculated.

[0014] The step of determining the reference respiration signal using a plurality of non-driven impedance signals may include determining either the strongest non-driven impedance signal among the plurality of impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal.

[0015] Among the plurality of non-driven impedance signals, the strongest non-driven impedance signal may be the signal with the largest standard deviation among the plurality of non-driven impedance signals, the signal with the largest amplitude among the plurality of non-driven impedance signals, or other appropriate signals.

[0016] Similarly, the step of determining either the strongest non-driven impedance signal among the plurality of impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal may include, on the one hand, determining either the strongest impedance signal or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal according to the correlation coefficient between the strongest impedance signal and a subset of the plurality of non-driven impedance signals, and on the other hand, according to the correlation coefficient between the first principal component signal and a subset of the plurality of non-driven impedance signals.

[0017] The present disclosure provides an electroanatomical mapping system including a respiration compensation module, wherein the respiration compensation module receives a plurality of non-driven impedance signals from a plurality of patch electrodes, determines a reference respiration signal using the plurality of non-driven impedance signals, calculates a polarity value of each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, calculates a scale factor of the non-driven impedance signal, multiplies the polarity value and the scale factor corresponding to each non-driven impedance signal within the subset of the plurality of non-driven impedance signals to calculate a plurality of weighted non-driven impedance signals, and sums the plurality of weighted non-driven impedance signals to calculate a composite respiration signal from the subset of the plurality of non-driven impedance signals.

[0018] The respiration compensation module may determine either the strongest non-driven impedance signal among the plurality of impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal.

[0019] The foregoing and other aspects, features, details, utilities and advantages of the present invention will become apparent from the following description and claims, and from the accompanying drawings. BRIEF DESCRIPTION OF THE DRAWINGS

[0020]

Figure 1

[0021]

Figure 2

[0022]

Figure 3A

Figure 3B

[0023]

Figure 4

[0024]

Figure 5

[0025] Although multiple embodiments are disclosed, further embodiments of the present disclosure will become apparent to those skilled in the art from the following detailed description, which shows and describes exemplary embodiments. Accordingly, the drawings and detailed description should be regarded as being essentially exemplary and not restrictive.

[0026] The present disclosure provides systems, devices, and methods for generating respiratory signals for use in electrophysiological procedures. For purposes of explanation, aspects of the present disclosure are described with reference to electrophysiological investigations performed in conjunction with high-density (HD) grid catheters, such as the Advisor™ HD Grid Mapping Catheter, Sensor Enabled™, of Abbott Laboratories (Abbott Park, Illinois), and electroanatomical mapping systems, such as the Ensite Precision™ cardiac mapping system or the Ensite™ X EP system, also of Abbott Laboratories. However, those of ordinary skill in the art will understand how to advantageously apply the teachings herein in other situations and / or with respect to other devices.

[0027] FIG. 1 is a schematic diagram of an exemplary electroanatomical mapping system 8 for performing a cardiac electrophysiological investigation by navigating a cardiac catheter, measuring electrical activity occurring in the heart 10 of patient 11, and three-dimensionally mapping the electrical activity and / or information related to or representative of the electrical activity so measured. For example, using system 8, an anatomical model of a patient's heart 10 can be created using one or more electrodes. For example, to create a diagnostic data map of a patient's heart 10, system 8 can be used to measure electrophysiological data at multiple points along the surface of the heart and store the measured data in association with the location information of each measurement point at which the electrophysiological data was measured.

[0028] As will be appreciated by those of ordinary skill in the art, system 8 typically identifies the position of an object, and in some aspects the orientation of the object, within a three-dimensional space and represents those positions as location information specified relative to at least one reference. This is referred to herein as "position identification."

[0029] For the sake of simplicity of illustration, patient 11 is depicted schematically as an ellipse. In the embodiment shown in FIG. 1, three sets of surface electrodes (e.g., patch electrodes) 12, 14, 16, 18, 19, 22 are applied to the surface of patient 11, defining in pairs three substantially orthogonal axes herein referred to as the x-axis (12, 14), the y-axis (18, 19), and the z-axis (16, 22). In other embodiments, the electrodes may be arranged in other configurations, for example, where a plurality of electrodes are on a particular body surface. As yet another option, the electrodes need not be on the body surface and may be disposed inside the body.

