Patent Description:
An implantable cardiac device, which can also be referred to more generally as an implantable medical device (IMD), often includes electrodes that enable the IMD to sense an electrocardiogram (ECG) or an intracardiac electrogram (IEGM), which can be referred to collectively as an ECG/IEGM, or more generally as a signal indicative of cardiac electrical activity. <FIG> is an idealized drawing of a portion of an ECG/IEGM signal <NUM> (which can also be referred to as an ECG/IEGM waveform <NUM>) that can be obtained using electrodes. One or more of the electrodes can be elements of a cardiac lead, for example, where such electrodes can be implanted within a patient's heart. It is also possible that electrodes be sub-cutaneous electrodes that are implanted external to a patient's heart. For another example, where the IMD is a leadless pacemaker, the electrodes can be located within, on, or near a housing of the leadless pacemaker. These are just a few examples of the type of electrodes that can be used to sense an ECG/IEGM.

Referring to <FIG>, each cycle of the ECG/IEGM signal <NUM> is shown as including a P wave, a QRS complex (including Q, R and S waves), a T wave and a U wave. The P wave is caused by depolarization of the atria. This is followed by an atrial contraction, during which expulsion of blood from the atrium results in further filling of the ventricle. Ventricular depolarization, indicated by the QRS complex, initiates contraction of the ventricles resulting in a rise in ventricular pressure until it exceeds the pulmonary and aortic diastolic blood pressures to result in forward flow as the blood is ejected from the ventricles. The Q, R, and S waves occur in rapid succession, and reflect a single event, and thus are usually considered together as the QRS complex. The Q wave is any downward deflection after the P wave. An R wave follows as an upward deflection, and the S wave is any downward deflection after the R wave. Ventricular repolarization occurs thereafter, as indicated by the T wave and this is associated with the onset of ventricular relaxation in which forward flow stops from the ventricles into the aorta and pulmonary arteries. Thereafter, the pressure in the ventricles falls below that in the atria at which time the mitral and tricuspid valves open to begin to passively fill the ventricles during diastole. Also shown in the exemplary ECG signal <NUM> is a U wave, which may not always be observed as a result of its small size, and which is thought to represent repolarization of the Purkinje fibers.

Also shown in <FIG> are various different intervals and segments that can be measured from an ECG/IEGM signal, such as the ECG/IEGM signal <NUM>. These various intervals and segments are examples of features of an ECG/IEGM signal. These include the PR interval, the QT interval, the RR interval, the PR segment, and the ST segment. The PR interval, which is sometimes referred to as the PQ interval, is the period that extends from the beginning of the P wave (the onset of atrial depolarization) until the beginning of the QRS complex (the onset of ventricular depolarization), and is normally between about <NUM> and <NUM> milliseconds (ms) in duration. The length and/or variability of the PR interval can be used to monitor for certain medical conditions, such as, but not limited to, heart block and pericarditis. The QT interval, which is the period that extends from the beginning of the Q wave until the end of the T wave, represents electrical depolarization and repolarization of the ventricles. A lengthened QT interval is a marker for the potential of ventricular tachyarrhythmias like torsades de pointes and a risk factor for sudden death. The RR interval is the period between R waves, or more generally, between QRS complexes, and is indicative of the heart rate (HR). For example, HR in beats per minute (bpm) can be determined by measuring a plurality of RR intervals, calculating an average RR interval, and dividing the number sixty (<NUM>) by the average RR interval. RR intervals can also be used to measure heart rate variability (HRV), which is the physiological phenomenon of variation in the time interval between heartbeats, which has been shown to be predictor of mortality after myocardial infarction. Additionally, a low HRV is believed to be an indicator of other conditions, such as congestive heart failure and diabetic neuropathy. HRV can be determined by calculating a measure of variance in RR intervals, such as, but not limited to, by calculating the standard deviation (SD), the root mean square of successive differences (RMSSD), or the standard deviation of successive differences (SDSD) of a plurality of consecutive RR intervals. The PR segment is the period that extends from the end of the P wave to the beginning of the QRS complex. PR segment abnormalities can be indicative of pericarditis or atrial ischemia. The ST segment is the period that extends from the end of the S wave (or the end of the QRS complex) to the beginning of the T wave, and is normally between about <NUM> and <NUM> in duration. A normal ST segment has a slight upward concavity. A flat, downsloping, or depressed ST segments, may indicate coronary ischemia. ST elevation may indicate transmural myocardial infarction. ST depression may be associated with subendocardial myocardial infarction, hypokalemia, or digitalis toxicity.

A fundamental operation of many types of IMDs is the ability to measure RR intervals, which enables the IMD to calculate HR and HRV, e.g., in the manners described above. Depending upon the specific type of IMD, RR intervals (and/or measures of HR calculated therefrom) can be used to monitor for arrhythmias, respond to arrhythmias, and/or the like. In order to measure RR intervals from an ECG/IEGM, an IMD should detect R-waves within an ECG/IEGM signal.

Conventionally, R-waves are detected by comparing the amplitude of an ECG/IEGM signal to a detection threshold, and detecting R-waves in response to a threshold crossing in a particular direction. Such a technique can be appreciated from <FIG>, which shows how the ECG/IEGM signal <NUM> can be compared to a detection threshold <NUM> represented by a dashed line. As can also be appreciated from <FIG>, an RR interval can be determined by measuring the interval between successive crossings of the detection threshold <NUM> in a particular direction. In other words, when the ECG/IEGM crosses the detection threshold <NUM> the interval between consecutive crossings (in the same direction) can be recorded and used to determine the RR interval. This interval can be used in various manners, including to make therapy decisions within the IMD.

When measuring RR intervals, or more generally detecting R-waves, it is important for the IMD to not mistaken one or more other morphological features of an ECG/IEGM signal as an R-wave. For example, it is important that an IMD not mistaken a T-wave for an R-wave, which can lead to the IMD mistakenly detecting very short RR intervals (or a very high HR). To avoid this, some IMDs adjust the amplitude of the detection threshold over time in relation to R-waves. This adjusting of the detection threshold is sometimes called automatic sensitivity control (ASC). The ASC feature of an IMD may, for example, automatically measure peak amplitudes of an ECG/IEGM signal and adapt the detection threshold (e.g., <NUM> in <FIG>) automatically. For a more specific example, after every detection threshold crossing indicative of an R-wave, a peak of the ECG/IEGM signal may be measured, and the detection threshold can be adjusted based on how close the peak is to the detection threshold. This may involve, for example, decreasing the detection threshold if the peak is much greater than the detection threshold, and increasing the detection threshold if the peak is only slightly greater than the detection threshold.

<CIT> discloses an automatic sensing system for an implantable cardiac rhythm management device. The system comprises a variable gain amplifier and associated filters where the gain of the amplifier is adjusted as a function of the peak amplitude of a cardiac depolarization signal (either a P-wave or an R-wave) and especially the relationship of the peak value to a maximum value dictated by the circuit's power supply rail. The trip point comparator has its trip point adjusted as a function of the difference between the detected peak value of the signal of interest and the peak value of noise not eliminated by the filtering employed.

Certain embodiments of the present technology relate to implantable medical devices (IMDs), and methods for use therewith, for dynamically controlling sensitivity associated with detecting R-waves without dynamically adjusting a detection threshold, and more specifically, while maintaining a fixed detection threshold. Such a method can include sensing an analog signal indicative of cardiac electrical activity, converting the analog signal indicative of cardiac electrical activity to a digital signal indicative of cardiac electrical activity, and detecting R-waves by comparing the digital signal indicative of cardiac electrical activity to a fixed detection threshold to thereby detect threshold crossings that corresponds to R-waves. The fixed detection threshold can be specified, e.g., during manufacture of an IMD that implements the method, during calibration of an IMD that implements the method, prior to implantation of an IMD that implements the method, or after implantation of the IMD that implements the method. The method further includes selectively adjusting a gain applied to the digital signal indicative of cardiac electrical activity to thereby selectively adjust a sensitivity associated with the detecting R-waves, while maintaining the fixed detection threshold.

In accordance with certain embodiments, the analog signal indicative of cardiac electrical activity is converted to the digital signal indicative of cardiac electrical activity is performed using an N-bit analog-to-digital converter (ADC) and a multiplier. The ADC accepts the analog signal indicative of cardiac electrical activity and outputs an N-bit digital signal. A course gain factor specifies which M-bits, of the N-bit digital signal output by the N-bit ADC, are provided to the multiplier, where M < N. In accordance with certain embodiments, the gain applied to the digital signal indicative of cardiac electrical activity is selectively adjusted, at least in part, by selectively adjusting the course gain factor. Additionally, or alternatively, a fine gain factor specifies a value that the M-bits, provided to the multiplier, are multiplied by to produce the digital signal indicative of cardiac electrical activity that is compared to the fixed detection threshold to detect R-waves. In accordance with certain embodiments, the gain applied to the digital signal indicative of cardiac electrical activity is selectively adjusted, at least in part, by selectively adjusting the fine gain factor.

Certain methods involve detecting peak amplitudes of the digital signal indicative of cardiac electrical activity, by detecting a peak amplitude of the digital signal indicative of cardiac electrical activity within a window following a threshold crossing, each time the comparing results in a threshold crossing that corresponds to an R-wave. In such embodiments, the selectively adjusting the fine gain factor can be based on the peak amplitudes of the digital signal indicative of cardiac electrical activity. This can involve, for example, adjusting the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being outside a specified range, and not adjusting the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being within the specified range. The adjusting the fine gain factor (in response to the peak amplitude of the digital signal indicative of cardiac electrical activity, being outside the specified range) can involve decreasing the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being above the specified range, and increasing the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being below the specified range.

An IMD that implements an above summarized method can, for example, measure RR intervals based on detected R-waves, monitor for one or more types of arrhythmias based on the RR intervals, and triggering an action in response to an arrhythmia being detected. Depending upon the specific arrhythmia detected, and the specific IMD, the action can involve stimulation therapy, including pacing, cardioversion and/or defibrillation stimulation, but is not limited thereto.

An IMD according to an embodiment of the present technology can include a plurality of electrodes, a sense amplifier, an analog-to-digital converter (ADC), adjustable gain circuitry downstream of the ADC, a comparator downstream of the adjustable gain circuitry, and a controller. The sense amplifier, which can be coupled to a pair of the electrodes, is configured to output an analog signal indicative of cardiac electrical activity, such as an ECG/IEGM signal. The ADC is configured to convert the analog signal indicative of cardiac electrical activity to a digital signal indicative of cardiac electrical activity. The adjustable gain circuitry is configured to adjust a gain applied to the digital signal indicative of cardiac electrical activity. The comparator is configured to detect R-waves by comparing the digital signal indicative of cardiac electrical activity to a fixed detection threshold to thereby detect threshold crossings that corresponds to R-waves. The controller is configured to selectively adjust the gain applied by the adjustable gain circuitry to thereby selectively adjust a sensitivity associated with the comparator detecting R-waves by detecting threshold crossings that corresponds to R-waves.

In accordance with certain embodiments, the ADC comprise an N-bit ADC, and the adjustable gain circuitry comprises an M-bit selector downstream of the N-bit ADC and a multiplier downstream of the M-bit selector. The M-bit selector can be configured to select which M-bits of an N-bit digital signal output by the N-bit ADC is provided to the multiplier, where M < N. The multiplier can be configured to multiplying the M-bits, selected by the M-bit selector, by a value provided to the multiplier by the controller. In accordance with certain embodiments, the controller is configured to perform one or more course gain adjustments by changing which M-bits of the N-bit digital signal output by the N-bit ADC is provided to the multiplier. Additionally, or alternatively, the controller can be configured to perform one or more fine gain adjustments by selectively changing the value, provided to the multiplier, that is multiplied by the M-bits selected by the M-bit selector. It would also be possible to not have an M-bit selector between the ADC and the multiplier, in which case, the adjustable gain circuitry can comprise the multiplier downstream of the ADC, and the controller can be configured to adjust the gain applied by the adjustable gain circuitry by changing a value that is provided to the multiplier to multiply by a digital signal output by the ADC.

This summary is not intended to be a complete description of the embodiments of the present technology. Other features and advantages of the embodiments of the present technology will appear from the following description in which the preferred embodiments have been set forth in detail, in conjunction with the accompanying drawings and claims.

Embodiments of the present technology relating to both structure and method of operation may best be understood by referring to the following description and accompanying drawings, in which similar reference characters denote similar elements throughout the several views:.

As noted above in the Background, some IMDs use automatic sensitivity control (ASC) to adjust a detection threshold that is used for detecting R-waves (and RR intervals). In general, with conventional ASC the amplitude of the detection threshold is varied and the amplitude of a gain applied to an ECG/IEGM signal (that is being compared to the detection threshold) is not adjusted. This is mostly due to the complexity involved with supporting a wide range and resolution of gain for an ECG/IEGM sense amplifier.