[0030] In FIG. 1, the x-axis surface electrodes 12, 14 are applied to the patient along a first axis, such as the outer aspect of the patient's thoracic region (e.g., applied to the patient's skin under both arms), and may be referred to as the left and right electrodes. The y-axis electrodes 18, 19 are applied to the patient along a second axis that is substantially orthogonal to the x-axis, such as along the patient's inner thighs and neck, and may be referred to as the left leg electrode and the neck electrode. The z-axis electrodes 16, 22 are applied along a third axis that is substantially orthogonal to both the x-axis and the y-axis, such as along the patient's sternum and spine in the thoracic region, and may be referred to as the chest electrode and the back electrode. The heart 10 is between these pairs of surface electrodes 12 / 14, 18 / 19, 16 / 22.

[0031] According to an embodiment of the present disclosure, each surface electrode measures six signals, specifically three resistance (impedance) signals and three reactance signals. These signals may then be classified into three resistance / reactance signal pairs. As described below, one resistance / reactance signal pair reflects a drive value, and the other two resistance / reactance signal pairs reflect non-drive values (e.g., measurements of the electric field generated by other drive pairs, similar to that described below for electrode 17).

[0032] An additional surface reference electrode (e.g., an "abdominal patch") 21 provides a reference electrode and / or a ground electrode to the system 8. The abdominal patch electrode 21 may be an alternative to the fixed intracardiac electrode 31 described in more detail below. In an alternative embodiment, the surface reference electrode 21 may include, instead of or in addition to, a magnetic patient reference sensor - anterior (a "PRS-A") disposed on the patient's chest.

[0033] Furthermore, it should be understood that the patient 11 may have most or all of the leads of a conventional electrocardiogram ("ECG" or "EKG") system in predetermined positions. In a particular embodiment, for example, a standard set of 12 ECG leads may be utilized to sense the electrocardiogram of the patient's heart 10. This ECG information is available in the system 8 (e.g., may be provided as an input to the computer system 20). To the extent that ECG leads are well understood and for purposes of clarity of the figure, FIG. 1 shows only one lead 6 and its connection to the computer 20.

[0034] A representative catheter 13 having at least one electrode 17 is also shown. This representative catheter electrode 17 is referred to herein as a "roving electrode", a "moving electrode", or a "measuring electrode". Typically, a plurality of electrodes 17 on the catheter 13, or a plurality of electrodes 17 on a plurality of such catheters, are used. In one embodiment, for example, the system 8 may include 64 electrodes on 12 catheters disposed within the patient's heart and / or vascular system. In other embodiments, the system 8 may utilize a single catheter that includes a plurality (e.g., eight) of splines, each spline including a plurality (e.g., eight) of electrodes.

[0035] However, the foregoing embodiments are merely exemplary, and the number of electrodes used and / or the number of catheters may be arbitrary. For example, for the purposes of the present disclosure, FIG. 2 shows a portion of an exemplary multi-electrode catheter, such as the Advisor (trademark) HD grid mapping catheter, the Sensor Enabled (trademark), and in particular the HD grid catheter 13. The HD grid catheter 13 includes a catheter body 200 coupled to paddles 202. The catheter body 200 may further include first and second body electrodes 204, 206, respectively. The paddles 202 may include a first spline 208, a second spline 210, a third spline 212, and a fourth spline 214, which are coupled to the catheter body 200 by a proximal coupler 216 and to each other by a distal coupler 218. In one embodiment, the first spline 208 and the fourth spline 214 may be one continuous segment, and the second spline 210 and the third spline 212 may be another continuous segment. In other embodiments, the plurality of splines 208, 210, 212, 214 may be separate segments that are coupled to each other (e.g., by proximal coupler 216 and distal coupler 218, respectively). It should be understood that the number of splines included in the HD catheter 13 may be any number, and the configuration of four splines shown in FIG. 2 is merely exemplary.

[0036] As described above, the splines 208, 210, 212, 214 may include any number of electrodes 17, and FIG. 2 shows 16 electrodes 17 arranged in a 4×4 array. It should also be understood that the electrodes 17 may be evenly and / or unevenly spaced when measured along the splines 208, 210, 212, 214 and when measured between the splines 208, 210, 212, 214. For purposes of facilitating reference in this description, FIG. 3A provides alphanumeric labels for the electrodes 17.

[0037] As will be appreciated by those skilled in the art, each of two adjacent electrodes 17 defines a bipole. Thus, the 16 electrodes 17 on the catheter 13 specifically define a total of 42 bipoles, including 12 along the spline (e.g., between electrode 17a and electrode 17b, or between electrode 17c and electrode 17d), 12 across the spline (e.g., between electrode 17a and electrode 17c, or between electrode 17b and electrode 17d), and 18 diagonally across the spline (e.g., between electrode 17a and electrode 17d, or between electrode 17b and electrode 17c).