In accordance with certain embodiments of the present technology, rather than adjusting the threshold that is used for detecting R-waves (and RR intervals), one or more gains (which adjust the amplitude of the ECG/IEGM signal) is/are adjusted, and a fixed detection threshold is maintained. This provides for some circuit simplification and reduced complexity. Further, where an IMD already has an adjustable gain capability to allow for patient variability, such embodiments of the present technology can take advantage of a capability already built into the IMD.

<FIG> is a high level block diagram of circuitry that can be used to detect R-waves, and more specifically, can control sensitivity associated with the detecting R-waves while maintaining a fixed detection threshold. Referring to <FIG>, the IMD is shown as including a plurality of electrodes 304_1, 304_2,. 304_N, which can be referred to collectively as the electrodes <NUM>, or individually as an electrode <NUM>. An electrode configuration switch bank <NUM> is used to select which pair of the electrodes <NUM> is coupled to a sense amplifier <NUM>. The sense amplifier <NUM> outputs an analog signal indicative of cardiac electrical activity, e.g., an ECG/IEGM signal similar to the one shown in <FIG>. If an IMD includes only a pair of electrodes, as is the case with some leadless pacemakers, there is no need for the switch bank <NUM>. In other words, if an IMD includes only two electrodes, those two electrodes can always be directly coupled to the sense amplifier <NUM>.

Downstream of the sense amplifier <NUM> is an N-bit analog-to-digital converter (ADC) <NUM> that converts the analog signal indicative of cardiac electrical activity, which is output by the sense amplifier <NUM>, to a digital signal indicative of cardiac electrical activity. Downstream of the ADC <NUM> is gain circuitry <NUM> that can be used to adjust a gain applied to the digital signal indicative of cardiac electrical activity, which is output by the ADC <NUM>. In accordance with certain embodiments, the gain circuitry <NUM> includes both course gain circuitry and fine gain circuitry. In the embodiment shown in <FIG>, the course gain circuitry is an M-bit selector <NUM>, and the fine gain circuitry is a multiplier <NUM>. Both the M-bit selector <NUM> and the multiplier <NUM> can be controlled by a controller <NUM>.

The M-bit selector <NUM>, under the control of the controller <NUM>, selects which M-bits of the N-bit digital signal digital signal (output by the N-bit ADC) is provided to the multiplier, where M < N. In this manner, the M-bit selector either applies 1x gain, 2x gain, 4x gain, 8x gain, etc., to the digital signal indicative of cardiac electrical activity. For example, assume that N = <NUM>, and M = <NUM>, meaning the ADC <NUM> is a <NUM>-bit ADC and the M-bit selector <NUM> is an <NUM>-bit selector. Also assume that the least significant bit (LSB) is the <NUM>th bit and the most significant bit (MSB) is the <NUM>th bit of the <NUM>-bit output of the ADC <NUM>. Continuing with this example: 1x gain can be applied by selecting bits <NUM> through <NUM> of the <NUM>-bit output of the ADC <NUM>; 2x gain can be applied by selecting bits <NUM> through <NUM> of the <NUM>-bit output of the ADC <NUM>; 4x gain can be applied by selecting bits <NUM> through <NUM> of the <NUM>-bit output of the ADC <NUM>; 8x gain can be applied by selecting bits <NUM> through <NUM> of the <NUM>-bit output of the ADC <NUM>; and 16x gain can be applied by selecting bits <NUM> through <NUM> of the <NUM>-bit output of the ADC <NUM>. This is pictorially illustrated in <FIG>. In accordance with an embodiment, the M-bit selector is implemented using an M-bit register that accepts M-bits, but only output a selected N-bits of the M-bits. For a more specific example, the M-bit selector <NUM> can be implemented using an M-bit shift register and hardware that can apply a gain by selectively and appropriately shifting up to M-N bits. M-N bits (for 16x gain), M-N-<NUM> bits (for 8x gain), M-N-<NUM> bits (for 4x gain), etc. Again assume that N=<NUM>, and M=<NUM>, meaning the ADC <NUM> is a <NUM>-bit ADC and the M-bit selector <NUM> is an <NUM>-bit selector. In order to apply a 16x gain, the <NUM>-bit output of the ADC can be loaded into the M-bit shift register, and hardware can be used to arithmetically shift the data by M-N bits (i.e., by <NUM>-<NUM> = <NUM> bits); in order to apply a 8x gain, the <NUM>-bit output of the ADC can be loaded into the M-bit shift register and hardware can be used to arithmetically shift the data by M-N-<NUM> bits (i.e., by <NUM>-<NUM>-<NUM> = <NUM> bits); in order to apply a 4x gain, the <NUM>-bit output of the ADC can be loaded into the M-bit shift register and hardware can be used to arithmetically shift the data by M-N-<NUM> bits (i.e., by <NUM>-<NUM>-<NUM> = <NUM> bits); etc. This can be accomplished by always taking the N most significant bits from M after the arithmetic shift. These are all binary point shifts, so only factors of 2x gain are permissible with such an implementation. Other ways of implementing the M-bit selector are also possible and within the scope of the embodiments disclosed herein.

Referring again to <FIG>, the multiplier <NUM> multiplies the M-bits (selected by the M-bit selector) by a value provided to the multiplier <NUM> by the controller <NUM>. The value that the controller <NUM> provides to the multiplier <NUM> can be referred to as a "fine gain factor. " By contrast, a "course gain factor" that is generated by the controller <NUM> specifies which M-bits (of the N-bits output by the ADC <NUM>) are provided to the multiplier <NUM>. In accordance with certain embodiments, the fine gain factor can be a value between negative two inclusive, and positive two exclusive. In other words, the fine gain factors can be within the range of [-<NUM>,<NUM>). Other variations are also possible and within the scope of the embodiments described herein.

The signal <NUM> output by the multiplier <NUM>, which is a digital signal indicative of cardiac electrical activity <NUM> (to which gain has been applied) is provided to a comparator <NUM>. The comparator <NUM> compares the digital signal indicative of cardiac electrical activity <NUM> to a fixed detection threshold <NUM>, a value for which can be stored in a register <NUM>, or the like. In the embodiments shown, the digital signal indicative of cardiac electrical activity <NUM> is provided to the positive (+) terminal of the comparator <NUM>, and the fixed threshold <NUM> is provided to the negative (-) terminal of the comparator <NUM>. In such a configuration, the signal <NUM> output by the comparator <NUM> will be go HIGH whenever the digital signal indicative of cardiac electrical activity <NUM> (to which gain has been applied) exceeds the fixed detection threshold <NUM>, and will go LOW whenever the digital signal indicative of cardiac electrical activity <NUM> (to which gain has been applied) is below the fixed threshold <NUM>.

The controller <NUM> can perform course gain adjustments by changing which M-bits (of the N-bit digital signal) output by the N-bit ADC <NUM> is provided to the multiplier <NUM>. The controller <NUM> can perform fine gain adjustments by selectively changing the value, provided to the multiplier <NUM>, that is multiplied by the M-bits selected by the M-bit selector <NUM>. In accordance with certain embodiments, an appropriate course gain factor is selected once, e.g., during calibration, but the fine gain factor is dynamically adjusted over time by the controller <NUM> to provide automatic sensitivity control (ASC) after the IMD including the circuitry has been implanted within a patient. In other embodiments, both the course gain factor and the fine gain factor are dynamically adjusted by the controller <NUM> over time to provide automatic sensitivity control (ASC) after the IMD including the circuitry is implanted within a patient. In <FIG>, the multiplier <NUM> is shown as being downstream of the M-bit selector <NUM> within the gain circuitry <NUM>. In alternatively embodiments the order of the M-bit selector <NUM> and the multiplier <NUM> are reversed such that the M-bit selector <NUM> is downstream of the multiplier <NUM>, in which case a fine gain adjustment can be performed prior to a course gain adjustment.

The ADC <NUM> and the gain circuitry <NUM> can be referred to collectively as converter and adjustable gain circuitry <NUM>. The converter and adjustable gain circuitry <NUM> can include the ADC <NUM>, the M-bit selector <NUM>, and the multiplier <NUM>. In an alternative embodiment, the M-bit selector <NUM> can be eliminated, and the converter and adjustable gain circuitry <NUM> can include the ADC <NUM> and the multiplier <NUM>. In another alternative embodiment, the multiplier <NUM> can be eliminated, and the converter and adjustable gain circuitry <NUM> can include the ADC <NUM> and the M-bit selector <NUM>. For much of the remaining description is will be assumed that the converter and adjustable gain circuitry <NUM> includes the ADC <NUM>, the M-bit selector <NUM>, and the multiplier <NUM>. As noted above, the order of the M-bit selector <NUM> and the multiplier <NUM> can be reversed such that the multiplier <NUM> follows the ADC <NUM>, and the M-bit selector follows the multiplier <NUM>. However, for much of the remaining description is will be assumed that the M-bit selector <NUM> follows the ADC <NUM>, and the multiplier <NUM> follows the M-bit selector <NUM>, as is shown in <FIG>. Nevertheless, embodiments of the present technology also cover the alternative ordering of the multiplier <NUM> and the M-bit selector <NUM>.

Details are now provided as to how adjustments can be performed to the gain applied to digital signal indicative of cardiac electrical activity. Still referring to <FIG>, the digital signal <NUM> (indicative of cardiac electrical activity) that is output by the gain circuitry <NUM>, in addition to be provided to the digital comparator <NUM>, is also shown as being provided to a digital peak detector <NUM>. The signal <NUM> output by the comparator <NUM> is used to trigger the peak detector <NUM> to detect a peak amplitude of the digital signal <NUM> (indicative of cardiac electrical activity) within a temporal window following a threshold crossing (in a specific direction). Such a detected peak amplitude should correspond to a peak amplitude of an R-wave. In accordance with certain embodiments, the gain (e.g., the fine gain factor) is increased if the peak amplitude is much greater than the fixed threshold, and the gain is decreased if the peak amplitude is only slightly greater than the fixed threshold. There are various ways to achieve such functionality. For example, in an embodiment, the gain (e.g., the fine gain factor) is increased if the peak amplitude is more than X percent (e.g., more than <NUM>%) greater than the fixed threshold, the gain is decreased if the peak amplitude is less than Y percent (e.g., <NUM>%) greater than the fixed threshold, and the gain is not adjusted if the peak amplitude is between X and Y percent (e.g., between <NUM>% and <NUM>%) greater than the fixed threshold. Other variations are also possible and within the embodiments of the present technology. For example, an amount by which the gain is increased or decreased can depend upon a magnitude of a difference between the peak amplitude and the fixed threshold, wherein the greater the magnitude the greater the adjustment.

In <FIG>, the signal <NUM> output by the comparator <NUM>, which is indicative of R-wave detections, is shown as being provided to the controller <NUM>. This enables the controller <NUM> to determine RR intervals as well as HR, HRV, and the like. For example, the controller <NUM> can determine RR intervals by determining the time between successive leading edges of the pulses in the signal <NUM> output by the comparator <NUM>. The controller <NUM> can determine HR in beats per minute (bpm) by determining an average of a plurality of the RR intervals, and dividing the number sixty (<NUM>) by the average RR interval. The controller <NUM> can determine HRV by calculating a measure of variance in RR intervals, such as, but not limited to, by calculating the standard deviation (SD), the root mean square of successive differences (RMSSD), or the standard deviation of successive differences (SDSD) of a plurality of consecutive RR intervals.

Additional instances of the gain circuitry <NUM> are also shown in <FIG>, with the additional instances labeled 328_2 and 328_3. While the gain circuitry <NUM> applies a gain to the signal <NUM> to produce the gain adjusted signal <NUM>, which is used to detect R-waves, RR intervals, etc., the signal <NUM> can also be provided to the gain circuitry 328_2 and 328_3, as shown in <FIG>. The gain circuitry 328_2 can be used to apply an appropriate gain such that a gain adjusted signal 322_2 output therefrom is within an appropriate amplitude range used to store an ECG/IEGM signal, or segments thereof, in memory for later uploading and/or analysis. The gain circuitry 328_3 can be used to apply an appropriate gain such that a gain adjusted signal 322_3 output therefrom is within an appropriate amplitude range used to perform real time telemetry (and potentially real time display) of an ECG/IEGM signal. More generally, the circuitry <NUM> can include multiple gain channels that are used to apply various different adjustable gains to a digital signal indicative of cardiac electrical activity, which signal can also be referred to as an ECG/IEGM signal.

Each instance of the gain circuitry <NUM> shown in <FIG> was shown as including an M-bit selector <NUM> (used to perform course gain adjustments) and a multiplier <NUM> (used to perform fine gain adjustments). As noted above, the order of the M-bit selector <NUM> and the multiplier <NUM> can be reversed, such that the multiplier <NUM> is upstream of the M-bit selector. In alternative embodiments, one or more instances of the gain circuitry <NUM> (e.g., potentially all instances of the gain circuitry <NUM>) includes a multiplier <NUM> but does not include an M-bit selector <NUM>. In such an embodiment, a digital signal output by the ADC <NUM> can be provided directly to the multiplier <NUM>, and the controller can adjust the gain applied to the digital signal indicative of cardiac electrical activity by changing the value that the controller provides to the multiplier <NUM>. Where there are multiple instances of gain circuitry <NUM>, and thus multiple instances of the multiplier <NUM>, the controller <NUM> can provide different values to the different multipliers <NUM> to apply different gains to generate multiple digital signals indicative of cardiac electrical activity.