[0038] For ease of reference in the description, FIG. 3B provides alphanumeric labels for the bipoles along the spline and across the spline. The alphanumeric labels for the diagonal bipoles are omitted in FIG. 3B, but this is for the sole purpose of clarity of illustration. It is clearly intended that the teachings herein apply also to the diagonal bipoles.

[0039] Any of the bipoles may be used to generate a bipolar potential map by techniques well known to those skilled in the art. Further, these bipolar potential maps may be combined (e.g., linearly combined) to generate a potential map including activation timing information in any direction in the plane of the catheter 13 by calculating the electric field loops of the electrode groups. U.S. Patent Application Publication No. 2018 / 0296111 (Publication No. 111), which is incorporated herein by reference as if fully set forth herein, discloses details of calculating the electric field loops of electrode groups on an HD grid catheter. These potential maps are referred to herein as "omnipolar potential maps", and their corresponding directions are referred to herein as "omnipoles" or "virtual bipoles".

[0040] In any case, catheter 13 may be used to simultaneously collect a plurality of electrophysiological data points of various bipoles defined by electrodes 17 provided thereon, and each of such electrophysiological data points includes both location-specific information (e.g., the position and orientation of the selected bipole) and the potential map signal of the selected bipole. For purposes of explanation, the method according to the present disclosure will be described with reference to the individual electrophysiological data points collected by catheter 13. However, it should be understood that the teachings herein are applicable to a plurality of electrophysiological data points collected by catheter 13, either serially and / or in parallel.

[0041] Catheter 13 (or a plurality of such catheters) is typically introduced into a patient's heart and / or vascular system via one or more introducers in a well-known procedure. In fact, various approaches for introducing catheter 13 into a patient's heart, such as a transseptal approach, should be well known to those skilled in the art, and thus need not be further described herein.

[0042] Since each electrode 17 is within the patient, the system 8 can simultaneously collect the position data of each electrode 17. Similarly, electrophysiological data (e.g., endocardial potential map) can be collected from the heart surface using each electrode 17. Those skilled in the art should be well aware of various modalities for obtaining and processing electrophysiological data points (including, for example, both contact electrophysiological mapping and non-contact electrophysiological mapping), and thus no further explanation is necessary for understanding the technology disclosed herein. Similarly, various techniques well known in the art can be used to generate a graphic display of the heart shape and / or the electrical activity of the heart from a plurality of electrophysiological data points. Further, to the extent that those skilled in the art understand the method of creating an electrophysiological map from electrophysiological data points, that aspect will be described herein only to the extent necessary for understanding the present disclosure.

[0043] Returning now to FIG. 1, in some embodiments, an optional fixed reference electrode 31 (e.g., attached to the wall of the heart 10) is shown on the second catheter 29. For calibration purposes, this electrode 31 may be stationary (e.g., attached to or near the wall of the heart) or arranged in a fixed spatial relationship with respect to the roving electrode (e.g., electrode 17) and may thus be referred to as a "navigation reference" or "local reference." The fixed reference electrode 31 may be used in addition to or instead of the surface reference electrode 21 described above. In many cases, a coronary sinus electrode or other fixed electrode within the heart 10 can be used as a reference for measuring voltage and displacement. That is, as described below, the fixed reference electrode 31 may define the origin of the coordinate system.

[0044] Each surface electrode is connected to a multiplex switch 24, and surface electrode pairs are selected by software executed on a computer 20 that connects the surface electrodes to a signal generator 25. Alternatively, the switch 24 may be omitted, and a plurality (e.g., three) of signal generators 25 may be provided, one for each measurement axis (i.e., for each of the surface electrode pairs).

[0045] The computer 20 may be, for example, a conventional general-purpose computer, a special-purpose computer, a distributed computer, or any other type of computer. The computer 20 may include one or more processors 28, such as one central processing unit ("CPU") or a plurality of processing units generally referred to as a parallel processing environment, and the processor 28 may execute instructions for implementing the various aspects described herein.

[0046] Typically, to achieve catheter navigation within a biological conductor, three nominally orthogonal electric fields are generated by a series of driven and sensed electrical bipoles (e.g., by driving patch electrode pairs 12 / 14, 18 / 19, 16 / 22). Alternatively, these orthogonal electric fields may be decomposed, and any patch electrode pair can be driven as a dipole to perform effective electrode triangulation. Similarly, electrodes 12, 14, 18, 19, 16, 22 (or any number of electrodes) may be arranged in any other effective configuration to pass current to or sense current from electrodes within the heart. For example, a plurality of electrodes may be arranged on the patient's back, side, and / or abdomen. Further, such non-orthogonal ways enhance the flexibility of the system. With respect to any desired axis, the measured potentials across the roving electrodes resulting from a given set of drive (source-sink) configurations can be algebraically combined to obtain the same effective potential as obtained by simply driving a uniform current along the orthogonal axes.