In the embodiments described above with reference to <FIG>, gain adjustments are performed in the digital domain, i.e., after the sensed analog signal indicative of cardiac electrical activity is converted to a digital signal indicative of cardiac electrical activity by the ADC <NUM>. This is beneficial because gain adjustments performed in the digital domain do not cause artefacts that often occur where gain adjustments are performed in the analog domain. Such artefacts, when they occur, can result in false detections of R-waves, i.e., in false positives. By performing the gain adjustments in the digital domain, such artefacts are avoided, and thus, the gain adjustments do not increase in the likelihood that false detections of R-waves will occur.

The circuitry <NUM> shown in <FIG> can be part of a pacemaker and/or implantable cardioverter defibrillator (ICD) to which are connected leads having electrodes, or part of a leadless pacemaker, or part of an implantable cardiac monitor that does not provide any therapy, but is not limited thereto. An exemplary IMD in which such circuitry <NUM> can be included is discussed below with reference to <FIG> and <FIG>.

The high level flow diagram of <FIG> will now be used to summarize methods for dynamically controlling sensitivity associated with detecting R-waves, in accordance with various embodiments of the present technology. Referring to <FIG>, step <NUM> involves sensing an analog signal indicative of cardiac electrical activity. Step <NUM> involves converting the analog signal indicative of cardiac electrical activity to a digital signal indicative of cardiac electrical activity. Step <NUM> involves detecting R-waves by comparing the digital signal indicative of cardiac electrical activity to a fixed detection threshold to thereby detect threshold crossings that corresponds to R-waves. Step <NUM> involves selectively adjusting a gain applied to the digital signal indicative of cardiac electrical activity to thereby selectively adjust a sensitivity associated with the detecting R-waves, while maintaining the fixed detection threshold.

Step <NUM> can be performed, e.g., using a pair of electrodes (<NUM> in <FIG>) and a sense amplifier (<NUM> in <FIG>), which are used to sense an ECG/IEGM signal.

Step <NUM> can be performed, e.g., using an N-bit ADC (e.g., <NUM> in <FIG>) that accepts the analog signal indicative of cardiac electrical activity and outputs an N-bit digital signal. In accordance with certain embodiments, a course gain factor specifies which M-bits, of the N-bit digital signal output by the N-bit ADC, are used when comparing the digital signal indicative of cardiac electrical activity to the fixed detection threshold to thereby detect threshold crossings that corresponds to R-waves, where M < N. Additionally, or alternatively, a fine gain factor can specify a value that the M-bits, provided to the multiplier, are multiplied by to produce the digital signal indicative of cardiac electrical activity that is compared to the fixed detection threshold to detect R-waves. In certain embodiments, the selectively adjusting the gain applied to the digital signal indicative of cardiac electrical activity is performed at step <NUM> by selectively adjusting the course gain factor. Additionally, or alternatively, step <NUM> can be performed by selectively adjusting the fine gain factor.

In accordance with certain embodiments, each time there is a threshold crossing in a particular direction (that is indicative of an R-wave detection), a peak amplitude of the digital signal indicative of cardiac electrical activity is detected within a window following the threshold crossing. In such embodiments, the fine gain factor can be selectively adjusted based on the peak amplitudes of the digital signal indicative of cardiac electrical activity. In certain embodiments this involves a controller (e.g., <NUM> in <FIG>) changing a value that it provides to a multiplier (e.g., <NUM> in <FIG>).

Selectively adjusting the fine gain factor, in accordance with certain embodiments, involves adjusting the fine gain factor in response to a peak amplitude (of the digital signal indicative of cardiac electrical activity) being outside a specified range, and not adjusting the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being within the specified range. More specifically, this can involve decreasing the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being above the specified range, and increasing the fine gain factor in response to the peak amplitude (of the digital signal indicative of cardiac electrical activity) being below the specified range. Other variations are also possible and within the scope of the embodiments described herein.

An exemplary IMD that can include the circuitry <NUM> discussed above with reference to <FIG>, and that can be used to perform the methods summarized with reference to <FIG>, will now be discussed below with reference to <FIG> and <FIG>.

Referring to <FIG>, an exemplary IMD <NUM> (also referred to as a pacing device, a pacing apparatus, a stimulation device, an implantable device or simply a device) is in electrical communication with a patient's heart <NUM> by way of three leads, <NUM>, <NUM> and <NUM>, suitable for delivering multi-chamber stimulation. While not necessary to perform embodiments of the present technology, the exemplary IMD <NUM> can also be capable of delivering shock therapy.

To sense atrial cardiac signals and to provide right atrial chamber stimulation therapy, the IMD <NUM> is coupled to an implantable right atrial lead <NUM> having at least an atrial tip electrode <NUM>, which typically is implanted in the patient's right atrial appendage. To sense left atrial and ventricular cardiac signals and to provide left-chamber pacing therapy, the IMD <NUM> is coupled to a "coronary sinus" lead <NUM> designed for placement in the "coronary sinus region" via the coronary sinus for positioning a distal electrode adjacent to the left ventricle and/or additional electrode(s) adjacent to the left atrium. As used herein, the phrase "coronary sinus region" refers to the vasculature of the left ventricle, including any portion of the coronary sinus, great cardiac vein, left marginal vein, left posterior ventricular vein, middle cardiac vein, and/or small cardiac vein or any other cardiac vein accessible by the coronary sinus.

Accordingly, an exemplary coronary sinus lead <NUM> is designed to receive left atrial and ventricular cardiac signals and to deliver left atrial and ventricular pacing therapy using at least a left ventricular tip electrode <NUM>, left atrial pacing therapy using at least a left atrial ring electrode <NUM>, and shocking therapy using at least a left atrial coil electrode <NUM>. The present technology may of course be practiced with a coronary sinus lead that does not include left atrial sensing, pacing or shocking electrodes.