[0047] Thus, with respect to a ground reference such as abdominal patch 21, any two of the patch electrodes 12, 14, 16, 18, 19, 22 may be selected as the source and drain of a dipole, and the non-excited electrodes measure the voltage with respect to the ground reference. The roving electrode 17 disposed within the heart 10 is also exposed to the electric field from the current pulse and may likewise be measured with respect to a ground reference such as abdominal patch 21. In practice, the catheter within the heart 10 may include more or fewer electrodes than the 16 shown, and each electrode potential may be measured. As described above, at least one electrode may be fixed to the inner surface of the heart to form a fixed reference electrode 31, which is likewise measured with respect to a ground reference such as abdominal patch 21 and may be defined as the origin of the coordinate system in which the system 8 measures position. The position of the roving electrode 17 within the heart 10 may be identified using all of the data sets from each of the surface electrodes, internal electrodes, and virtual electrodes.

[0048] The measured voltage is used by system 8, and the position of the electrode within the heart, such as roving electrode 17, relative to a reference position, such as reference electrode 31, may be determined in three-dimensional space. That is, the voltage measured at reference electrode 31 may be used to define the origin of the coordinate system, and the voltage measured at roving electrode 17 may be used to represent the position of roving electrode 17 relative to the origin. In some embodiments, the coordinate system is a three-dimensional (x, y, z) Cartesian coordinate system, although other coordinate systems, such as polar coordinate systems, spherical coordinate systems, cylindrical coordinate systems, etc., are also contemplated.

[0049] As is apparent from the foregoing description, the data used to determine the position of the electrode within the heart is measured while the surface electrode pair is applying an electric field to the heart. The electrode data may also be used, for example, to generate a respiration correction value used to improve the raw position data of the electrode position, as described in U.S. Patent No. 7,263,397, which is hereby incorporated by reference in its entirety. Also, the electrode data may be used, for example, to correct for impedance changes in the patient's body, as described in U.S. Patent No. 7,885,707, which is hereby incorporated by reference in its entirety.

[0050] Thus, in a representative embodiment, system 8 first selects a set of surface electrodes and then drives them with a current pulse. While the current pulse is being delivered, electrical activity, such as the voltage measured at at least one of the remaining surface electrodes and the in-vivo electrodes, is measured and stored. Correction of artifacts, such as respiration and / or impedance shifts, may be performed as described above.

[0051] In some aspects of the present disclosure, system 8 may be a hybrid system incorporating both an impedance-based localization function (e.g., the function described above) and a magnetic-based localization function. Thus, for example, system 8 may further include a magnetic source 30 coupled to one or more magnetic field generators. For clarity, only two magnetic field generators 32, 33 are shown in FIG. 1, but additional magnetic field generators (e.g., a total of six magnetic field generators defining three substantially orthogonal axes similar to the axes defined by patch electrode sets 12, 14, 16, 18, 19, 22) may be used without departing from the scope of the present teachings. Similarly, those skilled in the art will understand that catheter 13 may include one or more magnetic position sensors (e.g., coils) to determine the position of catheter 13 within the magnetic field generated in this manner.

[0052] In some embodiments, system 8 is the EnSite™ Velocity™, EnSite Precision™, or EnSite™ X cardiac mapping and visualization system of Abbott Laboratories. However, for example, other localization systems including the RHYTHMIA HDX™ mapping system of Boston Scientific (Marlborough, Massachusetts), the CARTO navigation and localization system of Biosense Webster (Irvine, California), the AURORA® system of Northern Digital (Waterloo, Ontario), the NIOBE® magnetic navigation system of Stereotaxis (St. Louis, Missouri), and the MediGuide™ technology of Abbott Laboratories may also be used in connection with the present teachings.

[0053] The following location identification and mapping systems (all of which are hereby incorporated by reference in their entirety) can also be used with the present invention: U.S. Patent Nos. 6,990,370; 6,978,168; 6,947,785; 6,939,309; 6,728,562; 6,640,119; 5,983,126; and 5,697,377.