The IMD <NUM> is also shown in electrical communication with the patient's heart <NUM> by way of an implantable right ventricular lead <NUM> having, in this embodiment, a right ventricular tip electrode <NUM>, a right ventricular ring electrode <NUM>, a right ventricular (RV) coil electrode <NUM>, and an SVC coil electrode <NUM>. Typically, the right ventricular lead <NUM> is transvenously inserted into the heart <NUM> so as to place the right ventricular tip electrode <NUM> in the right ventricular apex so that the RV coil electrode <NUM> will be positioned in the right ventricle and the SVC coil electrode <NUM> will be positioned in the superior vena cava. Accordingly, the right ventricular lead <NUM> is capable of receiving cardiac signals and delivering stimulation in the form of pacing and shock therapy to the right ventricle. It will be understood by those skilled in the art that other lead and electrode configurations such as epicardial leads and electrodes may be used in practicing the technology. For example, only a single lead, or only two leads, may be connected to the IMD. It should also be understood that the IMD can alternatively be a leadless device, such as an implantable monitor and/or a leadless pacer. The various electrodes shown in and described with reference to <FIG> can be specific implementations of the electrodes <NUM> discussed above with reference to <FIG>.

As illustrated in <FIG>, a simplified block diagram is shown of the multi-chamber implantable device <NUM>, which is capable of treating both fast and slow arrhythmias with stimulation therapy, including pacing, cardioversion and defibrillation stimulation. While a particular multi-chamber device is shown, this is for illustration purposes only, and one of skill in the art could readily duplicate, eliminate or disable the appropriate circuitry in any desired combination to provide a device capable of treating the appropriate chamber(s) with pacing, cardioversion and defibrillation stimulation.

The housing <NUM> for the IMD <NUM>, shown schematically in <FIG>, is often referred to as the "can", "case" or "case electrode" and may be programable to electrically act as the return electrode for all "unipolar" modes. The housing <NUM> may further be used as a return electrode alone or in combination with one or more of the coil electrodes, <NUM>, <NUM> and <NUM>, for shocking purposes. The housing <NUM> further includes a connector (not shown) having a plurality of terminals, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, and <NUM> (shown schematically and, for convenience, the names of the electrodes to which they are connected are shown next to the terminals). As such, to achieve right atrial sensing and pacing, the connector includes at least a right atrial tip terminal (AR TIP) <NUM> adapted for connection to the atrial tip electrode <NUM>.

To achieve left atrial and ventricular sensing, pacing and shocking, the connector includes at least a left ventricular tip terminal (VL TIP) <NUM>, a left atrial ring terminal (AL RING) <NUM>, and a left atrial shocking terminal (AL COIL) <NUM>, which are adapted for connection to the left ventricular ring electrode <NUM>, the left atrial tip electrode <NUM>, and the left atrial coil electrode <NUM>, respectively.

To support right ventricle sensing, pacing and shocking, the connector further includes a right ventricular tip terminal (VR TIP) <NUM>, a right ventricular ring terminal (VR RING) <NUM>, a right ventricular shocking terminal (RV COIL) <NUM>, and an SVC shocking terminal (SVC COIL) <NUM>, which are adapted for connection to the right ventricular tip electrode <NUM>, right ventricular ring electrode <NUM>, the RV coil electrode <NUM>, and the SVC coil electrode <NUM>, respectively.

At the core of the IMD <NUM> is a programmable microcontroller <NUM> which controls the various types and modes of stimulation therapy. As is well known in the art, the microcontroller <NUM> typically includes a microprocessor, or equivalent control circuitry, designed specifically for controlling the delivery of stimulation therapy and can further include RAM or ROM memory, logic and timing circuitry, state machine circuitry, and I/O circuitry. Typically, the microcontroller <NUM> includes the ability to process or monitor input signals (data) as controlled by a program code stored in a designated block of memory. The details of the design of the microcontroller <NUM> are not critical to the present technology. Rather, any suitable microcontroller <NUM> can be used to carry out the functions described herein. The use of microprocessor-based control circuits for performing timing and data analysis functions are well known in the art. In specific embodiments of the present technology, the microcontroller <NUM> performs some or all of the steps associated with arrhythmia detection. The microcontroller <NUM> can be used to implement that controller <NUM> discussed above with reference to <FIG>.

Representative types of control circuitry that may be used with the technology include the microprocessor-based control system of <CIT>) and the state-machines of <CIT>) and <CIT>). For a more detailed description of the various timing intervals used within the pacing device and their inter-relationship, see <CIT>).

An atrial pulse generator <NUM> and a ventricular pulse generator <NUM> generate pacing stimulation pulses for delivery by the right atrial lead <NUM>, the right ventricular lead <NUM>, and/or the coronary sinus lead <NUM> via an electrode configuration switch <NUM>. It is understood that in order to provide stimulation therapy in each of the four chambers of the heart, the atrial and ventricular pulse generators, <NUM> and <NUM>, may include dedicated, independent pulse generators, multiplexed pulse generators, or shared pulse generators. The pulse generators, <NUM> and <NUM>, are controlled by the microcontroller <NUM> via appropriate control signals, <NUM> and <NUM>, respectively, to trigger or inhibit the stimulation pulses.

The microcontroller <NUM> further includes timing control circuitry <NUM> which is used to control pacing parameters (e.g., the timing of stimulation pulses) as well as to keep track of the timing of refractory periods, noise detection windows, evoked response windows, alert intervals, marker channel timing, etc., which is well known in the art. Examples of pacing parameters include, but are not limited to, atrio-ventricular delay, interventricular delay and interatrial delay.

The switch bank <NUM> includes a plurality of switches for connecting the desired electrodes to the appropriate I/O circuits, thereby providing complete electrode programmability. Accordingly, the switch <NUM>, in response to a control signal <NUM> from the microcontroller <NUM>, determines the polarity of the stimulation pulses (e.g., unipolar, bipolar, etc.) by selectively closing the appropriate combination of switches (not shown) as is known in the art.

Atrial sensing circuits <NUM> and ventricular sensing circuits <NUM> may also be selectively coupled to the right atrial lead <NUM>, coronary sinus lead <NUM>, and the right ventricular lead <NUM>, through the switch <NUM> for detecting the presence of cardiac activity in each of the four chambers of the heart. Accordingly, the atrial (ATR. SENSE) and ventricular (VTR. SENSE) sensing circuits, <NUM> and <NUM>, may include dedicated sense amplifiers, multiplexed amplifiers, or shared amplifiers. The switch <NUM> determines the "sensing polarity" of the cardiac signal by selectively closing the appropriate switches, as is also known in the art. In this way, the clinician may program the sensing polarity independent of the stimulation polarity.

Each sensing circuit, <NUM> and <NUM>, preferably employs one or more low power, precision amplifiers with programmable gain and/or automatic gain control, bandpass filtering, and a threshold detection circuit, as known in the art, to selectively sense the cardiac signal of interest. The automatic gain control enables the IMD <NUM> to deal effectively with the difficult problem of sensing the low amplitude signal characteristics of atrial or ventricular fibrillation. Such sensing circuits, <NUM> and <NUM>, can be used to determine cardiac performance values used in the present technology. Alternatively, an automatic sensitivity control circuit may be used to effectively deal with signals of varying amplitude.