[0054] Aspects of the present disclosure relate to the generation of a composite respiratory signal for use by system 8 (e.g., for respiratory gating, motion compensation, and / or detection of irregular respiration). Accordingly, system 8 may include a respiratory compensation module 58.

[0055] Referring to flowchart 400 of representative steps presented in FIG. 4, one exemplary method according to aspects of the present disclosure will be described. In some embodiments, for example, flowchart 400 may represent some exemplary steps that may be performed by the electroanatomical mapping system 8 of FIG. 1 (e.g., by processor 28 and / or respiratory compensation module 58). It should be understood that the representative steps described below can be implemented either in hardware or in software. For the sake of convenience of explanation, in this specification, the term "signal processor" is used in describing either the hardware-based implementation or the software-based implementation of the teachings herein.

[0056] At block 402, a plurality of non-driven impedance signals are received from patch electrodes 12, 14, 16, 18, 19, 22. In particular, it is desirable to use non-driven impedance signals from chest electrodes, back electrodes, left electrodes, right electrodes, and neck electrodes, and according to the present teachings, non-driven signals from the left foot electrode are not normally used. Similarly, since reactance signals are more sensitive to disturbances, non-driven reactance signals are mostly ignored when applying the present teachings.

[0057] In block 404, a reference respiration signal (also called a base respiration signal) is typically determined using the non-driven impedance signals measured by body surface electrodes 12, 14, 16, 19, 22 (excluding, for example, the left foot electrode 18) received in block 402. Various approaches for determining the reference respiration signal are conceivable.

[0058] According to a particular embodiment of the present disclosure, the strongest non-driven impedance signal among the plurality of non-driven impedance signals received in block 402 is determined as the reference respiration signal. Those skilled in the art will understand that the term "strongest non-driven impedance signal" can be defined in various ways, although by way of example only, as a non-driven impedance signal having a maximum standard deviation or a non-driven impedance signal having a maximum amplitude.

[0059] In another embodiment of the present disclosure, among the plurality of non-driven impedance signals received in block 402, a first principal component signal that can be calculated using singular value decomposition is determined as the reference respiration signal.

[0060] It is contemplated that whether to use the strongest non-driven impedance signal as the reference respiration signal or the first principal component signal may be selected based on the correlation coefficient. That is, once the strongest non-driven impedance signal is identified, the correlation coefficient may be calculated between the strongest non-driven impedance signal and the remaining non-driven impedance signals received in block 402. The average of the absolute values of these correlation coefficients may also be calculated.

[0061] Similarly, once the first principal component signal is calculated, the correlation coefficient may be calculated between the first principal component signal and the non-driven impedance signals received in block 402. Here too, the average of the absolute values of these correlation coefficients may be calculated. The signal with the larger average of the absolute values of the correlation coefficients among the strongest non-driven impedance signal and the first principal component signal may be selected as the reference respiration signal.

[0062] At block 406, a polarity value is calculated for one selected non-driven impedance signal among the non-driven impedance signals received at block 402. The polarity value is used to align the selected non-driven impedance signal with the reference respiration signal, and is typically +1 when the selected non-driven impedance signal has a positive correlation with the reference respiration signal, and -1 when the selected non-driven impedance signal has a negative correlation with the reference respiration signal.

[0063] At block 408, a scale factor is calculated for the selected non-driven impedance signal. The scale factor normalizes the selected non-driven impedance signal, for example, by dividing the selected non-driven impedance signal by its signal range.

[0064] Block 410 initiates a loopback to blocks 406 and 408 to calculate the polarity values and scale factors for any other non-driven impedance signals. It should be clearly understood that it is not necessary to use all non-driven impedance signals, and fewer non-driven impedance signals than all non-driven impedance signals may be used. In particular, in aspects of the present disclosure, only non-driven signals from the left patch electrode, right patch electrode, front patch electrode, and rear patch electrode, excluding the neck patch electrode and left leg patch electrode, are used.

[0065] When there are no more non-driven impedance signals to process, the flow proceeds from "NO" in block 410 to block 412, where a composite respiration signal is calculated. Specifically, each non-driven impedance signal is multiplied by its polarity value and scale factor to calculate a weighted non-driven impedance signal. Then, all of the weighted non-driven impedance signals are summed to calculate the composite respiration signal. Mathematically, it is expressed as follows.

Equation

[0066] In block 414, the polarity value of the composite respiration signal (referred to as "polarity value - composite" in this specification) is calculated. Various approaches are conceivable.

[0067] For example, in embodiments of the present disclosure where the PRS - A signal is available, the polarity value - composite may be calculated based on the correlation between the composite respiration signal and the PRS - A signal. More specifically, when the absolute correlation between the composite respiration signal and the PRS - A signal exceeds a preset threshold such as about 75%, the polarity value - composite is determined such that the polarity of the composite respiration signal is the same as the polarity of the PRS - A signal (e.g., +1 for positive correlation and -1 for negative correlation).

[0068] When the PRS - A signal is not available, or when the correlation between the composite respiration signal and the PRS - A signal does not exceed the preset threshold, various heuristic methods may be employed to calculate the polarity value - composite. In particular, the polarity value - composite may be determined according to the assumption that the expiration phase of the composite respiration signal is longer than the inspiration phase. Therefore, the polarity value - composite may be determined such that one or more of the following conditions are true (e.g., as either +1 or -1). · The minimum value of the composite respiration signal is closer to the average value of the composite respiration signal than the maximum value of the composite respiration signal · The average value of the troughs of the composite respiration signal is closer to the average value of the composite respiration signal than the average value of the peaks of the composite respiration signal · The time interval between when the composite respiration signal first crosses zero downward and then crosses zero upward is longer than the time interval between when the composite respiration signal first crosses zero upward and then crosses zero downward

[0069] In block 416, a scale factor of the composite respiration signal (i.e., “scale factor - composite”) is calculated. Similar to the scale factor calculated in block 408, the scale factor - composite normalizes the composite respiration signal, such as by dividing the composite respiration signal by its signal range.

[0070] Once the polarity value - composite and the scale factor - composite are calculated, the composite respiration signal at any given time t (i.e., the real - time composite respiration signal) may be calculated as the sum of all weighted non - driven impedance signals at time t multiplied by the polarity value - composite and the scale factor - composite (block 418). Mathematically, it is as follows.

Equation

[0071] This real - time composite respiration signal can be used for any desired purpose (such as gating the collection of electrophysiological data, respiratory correction, detection of irregular respiration, etc.).

[0072] Also, the real - time composite respiration signal may be high - pass filtered, for example, using a filter with a cut - off frequency of about 0.02 Hz, before being used. One suitable high - pass filter cascades two exponentially weighted moving average filters.

Equation

[0073] Here, α is approximately 0.002. However, if the amplitude of the real-time composite respiration signal changes suddenly, the high-pass filter may be disadvantageous. Therefore, in an alternative embodiment of the present disclosure, the real-time composite respiration signal is further normalized, particularly shifted, until the end of the exhalation phase. This is achieved, for example, by subtracting the running minimum of the composite respiration signal from the composite respiration signal.

[0074] FIG. 5 shows a composite respiration signal 500 according to the foregoing teachings. For comparison, FIG. 5 also shows a respiration signal derived from drive impedance signals from a left patch electrode and a right patch electrode according to an existing respiration signal approach. Further, FIG. 5 shows a gating signal 504 based on the respiration signal 500.

[0075] Although several embodiments have been described in some detail above, those skilled in the art can make many changes to the disclosed embodiments without departing from the spirit or scope of the present invention.

[0076] For example, the teachings herein may be applied in real time (e.g., during an electrophysiological investigation) or in post-processing (e.g., to electrophysiological data points collected during a previously performed electrophysiological investigation).

[0077] All references to directions (e.g., up, down, upward, downward, left, right, leftward, rightward, apex, bottom, upper side, lower side, downward, vertical, horizontal, clockwise, and counterclockwise) are used for identification purposes only to assist the reader's understanding of the present invention and do not particularly limit the position, orientation, or use of the present invention. References to connections (e.g., attached, coupled, connected, etc.) should be construed broadly and may include intermediate members between the connections of the elements and relative movement between the elements. Therefore, references to connections do not necessarily imply that two elements are directly connected and in a fixed relationship with each other.