The outputs of the atrial and ventricular sensing circuits, <NUM> and <NUM>, are connected to the microcontroller <NUM> which, in turn, are able to trigger or inhibit the atrial and ventricular pulse generators, <NUM> and <NUM>, respectively, in a demand fashion in response to the absence or presence of cardiac activity, in the appropriate chambers of the heart. The sensing circuits, <NUM> and <NUM>, in turn, receive control signals over signal lines, <NUM> and <NUM>, from the microcontroller <NUM> for purposes of measuring cardiac performance at appropriate times, and for controlling the gain, threshold, polarization charge removal circuitry (not shown), and timing of any blocking circuitry (not shown) coupled to the inputs of the sensing circuits, <NUM> and <NUM>. The sensing circuits can be used, for example, to acquire IEGM signals.

For arrhythmia detection, the IMD <NUM> includes an arrhythmia detector <NUM> that utilizes the atrial and ventricular sensing circuits, <NUM> and <NUM>, to sense cardiac signals to determine whether a rhythm is physiologic or pathologic. The timing intervals between sensed events (e.g., P-waves, R-waves, and depolarization signals associated with fibrillation) are then classified by the microcontroller <NUM> by comparing them to a predefined rate zone limit (i.e., bradycardia, normal, low rate VT, high rate VT, and fibrillation rate zones) and various other characteristics (e.g., sudden onset, stability, physiologic sensors, and morphology, etc.) in order to assist with determining the type of remedial therapy that is needed (e.g., bradycardia pacing, anti-tachycardia pacing, cardioversion shocks or defibrillation shocks, collectively referred to as "tiered therapy"). The arrhythmia detector <NUM> can be implemented within the microcontroller <NUM>, as shown in <FIG>. Thus, this detector <NUM> can be implemented by software, firmware, or combinations thereof. It is also possible that all, or portions, of the arrhythmia detector <NUM> can be implemented using hardware. Further, it is also possible that all, or portions, of the arrhythmia detector <NUM> can be implemented separate from the microcontroller <NUM>.

The stimulation device <NUM> is also shown as including a pacing controller <NUM>, which can adjust a pacing rate and/or pacing intervals. The pacing controller <NUM> can be implemented within the microcontroller <NUM>, as shown in <FIG>. Thus, the pacing controller <NUM> can be implemented by software, firmware, or combinations thereof. It is also possible that all, or portions, of the pacing controller <NUM> can be implemented using hardware.

The accelerometer <NUM> of the IMD <NUM> can be or include, e.g., a MEMS (micro-electromechanical system) multi-axis accelerometer of the type exploiting capacitive or optical cantilever beam techniques, or a piezoelectric accelerometer that employs the piezoelectric effect of certain materials to measure dynamic changes in mechanical variables (e.g., acceleration, and/or vibration), but is not limited thereto. Depending upon implementation, the accelerometer <NUM> can be used to detect posture and/or motion of a patient in which an IMD <NUM> including the accelerometer <NUM> is implanted.

Additionally, the IMD <NUM> is shown as including a gain adjustor <NUM>. In accordance with certain embodiments of the present technology, the gain adjustor <NUM> can be used adjust a gain applied an ECG/IEGM signal sensed using a pair of electrodes to thereby selectively adjust a sensitivity associated detecting R-waves by detecting crossings of fixed detection threshold. Additional details of the operation of the gain adjustor <NUM>, according to various embodiments of the present technology, can be appreciated from the above discussion of <FIG>. The gain adjustor <NUM> can be implemented within the microcontroller <NUM>, as shown in <FIG>. Thus, the gain adjustor <NUM> can be implemented by software, firmware, hardware, or combinations thereof. It is also possible that all, or portions, of the gain adjustor <NUM> can be implemented using dedicated hardware, such as using an application specific integrated circuit (ASIC). More generally, the gain adjustor <NUM> can be implemented by a controller, wherein the controller may be a microcontroller (e.g., <NUM>), or an ASIC, but is not limited thereto.

Still referring to <FIG>, cardiac signals and/or other signals can be applied to the inputs of an analog-to-digital (A/D) data acquisition system <NUM>. The data acquisition system <NUM> is configured to acquire intracardiac electrogram signals, convert the raw analog data into a digital signal, and store the digital signals for later processing and/or telemetric transmission to an external device <NUM>. The data acquisition system <NUM> is coupled to the right atrial lead <NUM>, the coronary sinus lead <NUM>, and the right ventricular lead <NUM> through the switch <NUM> to sample cardiac signals across any pair of desired electrodes.

The data acquisition system <NUM> can be coupled to the microcontroller <NUM>, or other detection circuitry, for detecting an evoked response from the heart <NUM> in response to an applied stimulus, thereby aiding in the detection of "capture". Capture occurs when an electrical stimulus applied to the heart is of sufficient energy to depolarize the cardiac tissue, thereby causing the heart muscle to contract. The microcontroller <NUM> detects a depolarization signal during a window following a stimulation pulse, the presence of which indicates that capture has occurred. The microcontroller <NUM> enables capture detection by triggering the ventricular pulse generator <NUM> to generate a stimulation pulse, starting a capture detection window using the timing control circuitry <NUM> within the microcontroller <NUM>, and enabling the data acquisition system <NUM> via control signal <NUM> to sample the cardiac signal that falls in the capture detection window and, based on the amplitude, determines if capture has occurred. The data acquisition system <NUM> may also be used to acquire signals produced by the sensors <NUM> and/or <NUM>, and may convert analog signals produced by such sensor to digital signals. It is also possible that the sensors <NUM> and/or <NUM> output digital signals. The implementation of capture detection circuitry and algorithms are well known. See for example, <CIT>); <CIT>); <CIT>); <CIT>); and <CIT>). The type of capture detection system used is not critical to the present technology.

The converter and adjustable gain circuitry <NUM>, described above with reference to <FIG>, can be used to implement the data acquisition system <NUM> in <FIG>, or vice versa. It would also be possible for the sense amplifier <NUM> in <FIG> to be used to implement the sensing circuit <NUM> and/or <NUM> in <FIG>, or vice versa. Further, the electrode switch configuration bank <NUM> in <FIG> can be used to implement the electrode configuration switch <NUM> in <FIG>, or vice versa. Similarly, the various electrodes shown in <FIG> and discussed with reference to <FIG> and <FIG> can be the electrodes 304_1. 304_n in <FIG>, or vice versa. Further, the controller <NUM> in <FIG> can be used to implement the microcontroller <NUM> in <FIG>, or vice versa.

The microcontroller <NUM> is further coupled to the memory <NUM> by a suitable data/address bus <NUM>, wherein the programmable operating parameters used by the microcontroller <NUM> are stored and modified, as required, in order to customize the operation of the IMD <NUM> to suit the needs of a particular patient. Such operating parameters define, for example, pacing pulse amplitude, pulse duration, electrode polarity, rate, sensitivity, automatic features, arrhythmia detection criteria, and the amplitude, waveshape and vector of each shocking pulse to be delivered to the patient's heart <NUM> within each respective tier of therapy. The memory <NUM> can also be used to store data relating to one or more magnetic field thresholds, and other information that can be utilized in embodiments of the present technology described herein.

The operating parameters of the IMD <NUM> may be non-invasively programmed into the memory <NUM> through a telemetry circuit <NUM> in telemetric communication with an external device <NUM>, such as a programmer, transtelephonic transceiver, or a diagnostic system analyzer. The telemetry circuit <NUM> can be activated by the microcontroller <NUM> by a control signal <NUM>. The telemetry circuit <NUM> advantageously allows intracardiac electrograms and status information relating to the operation of the device <NUM> (as contained in the microcontroller <NUM> or memory <NUM>) to be sent to the external device <NUM> through an established communication link <NUM>. The telemetry circuit <NUM> can also be used to trigger alarms or alerts of the external device <NUM>, or to instruct the external device <NUM> to notify a caregiver regarding detection of various episodes, occurrences and changes in conditions that are detected using embodiments of the present technology.

For examples of such devices, see <CIT>); <CIT>); and <CIT>).

The IMD <NUM> additionally includes a battery <NUM> which provides operating power to all of the circuits shown in <FIG>. If the implantable device <NUM> also employs shocking therapy, the battery <NUM> should be capable of operating at low current drains for long periods of time, and then be capable of providing high-current pulses (for capacitor charging) when the patient requires a shock pulse. The battery <NUM> should also have a predictable discharge characteristic so that elective replacement time can be detected.

As further shown in <FIG>, the IMD <NUM> is also shown as having an impedance measuring circuit <NUM> which is enabled by the microcontroller <NUM> via a control signal <NUM>. The known uses for an impedance measuring circuit <NUM> include, but are not limited to, lead impedance surveillance during the acute and chronic phases for proper lead positioning or dislodgement; detecting operable electrodes and automatically switching to an operable pair if dislodgement occurs; measuring respiration or minute ventilation; measuring thoracic impedance for determining shock thresholds and heart failure condition; detecting when the device has been implanted; measuring stroke volume; and detecting the opening of heart valves, etc. The impedance measuring circuit <NUM> is advantageously coupled to the switch <NUM> so that any desired electrode may be used. The impedance measuring circuit <NUM> is not critical to the present technology and is shown only for completeness.

In the case where the IMD <NUM> is also intended to operate as an implantable cardioverter/defibrillator (ICD) device, it must detect the occurrence of an arrhythmia, and automatically apply an appropriate electrical shock therapy to the heart aimed at terminating the detected arrhythmia. To this end, the microcontroller <NUM> further controls a shocking circuit <NUM> by way of a control signal <NUM>. The shocking circuit <NUM> generates shocking pulses of low (up to <NUM> Joules), moderate (<NUM>-<NUM> Joules), or high energy (<NUM> to <NUM> Joules), as controlled by the microcontroller <NUM>. Such shocking pulses are applied to the patient's heart <NUM> through at least two shocking electrodes, and as shown in this embodiment, selected from the left atrial coil electrode <NUM>, the RV coil electrode <NUM>, and/or the SVC coil electrode <NUM>. As noted above, the housing <NUM> may act as an active electrode in combination with the RV electrode <NUM>, or as part of a split electrical vector using the SVC coil electrode <NUM> or the left atrial coil electrode <NUM> (i.e., using the RV electrode as a common electrode).

The above described IMD <NUM> was described as an exemplary pacing device. One or ordinary skill in the art would understand that embodiments of the present technology can be used with alternative types of implantable devices. Accordingly, embodiments of the present technology should not be limited to use only with the above described device.

As noted above, embodiments of the present technology may also be used with a leadless pacemaker, or with an implantable cardiac monitor that does not provide any therapy. Exemplary leadless pacemakers are described in <CIT>) and <CIT>) An implantable cardiac monitor that does not provide any therapy can, for example, store information indicative of R-waves, HR, HRV, as well as ECG/IEGM segments, and such stored information can be uploaded to an external device for analysis and or display.

Embodiments of the present technology describe above generally pertain to IMDs, and methods for use therewith. Such embodiments of the present technology have been described above with the aid of functional building blocks illustrating the performance of specified functions and relationships thereof. The boundaries of these functional building blocks have often been defined herein for the convenience of the description. Any such alternate boundaries are thus within the scope and spirit of the claimed technology. For example, it would be possible to combine or separate some of the steps shown in <FIG>. For another example, it is possible to change the boundaries of some of the blocks shown in <FIG> and <FIG>.

It is noted that the term "base on", as used herein, should be interpreted as meaning based at least in part on, unless stated otherwise. In other words, where a decision is based on something, that decision can also be based on additional things. By contrast, where a decision is based solely on something, that decision is not also based on additional things.

Claim 1:
A device capable of dynamically controlling sensitivity associated with detecting R-waves, the device comprising:
a plurality of electrodes (<NUM>);
a sense amplifier (<NUM>) configured to be coupled to a pair of the electrodes and configured to output an analog signal indicative of cardiac electrical activity;
an analog-to-digital converter (ADC) (<NUM>) configured to convert the analog signal indicative of cardiac electrical activity to a digital signal indicative of cardiac electrical activity;
adjustable gain circuitry (<NUM>) downstream of the ADC (<NUM>) and configured to adjust a gain applied to the digital signal indicative of cardiac electrical activity;
a comparator (<NUM>) downstream of the adjustable gain circuitry (<NUM>) and configured to detect R-waves by comparing the digital signal indicative of cardiac electrical activity to a fixed detection threshold to thereby detect threshold crossings that corresponds to R-waves; and
a controller (<NUM>) configured to selectively adjust the gain applied by the adjustable gain circuitry (<NUM>) to thereby selectively adjust a sensitivity associated with the comparator (<NUM>) detecting R-waves by detecting threshold crossings that corresponds to R-waves.