[0078] All matters included in the above description or shown in the attached drawings are intended to be construed only as illustrative and not as limiting. Without departing from the spirit of the invention defined in the appended claims, the details or structure may be changed. The following items are elements described in the claims at the time of international application. (Item 1) A method for generating a respiratory signal within an electroanatomical mapping system, the method comprising: the electroanatomical mapping system receiving a plurality of non-driven impedance signals from a plurality of patch electrodes; determining a reference respiratory signal using the plurality of non-driven impedance signals; for each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, calculating a polarity value of the non-driven impedance signal; calculating a scale factor of the non-driven impedance signal; calculating a composite respiratory signal from the subset of the plurality of non-driven impedance signals. (Item 2) Calculating the polarity value of the non-driven impedance signal includes: calculating a correlation coefficient between the non-driven impedance signal and the reference respiratory signal; calculating the polarity value based on the sign of the correlation coefficient. The method according to item 1. (Item 3) Calculating the scale factor of the non-driven impedance signal includes normalizing the non-driven impedance signal. The method according to item 1. (Item 4) Normalizing the non-driven impedance signal includes dividing the non-driven impedance signal by its signal range. The method according to item 3. (Item 5) Calculating the composite respiratory signal from the subset of the plurality of non-driven impedance signals includes: calculating a plurality of weighted non-driven impedance signals by multiplying each non-driven impedance signal within the subset of the plurality of non-driven impedance signals by a corresponding polarity value and scale factor; calculating the composite respiratory signal by summing the plurality of weighted non-driven impedance signals. The method according to item 1. (Item 6) The method according to item 1, further comprising determining a polarity value of the composite respiratory signal. (Item 7) The polarity value of the composite respiratory signal is determined such that the polarity of the composite respiratory signal corresponds to the polarity of the PRS-A signal. The method according to item 6. (Item 8) The method according to item 7, wherein the polarity value of the composite respiration signal is determined such that the polarity of the composite respiration signal corresponds to the polarity of the PRS-A signal only when the correlation coefficient between the composite respiration signal and the PRS-A signal exceeds a preset threshold value. (Item 9) The method according to item 8, wherein the preset threshold value is 75%. (Item 10) The method according to item 6, wherein the polarity value of the composite respiration signal is determined according to the assumption that the expiration duration of the composite respiration signal exceeds the inspiration duration of the composite respiration signal. (Item 11) The method according to item 10, wherein the polarity value of the composite respiration signal is determined such that the minimum value of the composite respiration signal is closer to the average of the composite respiration signal than the maximum value of the composite respiration signal. (Item 12) The method according to item 10, wherein the polarity value of the composite respiration signal is determined such that the average of the troughs of the composite respiration signal is closer to the average of the composite respiration signal than the average of the peaks of the composite respiration signal. (Item 13) The method according to item 10, wherein the polarity value of the composite respiration signal is determined such that the time interval between when the composite respiration signal first crosses zero downward and then crosses zero upward is longer than the time interval between when the composite respiration signal first crosses zero upward and then crosses zero downward. (Item 14) The method according to item 6, further comprising calculating a composite scale factor for normalizing the composite respiration signal. (Item 15) Determining the reference respiration signal using the plurality of non-driven impedance signals includes determining either the strongest non-driven impedance signal among the plurality of impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal, according to the method of item 1. (Item 16) The method according to item 15, wherein the strongest non-driven impedance signal among the plurality of non-driven impedance signals is the signal having the largest standard deviation among the plurality of non-driven impedance signals. (Item 17) The method according to item 15, wherein the strongest non-driven impedance signal among the plurality of non-driven impedance signals includes identifying the signal having the largest amplitude among the plurality of non-driven impedance signals. (Item 18) Determining either the strongest non-driven impedance signal among the plurality of impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal includes, on the one hand, defining either the strongest impedance signal among the plurality of non-driven impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal according to the correlation coefficient between the strongest impedance signal and the subset of the plurality of non-driven impedance signals and, on the other hand, according to the correlation coefficient between the first principal component signal and the subset of the plurality of non-driven impedance signals. The method according to item 15. (Item 19) An electroanatomical mapping system, Receiving a plurality of non-driven impedance signals from a plurality of patch electrodes, Determining a reference respiration signal using the plurality of non-driven impedance signals, For each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, Calculating a polarity value of the non-driven impedance signal, Calculating a scale factor of the non-driven impedance signal, Calculating a plurality of weighted non-driven impedance signals by multiplying each non-driven impedance signal within the subset of the plurality of non-driven impedance signals by a corresponding polarity value and a scale factor, and calculating a composite respiration signal from the subset of the plurality of non-driven impedance signals by summing the plurality of weighted non-driven impedance signals. An electroanatomical mapping system comprising a respiration compensation module configured to perform the above operations. (Item 20) The respiration compensation module determines the reference respiration signal as either the strongest non-driven impedance signal among the plurality of impedance signals or the first principal component signal among the plurality of non-driven impedance signals. The system according to item 19.

Claims

1. A method for generating a respiratory signal within an electroanatomical mapping system, the method comprising: the electroanatomical mapping system receiving a plurality of non-driven impedance signals from a plurality of patch electrodes; determining a reference respiratory signal using the plurality of non-driven impedance signals; for each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, calculating a polarity value of the non-driven impedance signal; calculating a scale factor of the non-driven impedance signal; calculating a composite respiratory signal from the subset of the plurality of non-driven impedance signals, including: calculating the composite respiratory signal from the subset of the plurality of non-driven impedance signals includes: calculating a plurality of weighted non-driven impedance signals by multiplying each non-driven impedance signal within the subset of the plurality of non-driven impedance signals by a corresponding polarity value and a scale factor; calculating the composite respiratory signal by summing the plurality of weighted non-driven impedance signals.

2. Calculating the polarity value of the non-driven impedance signal includes: calculating a correlation coefficient between the non-driven impedance signal and the reference respiratory signal; calculating the polarity value based on the sign of the correlation coefficient. The method according to claim 1.

3. Calculating the scale factor of the non-driven impedance signal includes normalizing the non-driven impedance signal. The method according to claim 1.

4. Normalizing the non-driven impedance signal includes dividing the non-driven impedance signal by its signal range. The method according to claim 3.

5. The method according to claim 1, further comprising determining a polarity value of the composite respiratory signal.

6. The polarity value of the composite respiratory signal is determined such that the polarity of the composite respiratory signal corresponds to the polarity of the PRS-A signal. The method according to claim 5.

7. The polarity value of the composite respiratory signal is determined such that the polarity of the composite respiratory signal corresponds to the polarity of the PRS-A signal only when a correlation coefficient between the composite respiratory signal and the PRS-A signal exceeds a preset threshold. The method according to claim 6.

8. The method according to claim 7, wherein the preset threshold value is 75%.

9. The method according to claim 5, wherein the polarity value of the composite respiration signal is determined according to the assumption that the expiration duration of the composite respiration signal is longer than the inspiration duration of the composite respiration signal.

10. The method according to claim 9, wherein the polarity value of the composite respiration signal is determined such that the minimum value of the composite respiration signal is closer to the average of the composite respiration signal than the maximum value of the composite respiration signal.

11. The method according to claim 9, wherein the polarity value of the composite respiration signal is determined such that the average of the troughs of the composite respiration signal is closer to the average of the composite respiration signal than the average of the peaks of the composite respiration signal.

12. The method according to claim 9, wherein the polarity value of the composite respiration signal is determined such that the time interval between the time when the composite respiration signal first crosses zero downward and the time when it then crosses zero upward is longer than the time interval between the time when the composite respiration signal first crosses zero upward and the time when it then crosses zero downward.

13. The method according to claim 5, further comprising calculating a composite scale factor for normalizing the composite respiration signal.

14. Determining the reference respiration signal using the plurality of non-driven impedance signals includes determining either the strongest non-driven impedance signal among the plurality of non-driven impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal, according to the method of claim 1.

15. The method according to claim 14, wherein the strongest non-driven impedance signal among the plurality of non-driven impedance signals comprises the signal having the largest standard deviation among the plurality of non-driven impedance signals.

16. The method according to claim 14, wherein the strongest non-driven impedance signal among the plurality of non-driven impedance signals comprises identifying the signal having the largest amplitude among the plurality of non-driven impedance signals.

17. Determining either the strongest non-driven impedance signal among the plurality of non-driven impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal includes, on the one hand, the correlation coefficient between the strongest impedance signal and the subset of the plurality of non-driven impedance signals, and on the other hand, the correlation coefficient between the first principal component signal and the subset of the plurality of non-driven impedance signals, and defining either the strongest impedance signal among the plurality of non-driven impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal according thereto. The method according to claim 14.

18. An electroanatomical mapping system, Receiving a plurality of non-driven impedance signals from a plurality of patch electrodes, Determining a reference respiration signal using the plurality of non-driven impedance signals, For each non-driven impedance signal within a subset of the plurality of non-driven impedance signals, Calculating a polarity value of the non-driven impedance signal, Calculating a scale factor of the non-driven impedance signal, Calculating a plurality of weighted non-driven impedance signals by multiplying each non-driven impedance signal within the subset of the plurality of non-driven impedance signals by a corresponding polarity value and scale factor, and calculating a composite respiration signal from the subset of the plurality of non-driven impedance signals by summing the plurality of weighted non-driven impedance signals. An electroanatomical mapping system comprising a respiration compensation module configured as such.

19. The system according to claim 18, wherein the respiration compensation module determines the reference respiration signal as either the strongest non-driven impedance signal among the plurality of non-driven impedance signals or the first principal component signal among the plurality of non-driven impedance signals as the reference respiration signal.

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